1 Introduction

Versatile devices and materials have been developed for the precise diagnosis and treatment of neurological diseases, such as neurodegenerative disorders (e.g., Parkinson’s disease and Alzheimer’s disease), neuromuscular diseases, and spinal cord and peripheral nerve injuries [1,2,3,83]. These stabilized conductive polymers are easily mixed with the hydrogel precursor solution and used to fabricate conductive hydrogels (Fig. 1f) [77,78,79,80,81,82,83,84,85]. Conductive polymer additives form interconnected conductive pathways in the hydrogel matrix, enabling the transfer of electrons through it. Similar to metal- and carbon-based additives, the fraction of conductive polymers plays a critical role in the electrical and mechanical properties of conductive hydrogels, and higher fractions increase the conductivity, modulus, and viscosity.

2.2 Conductive network formation

Although the addition of conductive materials provides a certain degree of electrical properties to the hydrogels, they still exhibit a much lower conductivity (< 1 S/m) than the original conductive materials (> 100 S/m). This is because numerous small isolated conductive domains covered by an insulating domain are formed in the hydrogels, inhibiting charge transfer over the matrix. Herein, we discuss a fabrication strategy to form a well-connected percolating conductive network.

2.2.1 Pure conductive polymer hydrogel

Fabricating conductive hydrogels with only conductive polymers minimizes the fraction of insulating domains and enhances the continuity of the conductive domains in the hydrogel matrix. However, conductive polymers are typically difficult to use as hydrogel backbone polymers because of their hydrophobic nature. They form agglomerations instead of uniform networks; therefore, hydrogels with high water content and low modulus are difficult to be achieved.

Recently, various pure conductive polymer hydrogel fabrication methods have been reported using PEDOT:PSS with additives such as acids [86], ionic liquids [87], and secondary dopants (e.g., dimethyl sulfoxide (DMSO) [88,89,90], ethylene glycol [90], and 4-dodecylbenzenesulfonic acid (DBSA) [91, 92]). Frequently, the additive support conformation changes from PEDOT:PSS colloidal particles to physically interconnected networks during the dry-annealing process, and more interconnections between the PEDOT polymers are generated than in the process without additives (Fig. 2a). After washing and re-swelling this PEDOT:PSS film, PEDOT:PSS hydrogels with high stability and conductivity are formed [86,87,88,89,90]. In contrast, PEDOT:PSS films without additives exhibit fragmentation [88] and conductive path failure because of gap generation after swelling [87]. Additionally, an increase in the PEDOT to PSS ratio by the removal of PSS, which means a decrease in the insulating domain in the hydrogel matrix when using acid and ionic liquids as additives, has been reported [86, 87]. The concentration of the additive significantly affects the hydrogel properties, and an optimization process is required. After the optimization process, the PEDOT:PSS hydrogel frequently exhibits ultrahigh conductivity (conductivity > 800 S/m), high water content (> 80 wt%), and low modulus (< 5 MPa for DMSO and < 50 kPa for others) [86,87,88,89,90]. Moreover, in situ conductive hydrogel fabrication can be conducted using DBSA as an additive, which is useful for injection and printing (Fig. 2b) [91, 92].

Fig. 2
figure 2

Illustration of conductive network formation strategies for high conductivities. a, b) Pure conductive polymer hydrogel fabrication using an additive. a DMSO addition supports formation of physically interconnected PEDOT:PSS network during dry-annealing. Reproduced with permission from [88], copyright Springer Nature, 2020. b DBSA micelle addition causes rearrangement of PEDOT:PSS colloidal particles and formation of a physically interconnected conductive hydrogel network within few minutes. Reproduced with permission from [91], copyright John Wiley and Sons, 2019. c Percolating conductive network hydrogel fabrication from a pure conductive polymer hydrogel. The monomer, crosslinker, and initiator can be infiltrated into the pure conductive polymer hydrogel and cause secondary network formation while maintaining primary percolating conductive network. Reproduced with permission from [93], copyright Springer Nature, 2018. d Fabrication of a percolating conductive network hydrogel from non-conductive hydrogel. PEDOT:PSS can be dispersed in the precursor solution. After crosslinked PVA network formation using glutaraldehyde (GA), acid treatment causes the rearrangement of PEDOT:PSS colloidal particles and formation of a secondary percolating conductive network. Reproduced with permission from [99], copyright John Wiley and Sons, 2022. e Spatial separation of conductive nanocomposite using microgel. Ag nanoparticles are only produced on the surface of the microgel through in situ reduction and a densely packed Ag nanoparticle percolating network is produced after microgel assembly. Reproduced with permission from [100], copyright John Wiley and Sons, 2019. f Spatial separation of conductive nanocomposite using freezing. CNT and GO are located in the surrounding region of ice crystal during freezing. This improves the percolating conductive network and can be maintained by crosslinking at ambient conditions. Reproduced with permission from [102], copyright John Wiley and Sons, 2022

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2.2.2 Percolating conductive nanocomposite network formation

Pure conductive polymer hydrogels have many advantages owing to their excellent electrical properties, but they exhibit low stretchability and difficulty in handling without fracturing. Treatment with Triton X-100 to enhance the stretchability of pure conductive polymer hydrogels has been reported, but the stretchability was still insufficient (< 60%) [90]. In contrast, conductive materials with non-conductive polymers, such as poly(vinyl alcohol) (PVA) and polyacrylic acid (PAAc), typically exhibit high stretchability (> 100%) and stable mechanical properties but lack electrical properties.

Recently, various methods have been reported to balance the electrical and mechanical properties of conductive hydrogels. The main objective of these methods is to improve the interconnections between conductive nanocomposites dispersed in a non-conductive polymer hydrogel matrix and to generate percolating conductive nanocomposite networks. This percolating conductive network minimizes the conductive path failure by insulating the domain and increasing the conductivity of the matrix. The first involves the fabrication of a conductive hydrogel from a pure conductive polymer hydrogel (Fig. 2c). Monomers such as acrylic acid can penetrate a pure conductive polymer hydrogel, and an interpenetrated non-conductive network is generated after polymerization. This process affects the conductive polymer network but maintains a percolating conductive network with high conductivity (> 20 S/m) and improved mechanical properties (stretched over 100%) [93, 94]. In other methods, percolating conductive networks are generated from conductive nanocomposites. When using conductive polymers, percolating conductive networks can be generated and improved by the addition of ethylene glycol [95, 96], crosslinkers such as phytic acid, which act as anchoring points for conductive polymers [18, 97, 98], and acid treatment (Fig. 2d) [99]. Such chemical treatment for formation of percolation path in the hydrogel results in volumetric shrinkage of the hydrogels capable of exhibiting a more densely packed structure [99]. That is, this shrinkage can increase the ratio of the conductive domain to the non-conductive domain (e.g., non-conductive polymers, water), which helps achieving high conductivity. In addition, micropatterning strategies have been reported to generate percolating conductive networks using metal nanoparticles, CNT, and GO. First, microgels (hydrogel microparticles) are used to embed conductive nanocomposites on a surface (Fig. 2e) [100, 101]. After the microgels are assembled, a spatially defined percolating conductive network is formed through the microgel surfaces. Therefore, conductive hydrogels fabricated using this method exhibit higher conductivity than normal bulk conductive hydrogels [100]. The second strategy uses spatial rearrangement of conductive nanocomposites through freezing (Fig. 2f) [102]. During the freezing of the aqueous solution, ice crystals are formed, and conductive nanocomposites are located in the region surrounding the ice crystals. A percolating conductive network is formed by this special arrangement and is maintained by crosslinking after removal of the ice crystals [102]. Additionally, the amounts of metal nanoparticles, CNT, and GO are frequently limited during conductive hydrogel fabrication owing to their intrinsic mechanical properties. In contrast, liquid metals such as EGaIn are relatively free from this problem owing to their intrinsic softness and stretchability. Recently, a strategy to assemble liquid metal particles in a hydrogel polymer matrix with ultrahigh conductivity (> 1,000,000 S/m), high stretchability (> 700%), and low modulus (< 200 kPa), even after the addition of 74.4 v/v% liquid metal, was reported [37]. After embedding the liquid-metal particles into the polymer matrix, the acoustic field generates liquid-metal nanoparticles at the particle surfaces. These nanoparticles interconnect with nearby particles to form percolating conductive networks.

3 Printing techniques for conductive hydrogel ink

Conventional methods for fabricating conductive hydrogel structures by molding, solvent treatment, and annealing are limited by low resolution, poor interfacial bonding, complex post-processing steps, and environmental hazards. Hence, 3D printing offers a potential solution to overcome these challenges by enabling precise control over the shape, size, and functionality of conductive hydrogel structures. Particularly, considering that one of the most important applications using 3D printing is tissue engineering, the biological tissues have intrinsic function based on coherent cell-to-cell communications within 3D objects geometries. In comparison with 2D structures, such 3D geometries distinctively affect cellular behavior, such as migration, differentiation, and proliferation [103]. Additionally, recent approaches for fabricating 3D bioelectronics to stably stimulate the tissues and record their electrophysiological signals can be deliberately achieved through versatile 3D printing techniques [104]. These 3D printings in bioelectronics field offer freedom of design in 3D space [89] and facile and scalable formation of dense interlayer with electrical connectivity that is crucial for neural electrode performance [105].

Recent advancements in 3D hydrogel-printing technology have enabled the precise fabrication of conductive hydrogels with complex geometries, creating new avenues for a wide range of biomedical applications. To ensure successful printing of the construct, an optimal hydrogel ink that is tailored to a specific printing technique should be prepared. The properties of the hydrogel ink are governed by the type and degree of polymer interactions within the network; therefore, design factors, such as resolution and crosslinking methods, are of utmost importance. For the ink to solidify into the intended architecture, the mechanical and chemical properties of both the precursors and post-print applications should be considered. Therefore, the selection of an appropriate crosslinking method is critical. Current state-of-the-art printing mechanisms can be divided into two categories: viscoelasticity-dependent and static-state printing (Fig. 3). In this section, the various crosslinking mechanisms and printing techniques required to obtain the desired mechanical and chemical properties of the printed constructs are discussed.

Fig. 3
figure 3

Schematic of 3D printing techniques to fabricate conductive hydrogels. a Viscoelastic-dependent printing of conductive hydrogel ink containing conductive particles or conductive polymers (inkjet or extrusion). b Static-state printing via photopolymerization of conductive hydrogel ink (DLP or SLA)

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3.1 Viscoelasticity-dependent printing

Hydrogels are soft, tissue-like materials with a high-water content, which contributes to their unique viscoelastic characteristics and shear-thinning behavior. Thus, the non-Newtonian trait of a hydrogel, in which the viscosity decreases with increasing shear or stress, enables the movement of the hydrogel through smaller confined spaces such as needles, thereby enabling the hydrogel to be extrudable for use in injection ink. Viscoelasticity-dependent printing techniques involve inkjet printing and extrusion-based printing, which rely on the viscosity of the ink to control the printing process.

3.1.1 Inkjet printing

Inkjet printing is a popular method in which small ink droplets are used to create high-resolution structures on various substrates. The inkjet printing of hydrogels depends on the small droplets formed by the pressure pulse at the nozzle. The size and velocity of the droplets are controlled by adjusting the pressure pulse and nozzle size. Inkjet printing has several advantages over other printing techniques, including the ease of hydrogel preparation and suitable viscosity. This reduces the complexity of the printing process and makes it more cost-effective for bioelectric applications, including tissue engineering [106] and biosensing [107,108,109]. Therefore, these viscoelasticity-dependent inkjet printing can be particularly viable when cells are encapsulated within the ink to be used as a bioink [

6 Outlook

In conductive hydrogel printing for versatile neural engineering, it is generally understood that the quality of 3D printed structures largely depends on the chemical composition of the hydrogel. Therefore, while many studies have focused on materials associated with 3D printable structures, studies related to selection of optimal printing strategies have not been considered as significant. A few novel engineering methods for printing the structures of conductive hydrogels include combined strategies or augmentation of conventional methods. Orthogonal photochemistry-assisted printing (OPAP), developed by Wei et al., combines 3D extrusion printing with visible-light photogelation to print TCHs that can be used for hydrogel arrays [189]. Ahn et al. utilized air-pressure-assisted pen-nib printing for novel approach to printing electronics [231]. Silva et al. introduced a unique electro-assisted hydrogel deposition of PEDOT and alginate droplets based on electrochemical reactions (Fig. 8a) [232]. Peng et al. demonstrated the co-axial 3D printing of conductive and degradable GO-PPy-alginate chips for skin neuronal differentiation. The shell ink (GPA-SDF-1) was fabricated by anchoring SDF-1 chemokine to GO-PPy in alginate ink while the core ink (CGP/bFGF-pDNAs) was made by conjugating GO with PEI (polyetherimide), transfected with bFGF-pDNA (plasmid DNA), and crosslinked with MMP-2 (matrix metalloproteinase-2) sensitive peptide sequences. The core–shell structure from co-axial printing was stimulated with bioelectrical signal and showed differentiation and maturation of MSCs (mesenchymal stem cells) to functional neuronal cells (Fig. 8b) [217]. Beyond unconventional printing techniques described above, the conductive materials to be loaded in the hydrogel ink can be newly designed for further neural engineering. Using laser irradiation, graphitization of various organic materials (e.g., polydopamine and lignin) and thermal annealing of PEDOT:PSS can be induced [35, 233, 234]. Furthermore, photothermal energy caused by laser irradiation can generate selective conductive path within the hydrogel inks, realizing delicate 3D printing of bioelectronic devices [35, 234]. Regarding these, Miyakoshi et al. developed a hydrogel capacitor through laser-induced lignin graphitization in agarose hydrogel, and Won et al. reported PEDOT:PSS patterning based on laser induced photothermal energy from gold nanoparticles [35, 234]. To sum up, the 3D printing of conductive hydrogels has recently led to remarkable advances in the fabrication of complex and functional structures. Novel 3D printing strategies can solve the limitations of conventional methods such as printing resolution, biocompatibility, and functionality. These printing strategies have the potential to revolutionize and ultimately contribute to the development of a toolbox for designing conductive hydrogels for bioelectronics as well as advanced tissue engineering applications. In the near future, such 3D printed conductive hydrogels would be promising as a tool to electrically stimulate on neural tissues, record their electrophysiological signals with high signal-to-noise ratio and provide neural tissue-mimetic scaffold with improved therapeutic efficacy. Toward clinical translation, the 3D-printed conductive objects, or the conductive hydrogel inks by themselves might allow future close-looped neuroprosthesis capable of stimulation, recording, and restoration of the electrophysiological signals from human body stimulate and can be utilized as therapeutic platform for spinal cord and peripheral nerve.

Fig. 8
figure 8

Unconventional printing techniques for conductive hydrogels. a Electro-assisted printing via controlled electrochemical reactions: (i) printing mechanism and (ii) printing set-up with P/G-stat (potentiostat/galvanostat). Reproduced with permission from [232], copyright Springer Nature, 2022. b 3D co-axial printed chip for skin nerve regeneration in wound: (i) ink materials and preparation, (ii) co-axial printing of core–shell microfiber. Reproduced with permission from [217], copyright John Wiley and Sons, 2020

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