Highly Stretchable CoreâSheath Fibers via Wet-Spinning for Wearable

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Highly-Stretchable Core-Sheath Fibers via Wet-Spinning for Wearable Strain Sensors Zhenhua Tang, Shuhai Jia, Fei Wang, Changsheng Bian, Yuyu Chen, Yonglin Wang, and Bo Li ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.7b18677 • Publication Date (Web): 31 Jan 2018 Downloaded from http://pubs.acs.org on February 1, 2018

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ACS Applied Materials & Interfaces

Highly-Stretchable Core-Sheath Fibers via Wet-Spinning for Wearable Strain Sensors Zhenhua Tang, Shuhai Jia*, Fei Wang, Changsheng Bian, Yuyu Chen, Yonglin Wang, and Bo Li School of Mechanical Engineering, Xi’an Jiaotong University, Xi’an 710049, P. R. China *Corresponding authors. E-mail: [email protected]

ABSTRACT

Lightweight, stretchable, and wearable strain sensors have recently been widely studied for the development of health monitoring systems, human machine interfaces, and wearable devices. Herein, highly stretchable polymer elastomer-wrapped carbon nanocomposite piezoresistive core-sheath fibers (CSFs) are successfully prepared using a facile and scalable one-step coaxial wet-spinning assembly approach. The carbon nanotube-polymeric composite core of the stretchable fiber is surrounded by an insulating sheath, similar to conventional cables, and shows excellent electrical conductivity with a low percolation threshold (0.74 vol%). The core-sheath elastic fibers are used as wearable strain sensors, exhibiting ultra-high stretchability (above 300%), excellent stability (>10000 cycles), fast response, low hysteresis, and good washability. Furthermore, the piezoresistive core-sheath fiber possesses bending-insensitiveness and negligible torsion-sensitive properties, and the strain sensing performance of piezoresistive fibers

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maintains a high degree of stability under harsh conditions. Based on this high level of performance, the fiber-shaped strain sensor can accurately detect both subtle and large scale human movements by embedding it in gloves, garments, and directly attaching it to the skin. The current results indicate that the proposed stretchable strain sensor has many potential applications in health monitoring, human–machine interfaces, soft robotics, and wearable electronics.

KEYWORDS: wet-spinning; carbon nanotubes; core-sheath fibers; wearable strain sensors; motion detection 1. Introduction Soft, stretchable and wearable sensing devices will impact and advance diverse fields, including human–machine interaction, human health monitoring, wearable electronics, and soft robotics.1-4 Recently, a number of soft and stretchable strain sensors have been developed, based on various transduction mechanisms,5 such as capacitive,6 piezoelectric,7 triboelectric,8 and piezoresistive sensing.9 In particular, piezoresistive sensing mechanisms involving the transduction of strain/stress into electrical resistance signals have gained considerable attention, primarily because of their easy signal collection capability, simple sensor architecture, facile fabrication and low cost. However, conventional metal foil strain gauges are unsuitable as sensing components in wearable sensors because of limitations related to their poor mechanical compliance, large hysteresis, and narrow sensing range (5%).10,11 As alternative piezoresistive materials, nanomaterials have been shown to be promising building blocks for innovative strain sensors with enhanced performance, and high-performance sensor devices based on several representative nanostructures such as carbon nanotubes (CNTs),12 graphene,13 nanowires,14 and nanoparticles15 have been developed. Specifically, CNTs and graphene with extraordinary


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electrical and mechanical properties have been extensively studied for stretchable strain sensing applications.12,13,16 To create the desired sensing geometry, many approaches have been employed to date, including vacuum filtration,17 printing,18,19 dipping,13,20 spin-coating,21 dipcoating,22 depositing,23 casting,24 writing,25 micro-channel molding,26,27 and lamination techniques.28 However, although each of these methods is effective for creating sensors, drawbacks such as time-consuming, complex process, expensive instruments, or lack of manufacturing scalability have hindered their widespread application. In addition, the geometry of the majority of the sensors is planar laminated construction, which may limit their practical and wide applications.29 For example, the monolithic film format of many sensing devices may not fully meet the requirements of high flexibility and conformability. Fiber-shaped sensors hold great promise for the development of high-performance stretchable strain sensors, and fiber-like sensors can provide excellent wearable properties that mimic regular textiles and withstand mechanical deformations like bending, folding, twisting, and stretching. In addition, these sensors can be attached conformably onto uneven surfaces, including human skin.30 Recently, a variety of flexible piezoresistive fibers have been developed as wearable electronic systems. For example, Bautista-Quijano and colleagues reported a fiber-shaped strain sensor based on melt-spun polycarbonate/multiwall carbon nanotube (MWCNT) monofilament fiber, which exhibited a narrow strain range that restricts its applications in monitoring large deformations.31 Cheng et al. constructed a graphene-based fiber by winding and repetitious dipcoating, which exhibited a relatively wide strain sensing range (about 100%).22 Park et al. fabricated a graphene strain sensor based on stretchable yarns using a layer-by-layer assembly method that was low-cost and solution-processable.32 Wang et al. demonstrated core-sheath fibers fabricated by ultralong CVD-grow graphene bundles from a copper wire template, then


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coated it with a thin layer of poly(vinyl alcohol) (PVA). The core-sheath graphene fiber showed a highly sensitive response to bending and stretching, with a gauge factor of ~5.02 under a strain of 1–6.3%.33 In addition, a small number of researchers have fabricated fiber-shaped sensors by injecting liquid metal alloys or ionic fluids into hollow elastomeric fibers.27,34 Although various fibrous sensing devices have been demonstrated to have beneficial characteristics (lightweight, soft, and flexible), the key issues for wearable electronics, including stability, washability, scalability, reproducibility, sensitivity, and mechanical stretchability, still need to be addressed. Unfortunately, many of the methods for fabricating the fiber-like sensors mentioned above are more complex, time-consuming and uncontrollable.32,33,35 In particular, their applications to wearable devices are impeded by their limited stretchability (e.g., less than 100%),31,33 and poor washability. Moreover, most previously reported fiber-like sensors are exposed, involving the risk of short-circuiting, and are sensitive to a range of external stimuli including torsion, bending, and humidity rather than to specific mechanical stimuli (strain/stress),22,36 which poses great difficulty for real-world applications. Thus, the development of a versatile, cost-effective, massproduction approach for the fabrication of highly stretchable, durable, and washable fiber-shaped strain sensors would be a valuable innovation. In the current study, we developed a one-step coaxial wet-spinning assembly strategy to fabricate silicone elastomer-protected MWCNT-based core-sheath fiber (CSF) strain sensors with high stretchability (above 300%), high sensitivity (at large strain range, gauge factor = 1378), high reproducibility (>10000 cycles), fast signal response, low hysteresis, and excellent washability. Although wet-spinning strategy has been utilized to spin fiber-like nanocomposites,37,38 a coaxial wet-spinning technology has not previously been developed for preparing stretchable coaxial fibers from viscous silicone elastomers, because of the extreme difficulty of selecting an


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appropriate coagulation bath and spinning precursor. The currently proposed methodology is facile, controllable and suitable for mass manufacturing. Furthermore, the obtained CSF strain sensor possesses outstanding bending-insensitiveness and slight torsion-sensitive properties, and the strain sensing performance of CSFs maintained stability in an ambient environment. To demonstrate the practical application of the flexible fibers, we integrated the CSFs into wearable sensors and successfully detected both delicate motion (such as respiratory activity and muscle motion) and large movements (such as joint bending) of the human body. The combination of the fast yet scalable coaxial wet-spinning assembly approach and the reliable performance of fibershaped strain sensors may pave the way for the development of safe wearable electronics. 2. Results and Discussion 2.1. Materials System. For spinning elastic fibers, we modified a platinum-catalyzed silicone rubber, Ecoflex 00-30 (Smooth on), to fabricate both the core and sheath matrix. The sheath spinning ink was pure silicone elastomer solution. The inner core spinning ink was prepared by dispersing the required amount of MWCNTs into the silicone elastomer solution. Other conductive fillers, such as carbon black and graphene, could also be used. Scanning electron microscopy (SEM) imaging showed the MWCNTs, which exhibited an average diameter of ~12.27 nm (Figure 1a and b). The quality of these CNTs was characterized by Raman spectroscopy with a laser excitation wavelength of 633 nm. Two characteristic peaks of MWCNT at 1320 and 1580 cm-1, corresponding to the D and G band, respectively, are observed (Figure 1c). The intensity ratio of D to G band (ID/IG) turned out to be 1.52, indicating a high level of impurity and defect density in the as-purchased MWCNTs.


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Spinning inks developed for wet-spinning must exhibit shear-thinning behavior to enable efficient flow through fine deposition nozzles. Figure 1d shows the rheological behavior of viscoelastic inks containing different concentrations of CNT. These viscoelastic inks exhibit strong shear thinning behavior without the need for other rheology modifiers, in which the apparent viscosity decreases to 10 Pa•s when the shear rate increases to 103 s-1, allowing it to be extruded through the coaxial nozzles. The viscosity increased with increasing CNT loading in all frequency ranges, especially in the lower frequency range of 0.1 Hz to 1 Hz (Figure 1d). This suggests that the mobility of the silicone elastomer macromolecular chains was restrained by the addition of CNT, causing higher viscosity. Furthermore, oscillatory measurements were carried out to assess the viscoelastic properties of the spinning formulations. The CNT-filled silicone elastomer possesses a yield stress that allows the ink to spin at modest pressures, as well as exhibiting a sufficiently high shear elastic modulus to enable the ink to keep its filamentary shape after it exits the coaxial spinneret (Figure 1e). Finally, when choosing an appropriate coagulation bath for the wet-spinning, several criteria must be taken into account. First, the extruded viscoelastic inks must be rapidly vulcanized in the coagulation bath. Second, the coagulation bath must facilitate patterning the desired filaments without breakup. To meet these requirements, we used higher viscosity silicone oil (1000 cs) to work as a coagulation bath solution.


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Figure 1. Basic characteristics of spinning materials. (a) Scanning electron microscope (SEM) image of the MWCNTs. (b) Histogram showing diameter of MWCNTs. The mean diameter was 12.27 nm. (c) Raman spectrum of MWCNTs. (d) The apparent viscosity versus shear rate and (e) The shear elastic modulus versus shear stress for the spinning ink. 2.2. Fabrication and Characterization of Stretchable Coaxial Fibers. Figure 2a shows a schematic of our coaxial wet-spinning strategy (an optical image of the corresponding equipment is shown in Figure S1a, Supporting Information) of the elastomeric core-sheath fibers. The fiber consists of two concentric layers of a CNT-silicone elastomer and a pure silicone elastomer that serve as the resistive sensing element and electrically insulative layer, respectively. Briefly, viscoelastic CNT-silicone elastomer ink from the inner channel and viscoelastic silicone


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elastomer fluid from the outer channel were synchronously injected into the silicone oil bath (~110 °C). Silicone elastomer wrapped CNT-silicone elastomer hybrid fibers were then solidified in the coagulation bath for several minutes to complete the vulcanization process, followed by ethanol washing and winding onto a spool (see Movie S1, Supporting Information). Further details of the fabrication process are given in the Experimental Section (Section 4). Consequently, several meters of CSF were easily produced in ~10 min (Figure 2b), demonstrating the potential for large-scale productivity of the multifunctional coaxial fibers. The obtained CSFs were immersed in ethanol, and showed the clear core-shell structure as designed (Figure 2c). Moreover, the produced CSFs exhibited excellent stretchability (Figure 2d and e) and superior mechanical properties (Figure 2f). Interestingly, the highly flexible CSFs were successfully tied into various knots, such as the million characters knot (Figure 2g) and the overhand knot (Figure 2h). No evident fractures were observed during knotting, indicating the excellent knitting feasibility and conformability of the CSFs. A cross-section SEM of the CSFs is shown in Figure 2i, illustrating the cylindrical shape of the fabricated coaxial yarn with an outer diameter of 1.71 mm and core diameter of 1.01 mm. Figure 2j shows a higher magnification SEM image of the CNTs in the core layer of CSF, which exhibits a uniform distribution. The uniformity of MWCNTs in the matrix had a marked influence on their electrical performance. Figure S2 shows the strong interface bonding between the core layer and the sheath layer of the coaxial fiber because of their similar components, which effectively avoided delamination in the interface area during mechanical deformation. Energy dispersive spectroscopy (EDS) analysis shows that the core of the fiber is mainly composed of C, O, and Si, as shown in Figure S3 of the Supporting Information. Thus, MWCNTs were found throughout the inner fracture surface.


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The dimensions of the resulting coaxial fibers are dictated by the relative nozzle sizes and the respective flow rates of the ink in each layer. In this study, the inner diameters of the CSFs were tuned from ~0.3 mm to ~1.2 mm by changing the outflow velocity of the inner ink, and the wall thicknesses were controlled in a range from ~200 µm to ~600 µm by changing the flow rates of the shell nozzle with a determined coaxial nozzle.


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Figure 2. The wet-spinning procedure and the characterization of the coaxial fibers. (a) Schematic illustration of the coaxial spinning process for the highly stretchable fibers. (b) Photograph of the flexible fiber collected on a perspex bar. (c) The stretchable fibers were immersed in ethanol, showing the clear core-shell structure. (d, e) Coaxial fiber strained to 0% and ≈150% strain, respectively. (f) Photograph of a knotted CSF with an outer diameter of 1.7 mm carrying a weight of 200 g. (g) Photograph of a CSF woven million characters knot. (h) SEM image of a CSF overhand knot. (i) The cross-sectional SEM image of the CSF with 2 wt% MWCNT showing the clear concentric construction. (j) SEM image showing the core of the CSF with uniform distribution of MWCNTs. Scale bar in (h) and (i) are 400 µm. Scale bar in (j) is 20 µm. 2.3. Electromechanical Performance of CSF Strain Sensors. To understand the strain sensing behavior of the generated CSFs, we investigated their electrical properties. CSFs with different MWCNT loading levels (mass contents of 1, 1.5, 2, 2.5, 3, and 3.5 wt%, corresponding to volume concentrations of 0.51, 0.76, 1.02, 1.28, 1.55, and 1.81 vol%) were prepared using a wetspinning technique. For the experimental results of the percolation threshold reported in weight percentages, conversion to volume percentage was performed by assuming that the densities of MWCNTs and polymer were 2.1 g/cm3 and 1.07 g/cm3, respectively. Figure 3 shows the dependence of volume resistivity of CSF composites on the volume content of MWCNTs. It can be seen that a rapid increase in the conductivity of several orders of magnitude took place when the MWCNTs content increases, indicating the formation of a continuous conductive network. According to statistical percolation theory, the MWCNT-based nanocomposite conductivity scales with filler volume fraction φ can be described as


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   0 (  c )t

(1)

Where σ represents the electrical conductivity of CSF at a given filler loading, σ0 is a scaling factor, φ is the volumetric fraction of CNTs, and φc and t are the percolation threshold and exponent, respectively.39 The critical exponent reflects the dimensionality of the conductive networks in the CSFs. This equation fits our experimental data well, giving φc = 0.74 vol% and t = 1.84. Generally, the t of a two-dimensional system is approximately 1.1–1.3, and the t of a three-dimensional system is between 1.6–2.0.40 This indicates that a three-dimensional conductive network of these nanocomposites was achieved by contact among the CNTs. Furthermore, the percolation threshold of CSF was lower than that of CNT/silicone rubber composites.41 Such a low percolation threshold is likely to be caused by the extremely high

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Figure 3. Electrical conductivity of CSF as a function of MWCNT volume fraction. Inset shows the fitted results of experimental data of CSF according to the percolation law. Based on the findings described above, we further investigated the electrical performance of the CSF strain sensors with CNT content of 2 wt% and 3 wt% under uniaxial tensile loadings. The


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sensors were denoted as CSF-2 and CSF-3, corresponding to CNT concentration of 2 wt% and 3 wt%, respectively. Figure 4a shows the relative resistance change (ΔR/R0=(R-R0)/R0, where R is the measured resistance and R0 is the original resistance before stretching) as a function of tensile strain (quasi-static loading). From this figure it can be seen that with higher CNT loading, CSF exhibited slower change in electrical resistance under strain. In addition, once the applied strain reached a threshold value, the separation between MWCNTs was so large that the electrically percolating network broke down. The composites transfer from semiconductors to electrically insulating materials. This behavior occurs in highly deformable polymers and has been named “electrical depercolation”.42 In addition, the sensing range (strain measurement capability) of the CSFs increased as the CNT content increased, because higher CNT loading resulted in denser CNT conductive networks with more junctions or entanglements between CNTs, which was more resistant to deformation. Furthermore, in all cases, the resistance was found to exhibit four different stages of development. The four regimes in the normalized resistivity of CSF-3 are shown in Figure S4 (supporting Information). After a small plateau or increase because of the network breakage,43 the resistance measured at low strain levels consistently decreased linearly with increasing strain (negative piezoresistivity) in area 1. As the strain was increased further (area 2), the resistance reduced more slowly and eventually exhibited a minimum at a critical piezoresistivity transition strain (εc). At strain levels beyond εc in area 3, the resistance increased slowly with increasing strain (positive piezoresistivity), whereas an exponential increase in electrical resistance was observed at a large strain range (area 4). In area 4, a higher increase in the rate of resistance change at higher strains was clearly observed, which may be attributed to the destruction of the conductive networks in the matrix. While similarly piezoresistive behavior has been previously reported for graphene- and CNT-based polymer composites,44-46 the peculiar


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response at such a low strain observed in the present CNT based system can be attributed to the reorientation (rotation and uncurling) of high-aspect-ratio CNTs at low strains and the corresponding strain-induced the slippage between conductive fillers. During the process of stretching/releasing, the conductive network is destructed and reconstructed by the movement of polymer segments, leading to changes in the resistance of the nanocomposite.40,47 According to other CNT-based nanocomposites, the changes in electrical resistance during stretching can be described by the deformation of the CNT conductive networks and the increase in the tunnel distance between adjacent conductive particles. At large strains, the second behavior (tunneling conduction) may be dominant.42,48 To better understand the observed electromechanical properties mentioned above, a mechanism of the evolution of conductive phase morphology during different stretching stages is illustrated in Figure S5 (Supporting Information). In the original state (Figure S5a), CNTs are curled and exhibit a uniform distribution in the polymer matrix. At a low level of strain (ε < εc, in Figure S5b), the mean interparticle separation decreases due to the Poisson effect (extrusion or compression in the perpendicular direction) and rotation of CNTs that are out of plane, leading to improved particle alignment and greater electrical conductivity (negative piezoresistivity). At much higher strains (in Figure S5d), the destruction and reconstruction of the conductive CNT networks coexist in this stretching stage. However, the destruction of aligned CNT pathways significantly dominates, so that conductivity pathways are broken and the resistance exhibits an obvious increase (positive piezoresistivity) in area 3 and 4. At the critical state (ε = εc, in Figure S5c), the destruction and reconstruction of the conductive CNT networks are balanced, leading to the minimum level of resistivity.


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In the piezoresistive strain sensor, gauge factor (GF) is a characteristic parameter representing the sensitivity of the strain sensors, and can be derived from (ΔR/R0)/ε, where ε is the strain. The highly stretchable CSF-2 strain sensor exhibited GF = −0.063 below the critical strain (about 25%), GF = 0.68 for linear region (50%–100%), and GF = 1378 at maximum stretching up to 330%. Similarly, the CSF-3 strain sensor exhibited GF = −0.45 below the critical strain (approximately 110%), GF = 0.44 for the linear region (150%–300%) and GF = 153 at maximum stretching up to 600%. Generally, the piezoresistive composites with filler concentrations around the percolation threshold should have higher sensitivity to strain, as explained by percolation theory. Therefore, the GF with a CNT loading content of 2 wt% is much higher than that of a composite with higher filler concentration. Furthermore, it is evident that higher strain sensitivity was obtained at a high level of strain; the distinct sensitivity was attributed to the large amount of breakdown in the conductive network under higher strain amplitude, causing larger variation in electrical resistance. These results also indicate that the sensitivity (GF) and sensing range of stretchable CSFs can be easily tuned by controlling the filler concentration. We further compared the gauge factors of our sensors with the maximum values reported in the literature, as shown in Figure 4b. This comparison revealed that the proposed coaxial fiber-shaped sensor provides one of the largest working ranges of strain (>300%), but also offers a high GF of 1378 (see a detailed comparison in Table S1, Supporting Information). Although previous reported strain sensors achieved higher GF, they were typically limited by lower stretchability (ε 50% is required.49,50 For example, a previous study reported that a graphene-based strain sensor had a very high GF (about 1000) but very limited strain sensing range of 0–2%,51 limiting its practical applications for large-scale motion detection. On the contrary, strain sensors with a wide workable sensing range (280%)


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have been prepared using carbon nanotube film. However, the GF of this system was only 0.06 within a strain range of 60%–200%.52 It should be noted that the CSF strain sensors discussed in the following section are CSF-2, except for those specifically defined.

Figure 4. Strain sensing properties of the CSF strain sensor. (a) Resistance change (ΔR/R0) as a function of the applied strain with different CNT content, i.e., 2 wt% and 3 wt%. (b) Comparison


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of the gauge factor and maximum sensing range of the CSF strain sensors with those of other piezoresistive sensors reported in the literature. (c) Relative resistance changes versus tensile strain of 50%, 100%, 150%, 200%, and 250% at a frequency of 0.5 Hz. (d) Relative resistance variations versus a tensile strain of 150% at frequencies of 0.05 Hz, 0.1 Hz, 0.2 Hz, 0.5 Hz, and 1 Hz. (e) Time response of the strain sensor upon applying a quasi-transient step strain form ε = 0% to ε =100%. (f) Magnified sensor responses extracted from (e) to show response time. (g) Relative change in resistance under an isosceles trapezoid strain profile. The magnitudes of the respective peak strains are 100%, 150%, and 200%. (h) Hysteresis performance of CSF sensors (strain rate: 10% s-1). (i) Relative change of resistance under repeated stretching/releasing cycles with 100% strain at a frequency of ~0.25 Hz for 10 000 cycles, demonstrating the durability of the CSF sensor. To confirm the reliability and wide detection range of the sensor, dynamic tensile with different amplitude (50%–250%) at a frequency of 0.5 Hz were applied to the sensor. As shown in Figure 4c, the relative change in resistance showed a stable and repeatable signal without obvious degradation, suggesting that the strain sensor can operate within a very wide strain range. Furthermore, the response of the CSF strain sensor under cyclic tensile strain at a frequency range of 0.05 Hz to 1 Hz was investigated. Figure 4d shows the relative change of the resistance for the CSF sensor, which exhibited frequency dependence within the test frequency range (0.05 Hz–1 Hz) when strain was applied. From this figure it can be seen that the responses of the CSF sensor match perfectly with the tested signals, and the relative resistance change increases with an increase of the frequency of the applied strain. It is well known that a higher strain rate causes greater stress in the materials because the mobility reduction of polymer chains under a high


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strain rate leads to an increase of inherent stiffness, which leads to an increase in the amplitude of relative resistance change at high frequencies.53 Next, we investigated the response of the proposed sensor to a quasi-transient step strain input. The fiber-like sensor was stretched to 100% strain at a loading speed of 500% s-1, held at 100% strain for 5 s, and then released to 0% strain state as the same rate. Figure 4e shows the relative resistance change during the first 10 cycles. During quasi-transient step testing, the CSF sensors showed overshooting in response to acceleration, followed by relaxation back to a plateau value. Similar behavior has been previously reported.29,39 The sharp overshoots associated with acceleration can be attributed to the viscoelasticity of the silicone elastomer composite. Finally, upon relaxation, the strain sensors did not recover their original resistance value. Resistance at zero strain increased irreversibly during the initial loading cycle, because of the irreversible destruction of the conductive networks reducing the number of conductive pathways. In general, the degree of irreversible deformation depends on the magnitude of the maximum strain applied. This behavior may be attributed to the partial loss of particle-particle interactions, as suggested by the Mullins effect in filled rubber.54,55 Additionally, Figure 4f shows the response time of the CSF sensors. From this figure it can be seen that the response time was less than 300 ms (step strain of 100%, loading speed of 500% s-1). If the delay of the measurement system was considered, the actual response time for the sensor itself would be even shorter than this value. The fast response would be expected to facilitate the real-time monitoring of intricate body activities, such as running. Figure 4g shows the CSF strain sensor was subjected to a tensile cyclic test with an isosceles trapezoid profile of which the peak strain values were 100%, 150%, and 200%. From this figure it can be noted that the relative change of resistance was wellmatched to the strain profile.


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Furthermore, hysteresis characteristic is also a significant consideration in most applications of strain sensors. Figure 4h exhibits the hysteresis curve of the uniaxial tensile test measured during loading and unloading cycles under different levels of applied strain (100%, 150%), indicating that the slight hysteresis of resistance during the stretching–releasing process, which may be attributed to the viscoelasticity of the silicon elastomer. In the present case, the outstanding hysteresis performance of the CSF strain sensors was achieved by the strong interaction between the polymer chains and the MWCNTs. These findings compare favorably with previously reported results.56 In addition, the durability and stability of the CSF strain sensors were investigated, owing to their importance in practical applications. To demonstrate the durability and repeatability of the CSF sensors, about 10000 cycles (more than 10 hours) of 100% strain at a frequency of ~0.25 Hz were measured for the stretchable strain sensor, and the relative change of resistance was recorded in real-time, as shown in Figure 4i. As shown in the enlarged insets, the resistance exhibited good repeatability in the 10000 cycles, indicating that the CSF strain sensors had a long working life and high stability. 2.4. Electrical Characteristics of CSFs Under Bending and Torsion. Investigating the bending and twisting sensing properties for strain sensors is important for practical applications, owing to bending and torsion are ubiquitous in engineering practice. We first examined the effects of bending-induced strain on the electromechanical performance of the CSFs. As shown in Figure 5a, the resistance change of the composite fiber was less than 1% under mechanical bending with curvature ranging from 0 to 1/2.5 mm-1, suggesting that there was no severe breakdown of the conductive pathways inside the specimen upon bending. Torsion on the CSF is a more complex mechanical deformation, and is commonly observed in the movement of large joints of human, natural muscles, and even artificial joints of robots. Here, the torsion level was


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defined as the ratio of the twist angle and the length of the stretchable fiber. Figure 5b shows the curve of the electrical resistance change versus the torsion level. With a torsion level from 0 rad m-1 to 251 rad m-1(1 turn), the electrical resistance change is negligible and less than 5% at a torsion level of 251 rad m-1. Thus, CSF also exhibited remarkable stability against slight twisting. These results suggest that the CSF strain sensor was able to delicately sense strain signals.

Figure 5. Bending and torsion sensing performance of CSF. (a) Relative resistance change of the fibrous sensor as a function of curvature. (b) Relative resistance change versus torsion level up to 2000 rad m-1(~8 turns). Inset: structural illustration of the twisting CSF. 2.5. Environmental Stability and Washability of the CSF Sensors. Environmental factors such as humidity and temperature can affect the performance of sensors. Therefore, investigating the effect of these external factors is necessary for meeting the requirements of long-term environmental stability. In the current study, the response of the CSF sensor as a function of


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relative humidity (RH) was achieved by exposing the CSF inside the closed vessels with different RH environments for the uptake of water molecules. A schematic diagram of the experimental setup for humidity exposure is shown in Figure S6. The relative change in resistance of the CSF exposed to varying RH is shown in Figure 6a. The relative electrical resistance variation exhibited no significant difference as the RH increased from 10% to 90%. Furthermore, Figure 6b shows the real-time measurements of the resistance of CSF as RH was adjusted between 65% and 90% for six cycles (a longer exposure interval of 10 min was performed for reaching saturation). From this figure it can be seen that the electrical properties were less affected by humidity. The verified humidity stability of CSFs can be explained as follows. The conductive core was fully sealed by superhydrophobic silicone rubber, preventing water molecules from absorbing onto the CSF surface. Thus, the insulating layer process effectively protected the CNTs that interacted with external water molecules, thereby reducing the humidity sensitivity of CNTs. In addition, we also investigated the effect of temperature on the electrical properties of the CSF strain sensor. The zero-strain resistance change under temperature can be calculated by measuring the difference between the resistance at a defined temperature and resistance under ambient conditions. A CSF sensor was placed into a constant temperature chamber while its temperature was gradually increased from room temperature to 100 °C. The effect of temperature on the electrical resistance of the CSF strain sensor is shown in Figure 6c, indicating that the resistance of the strain sensor decreased with increasing environmental temperature from 25 °C to 100 °C (negative temperature coefficient; NTC). Similar behavior has been reported for CNTbased nanocomposites.57,58 This behavior can be explained by the semi-conductive charge transport behavior of the CNTs.59 Meanwhile, previous studies have shown that the charge


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transport in CNT thin film is dictated by a variable range hopping (VRH) mechanism, which facilitates the charge carrier mobility and reduces the resistance at high temperature.60 Nevertheless, the CSF sensor is designed for wearable uses, normally functioning within a relatively small temperature range, from room temperature to body temperature. Therefore, the small effect of temperature on the CSF sensor can be neglected for wearable applications. These results indicate that the perfect stability of the CSF strain sensor in response to both humidity and temperature changes. For practical applications, further work is required to address the effect of relative humidity and temperature on the sensitivity and mechanical properties of this strain sensor. In addition, we assessed the washability of the fibrous sensor, which is an important characteristic for wearable electronics on textiles. The resistance changes of CSF after repeated washing were recorded (Figure S7, Supporting Information). As shown in the figure, the relative change of resistance increased only slightly after the first washing cycle. These results indicate that the performance of CSF sensors was reliable, supporting their broad application in wearable electronics.

Figure 6. Electrical properties of the CSF under different external environmental conditions. (a) Relative resistance change as a function of humidity. RH increased from 11% to 95%. (b) Relative resistance change of the CSF under changing RH (variation between 65% and 95%). (c)


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Electrical resistance variation of the CSF as a function of temperature. The temperature increased from 25 °C to 100 °C. 2.6. Application as Wearable Sensors for Human Motion Detection. To demonstrate the utility of CSFs in wearable devices, we mounted them onto textiles by either sewing or weaving, and tested their capacity to detect human motion. We selected ~1 mm for the inner diameter and ~0.35 mm for the wall thickness as the dimensions of the fabricated CSFs for the subsequent experiments. Owing to their excellent stretchability, high sensitivity and wide sensing range, we demonstrated the applications of CSF sensors for detection of both subtle motion and large motion. In addition, the CSF with a Young’s modulus of ~0.49 MPa (Figure S8) exhibits mechanical compliance as high as that of the human skin and meets the requirements for human motion detection. The ability to detect very small movements, such as pulse, respiration, and muscle movements, is important for electronic skin sensors. To explore the application for human physical signal detection, the CSF strain sensor was directly attached to the artery of the wrist of a volunteer subject using medical tape, for the detection of pulse (see Figure 7a). Figure 7b shows the realtime resistance change of the sensor during relaxation. The results revealed stable and regular pulse signals with a frequency of 67 beats min-1, indicating the excellent sensitivity of the strain sensor. Respiration is an important physiological signal for human health. We demonstrated that the CSF sensor functioned as a resistive belt breathing sensor with high sensitivity. The respiration sensor employed a CSF connected to a high durability latex rubber band which can be worn over clothing at the abdomen (Figure 7c). Figure 7d shows the relative resistance change of the strain sensor which worked as a respiration sensor (at rest). The upward section in the sensor output signal represents inhalation and the downward section of the output curve


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represents exhalation. The sensor produced repeatable and regular signal patterns at rest with a respiration rate of 20 breaths min-1. It is well known that the normal respiration rate for an adult during relaxation is 12–20 breaths min-1; a respiration rate under 12 or over 25 breaths min-1 while resting is considered abnormal. Furthermore, a CSF sensor was attached to the neck to detect the muscle motion near the throat (Figure 7e). When the subject swallowed saliva, resistance changes of the sensor were detected. As shown in Figure 7f, similar resistance changes were produced when the subject swallowed saliva repeatedly, indicating the high sensitivity and stability of the sensor. In addition, the fiber-shaped strain sensors were attached at the corner of the eye (see Figure 7g) to monitor the subtle muscle motion induced by blinking. As shown in Figure 7h, the resistance variation of the sensors induced by blinking could be precisely recorded. The fast and sensitive monition of subtle motion supports the potential applications in human–machine interaction and health monitoring for this type of wearable sensor. Furthermore, to demonstrate the potential of the CSF sensors in large-scale human motion, the CSF sensors were woven into the index finger of fabric gloves using a sewing method. Figure 7i shows a glove assembled with a CSF sensor, which can detect real-time motion of the index finger. When a volunteer subject wore the glove and repeatedly bent the index finger, the response of the CSF sensors to finger motion could be clearly observed, as shown in Figure 7j. From this figure, it can be noted that the resistance of the CSF sensor increased rapidly during bending of the finger (see Movie S2, Supporting Information). In addition, when the finger was held in a bent state, a stepwise response was observed. The resistance signal of the wearable sensor showed a consistent pattern, indicating high levels of stability and repeatability,


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suggesting that the glove system with CSF sensors could be used for fine-motion control in robotic systems and other virtual reality devices.

Figure 7. Application of the CNT-based fibrous sensor as a wearable device for use with humans. (a) Photograph of the CSF sensor attached to the wrist. (b) Arterial pulse waves under normal conditions. (c) Schematic illustration showing a breath sensor fabricated by attaching a CSF strain sensor directly above the abdomen. (d) Relative resistance change of the fibrous sensor over time during human breath monitoring. (e) Photograph of the sensor comfortably attached to the throat of a person. (f) Normalized change of the resistance of the wearable sensor over time during saliva swallowing. (g) Optical images of eye-opening (bottom) and eye-closing (top). (h)


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Relative resistance change of the sensor showing the small muscle movement caused by blinking. (i) Photograph of a wearable sensor assembled on a commercial glove. (j) Response of wearable sensor in monitoring finger bending. Insets show photographs of index finger motion. (k) Optical image of a strain sensor attached to the wrist. (l) Response of the wearable sensor to the cyclic motion of the wrist. (m) Sensor responses under repeated dynamic flexion and straightening motion of the elbow. (n) Sensor response to sequential flexion motion of the elbow. (o) Responsive curves of CSF sensors on the knee during flexing–extending motions. Inset shows the CSF strain sensor placed directly above the knee of a volunteer subject. We also wove a CSF strain sensor into a commercial wristlet over the wrist joint (Figure 7k) to detect wrist motion. Such a sensor was able to easily detect wrist extension and flexion. Figure 7l shows the relative resistance change of the wearable CSF sensor in response to periodic motion of the wrist. As seen in Figure 7l, the normalized change of the resistance over several bending cycles demonstrates a high degree of reproducibility. In addition, the proposed wearable CSF sensor was also able to precisely detect elbow motion, as shown in Figure 7m. Moreover, testing was conducted during performance of a stepped bending sequence in which the elbow was bent and held for a few seconds at each position. Figure 7n shows the relative resistance change of the CSF sensor at different bending stages over time. It can be seen that the resistance of the CSF sensor increased as the bending angle increased. Thus, the CSF strain sensor was capable of distinguishing the different bending angles of the joints, demonstrating that the system is able to both detect and quantify human joint bending. Furthermore, when the CSF sensor was used to further monitor other larger types of human motion, such as knee bending, stable and repeatable responses were also observed, as shown in Figure 7o. The CSF strain sensor was able to accurately detect motion over several cycles, with resistance increasing with bending angle.


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Long-term accurate monitoring of human joint motion could have important applications in medicine and rehabilitation. These results indicate that the proposed flexible fibrous strain sensors could be used for detecting motion of the upper and lower limbs. Their small size, large stretchability, and stability make them a suitable candidate for applications in these fields. Overall, these findings demonstrate that the proposed sensors can be successfully used for monitoring both subtle human motion (such as pulse and respiration) as well as vigorous motion (such as finger and elbow movements), suggesting that the wearable and stretchable fiber could have a range of potential applications in human–machine interaction, health monitoring, and soft robotics in the near future. 3. Conclusions We presented a one-step wet-spinning assembly approach for continuously spinning polymer elastomer-wrapped carbon nanocomposite core-sheath fibers (CSFs) which can be used directly as stretchable, washable, and safe strain sensors. The fabrication strategy was facile, low-cost, ultrafast, and scalable for preparation of polymer-based core-sheath fibers which have not previously been produced using continuous wet-spinning. The proposed CSF strain sensors exhibited a tolerable strain of more than 300%, and GF of 1378. The flexible fiber-shaped strain sensor exhibited a suitable trade-off between the contrary properties of sensitivity and stretchability, enabling it to detect human motion at different scales. Importantly, the CSF strain sensor had high durability (>10000 cycles), fast response, low hysteresis, excellent washability, and bending, torsion and humidity stability. Furthermore, we demonstrated that the CSF strain sensors could be used for monitoring both subtle motions, such as muscle motion and respiration, and large movements, such as bending of joints, suggesting a wide range of potential


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applications in flexible and stretchable devices. To the best of our knowledge, this is the first study utilizing a wet-spinning technique for fabricating stretchable silicone polymer fibers. Moreover, this facile fabrication technology and unique structure could be extended to fabricate other functionalized composites with useful properties. Experimental Section Materials and Characterizations. The silicone elastomer was prepared from Ecoflex 00-30 Cure Part A and B, and Slo-Jo Platinum Silicone Cure Retarder, all obtained from Smooth-On, Inc. The silicone cure retarder is an additive designed to extend the pot life of silicone rubber to maximize spinning time. The vulcanized elastomeric matrix exhibits large stretchability (up to 900%) and low modulus, making it ideally suited for wearable and stretchable electronics. MWCNTs with 98% purity and 10–30 µm length were supplied by Chengdu Organic Chemicals Co. Ltd. of the Chinese Academy of Science and used without further purification. Dimethyl silicone oil (1000 cs) was used as the coagulation agent. SEM images were recorded with a field emission scanning electron microscope (FE-SEM, Zeiss GenimiSEM 500). Energy Dispersive Spectroscopy (EDS) signals were collected using detectors from Oxford Instruments. The quality of MWCNTs was characterized by Raman spectroscopy HR800 (Jobin Yvon Horiba) with a laser excitation wavelength of 633 nm. Preparation of the Spinning Solutions. First, the viscous silicone elastomer solution was synthesized by mixing Part A and Part B with a weight ratio of 1:1. It is noteworthy that the retarder is added into Part B by weight to Part B (1 wt%) before adding Part A. A series of spinning inks of CNTs in silicone elastomer (1–3 wt%) was then prepared by dissolving the required amount of CNTs in silicone elastomer solution, followed by mixing for 10 min. Finally,


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the mixtures were degassed in a vacuum for an additional 5 min to obtain homogeneous spinning inks. The final spinning formulations were loaded into the syringe without introducing air for wet-spinning. Rheological Characterization. The rheological properties of the inks used were determined using a controlled-stress rheometer (MCR302, Anton Paar, Austria) with a 25 mm diameter parallel plate geometry for suspensions, with CNT content ranging from 1 to 3 wt%. Viscometry measurements were performed in a frequency range of 0.1–1000 Hz, at a temperature of 25 °C, while oscillatory measurements were carried out at a frequency of 1 Hz within the stress range of 0.1–1000 Pa. All rheological measurements of the inks were performed immediately after homogenization. Coaxial Wet-Spinning. The highly stretchable coaxial fibers were fabricated using the wetspinning method, as shown in Figure 1a. The coaxial spinneret was fabricated by inserting a 17G stainless steel needle (inner diameter 1.07 mm, outer diameter 1.47 mm) into a 13G stainless steel needle (inner diameter 1.9 mm, outer diameter 2.4 mm). Proper coaxial alignment of the core needle was achieved manually and the two lengths of needle were fastened and sealed with rubber stoppers. Two independent syringe pumps (LSP01-1A, Longer Precision Pump Co., Ltd. China) were used to control the extrusion of core/shell channels. The coagulation bath was filled with hot silicone oil (approximately 110 °C). First, the CNT-silicone elastomer inks and pure silicone elastomer solutions were transferred to two injection syringes connected with the inner and outer spinneret, respectively. The typical extruded velocities of the core and sheath layers were set at 150 µL min-1 and 100 µL min-1 respectively. As the spinning solutions exited the coaxial nozzle, the vulcanization process began immediately, after which it continued to vulcanize and be stretched by gravity. After full coagulation for 2 min (in hot silicone oil), the


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wet-spun fibers were immersed in a washing bath (ethanol), then wound onto a spool. The collection speed was set to keep the balance between fiber production and collection speed without applying any stretch to the fibers. Finally, the resulting fibers were further washed with ethanol and deionized water, followed by vacuum oven drying overnight. Electromechanical Measurements of CSF Strain Sensors. Copper wires were connected to the two ends of the CSF as external electrodes with the help of silver paste and silicone adhesive (Sil-Poxy, Smooth-On, Inc.). A detailed description of the sensor assembly process is illustrated in Figure S9, Supporting Information. The gauge length between the copper wires was ~13 mm. The silicone adhesive was used to cover the silver electrodes to avoid mechanical failure between the stretchable fiber and rigid electrodes. The fibrous sensor was placed in a computercontrolled, homemade motorized actuating system for the stretching test, with one side of the sample held by a fixed stage, and the other side held by a movable stage. Electrical resistance measurements of CSFs were carried out in a two-probe configuration using a source meter (Keithley 2450). Stability testing was performed up to 10000 cycles with various levels of strain, from 0 to 100%. To examine the bending behavior of the series composites, specimens were bent around an insulating cylindrical object of known radius while electrical resistance was measured. Various constant humidity environments were achieved using humidity bottles containing saturated solutions of different metal salts. LiCl, MgCl2, Mg(NO3)2, NaCl, KCl, and KNO3 were sealed in glass bottles at 25 °C to obtain RH conditions of 11%, 33%, 54%, 75%, 86%, and 94%, respectively. The relative humidity was checked by a commercial digital humidity meter (MS6508). Temperature response was conducted in an environmental testing system (YOMA


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MHU-02, China). The temperature range of the system was −70–150 °C, and the precision of temperature measurement was ±0.5 °C. The sensor was washed in a commercially available detergent solution with mechanical stirring (at 500 rpm) for 10 min. After drying, the resistance signals of the CSF sensor were measured.

Supporting Information. The Supporting Information is available free of charge on the ACS Publications website at DOI: XXX. Figure S1. (a) Photograph of the experimental setup for coaxial spinning. (b) Picture of a coaxial two-capillary spinneret (17G/13G). Figure S2. (a) Cross-sectional optical microscope image of the CSF (2 wt%). (b) Magnified SEM image of the core sheath interface of CSFs. Figure S3. (a) High-resolution SEM image and (b) EDS spectrum of CNT-based fiber. Figure S4. Definition of the four stages in the relative change of resistance versus strain curve of CSF-3. Figure S5. Schematic illustration showing the evolution of the conductive phase morphology during different stretching stages. (a) Initially unstrained. (b) The stretching state lower than the critical strain. (c) The stretching strain up to critical strain. (d) At high strain state above critical strain. Figure S6. Schematic diagram of humidity investigation experimental setup.


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Figure S7. Relative resistance change of CSF as a function of washing cycles. Figure S8. The typical stress−strain curve of CSF. Figure S9. Schematic illustration of sensor assembly process. Table S1. Selected parameters extracted from our work and the recently reported flexible and stretchable piezoresistive strain sensors. Movie S1. Coaxial wet-spinning assembly process of CSFs.(.avi) Movie S2. Wearable sensor in real-time monitoring of finger motion.(.avi) AUTHOR INFORMATION Corresponding Author *E-mail: [email protected] Notes The authors declare no competing financial interest. Funding Sources The National Natural Science Foundation of China-Shanxi Provincial Government Coal-Based Low Carbon Joint Fund (Grant No. U1510114) and the National Natural Science Foundation of China (Grant No. 51575437).

ACKNOWLEDGMENTS


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This work was financially supported by the National Natural Science Foundation of ChinaShanxi Provincial Government Coal-Based Low Carbon Joint Fund (Grant No. U1510114) and the National Natural Science Foundation of China (Grant No. 51575437). The authors also would like to thank Wen Wang (School of Science, XJTU) for help in rheological measurements, Liangquan Zhu for helping with the experimental tests, and Zijun Ren for SEM images. REFERENCES (1) Gong, S.; Lai, D.; Wang, Y.; Yap, L. W.; Si, K. J.; Shi, Q.; Jason, N. N.; Sridhar, T.; Uddin, H.; Cheng, W. Tattoolike Polyaniline Microparticle-Doped Gold Nanowire Patches as Highly Durable Wearable Sensors. ACS Appl. Mater. Interfaces 2015, 7 (35), 19700−19708. (2) Amjadi, M.; Kyung, K.-U.; Park, I.; Sitti, M. Stretchable, Skin-Mountable, and Wearable Strain Sensors and Their Potential Applications: A Review. Adv. Funct. Mater. 2016, 26 (11), 1678−1698. (3) Yao, H. B.; Ge, J.; Wang, C. F.; Wang, X.; Hu, W.; Zheng, Z. J.; Ni, Y.; Yu, S. H. A Flexible and

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