Multi-role conductive hydrogels for flexible transducers regulated by MOFs for monitoring human activities and electronic skin functions

Abstract

Metal organic frameworks (MOFs) have garnered significant attention in the development of stretchable and wearable conductive hydrogels for flexible transducers. However, MOFs used in hydrogel networks have been hampered by low mechanical performance and poor dispersibility in aqueous solutions, which affect the performance of hydrogels, including low toughness, limited self-recovery, short working ranges, low conductivity, and prolonged response–recovery times. To address these shortcomings, a novel approach was adopted in which micelle co-polymerization was used for the ex situ synthesis of Zn-MOF-based hydrogels with exceptional stretchability, robust toughness, anti-fatigue properties, and commendable conductivity. This breakthrough involved the ex situ integration of Zn-MOFs into hydrophobically cross-linked polymer chains. Here the micelles of EHDDAB had two functions, first they uniformly dispersed the Zn-MOFs and secondly they dynamically cross-linked the polymer chains, profoundly influencing the mechanical characteristics of the hydrogels. The non-covalent synergistic interactions introduced by Zn-MOFs endowed the hydrogels with the capacity for high stretchability, high stress, rapid self-recovery, anti-fatigue properties, and conductivity, all achieved without external stimuli. Furthermore, hydrogels based on Zn-MOFs can serve as durable and highly sensitive flexible transducers, adept at detecting diverse mechanical deformations with swift response–recovery times and high gauge factor values. Consequently, these hydrogels can be tailored to function as wearable strain sensors capable of sensing significant human joint movements, such as wrist bending, and motions involving the wrist, fingers, and elbows. Similarly, they excel at monitoring subtle human motions, such as speech pronunciation, distinguishing between different words, as well as detecting swallowing and larynx vibrations during various activities. Beyond these applications, the hydrogels exhibit proficiency in distinguishing and reproducing various written words with reliability. The Zn-MOF-based hydrogels hold promising potential for development in electronic skin, medical monitoring, soft robotics, and flexible touch panels.

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1. Introduction

Flexible transducers have garnered significant attention for their broad range of potential applications, including electronic skin,^1–3 human movement detection,^4–6 health monitoring,^7–9 and energy storage.^10–12 These transducers can differentiate between biological and mechanical activities, transmitting feedback through changes in their electrical output.¹³ However, existing sensor materials face limitations such as restricted operational range, prolonged response time, and low conductivity degradation, hindering their integration into wearable devices.^14–16 Thus, there is a pressing need for highly flexible sensors capable of maintaining conductivity after enduring numerous deformations, enabling the creation of more complex and adaptable wearable devices.

Hydrogels, characterized as a class of soft water-based materials, present substantial promise for incorporation into wearable sensors owing to their exceptional biocompatibility, elevated stretchability, and customizable conductivity.^17–19 To make hydrogels suitable for a diverse array of applications, researchers have implemented a variety of techniques, including chemical cross-linking,^20–22 composite hydrogels,^23–25 double-network hydrogels^26–28etc. Yet, for hydrogels to be appropriate for flexible transducers, they must possess formidable mechanical resilience and robust self-recovery to facilitate large-range strain sensing and enduring cycling stability. Consequently, scientists have made intensive efforts thus far to create hydrogels endowed with fast self-recovery and exceptional toughness, achieved by incorporating dynamic non-covalent bonds, which helps to effectively dissipate energy.^29,30

A range of highly stretchable and durable hydrogels have been developed by integrating reversible non-covalent interactions such as ionic and electrostatic bonds,^31–33 hydrogen bonding,^34–36 crystallization,^37–39 and hydrophobic interactions.^16,40,41 Composite hydrogels, integrating reversible non-covalent interactions, have shown exceptional elongation and enhanced electrical conductivity. Nanoscale fillers like clay, liquid metals, and cellulose nanocrystals improve hydrogel properties through electrostatic attraction or hydrogen bonds.^42–45 However, challenges such as weak interactions and low dispersibility result in hydrogels with low mechanical strength and hysteresis.

Hydrophobic materials like multiwalled carbon nanotubes (MWCNTs), graphene, and metal organic frameworks (MOFs) have been utilized in polymer composites. MOFs, with their crystalline structures and high surface areas, present opportunities for enhancing hydrogel properties.^13,46,47 However, their low dispersibility and weak mechanical performance pose challenges. Efforts to improve water solubility and interfacial interactions are crucial for achieving homogeneous conductivity and mechanical stability. Improving water solubility by introducing a hydrophilic group or boiling water is a widely adopted approach.^48,49 However, such modifications may alter the structure and properties of MOFs. Additionally, weak interfacial interactions between MOFs and the hydrogel network can result in easy movement of fillers within the hydrogel under external stimuli, potentially compromising the deformable network and resulting in insufficient mechanical properties. Furthermore, like traditional hydrogels, conductivity responses to external mechanical forces typically exhibit hysteresis due to their high viscoelasticity.^50–52 Enhancing the sensitivity and reliability of MOF-hydrogel-based transducers is important. Therefore, a straightforward and efficient method for integrating MOFs into hydrogels with outstanding performance is crucial to meet practical applications in flexible transducers and skin-like materials. If MOFs can be uniformly dispersed in the hydrogel network and serve as robust cross-linking sites, this would markedly enhance interactions between MOFs and the hydrogel network, subsequently elevating the mechanical properties of the conductive hydrogel and preventing the displacement of MOFs when subjected to external mechanical stimuli.

A one-pot approach was applied to fabricate Zn-MOF-based conductive hydrogels, introducing Zn-MOFs via an ex situ method into a hydrophobically crosslinked hydrogel network. The ethylhexadecyldimethylammonium bromide (EHDDAB) surfactant acted as a dynamic linker, facilitating uniform dispersion of Zn-MOFs and enhancing mechanical performance through non-covalent interactions. The resulting hydrogels exhibited good conductivity, high mechanical performance, cyclic stability, and sensitivity, making them ideal for flexible transducers. These transducers can monitor large and physiological movements, serve as pressure sensors, and mimic electronic skin for various activities. The addition of Zn-MOFs into hydrogels offers a promising solution for overcoming the limitations of traditional sensor materials, paving the way for the development of advanced flexible wearable devices with enhanced functionality and performance.

2. Experimental

2.1. Chemicals

The chemicals utilized in this work were of analytical grade and were employed without any chemical modifications. Ethylhexadecyldimethylammonium bromide (EHDDAB), dodecyl methacrylate (DDMA), and acrylamide (AM) were obtained from Acros Organics. Sodium chloride (NaCl), 2-(dimethylamino)ethyl acrylate methochloride (DMAEMC), tetramethylethylenediamine (TEMED), and ammonium persulfate (APS) were acquired from Sigma Aldrich, while pre-synthesized zinc metal organic frameworks (Zn-MOFs) were also utilized (Text 1 and Fig. S1–S3, ESI†).

2.2. Fabrication of MOF-based hydrogels

To synthesize the hydrogels, a single-step method was used with a free radical micelle copolymerization by applying APS as the radical source. Initially, 0.6 g of EHDDAB, a cationic surfactant, and 0.2 g of NaCl were dissolved in 10 ml of triple distilled water at 50 °C and 600 rpm. The solution was then cooled to room temperature before adding 0.2 ml of the hydrophobic monomer DDMA and stirring for an additional 20 minutes. Subsequently, the solution was mixed with varying amounts (10 mg, 20 mg, and 30 mg) of Zn-MOFs, stirring at 1200 rpm until complete dispersion and the attainment of a homogeneous solution. Next, 1.9 g of AM, 100 μl of a cationic monomer DMAEMC, and 50 mg of APS were added, and the solution was stirred until homogeneity was achieved. Finally, 5 μl of TEMED was added and stirred for an additional 30 seconds. The resulting solution was transferred to a plastic mold and placed in an oven at 50 °C for 1 hour to obtain hydrogels. After 1 hour, the hydrogels were cooled at ambient temperature and stored in air-sealed sample bags for further use. In order to investigate the effect of MOFs, a pure hydrogel without MOFs was also prepared. Samples were labeled as Z0, Z1, Z2, and Z3 with numerical values indicating the amounts of MOFs used.

2.3. Functional and morphological characterization

The functional and morphological characteristics of the prepared p(DDMA-AM-DMAEMC) Zn-MOFs hydrogels were assessed using Fourier transform infrared spectroscopy (FTIR) and scanning electron microscopy (SEM). Prior to characterization, the hydrogels underwent a 24-hour freeze-drying process. FTIR analysis was conducted using a Bruker Corp Vertex-70V spectrometer, while SEM imaging was performed using a Coxem Corp. EM-30N.

2.4. Mechanical properties

The mechanical evaluation utilized a Lloyd Instruments’ universal testing machine (LR 10k) equipped with a 10 kN load cell. Tensile and tensile cyclic tests were employed, involving the cutting of prepared hydrogels into rectangular shapes with known dimensions. During testing, a constant speed of 50 mm min⁻¹ was maintained. From the tensile tests, parameters such as toughness, Young's modulus, and fracture stress–strain were derived, while cyclic tests provided data for calculating the dissipated energy.

2.5. Rheological study

Strain and frequency sweep tests were conducted in the rheological study using an Anton Parr Physica 302 instrument, equipped with a measuring tool (16 290) with a diameter of 25 mm. The tests were carried out at a fixed temperature of 25 °C. The strain range spanned from 0.01 to 1000%, while the frequency ranged from 0.1 to 100 rad s⁻¹. A fixed amplitude of 5% was maintained for the frequency sweep tests. Circular hydrogel samples with a diameter of 13 mm were used for both strain and frequency tests.

2.6. Conductivity measurement

The conductivity of the hydrogels was determined using an LCR meter obtained from Hioki (Japan). Initially, the resistivity of the hydrogels was computed using the formula ρ = RA/L, with ρ denoting the resistivity, R representing the resistance, A denoting the cross-sectional area of the hydrogel, and L indicating the length of the hydrogel. The reciprocal of the resistivity yielded the conductivity (σ), i.e., σ = 1/ρ.

2.7. Strain sensitivity and human epidermis sensor

For strain sensing and human monitoring, the p(DDMA-AM-DMAEMC) Zn-MOF hydrogel Z2 was employed. In the strain sensing test, a rectangular piece of hydrogel was connected to a DC voltage source (3 volts) manually. The strain sensing and related parameters were then determined using an auto lab (electrochemical work station), and data were recorded in a current vs. time graph. These data were subsequently converted into resistances using Ohm's law, resulting in relative resistance (ΔR/R₀ (%)). To monitor human activities both large and subtle movements were used, the hydrogel (Z2) electrode was fabricated by covering both ends with copper sheets, which were connected to the instrument via wires. The hydrogel was then utilized to monitor various human activities. In writing tests, the hydrogel Z2 sensor was created by placing the hydrogel between two copper sheets, and was employed to assess the sensing capabilities of the transducer.

2.8. Statement

Informed consent was obtained for the experiments involving human participants.

3. Results and discussion

3.1. Synthesis

Initially Zn-MOFs composed of zinc through the hydrothermal method were prepared. This synthesis involved using Zn(NH₃)₂ as a source of metallic ions, with TMA acting as an organic linker. However, the challenge encountered was the poor dispersibility of these MOFs, limiting their utility within the hydrogel network. To address this dispersibility issue, we introduced the surfactant EHDDAB as a dispersing agent, effectively adding in the dispersion of the Zn-MOFs. After successful preparation of Zn-MOFs, Zn-MOF-based hydrogels were developed through ex situ micelle co-polymerization. This process involved combining Zn-MOFs, AM, DMAEMC, and DDMA monomers, with APS serving as a radical source to initiate polymerization. Our approach began with the creation of cationic micelles in water using EHDDAB as the surfactant, NaCl, and incorporation of the hydrophobic monomer DDMA, which dissolved within the micelles, serving as dynamic crosslinking points.^41,53 In the absence of EHDDAB, the Zn-MOFs could not disperse, leading to MOFs settling in the solution, as indicated in Fig. 1(a). Notably, some of the surfactant molecules engaged in hydrophobic interactions with the Zn-MOFs, while others formed physical interactions, resulting in a stable and uniform dispersion solution (Fig. 1(b)). Upon introducing AM, DMAEMC, and APS, initiation was performed through heat, causing the monomers to polymerize. During this polymerization process, the DDMA polymer molecules within the micelles underwent polymerization, effectively stabilizing the micelles. These DDMA polymer segments formed copolymers with AM and DMAEMC, serving as crucial driving forces and dynamic crosslinking points, maintaining the structural integrity of the hydrogel network, which was further reinforced by the presence of the Zn-MOFs. Fig. 1 provides a clear visualization of the interactions involved in the hydrogel formation. It is noteworthy that the Zn-MOFs were crosslinked with TMA, which carried a negative charge. In contrast, the monomer DMAEMC and the surfactant EHDAB bear positive charges, allowing them to interact non-covalently and ultimately enhancing the mechanical performance of the hydrogel dispersion.

Fig. 1 Diagrammatical representation of the hydrogel structure and different interactions involved in Zn-MOF-based conductive hydrogels. It can be noted that the Zn-MOFs without surfactant could not be dispersed and settled down (a), while in the surfactant solutions (EHDDAB), Zn-MOFs were uniformly and homogeneously dispersed (b).

To ascertain the presence of Zn-MOFs and to probe the chemistry between Zn-MOFs and polymer chains, FTIR spectroscopy was employed as a diagnostic tool, as illustrated in Fig. S4 (ESI†). Notably, in Z0, a wide absorption band at 3365 cm⁻¹ corresponds to the stretching vibration of NH₂ and ammonium quaternary groups.^54,55 Additionally, two distinct sharp bands at 2919 and 2854 cm−¹ signify the stretching vibrations of CH₃ and CH₂ within the hydrogel matrix. The sharp peak at 1657 cm⁻¹ is attributed to the stretching vibrations of C=O in the polymer chain associated with monomers. Moreover, the peak at 1451 cm⁻¹ indicates the bending of C–H.⁵⁶ Upon the introduction of Zn-MOFs into the Z2 hydrogel, a noticeable red shift is observed in the NH₂ group, with the broad peak shifting from 3365 cm⁻¹ to 3336 cm⁻¹. These shifts validate the interaction between the Zn-MOFs and the polymer chains.⁵³ Next the disappearance of peak around 1700–1740 cm⁻¹ of carboxylate groups evidenced the interaction between the Zn-MOFs and polymer chains. In order to look deep into the structure of the hydrogels the SEM analysis was performed for Z0 and Z2 hydrogels and the results are displayed in Fig. S5 (ESI†). It can be noticed from the figure that the Z0 hydrogel has a porous bumpy surface, and after the addition of Zn-MOFs the pore size decreases and compactness of the hydrogel structure increases. These changes in the hydrogel structure are due to the multiple crosslinking points that come with the addition of the Zn-MOFs, which is beneficial for the high mechanical performance and the conductivity of the hydrogel system.

3.2. Mechanical performance

The mechanical characteristics of the hydrogels were evaluated using manual and visual observation by humans (Fig. 2(a)–(g)). It can be observed in Fig. 2(a)–(c) that the hydrogel Z2 exhibits excellent extensibility under different conditions such as curling, knotting, and curl-knotting. Additionally, Fig. 2(d)–(f) demonstrate that the hydrogel possesses outstanding resistance to cuts; even a sharp cutter failed to cut the hydrogel, and no dents were observed on its surface. These resistance properties stem from the multiple non-covalent bonding points within the hydrogel, contributing to its resilience. Next, the hydrogel was subjected to full force to assess its shape recovery properties (Fig. 2(g)). Remarkably, the hydrogel exhibited significant shape memory, regaining its original shape after the applied force was removed.

Fig. 2 The photographs of the hydrogels having excellent mechanical stability in stretching form (a) curl, (b) knot, and (c) curl-knot, (d) and (e) represent the stretching capabilities and resistance to different forms of cuts, (f) resistance towards knife cut and (g) shape memory.

For quantitative measurement of the mechanical performance, tensile and cyclic tests were conducted on Zn-MOF-based hydrogels to evaluate toughness, Young's modulus, and dissipated energy. Tensile tests were performed on all hydrogels, and their results are depicted in Fig. 3(a)–(c). Fig. 3(a) demonstrates that the Zn-MOF content greatly influences the tensile properties of the hydrogel system. Fracture stress and fracture strain values for the Z0 hydrogel were 0.26 MPa and 1139%, respectively (Fig. 3(b)). Z1 and Z2 hydrogels exhibited positive variations with fracture stress values of 0.32 MPa (Z1), 0.67 MPa (Z2), and fracture strain values of 1313% (Z1), 1448% (Z2). This positive impact of Zn-MOFs is attributed to the crosslinking density, enhancing non-covalent interactions with polymer chains, thus boosting fracture stress and strain. However, further addition of Zn-MOFs in the Z3 hydrogel inversely impacted fracture strain (1121%), while fracture stress increased to 0.8 MPa due to higher crosslinking density. This suggests that the optimal Zn-MOF content for high mechanical performance in the current system is 20 mg. Similar to stress–strain values, Zn-MOFs positively affected toughness and Young's modulus (determined at 75% strain). Z0 hydrogel exhibited toughness and Young's modulus of 119 kJ m⁻³ and 0.2 kPa, respectively. These parameters linearly increased with Zn-MOF content, reaching maximum values of 475 kJ m⁻³ and 1 kPa, respectively, for the Z3 hydrogel. These results confirm the remarkable impact of Zn-MOFs on the mechanical performance, attributed to physical interactions between Zn-MOFs and polymers.

Fig. 3 Demonstration of the mechanical performance of hydrogels using UTM. (a) Tensile test, (b) estimated fracture stress and fracture strain, (c) calculated toughness and young's modulus, (d) cyclic loading unloading test at constant 800% up strain, (e) dissipated energy, and (f), (g) cyclic loading unloading at various strains and dissipated energies.

Among the four hydrogel systems, the Z2 hydrogel demonstrated superior mechanical performance and consequently was chosen for further study. Cyclic loading–unloading tests were conducted to determine the dissipated energy and fatigue performance. Initial tests were performed through ten cycles at a constant 800% strain and the dissipated energy was found from the area between the cycles (Fig. 3(d)–(e)). The dissipated energy for the first cycle exceeded that of the subsequent nine cycles, measuring 35.7 kJ m⁻³ at a stress value of 0.38 MPa. Subsequent cycles displayed decreasing dissipated energy, with values of 10.1 kJ m⁻³ and 8.4 kJ m⁻³ for the second and third cycles, respectively, followed by stress values of 0.34 MPa and 0.32 MPa. The diminishing dissipated energy is attributed to the alignment of polymer coils and physical entanglements with the applied stress during the first cycle. No significant changes in dissipated energy and stress were observed after the third cycle. To further clarify the toughening mechanism, cyclic loading–unloading tests were performed at different strains (100%, 300%, 500%, 700%, and 900%), results are depicted in Fig. 3(f) and (g). Notably, at 100% strain, dissipated energy was only 0.56 kJ m⁻³, with no residual strain, indicating high elasticity. In the range from 0 to 100% strain, the fabricated hydrogel behaves similar to the human skin, which reaches up to 75%. Dissipated energy increased from 2.8 kJ m⁻³ at 300% strain to 25.7 kJ m⁻³ at 900% strain. Furthermore, the dissipated energy demonstrated a direct relationship with applied strain. These results confirm the presence of reversible non-covalent interactions within the hydrogel network that temporarily crack under stress, regaining their original shape upon stress release. These reversible interactions play a crucial role in the hydrogel network's toughening mechanism.

3.3. Rheological study

The oscillatory rheological experiments for Z0 and Z2 hydrogels were carried out at a constant temperature of 25 °C to provide a strong proof about the effect of Zn-MOFs on the mechanical and rheological characteristics of the hydrogels. The first step involved performing a strain sweep between the strain ranges of 0.01 and 1000% at a fixed frequency of 10 rad s⁻¹. The strain-dependent viscoelastic responses of hydrogels Z0 and Z2 are shown in Fig. 4(a). The response of both hydrogels was linear within the 0.01 to 21% strain, and within this range, the storage modulus (G′) has higher values then loss modulus (G′′). This 21% strain is critical strain and the deformation in this range is reversible, next the higher value of G′ also suggesting the elastic behavior of the hydrogels. Fig. 4(b) shows the hydrogels’ frequency-dependency in the frequency range from 0.1 to 100 rad s⁻¹. In this instance, the hydrogels behaved like elastic solids since the G′ of Z0 and Z2 both exceeded the G′′ for both samples. Notably, the G′ of Z2 was higher than that of Z0, demonstrating that Zn-MOFs contribute to more crosslinking within the hydrogel network, and these crosslinking points caused the improvement of G′ in Z2. Furthermore, the G′ and G′′ of Z0 and Z2 both increased at a frequency around 63 rad s⁻¹, indicating an improvement in the hydrogel's elastic properties at lower frequencies. But beyond 60 rad s⁻³ up to 100 rad s⁻³, a modest decline in modulus was noticed, indicating a reversal in the elastic nature within this frequency range because of significant shear inside the hydrogel network. The damping factor (tan δ) is illustrated in Fig. 4(c). It shows that the tan δ of Z0 was higher than that of Z2 across all frequency ranges, indicating that Z2 behaves more elastically than Z0. It shows that Zn-MOFs exhibit viscoelastic behavior that is frequency-dependent, with tanδ decreasing at low frequencies. The elastic feature was however, found to be slightly less pronounced at high frequencies due to an increase in the tan δ. The improvement in rheological properties show that Zn-MOFs were successfully integrated into the hydrogel network, adding to the number of crosslinking sites between the polymer chains, Zn-MOFs, and surfactant.

Fig. 4 Rheological properties of the hydrogels. (a) Strain sweep test, (b) frequency sweep test, and (c) damping factors. These tests suggest that the rheological properties were successfully increased by Zn-MOFs.

3.4. Conductivity and strain sensitivity of hydrogels

The conductivity of the prepared hydrogels was estimated through an LCR meter, and the estimated conductivity values are given in Fig. S6 (ESI†). The virgin hydrogel (Z0) exhibited a conductivity value of 0.68 S m⁻¹, which was attributed to the presence of NaCl salt that ionized into Na⁺ and Cl⁻ in the solution. It is evident that there were interactions between Cl⁻ and positive EHDDAB, but the fundamental conductivity in Z0 was primarily due to Na⁺ and Cl⁻. Fig. S6 (ESI†) clearly shows that conductivity has a direct relationship with the Zn-MOFs. With 10 mg of Zn-MOFs in the hydrogel network (Z1), the conductivity increased to 0.79 S m⁻¹. The conductivity gradually increased with increasing amounts of Zn-MOFs, reaching a maximum value of 0.96 S m⁻¹ for the Z3 hydrogel. This increase indicates that Zn-MOFs positively impact conductivity, which can be attributed to the metallic and conductive nature of zinc. Additionally, the porous nature of the hydrogels contributes to conductivity by providing a route for the conduction of conductive species. The excellent conductivity and high mechanical performance make the hydrogel networks ideal candidates for the development of flexible transducers.

To visually observe strain sensitivity of the transducer, a rectangular-shaped Z2 hydrogel was connected in series with an LED system, as shown in Fig. S7 (ESI†). As the hydrogel stretched, the intensity of the LED dimmed, gradually decreasing with increasing strain, indicating the hydrogel's sensitivity to applied strain. This behavior can be explained by the fact that stretching restricts the conduction of ions and electrons as it narrows the conduction pathways. Once the applied strain is removed, the LED intensity increases again, demonstrating the reversibility of the process.

Furthermore, the strain performance of flexible transducer was evaluated using an electrochemical workstation (Autolab). Fig. 5(a) shows that the hydrogels have the ability to sense low strains in the range from 0.1% to 50%, indicating the ultra-sensitivity of the transducer. Additionally, as shown in Fig. 5(b), these hydrogels can also sense large strains, starting from 100% and reaching up to 900%. The sensitivity of the hydrogel system was further studied through the gauge factor (GF), a quantitative parameter for flexible transducers and sensors. The GF values for different strain ranges are given in Fig. 5(c)–(e), determined according to the previously reported literature.⁵⁷ According to Fig. 5(c), the GF can be calculated using the relative resistance equation: ΔR/R₀ = 1.82ε + 0.0022ε², where ΔR/R₀ is the relative resistance and ε is the applied strain. The GF values were 3.65 for strains ranging from 0 to 200%, 6.45 for strains ranging from 201 to 550%, and 6.50 for strains ranging from 501 to 900%. Fig. 5(e) shows a direct relationship between the GF values and strain, with a total GF value of 16.56 for strains ranging from 0 to 900%. Similarly, Fig. 5(d) reveals a linear relationship between relative resistance and strain in the range from 0 to 150%, with an R² value of 0.99. These results suggest that the prepared hydrogels can be used as an artificial human skin, considering that human epidermis has a stretchability of 75%.⁵⁸ The response and recovery times were also recorded through stretching and release, and as shown in Fig. 5(f), the hydrogel system exhibited a fast response time of 80 ms and a recovery time of 100 ms. Furthermore, the fatigue performance was evaluated using multiple stretching and release cycles, with more than 700 cycles performed, as shown in Fig. 5(g). The results confirmed that the hydrogels have excellent fatigue performance, as there was no significant drop in resistance even after 700 cycles, indicating the stability and reliability of flexible transducer. These findings demonstrate that the designed hydrogels possess high conductivity, fast response–recovery times, excellent fatigue performance, and a wide range of strain sensitivity from 0.1% to 900%. These properties position the hydrogel system as a potential and efficient candidate for various flexible and wearable electronic devices.

Fig. 5 Demonstration of the strain sensitivity of the prepared sensor. (a) Sensor response to small strains from 0.1% to 50%, (b) sensor response to larger strains from 100% to 900%, (c)–(e) calculated gauge factors from 0 to 900% strains with a GF value of 16.56, revealing a large strain-sensitivity of the sensor, (f) response and recovery time of the sensor, and (g) stability of the sensor to multiple cycles.

3.5. Hydrogels as epidermis sensors

With their excellent sensing properties, including ultra-sensitivity, high fatigue performance, fast response–recovery, stability, and a wide range of strain sensitivity (0.1% to 900%), Zn-MOFs-based conductive hydrogels for flexible transducers have tremendous potential for application in wearable electronic devices for monitoring human activities. To demonstrate real-time and practical applications, the flexible transducer was applied to monitor human activities such as wrist, neck, finger, and elbow motions. For this purpose, the hydrogel strip was capped with a copper sheet and attached to various body parts using adhesive tape, connected to an electrochemical workstation via a copper wire. First, the flexible transducer was attached to the wrist, and as shown in Fig. 6(a), it successfully sensed wrist bending, straightening, and upward bending. Stable signals were obtained in all positions when the wrist was at rest. The original graph depicts the relationship between current and time. Throughout the recorded data, no negative current values were observed. Specific wrist movements, such as when the wrist is lowered, resulted in disruptions in conductivity. As the wrist slopes, the hydrogel stretches, increasing resistance and decreasing current flow. Conversely, when the wrist returns to its initial position, there is no noticeable change in current. However, upon upward wrist movement, the hydrogel relaxes further, reducing the resistance of conduction pathways and facilitating ion conduction. Consequently, an increase in current flow is observed, resulting in the detection of negative resistance values after conversion through Ohm's law. Next, the hydrogel was attached to the neck, and as demonstrated in Fig. 6(b), the transducer efficiently responded to flexion movements of the neck. Similarly, for simple bending and unbending motions of the finger and elbow, the hydrogel was attached to these body parts, and the transducer effectively responded to the motions, as shown in Fig. 6(c)–(e). Furthermore, the transducer was also capable of detecting complex bending motions of the finger and elbow at various angles, as illustrated in Fig. 6(d)–(f) and Videos V1, V2 (ESI†). As the finger and elbow were bent, a stepwise increase in resistance was detected with increasing angles. At fixed positions of the finger and elbow, the resistance values remained constant, indicating the stability of the sensor. These detections of large motions confirm the high stability, sensitivity, and reliability of the sensor. Theoretically, this sensor can also be applied to monitor larger human joint motions in real-time, such as the knee and shoulder etc.

Fig. 6 The epidermis sensor toward large human motion. (a) Writ bending, (b) neck up-down motions, (c) finger motion at a single angle, (d) finger motion at various angles, (e) elbow bending and un-bending motions at a single angle, and (f) bending and unbending of the elbow at various angles.

In addition to detecting large human motions, this flexible transducer is investigated for physiological movements, in which the transducer was attached to the larynx of a volunteer and different movements were examined. Fig. 7(a)–(c) demonstrates that the transducer precisely and repeatedly responds to different physiological activities, such as speaking the word “One” (Fig. 7(a)), coughing (Fig. 7(b)), and whistling (Fig. 7(c)), clearly differentiating between these subtle human activities. Similarly, the sensor also differentiates and repeatedly responds to other physiological activities, including swallowing saliva (Fig. 7(d)), drinking water (Fig. 7(e)), and chewing bubble gum (Fig. 7(e)). All these subtle activities are repeatable with similar patterns for each activity, which confirmed the robustness, consistency, and versatility of the hydrogel transducer for both large and subtle human motions, making it suitable for various wearable electronic devices.

Fig. 7 Sensor capabilities for physiological (small) activities. (a) Speaking word “One”, (b) coughing, (c) whistling, (d) swallowing of saliva, (e) drinking of water, and (f) chewing of bubble gum. For each subtle activity the sensor produces different responses, which reveals the reliability of the sensor.

Surprisingly, the Zn-MOFs-based hydrogels have the ability to sense and precisely respond to different handwriting, as demonstrated in Fig. 8. To investigate this property, the Z2 hydrogel was sandwiched between two copper electrodes and connected to a measuring instrument. It can be observed in Fig. 8(a)–(c) that the hydrogel transducer not only responds to different English letters, such as “a, b, and c,” but also generates the same response when the letters are repeated, as shown in Fig. 8(b). Additionally, the transducer can differentiate between lowercase and uppercase letters, as depicted in Fig. 8(a) and (b), with examples like “a A, b B, and c C.” It is important to note that the sensor response to the letter “c C” was the same in Fig. 8(a) and (b) because there was no significant difference in this case, confirming the stability, reliability, and repeatability of the sensor. Similarly, the transducer generates different responses when the word “Pakistan” was written. Moreover, the sensor was tested for its capability to detect other languages. The author wrote his name “Mansoor” in his native language (Urdu), and as seen in Fig. 8(d), the sensor successfully and repeatedly sensed the Urdu language. Finally, the sensor was used to detect complex signatures, as shown in Fig. 8(e). The sensor exhibited a similar response pattern for the complex signature, indicating that Zn-MOFs-based hydrogel transducer has an excellent ability to sense and differentiate between different letters, languages, and signatures. These features of the sensor to the writing make these materials strong candidates for mobiles, touch panels, and other electronic devices.

Fig. 8 Sensor response towards writing. (a), (b) Response and differentiation of small (a, b, and c) and capital letters (A, B, and C), (c) writing response of the sensor towards word “Pakistan”, and (d), (e) response and repetition of the sensor using Urdu (Mansoor) language and signature.

3.6. Hydrogels as pressure and biomimetic artificial skin sensors

Corresponding to strain and human activity detection, the prepared Zn-MOFs-based hydrogels have potential applications in pressure and artificial skin sensors due to their excellent performance and conductivity. As demonstrated in Fig. 9(a) and Video V3 (ESI†), the Z2 hydrogel responds to applied pressure by causing the LED to become brighter. The increase in LED's brightness is attributed to a decrease in the conduction path for the conductive species under pressure, allowing for easier movement and a resulting decrease in resistance within the hydrogel. Similarly, Fig. 9(b) and (c) reveals that the hydrogel successfully and repeatedly responds to low (0.2 to 2 kPa) and high pressures (10 to 60 kPa). From these results it can be concluded that the resistance is inversely related to the magnitude of pressure. Additionally, like the large strain sensitivity range, the hydrogel pressure sensor also has a wide range of pressure sensing capability (0.2 to 60 kPa). Furthermore, a wearable wristband was developed (Fig. 9(d) and (e)), and it can be observed that the wristband precisely monitors subtle touches of the human finger through applied pressure, with the peak intensity correlating to the pressure applied on the band's surface. With their large pressure ranges and high strain sensitivities, these hydrogels can be practically applied in various wearable and smart electronic devices. Next, the artificial skin properties were investigated as shown in photographs in Fig. 9(f)–(k), where the Z2 hydrogel is prepared in a specific mold. The hydrogel is applied to the finger (Fig. 9(f)), and various functions were performed, as shown in Fig. 9(g) and (h) and Video V4 (ESI†). It was observed that the hydrogel smoothly and successfully interacted with mobile screens without any delay. Similarly, the hydrogel is used for the fabrication of a metallic pen, as demonstrated in Fig. 9(i)–(k). Notably, the Z2 hydrogel functioned smoothly on the screen without causing damage or delays, and it could be used to write and draw various letters and diagrams, respectively. These results suggest that these hydrogels can mimic human skin activities and find applications in various electronic sensors as biomimetic artificial skins.

Fig. 9 Pressure and human skin-like response of the prepared hydrogels. (a) Photographs of the pressure sensor, (b) repeated response of the sensor at different pressures, (c) pressure sensing at different pressures, (d), (e) wearable wrist band and its pressure responses, (f)–(h) hydrogel like human skin, which performed different functions on a smartphone, and (i)–(k) hydrogel as a metallic pen for performing different human activities on a screen.

4. Conclusions

In summary, a novel method has been developed to create hydrogel-based flexible transducers to strengthen Zn-MOFs by incorporating them into a hydrophobically cross-linked hydrogel network. In this system, the Zn-MOFs act as robust fillers, efficiently linking the surfactant micelles and also establishing connections between the micelles and polymer chains. This results in hydrogels with high mechanical properties, including remarkable stretchability (1448%), high toughness (475 kJ m⁻³), and a Young's modulus of 1 kPa. These mechanical properties, attributed to non-covalent interactions, grant the hydrogels swift self-recovery and anti-fatigue performance at ambient temperatures, without the need for external stimuli. Additionally, the hydrogels exhibit remarkable strain sensitivity, showing a gauge factor of 16.56 for both small and large strains, along with ultrafast response and recovery times of 80 and 100 milliseconds, respectively. Consequently, due to their exceptional mechanical performance, excellent sensitivity, and rapid response–recovery capabilities, these hydrogels are well suited for fabricating strain transducers dedicated to monitoring human motion. This hydrogel transducer accurately tracks and differentiates between extensive human movements, as well as subtle physiological signals, including joint motions of the neck, wrist, finger, and elbow, speech, swallowing, pressure and writing. Beside this, the flexible transducer is also capable of mimicking the properties of epidermis and working as electronic skin. This innovative approach of MOF-based hydrogels holds promising potential in paving the way for the creation of flexible hydrogel sensors with customizable mechanical properties and enduring reliability.

Conflicts of interest

The authors declare no conflicts of interest.

Acknowledgements

The authors are thankful to the Cooperative Equipment Center at KOREATECH for assistance with FT-IR, SEM, EDS XRD and UTM analyses.

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Footnote

† Electronic supplementary information (ESI) available: Text 1: synthesis and characterization of Zn-MOFs. Text 2: explanation of Zn-MOFs FTIR. Text 3: discussion of XRD of Zn-MOFs. Fig. S1: FTIR spectra of the prepared Zn-MOFs. Fig. S2: XRD analysis of the prepared Zn-MOFs. Fig. S3: (a), (b) Illustration of the EDS analysis and (c) SEM study of the Zn-MOFs. Video V1: response of the sensor to finger motion at various angles. Video V2: response of the sensor to elbow motion at various angles. Video V3: pressure response of the sensor. Video V4: hydrogel response as biomimetic artificial skin. See DOI: https://doi.org/10.1039/d4tb00718b

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