3D printing of electrically conductive hydrogels for tissue engineering and biosensors – A review

Journal Pre-proofs Review article 3D Printing of Electrically Conductive Hydrogels for Tissue Engineering and Biosensors – A Review Thomas Distler, Aldo R. Boccaccini PII: DOI: Reference:

S1742-7061(19)30609-9 https://doi.org/10.1016/j.actbio.2019.08.044 ACTBIO 6336

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Acta Biomaterialia

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14 May 2019 18 August 2019 28 August 2019

Please cite this article as: Distler, T., Boccaccini, A.R., 3D Printing of Electrically Conductive Hydrogels for Tissue Engineering and Biosensors – A Review, Acta Biomaterialia (2019), doi: https://doi.org/10.1016/j.actbio. 2019.08.044

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3D Printing of Electrically Conductive Hydrogels for Tissue Engineering and Biosensors – A Review Thomas Distler, Aldo R. Boccaccini* Institute of Biomaterials, Department of Material Science and Engineering Friedrich-Alexander-University Erlangen-Nuremberg Cauerstr. 6, 91058 Erlangen, Germany *Corresponding Author E-mail: [email protected] Abstract: Electrically conductive biomaterials are gaining increasing interest owing to their potential to be used in smart, biosensoric and functional tissue-engineered scaffolds and implants. In combination with 3D printing technology, this class of materials might be one of the most advanced approaches towards future medical implants regarding potential functionalities and design possibilities. Conductive hydrogels themselves have been researched for potential sensoric and tissue engineering applications for more than a decade, while the 3D printing of such functional materials is still under early exploration. This review aims to provide a short insight into the most recent developments of 3D printable and electrically conductive hydrogels. It also provides a summary of the last few years of research in this field, with key scope on 3D printing for biomedical applications. The final literature search was conducted in May 2019, with the specific keywords ‘3D’, ‘printing’, ‘conductive’, ‘hydrogel’, ‘biocompatible’ and combinations of the latter, using advanced search in the databases Scopus®, Web of Science® (Web of Knowledge®) and Google Scholar®. A total of 491 results were gained, while 19 recent publications were identified with the above-mentioned criteria and keywords, which are the studies finally discussed in the paper. The key results have been summarised, and the remaining challenges in the field and the scope for future research activities have been discussed. Keywords: 3D printing, electrically conductive polymers, conductive hydrogels, smart biomaterials, bioplotting, biosensors, advanced tissue engineering strategies

1.

Introduction

Biomaterials research is continuously expanding towards more functional and advanced materials to tackle the current needs of healthcare [1]. A material functionality that provides promising control over functions, for example, drug release [2–6], stimulation of cells [7–9], tissue engineering [9–12], stimulation to restore body functions [13,14], capability to produce bioactuators [15–17] or in vivo biosensors to monitor tissue regeneration [18], is the electrical conductivity of biocompatible polymers [19,20]. These materials promise to provide a new

level of control over biomaterials’ behaviour when applied inside the human body. In particular, for tissue engineering and active prosthetic devices with sensoric properties, conductive polymers (CP) in form of hydrogels are under extensive investigation [21,22]. The similarity of hydrogels to the native tissue extracellular matrix (ECM) of the body, composed primarily of water and biomolecules, renders them excellent candidates for the attempt to mimic the natural cellular environment’s microstructure and mechanics. Electrically conductive hydrogels are highly relevant, especially in tissue engineering (TE) disciplines such as nervous or cardiac TE, where electrical conductivity is key for successful function. Although comprehensive reviews on CP hydrogels and the functionalisation possibilities to achieve conductive hydrogels exist [21,23], works with particular focus on the 3D printing of CP hydrogels have only recently begun to emerge from the scientific community [24–27]. This focused review provides a short overview about the current progress and advances in the exciting field of 3D printing of these novel, functional, electrically conductive biomaterials. Research in the field of 3D printing has been exponentially growing during the past decade (Fig. 1). Considering that with the aid of CP hydrogels, functional solutions such as sensoric patches [28–31], tissue-engineered grafts [24,26,32] or soft actuators [31,33] can be achieved, publications in this field have been steadily increasing in the past years (Fig.1, red curve). However, the number of publications in the combination of both

Fig. 1. Number of publications from 2000 to 2019 derived from Scopus® advanced literature search (https://www.scopus.com/) with keywords “3D Printing”, “Conductive” AND “Hydrogel”, “Electrically Conductive” AND “Hydrogel” and “Printing” AND “Conductive” AND “Hydrogel”, indicating trends of increasing number of publications in the past five years. While the number of publications on “3D printing” significantly increased, exceeding 200 publications in 2012, increasing number of publications is also observed on the topic “3D printing of electrically conductive hydrogels”, especially in recent years.

‘3D printing’ and ‘electrically conductive hydrogels’ has increased only recently. Although the number of publications on 3D printing of conductive hydrogels has been lower than the total number of publications in the general field of CP hydrogels, interest in the subject is evident from the relative increase in the number of papers published in the last 5 years. 2.

Origins of Conductivity in Electrically Conductive Hydrogels

Comprehensive reviews on the underlying mechanisms of conductivity in conductive hydrogels and their functionality have been reported recently [21,22,34]. In brief, the overall electrical conductivity of polymer hydrogels can be distinguished between ionic and electronic conductivity [21,22,34]. Conductive hydrogels for medical applications, often composites of CPs (e.g., polypyrrole (PPy), polyaniline (PANI), polyethylene dioxythiophene (PEDOT)[21]) and polysaccharides or proteins (e.g., alginates, gelatin/gelMA, collagen, chitosan), can be viewed as electrolyte networks of high water content with regard to the basic definition of superabsorbent polymers or hydrogels [35,36]. The electronic functionality of CPs relies on the presence of the conjugated pi (p) system (p-p stacking), which consists of overlapping p-orbitals forming p-bonds throughout the polymer backbone by polymerisation. Through this process, p-systems with p-electrons exist, in which electron redistribution can be triggered. At the molecular level, an alternating pattern of single and double bonds is created, often realised by, for example, the presence of aromatic rings, which allow charge transfer [21]. By either stacking of ring structures or implementing ideally perfectly conjugated pi-systems in the polymer, charge transfer can be allowed along its backbone (intra-chain charge transfer and charge hopping) or in between CP chains inside a hydrogel (inter-chain charge transfer). Hence, the final functionality of the CP network inside conductive hydrogels inherently relies on the capability of allowing charge transfer either between polymer chains or along distinct polymer chains [21,37]. It is therefore desired to form CP entities inside bulk hydrogels in a superstructure and ordered fashion, allowing inter- and intra-chain charge transfer throughout the hydrogel. Inherent to the definition of hydrogels, which often swell in electrolytes when applied for cell culture, an ionic contribution towards the total conductivity is present in the wet state. The electronic function of CP itself is often enhanced by dopants, referring to charged entities (ions, e.g. Cl-) or charged side groups (e.g. SO3-, PO4-) of the CP or other polymers participating in the oxidation reaction. Their purpose is to deteriorate the electronic conformation (oxidation degree) of the formed CP network, provoking unbound charge carriers by charge extraction or charge transfer to the polymer backbone (n-type reduction, p-type oxidation) [21]. Ideally, the conductivity of the conjugated system of CPs can be increased by the efficacy of the dopants to extract or induce electrons in the CP backbone or by the structure of the formed CP network to allow inter-chain charge transfer by stacking of the CP [21,38]. Some of the most prominent examples for CPs used in biomedical applications are polypyrrole (PPy), polyaniline (PANI) and polyethylene dioxythiophene (PEDOT) [21]. Dopants, which are exemplary in the case of polypyrrole, have been introduced to alter its oxidative state from undoped to polaron and bi-polaron states of PPy, resulting in the creation of holes and p-type conduction, increasing the conductivity from 2.15x10-3 S.cm-1 to 50x10-3 S.cm-1 [39,40]. Additionally, conductive

functionality can be introduced to hydrogels by the addition of inherently conductive particles like graphene and its derivatives or carbon nanotubes (CNTs). This approach has attracted attention in the last few years [24,26,29,32,41,42]. The following section briefly elucidates the different methods put forward to create electrical conductivity functionality inside polymer hydrogels. 3.

Conductivity Functionalisation of Hydrogels

From the review of the recent literature (Fig. 1), four different approaches to achieve conductive hydrogels were identified. The basic principles of introducing conductivity into a hydrogel network have been reviewed earlier [2123] and can be divided into the following approaches, which are briefly discussed in this section: 1) Conductivity by in situ polymerisation 2) Conductivity by postpolymerisation/coating 3) Conductivity by addition of conductive particles (composite approach) 4) Conductivity by ionic conductivity 3.1 Conductivity by in situ polymerisation This process describes the formation of a CP phase inside a, for example, polysaccharide hydrogel matrix by addition of monomers of the aimed CP (e.g. pyrrole, aniline, EDOT) and subsequent addition of an oxidiser (commonly used oxidisers are FeCl3, ammonium persulfate (APS)) in one reaction step. As a result, a CP matrix is formed simultaneous to the formation of the hydrogel. By this approach, a homogeneous polymerisation of the CP network can be triggered inside the hydrogel matrix. However, owing to the lack of solubility of many frequently used CPs and their monomers in water (pyrrole, polypyrrole, aniline, polyaniline), an emulsion of CP precursor monomers in the aqueous phase has to be assumed, hence leading to a potential particle synthesis of hydrophobic CP inside the aqueous hydrogel bulk network if no additional functionalisation of both polymers is employed [21]. To address the need for a homogeneous conductive hydrogel, Pan et al. (2012) [43] performed in situ polymerisation of PANI with the addition of phytic acid as a noncovalent crosslinker and dopant (Fig. 2) to form a conductive hydrogel network. Supercapacitor electrodes with high cyclic performance (~83% capacitance retention after 10,000 cycles) and a material conductivity of 0.23 S.cm-1 were fabricated. The production of these electrodes by ink-jet printing allowed a high-throughput process suitable for bioelectronic applications or electrodes for energy storage [43].

Fig. 2. Schematic illustration of the network formation of PANI/Phytic acid conductive hydrogel utilising phytic acid as a crosslinking agent and dopant (A). Hydrogel inside a glass vial (B), SEM and TEM micrographs of the conductive hydrogel (C–E) (adapted from Pan et al. (2012)[43]). Reproduced with permission from [43] (PNAS).

3.2 Conductivity by post-polymerisation In this approach, pre-formed hydrogels are immersed in either monomer solutions of the aimed CP or oxidant solution, and in a subsequent step, the conductive hydrogel is formed by immersion of either monomer-soaked hydrogels inside an oxidant solution [44,45] or oxidant-soaked hydrogels in monomer solutions [46], respectively. This process can also be employed to achieve an electrically CP coating on hydrogels or other materials [9,23]. Post-polymerisation of CPs offers a high versatility by being a potential tool to coat materials of different classes with the CPs of choice. Furthermore, hydrogels can be pre-fabricated, which can be advantageous, while the conductive functionalisation can be introduced in a subsequent step. Yet, potential diffusion limitation and hence inhomogeneous formation of conductive phase throughout the gel may be an issue of this process, which inherently relies on a diffusion step of either monomer or oxidant solutions inside/on the bulk material. As a result, to achieve homogeneous CPH networks, a processing parameter optimisation has to be performed [44–47].

Wu et al. (2016) [47] fabricated methacrylated gelatin (GelMA) hydrogels by a 3D printing µSLA technique (Fig. 3) employing a post-polymerisation approach. Gels were incubated in an oxidant solution to incorporate ammonium persulfate (APS) as an oxidising agent for subsequent polymerisation of aniline to PANI. APS-soaked gels were incubated in aniline solution to form polyaniline-GelMA hydrogels. The electrical resistance of 1.62 mm-thin discs was determined by cyclic voltammetry (CV) measurements, resulting in a value of 2.9 kΩ @ 0.1 Hz, while a resistance of 165.5 kΩ was measured using a 2-point probe measurement approach. Biocompatibility was verified using 10T1/2 cells, with the scaffold potential application in bioelectronic interfaces [47].

Fig. 3. Fabrication of GelMA/polyaniline (PANI) conductive hydrogels by a postpolymerisation process through immersion of 3D µSLA-printed GelMA scaffolds in an oxidant solution, with subsequent immersion in an aniline/hexane solution to form the conductive polyaniline phase (Wu et al. (2016) [47]). Reproduced with permission from [47] (Elsevier).

3.3 Conductivity by addition of conductive particles (composite approach) In this approach, conductive particles, for example, metals, carbons (graphene, CNTs) or CPs, are added to hydrogel precursor solutions. Following this, hydrogels are formed by polymerisation, while the introduction of the electrical conductivity is performed by the conductive particles. At certain filling degrees, percolating particle networks

can

form,

which

then

promote

electronic

conductivity

through

the

hydrogels.

Using the approach of particle addition, conductive hydrogels of relatively high conductivities (>20 S.cm-1) compared to those of other approaches were produced [29]. This is possible by the addition of materials like multiwalled carbon nanotubes (MWCNT) or graphene with inherently high conductivity [48,49]. In contrast to in situ and postpolymerisation processes, no additional washing step with the purpose to remove potential unreacted monomers or oxidisers is required regarding tissue engineering applications, as this approach does not rely on any

synthesis involving CP monomers or oxidisers. Hence, such systems can directly be applied in cell culture and in vitro studies [24,26,29,32], while other approaches may require additional washing. Often, nanoparticles are employed to allow a high surface area to expose functional groups or to achieve percolation. With toxicity being of major importance for all biomedical applications, toxicological evaluations regarding nanoparticles may have to be considered specifically, depending on the biomedical field of interest for the particular systems [50,51]. Shin et al. (2016) [41] bioplotted a conductive hydrogel composite from GelMA and DNA-coated MWCNT (Fig. 4). With an MWCNT content of 6 mg.ml-1, conductivities of 24 ± 1.8 S.cm-1 were achieved, measured using a 4-point probe measurement device on dried films [29]. For determining potential use as capacitors, CV and additional electrical impedance spectroscopy (EIS) were performed. Cardiac fibroblasts were seeded on printed hydrogel patterns, resulting in a high degree of cardiomyocyte maturation observed when applying the conductive hydrogel ink. The authors elucidated potential applications for flexible biosensors as well as tissue engineering for cardiac muscle repair [29].

Fig. 4. 3D-printed sensors and functional tissue engineering scaffolds from GelMA and DNA-functionalised multi-walled carbon nanotube (DNA-MWCNT) ink. (left, A, B) 2D-printed conductive paths and bendable sensors for potential flexible electronics application. (right, C, D) 3D-printed network of DNA-MWCNT-laden ink for cardiomyocyte stimulation with increased maturation observed when stimulation was applied (adapted from Shin et al. (2016)[29]). Reproduced with permission from [29] (Wiley).

3.4. Conductivity by ionic conductivity Hydrogels are highly water-swollen superabsorbent polymer networks, with a water content usually exceeding 90% [35,36]. In biomedical applications, hydrogels are often formed in electrolyte solutions (e.g. phosphatebuffered saline, buffered salt solutions), especially for tissue engineering applications. By this process, ions are incorporated introducing an inherent capability of ionic conductance to the gel. In an in vivo scenario at pH 7.4,

hydrogels would be immersed in body fluid, intrinsically consisting of ions (Ca2+, Cl-, Na+, PO4-, etc.), causing an increase in conductivity by their presence [21]. Ionic electrically conductive and 3D-printable hydrogels have been produced utilising this approach [27,28,31,33]. By relying on ionic conductivity, no additionally electrical CP phase like polypyrrole, PANI, PEDOT, CNTs or graphene needs to be incorporated. Consequently, transparent hydrogels can be achieved [28,31], while the absence of nanoparticulate phases (e.g. CNT) can be advantageous regarding biomedical applications owing to possible nanoparticle-related toxicity considerations [50]. However, CP-dopant systems like PEDOT:PSS can contribute to ionic conductivity by the presence of negatively charged PSS, enabling cation transport by acting as ion pumps. The stability of the conductivity of pure ionic systems will be dependent on the capability of the hydrogels to retain the incorporated ions. The application of ionically conductive hydrogel systems is often found in stretchable electronics, usually based on acrylamide polymers (PAAM), owing to their high stretching ability [27,28,31,33]. However, these materials and the use of LiCl as an ionic conductor must be considered regarding their toxicity when applied in vivo [52–54]. Cell responsiveness to electrically conductive substrates may differ between cases of being in contact with ionically conductive materials in comparison to cell adhesion on electrically CPs, which bare the potential to change redox states by application of electrical fields [55]. While the latter has been reported to influence cell adhesion [55], cell stimulation by electrical fields on electrically conductive

polypyrrole

substrates

was

reported

to

have

influence

on

cell

differentiation

[56].

Zhang et al. (2018) described a digital light-processed (DLP) polyacrylamide-PEGDA hydrogel that exhibited electrical conductivity by ionic conductivity through the addition of LiCl (Fig. 5) [27]. The UV-based curing in DLP was achieved by the addition of soluble 2,4,6-trimethylbenzoyl-diphenylphosphine oxide (TPO) nanoparticles. High-printing resolutions of up to 7 µm were achieved by this process, resulting in stretchable complex structures exhibiting high elasticity (ϵ = 1300%). Biocompatibility was assessed by seeding NH-3T3 fibroblast cells, although no toxic effects were reported [27]. The resulting hydrogels were printed on flexible substrates, allowing the

Fig. 5. Polyacrylamide (PAAM) – PEGDA digital light-processed (DLP) 3D-printed conductive hydrogel structures. Highly elastic and stretchable electrically functional gels can be fabricated using LiCl as a dopant (adapted from Zhang et al. (2018)[27]). Reproduced with permission from [27] (the Royal Society of Chemistry).

development of electrically conductive and bendable circuits promising for flexible biosensor applications.

4.

3D Printing Technologies for Electrically Conductive Hydrogel Processing

Various technologies have been utilised to fabricate three-dimensional (3D) structures from electrically conductive hydrogels or electrically conductive hydrogel precursors (Table 1). An overview of the basic 3D printing principles used in the reviewed literature is illustrated in Fig. 6, namely, bioplotting (Fig. 6. A), light-based approaches (Fig. 6. B) and ink-jet printing (Fig. 6. C). In 3D bioplotting, either conductive hydrogels or conductive hydrogel precursors are deposited through a nozzle on a substrate of interest (e.g. tissue culture well plate) utilising, for example, pressurised air or piston displacement (Table 1). Often relying on the shear thinning behaviour of the printed material, the applied pressure forces the hydrogel to flow, hence introducing shear forces, which allow material deposition in the desired patterns. A key challenge is to ensure the shape stability of the printed structures, caused by, for example, the recovery of the prior sheared hydrogel or by the crosslinking of the printed structure directly after printing (e.g. ionic, covalent by UV light). In light-based printing approaches such as stereolithography (SL/SLA), direct laser writing (DLW) or DLP, a light-reactive hydrogel precursor is either illuminated by a travelling beam of UV light or laser or by a projected pattern on the hydrogel precursor to induce curing in specific regions to form crosslinked patterned areas. Often a z-displacement of the precursor reservoir is performed to allow layer-bylayer deposition in the z-direction to fabricate 3D structures (Fig. 6). Acrylate polymer-based chemistry has been frequently used for such processes as photocrosslinkable hydrogels, combined with electrically CPs (Table 1). Using inkjet or piezo printing, microvolumes (µl) of either conductive hydrogel or conductive hydrogel precursor (e.g. monomer, oxidant) are deposited on a printing stage by deformation of a piezo element, leading to the deposition of small droplets (Fig. 6, C). The following section discusses examples and advantages of each technique utilised for 3D printing of electrically conductive hydrogels.

Fig. 6. 3D printing techniques to process electrically conductive hydrogels. (A) Bioplotting of conductive hydrogel ink containing conductive particles or conductive-polymer precursors (e.g. pyrrole (py), aniline (ani), edot) combined with protein, polysaccharides or synthetic polymer matrices (Table 1). (B) Light-based printing using light (e.g. UV)-curable hydrogel precursors in a digital light-processing (DLP) or stereo lithography (SLA/SL) setup, or alternatively by curing already 3D bioplotted hydrogels by light exposure. (C) Ink-jet printing by (i) direct deposition of conductive hydrogel or (ii) simultaneous deposition of conductive hydrogel precursor (monomer) and oxidant triggering conductive polymer formation during printing as demonstrated before [43].

4.1 Bioplotting A popular technique to process conductive hydrogels with suitable properties is 3D bioplotting. Some of the relevant previous studies in this field are discussed in this section. Methacrylated chitosan (ChiMA) was loaded with 3% (wt) graphene nanosheets (GNS) and bioplotted in an extrusion approach through a 200 µm-diameter nozzle to form grid-like, electrically conductive scaffolds [32]. Because of unsatisfactory rheological properties for room temperature printing, the GNS-ChiMA composite was printed inside a precipitation bath consisting of isopropanol, triggering strut solidification and sedimentation inside the bath due to the density differences between isopropanol (0.786 g.ml-1) and GNS-ChiMA composites (~1.2 g.ml-1) [32]. After 3D printing, network formation was induced by UV light exposure to initiate crosslinking, making the printed scaffolds stable in aqueous solution [32]. It was observed that with increasing graphene-nanosheet content, Young’s modulus and tensile strength increased from 3.2 ± 0.1 to 6.6 MPa and 683.5 kPa to 1510.4 ± 52 kPa, respectively, in comparison to pure ChiMA. The conductivity of dried films was determined by a 4-point probe measurement, leading to conductivities of up to 0.0025 S.cm-1 in the case of 3% (wt) graphene nanosheet-laden ChiMA. A viscosity increase was observed by the addition of graphene nanosheets (~3-4 Pa.s at a shear rate of 1/s) for bioplotting [32]. The UV-crosslinked structures were characterised in terms of biocompatibility by seeding L929 mouse fibroblasts. Håkansson et al. (2016) bioplotted a cellulose nanofibril and CNT composite [42]. Techniques for drying the printed hydrogels without structural loss of geometry were identified with optimal results by a combination of liquid nitrogen and freeze drying. Conductivities of 0.1 (dry) and 0.01 (wet, hydrogel) S.cm-2 were achieved (Table 1) [42]. In another approach, chitosan with in situ-polymerised polypyrrole (PPy) was 3D-printed using a supportive gelatin-APS layer [30]. After printing the hydrogels, the dissolution of the gelatin-APS layer was triggered. The APS release resulted in a second solidification of the hydrogels, allowing successful printing [30]. To form the conductive hydrogel, polypyrrole was in situ-polymerised using APS in the presence of double-bond-grafted chitosan (DCH) to form DCHPPy particles. Second, by the addition of acrylic acid (AA), N,N’-methylenebis-acrylamide (MBAA) and FeCl3 to those particles, poly(acrylic acid) (PAA) was formed, resulting in a final double-network chitosan hydrogels (CSH). On the basis of the presence of MBAA, self-healing properties of the hydrogels were identified to originate from reversible ionic interactions of Fe3+ with NH groups of pyrrole and carboxylic acid groups of PAA [30]. The electrical conductivity was measured using a 4-point probe, with conductivities of 20–70 S.cm-1. Darabi et al. (2017) highlighted that both ionic contributions by Fe3+ and electrical conductivity through PPy were present in their gel system, explaining the high conductivity achieved [30]. In addition to self-healing characteristics, the gels showed a high elasticity, allowing more than 1500% of strain before break. The highlighted applications for the hydrogels were in 3D print technology, wearable electronics and medical sensors [30].

85% wt

addition

ChitosanMA,

Bioplotti

Graphene

Particle

GrapheneNS

ng

3% wt

addition

CollagenMA,

UV-Bio

CNT

Particle

CNTb,alginate

printing

GelMA, PANI

µSLAe

a

/

/

addition PANI

Post

Cl-

polym. GelMA, DNA-

Bio

CNT

coated

plotting

Particle

/

0.0025

PEGDA, TPO

DLP

h

g

Ionic

/

printing

Li

+

upregulation of glial and

fibre

nerve guidance

neuronal genes by conductivity

4-Point,

6.6 ±

Tissue

0.4

Engineering

0.01 - 0.02

EISc,

43 -

Tissue

kΩ @ 5 Hzb

Scaffold

54e-3

Engineering

2.9 kΩ @

EIS,

13.7 -

Bioelectronic

0.1 Hz

2-Point,

(165.5 kΩ)

Hydrogel

24 ± 1.8

4-Point,

0.6 ±

Dry Film

0.1

Cl

/

15.2e

-3

/

13 e

-3

-

NPs PNIPAM-L-

Plotting

CNT

SLA

PEDOT:PSS

CNT

Particle

/

0.016–0.02

PSS

662 ± 100.6

addition

i

PEGDA, j

Particle addition

Ω.□

-1

Key Findings

  Neurogenic differentiation,





L929

  Increased mechanical and





HCAECd

  Exposure time dep. cell





10T1/2f

  SLA-printed GelMA possible, successful modification with

biomedical

PANI, biocompatible scaffolds 



Tissue

4-Point

Flexible

[57]

[32]

[24]

viability, optimal exposure time

interfaces, Flex. biosensors,

Ref

conductive properties

Cardio

  Cells seeded, integrated in

Myocytes

[47]

[29]

hydrogels, high degree of

engineering

Acrylamide-

a

hMSC

Engineering,

Film

addition

 

printed

MWCNT

PEDOT:PSS

Tissue

Parameters

ng

3 ± 0.4

Electrical Stimulation

PLGA

2-Point,

Cell Type

8

In-vitro Biocompatibility

/

Hydrogel

Particle

Application

Conductivity/S.cm-1

Graphene

Young´s Modulus / MPa

Dopant

Bioplotti

Material

Measurement Method

Modification

Graphene,

Printing

Conductivity

Method

Conductive Phase

Table 1 | 3D-printed conductive hydrogels for biomedical and tissue engineering applications including key results achieved, according to the recent literature

cardiomyocyte maturation 



NIH3T3

  Highly stretchable (ϵ = 1300 %)

electronics,

hydrogel-elastomer, high

biostructures

resolution (7 µm), biocomp. 



L929

  Self-healing, laponite





DRG



14.2 –

Wearable

probe

46.8

electronics

Cyclic

26.3 ±

Tissue

voltamm

4.2 –

engineering,

encapsulated in GelMA, ES by

etry

35.3 ±

structural support

surrounding condu. hydrogel,

1.4

for ele. stim. (ES)

incr. neural. diff markers

[27]

[75]

reinforced, adhesive 

Dorsal root ganglia (DRG) cells,

[73]

Chitosan methacrylate (MA) containing graphene nanosheets (NS), bcarbon nanotubes (CNT), impedance at 5 Hz, c electrical impedance spectroscopy (EIS), dhuman coronary artery endothelial cells (HCAEC), eµStereolithography (SLA), referring to a resolution

achievable in the micrometre range, fmurine embryo fibroblast cell line 10T1/2, g2,4,6- trimethylbenzoyl-diphenylphosphine oxide (TPO), hdigital light processing (DLP), iPNIPAM-Laponite-CNT hydrogels (PNIPAM-L-CNT), jpolystyrenesulfonate (PSS)

addition

ng

mC.cm-2

etry, Hydrogel

Tissue



Engineering,

NE-4C

m

NSC



Parameters



Electrical Stimulation

2.21 ±

Cell Type

/

In-vitro Biocompatibility

0.7

Particle

Hydrogel

voltamm

CNT

Application

Young´s Modulus/MPa

Bioprinti

1.1 ±

Conductivity/S.cm-1

MWCNT-NH2

Cyclic

0.12l

Dopant

SL-

Modification

PEGDA , k

Conductivity

Method

Conductive Phase

Printing Material

Measurement Method

Table 1 (continued) | 3D-printed conductive hydrogels for biomedical and tissue engineering applications including key results achieved, according to the recent literature

Key Findings

DC,

500 µA direct pulse stimulation

pulse

induced neuronal and

spinal cord

100 µA

oligodendroglial differentiation

repair

-1A

of NE-4C neural cells, increased

Ref [26]

TUJ1 and GFAP PANI, Phytic

Ink-jet

PANI

acid

In-situ

Phyti

polym.

c

0.23

4-Point,

/

Pellet

Bio

PPy

plotting

PEGDA, PPy

µSLA

PPy

In-situ

Fe3+







 

medical

acid DChitosann,

Biosensors,

4-Point,

0.774 -

3D printed

Hydrogel

0.237

wearable

0.013 -

2-Point,

0.6 -

Bioelectronics

3.5

Thin film,

1.4

Biomedical

4-Point,

/p

Bioelectronics

polym.

(ink-jet) and conductivity 





 

Post

MOo

polym. Bio

nanofibrils,

plotting

CNT

CNT

Particle addition

/

~0.1 0.01, wet vs dry

[30]

healing capacity hydrogels 





 

High-resolution printing,

[25]

mechanically tough honeycomb

mΩ.cm-1 Cellulose

Highly flexible (ϵ = 1500 %), conductive, 3D printable self-

Sensor PPy

[43]

supercapacitors, high-throughput

electrodes 25 - 70

High cycling performance as

structures 





 

Drying of hydrogel structure with

Thin film;

Biomed, not

sustained pore size, printing with

2-Point,

specific

2% nanocellulose

[42]

Hydrogel k PEGDA

combined with NH2-decorated multi-walled carbon nanotubes (MWCNT), lcharge capacity, mmurine neuroectodermal neuronal stem cells NE-4C NSC, ndouble-bond-decorated chitosan (DChitosan), omethyl orange (MO), pnot tested on hydrogels

printing

Dielectric

0.82 -

artificial



ted

1 MHz

Spectrom

21.86

tissue, soft

fabrication, tough/elastic, complex

eter

e-3

actuators/sen

geometries, macroscale 3D objects

sors, robotics

possible, production rate 3 cm.h-1

NPs AAM,TEMEDq

plotting

Ionic

/

/

, LiCl

0.10 ±

4-Point,

0.03r

Hydrogel

Soft sensors,

/s

(0.029 ±











Direct ink

PAAM

writing

Ionic

/

/

ΔC 45 pFt

maskless

Acid (PAA)

SL

PEGDA,

SL

CNT

MWCNT PAAM-

/ Particle

/ /

/ 21.6 Ω.m

Ionic

/

electronics

hydrogel, elastomeric device,

PEGDA

(PAAM), acrylamide (AAM), tetramethylethelendiamine (TEMED),

measurement

Sensor-motor

103 e

-3

rin

[31]

[28]





  Pneumatic-powered haptic displays

[33]

(PHD)

mechanics 0.6 -

long periodic

Thin film,

1.4

gating

2-Point



coupling, soft

19.7 –

Liquid sensor







  new optical maskless stereolithography,







  Liquid sensor based on swelling altering

[58]

pH sensors, optical microfibre sensor [74]

conductivity

231 –

4-Poing,

0.0569

Smart

222.5

Hydrogel

– 0.435

electronics,

Ω.□-1 tresistivity

73 -

29.5 /

  Highly elastic (up to 150% strain), transparent and conductive ionic

2-Point,

addition DLP

q polyacrylamide

Capacita Hydrogel

/

  Transparent and conductive, rapid

Ref

stretchable, sensoric functionality

nce,

(DIW) Polyacrylic

Key Findings

stretchable

0.004) Silicone,

Parameters

0.029 @

SiO2 Bio

Electrical Stimulation

Sulfa

Cell Type

/

Hydrogel PDMS,PAAM

In-vitro Biocompatibility

Young´s Modulus/MPa

Measurement Method

Conductivity/S.cm-1

Hydrogel

composite

Ionic

Application

SL-

Dopant

Ionic

Modification

Method

Conductivity

Printing Material

Conductive Phase

Table 1 (continued) | 3D-printed conductive hydrogels for biomedical and tissue engineering applications including key results achieved, according to the recent literature







  High-fidelity body signal acquisition

[76]

biosensors

correspondence to liquid salt solution conductivity of printed lines (in brackets), sviscoelastic behaviour was analysed of rheological modifier of the hydrogel precursor,

4.2 Direct laser writing (DLW), µStereolithography (µSLA) and light-based approaches A combination of CollagenMA with alginate and carboxylated-CNTs has been UV-bioprinted, with conductivity introduced by CNT addition [24]. The interest was to fabricate 3D-printed electrically conductive cardiac patches. CNT-laden alginate (25 and 50 mg CNT per mg alginate, final concentration: 2.5% (w/v)) was printed inside a 50 mM CaCl2 solution to form a printing framework [24]. After CaCl2 removal, cell-laden CollagenMA was printed between the CNT-alginate struts in an alternating manner. Human coronary artery endothelial cells (HCAECs) were printed inside CollagenMA, and the hydrogels were crosslinked using UV light (CollagenMA) and CaCl2 (alginate) [24]. A crosslinking time of 45 seconds was identified for sufficient crosslinking of CollagenMA, not compromising cell viability. This approach was hence a combination of bioplotting and UV light crosslinking-based printing. The CNT concentration of 10 mg CNT/mg CollagenMA showed no significant effect on HCAEC viability. The incorporation of CNT significantly decreased the impedance of the gels in comparison to pure CollagenMA at frequencies of 5 Hz [24]. Spread and viable cells were similarly observed for 0.05 mg CNT/mg alginate solutions. A stiffness in the range of 43–54 kPa over 20 days of incubation was achieved with similarity to the elastic modulus of the right ventricular myocardium of rats [24,59]. The results indicated no cytotoxic effects of CNTs at the used concentration, while the desired effect of forming an electrically conductive gel-CNT composite was achieved, not compromising cell viability but adding functionality to the gels. PEGDA-MWCNT and PEGDA-PPy hydrogels have both been light-assisted printed recently [25,26]. In the first approach, amine-decorated MWCNTs were added to form a particle-hydrogel composite with PEGDA. The material was printed in square-pore geometries using a customised stereolithography (SL) platform (355 nm UV laser, spot diameter 190 ± 50 µm) [25]. In the second approach, PPy was post-polymerised on DLP PEGDA (X-Y resolution 39 µm, light intensity 30 mW.cm-2) after printing (Fig. 7) [25]. In the case of PEGDA incorporating amine-decorated MWCNTs, 3D-printed scaffolds for neuronal tissue engineering applications were achieved [26]. The addition of 0.1% MWCNTs resulted in an increase in the Young’s modulus by 189%, while the water contact angle (20% to >40%) and conductivity were also increased in comparison to pristine PEGDA [26]. The electrical charge capacity increased significantly from 0.113 ± 0.09 mC.cm-2 to 2.21 ± 0.12 mC.cm-2 by addition of MWCNTs to bare PEGDA. It was found that the total neurite length of grown NE-4C neural stem cells (NSC) increased, with an increase in

Fig. 7. DLP-printed PEGDA with phenylbis(2,4,6-trimethylbenzoyl)phosphine oxide as a photoinitiator (PI) and methyl orange(MO) as a dopant for polypyrrole (PPy), formed by post-polymerisation after printing by subsequent incubation in both FeCl3 and Py solutions (adapted from Fantino et al. (2018)[25]). Reproduced with permission from [25] (Wiley).

neural differentiation factors (TUJ1, early-stage neural differentiation and GFAP, intermediate filament protein, astrocyte specific) observed when electrical stimulation (ES) was applied on MWCNT-laden scaffolds [26]. This result highlights the relevance of the ES of electrically conductive hydrogels to achieve positive effects on cell maturation. With main focus on biosensor applications, 3D DLP-printed PEGDA was electrically functionalised with PPy [25]. Methyl orange (MO) was utilised as a dopant. Resistivities of 0.013–3.5 mΩ.cm-1 were measured on thin films using the 2-point probe measurement method. It was possible to achieve mechanically tough structures utilising a honeycomb geometry, encouraging the possibility of shaping scaffolds through 3D printing as a key approach to yield toughened and mechanically tailored structures (Fig. 7) [25]. 4.3 Ink-Jet printing Pan et al. (2012) 3D-printed conductive hydrogels by sequential deposition of two distinct solutions consisting of CP monomer precursor plus dopant (aniline/phytic acid) (i) and oxidising agent (ammonium persulfate, APS) (ii) by ink-jet printing (Fig. 8) [43]. The merging of droplets led to the polymerisation of aniline to PANI by oxidation using APS. Using this approach, ink-jetting of in situ formed PANI/phytic acid hydrogels with a resolution of ~18 µm was achieved [43]. This approach highlights the possibility of utilising 3D printing technology to directly deposit chemical compounds required to trigger the polymerisation of CPs, achieving defined patterns of conductive

Fig. 8. Ink-jet-printed patterns of polyaniline (PANI)- phytic acid conductive hydrogels (adapted from Pan et al. (2012) [43]). Reproduced with permission from [43] (PNAS).

hydrogel in high spatial resolution.

5.

Discussion

Different material combinations have been employed to achieve 3D-printed electrically conductive hydrogels (Table 1). The highest conductivity among the reviewed materials was in the range of 25–70 S.cm-1 [30]. In the corresponding study [30], these values were derived from a 4-point probe measurement of swollen hydrogels containing Fe3+ and PPy in the wet state. The authors attributed the high electrical conductivity to the contributions from Fe3+ ionic and PPy electronic conductivity [30]. Because of the limitations of 4-point probe measurements to differentiate between these contributions, it is highlighted that wet-state gels might be assessed with spectroscopic techniques to identify the different contributions of ions and CPs to the overall conductivity of the hydrogels. By this approach, the efficacy of conductivity enhancement by the incorporation of CPs could be identified as well as comparability achieved among different studies using different hydrogel systems. The second highest conductivity of the hydrogels reported was 24 ± 1.8 S.cm-1, measured on dry films of MWCNTfunctionalised GelMA [29]. Through the assessment of dry films, the conductivity contribution can be related to the contribution of the MWCNT modification, excluding potential influences of ionic mobility in wet-state gels. However, drying of hydrogels may lead to an increasing percolation probability inside the material because of the removal of the liquid phase and shrinkage, which may change after re-hydration of the gels. Exemplary, Håkansson et al. (2016) assessed CNT-modified hydrogels using 4-point and 2-point probe methods as possible approaches to determine dry- and wet-state conductivity of the hydrogel composite material, respectively [42]. In general, different 3D-printable conductive hydrogels with conductivities ranging from 0.0025 S.cm-1 (graphene nanosheets [32]) to 0.1 S.cm-1 (LiCl [28], CNT [42]), 0.23 S.cm-1 (PANI [43]) and >20 S.cm-1 (PPy, Fe3+, wet-state gels [30]) were reported (Table 1). Hence, tunable conductivities utilising all possible functionalisation methods (in situ, postpolymerisation, particle addition, ionic) for 3D-printable conductive hydrogels can be engineered. This allows a potential selection of the material system dependent on the requirements of conductivity fitting the application of interest. Frequency-dependent electrical properties were assessed by CV and EIS techniques [24][26][47][31,33]. Different printable hydrogels showed impedances of 0.01–0.02 kΩ at 5 Hz (<0.02 kΩ at 0.1 Hz) for CNT-based and 2.9 kΩ at 0.1 Hz for PANI-based functionalisation [24][47]. MWCNT-NH2-functionalised PEGDA showed a charge capacity of 2.21 ± 0.12 mC.cm-2 [26]. Hydrogels with an ionic-based conductivity of up to 0.029 S.cm-1 @ 1 MHz were reported [31]. PAAM silicone hydrogels showed a capacitance of

ΔC

= 45 pF [33]. Because of the complexity of frequency-

dependent electrical behaviour and the differences in parameters reported in the different studies, direct comparison is barely possible. The differences in reported values and assessment relate to the final application of the gels, which range from sensor motor soft mechanics [33] and tissue actuators [31] to TE [24, 26, 31]. Regarding mechanical properties, 3D-printed conductive hydrogels with a wide range of mechanical properties have been reported. Highly stretchable gels of up to 1500% strain were achieved [30]. Interestingly, stiffness values ranging from 200 to 700 kPa for hydrogels allowing 1500% strain [30] to 13 kPa for hydrogels allowing 1300% strain [27] were identified. Hence, depending on the application of interest, different material systems are

available that combine the required electrical conductivity with suitable mechanical properties. 3D-printed conductive hydrogels covering a wide range of stiffnesses from ~13–54 kPa [24,27,47], 100–600 kPa [26] to 700– 1100 kPa [29] and 3–6.6 MPa [32] were reported. As a result, this technology allows the potential to select different 3D printable conductive hydrogels of various stiffnesses to meet the stiffness of the target tissue in tissue engineering applications, mimicking the mechanical properties of the native ECM in each case. The highest printing resolution reported was achieved by inkjet printing (18 µm) [42], followed by DLP/light-based approaches (details down to <100 µm [31] and 200 µm [25]) and 3D bioplotting (100–250 µm range [29]). In DLPbased printing, Fantino et al. (2018) functionalised printed structures electrically by post-polymerisation [25]. Interestingly, post-polymerisation of the structures did not influence the resolution of the printed structures [25]. Hence, this approach could be applied on different high-resolution 3D-printed structures without altering their resolution, having the potential of being highly promising to form high-resolution and conductive 3D-printed structures [25]. Two main applications have been considered in this review, namely, sensoric and tissue engineering applications. When applied as body movement sensors, conductive NP-based hydrogels could be enclosed inside supportive structures (e.g. PDMS) and applied on the skin outside the human body. As a result, potentially toxic components like NPs, oxidants or CP precursors can be sealed from direct human contact. This type of applications may require thus less toxicological consideration in comparison to in vivo applications. In contrast, for tissue engineering approaches aiming to yield tissue analogues or to facilitate and control cellular response by the electrically conductive hydrogels (scaffolds), the material systems require an in-depth toxicological evaluation. Hence, depending on the site and type of application (sensoric, outside the body vs tissue engineering, aiming for implantation), the materials have to fulfil different toxicological requirements. In addition, considering conductive hydrogels, potentially toxic products from chemical oxidation reactions (oxidant, monomers) must be considered for the final application of 3D-printed electrically conductive hydrogels. In the two main fields of application, the reviewed literature revealed advantages of 3D-printed ECHs, such as superior elasticity of conductive flexible hydrogels of up to 1300%-1500% strain [27, 30] in comparison to metalbased electronic circuits. Indeed, ionic-based ECHs have the potential to achieve transparency and conductivity in applications such as soft actuators [31]. In addition, ECHs exhibit the ability to trigger, control and influence cells in tissue engineering strategies through 3D-printed structures [26,29,57,73] such as electrically conductive substrates of stiffness similar to that of native tissues (Table 1). In the present literature search, it was not possible to identify a biofabrication approach for conductive hydrogels defined by simultaneous printing of cells encapsulated in an electrically conductive-hydrogel precursor. One inherent reason may be the mainly cytotoxic chemistry employed when forming conductive hydrogels by oxidative polymerisation using in situ and post-polymerisation approaches [21,22]. One challenge to be addressed is the identification of a chemical composition of CP hydrogels, which allows the formation of a homogeneous conductive matrix simultaneously to incorporating cells in the hydrogel precursor, not compromising cell viability.

One possible approach for the fabrication of cell-laden conductive hydrogels is conductive-particle-functionalised hydrogels [24,26][32][39]. The potential of functionalisation of these particles (e.g. amine-functionalised MWCNT) may allow a crosslinking to the bulk hydrogel [26]. This could be a valuable approach to retain particles in the polymer matrix by covalent crosslinking and to prevent their unwanted release. The authors highlighted that positive neuronal cell response occurred due to the electrostatic interaction of the potential positive charge of the amine-decorated MWCNTs with the negatively charged plasma membrane of neuronal cells [26][70]. Improved cellular adhesion and spreading capacity on positively charged surfaces due to electrostatic interaction with negatively charged ECM biomolecules relevant for osteoblast adhesion and spreading has been also described [71, 72]. CPs in the oxidised form feature a positively charged backbone with charge neutrality caused by incorporating anions [21]. The backbone charge can be altered by changing the CP redox state when applying an external electrical field [21]. Different effects of conductive hydrogels on cellular responses have been reported, and comprehensive reviews have been published on the subject, which summarise the effects of these materials on cell behaviour for biomedical applications [20–22]. Increased cell attachment has been associated with changes in surface charge, for example, by addition of polypyrrole-containing cations along its polymer backbone [60]. The potential to change the overall surface charge of CPs by changing surface redox state from oxidised to reduced was shown to be a tool, allowing significant influence on cell adhesion [20,55], being also useful for drug delivery [21]. By incorporation of conductive materials in hydrogels, homogeneously expanding conductive matrices can be achieved to optimise cell ES and signal transduction [22,60]. Moreover, ES has been shown to influence voltagegated Ca2+ channel activity [61], influencing cellular pathways and cell proliferation [62] in various cell types such as vascular cells [61], stem cells [63–65], fibroblasts [66,67] and chondrocytes [68,69]. Hence, ES of 3D-printed electrically conductive hydrogels may be a valuable tool to tailor charge distributions inside hydrogel networks and hence to control cellular growth and adhesion [29]. In TE applications, the correct assessment of scaffolds’ relevant electrical properties that correlate with cellular response is challenging. Considering that in in vitro and in vivo scenarios scaffolds perform in aqueous environments, the assessment of hydrogels under dry conditions does not lead to the realistic appraisal of the relevant material properties. In ES setups, it is therefore critical to be able to determine scaffold properties, such as conductivity, in the wet state as well as frequency-dependent electrical resistance (impedance) values. A variety of techniques has been reported to assess the electrical properties of 3D-printed conductive hydrogels (Table 1). Although techniques like four-point probe setups will yield conductivity values independent of scaffold geometry, their application may be most suitable for dry samples owing to its direct current (DC)-based measuring principle, which may be prone to errors when applied to wet, electrolyte swollen hydrogels. Two-point probe electrical resistivity measurements using two electrodes to assess the final 3D-printed scaffold geometry will allow gaining information of the electrical behaviour of the actual hydrogel construct. These measurements are relevant for scaffolds in direct ES setups utilising two electrodes to apply electric potentials. It is understood that the

mechanical properties of the native tissue vary depending on scale [77,78], while mechanical properties of scaffolds can significantly alter cellular response. Analogue to mechanical properties, differences of electrical properties in scale, for example, geometry-independent (microscopic) properties in comparison to bulk electrical properties (macroscopic) and their relative influence on cellular response, require further investigations. A deeper understanding of such effects will influence the choice of methodology to measure electrical properties to gain knowledge of the correlation of electrical properties of hydrogel scaffolds with cellular response. Indeed, the cellular response on ES has been shown to be at the microscopic level [61, 62]. Therefore, cells seeded on scaffold surfaces under wet conditions may be influenced by material properties on the micrometre range, which requires the assessment of wet scaffolds to yield geometry-independent electrical properties. The final application of the scaffold, for example, direct current ES versus alternating current ES, may determine the method of choice to characterise 3D-printed conductive hydrogels. Frequency-dependent electrical properties assessed by CV or EIS may allow a more precise assessment of the wet hydrogel behaviour.

6.

Final remarks and conclusions

We have reviewed the field of 3D printing of electrically conductive hydrogels for medical applications. The main applications of such materials in combination with 3D printing technology are in the fields of biosensors, flexible electronics and tissue engineering. From the increasing volume of literature in this field, it is clear that academic and industrial researchers have started to recognise the advantages and challenges in this area. The main functionalisation methods to obtain electrical conductivity inside hydrogels were presented, which were identified as composite approaches utilising conductive particles (CNTs, graphene), in situ or post-polymerisation of CPs (polypyrrole PPy, PANI, PEDOT) and ionically conductive hydrogels. Examples from recent literature (the past six years) were displayed, featuring 3D printing by light-based methods such as DLP using mainly methacrylation chemistry, 3D bioplotting and ink-jet printing. The measurement of conductivity was performed with various methodologies, ranging from 4-point probe measurement techniques to 2-point probe measurements, EIS and CV. It was shown that the conductivity of hydrogels can originate in both electronic and ionic contributions. We suggest that electrochemical spectroscopic techniques have to be employed to determine the conductivity in wet, conductive hydrogels. Given the complexity of the electrically conductive hydrogel systems, it is very important to understand the different phases of the specific conductive hydrogel and their relative contributions towards the total conductivity of the gels. The latter would allow a better comparison of different material systems, with the aim to highlight the most promising printing and functionalisation approaches for 3D-printed conductive hydrogels in the future. For the emerging field of biofabrication, that is, printing with cells, chemical functionalisation and composite approaches leading to cell-friendly, printable and electrically conductive systems must be investigated.

Acknowledgements This research was supported by the German Research Foundation (Deutsche Forschungsgemeinschaft, DFG) within the Collaborative Research Centre 1270 ELAINE on ELectrically Active ImplaNts. References [1]

A.K. Gaharwar, M.S. Detamore, A. Khademhosseini, Emerging Trends in Biomaterials Research, Ann. Biomed. Eng. 44 (2016) 1861–1862.

[2]

I. Stassen, T. Sloboda, G. Hambitzer, Membrane with controllable permeability for drugs, Synth. Met. 71 (1995) 2243–2244.

[3]

J.M. Pernaut, J.R. Reynolds, Use of conducting electroactive polymers for drug delivery and sensing of bioactive molecules. A redox chemistry approach., J. Phys. Chem. B. 104 (2000) 4080–4090.

[4]

B. Garner, A. Georgevich, A. Hodgson, L. Liu, G.G. Wallace, Polypyrrole-heparin composites as stimulusresponsive substrates for endothelial cell growth, J Biomed Mater Res. 44 (1999) 121–9.

[5]

M.R. Abidian, D.-H. Kim, D.C. Martin, Conducting-Polymer Nanotubes for Controlled Drug Release, Adv. Mater. 18 (2006) 405–409.

[6]

Y. Li, K.G. Neoh, E.T. Kang, Controlled release of heparin from polypyrrole-poly(vinyl alcohol) assembly by electrical stimulation, J. Biomed. Mater. Res. - Part A. 73 (2005) 171–181.

[7]

C.E. Schmidt, V.R. Shastri, J.P. Vacanti, R. Langer, Stimulation of neurite outgrowth using an electrically conducting polymer, Proc. Natl. Acad. Sci. 94 (2007) 8948–8953.

[8]

N. Gomez, J.Y. Lee, J.D. Nickels, C.E. Schmidt, Micropatterned polypyrrole: A combination of electrical and topographical characteristics for the stimulation of cells, Adv. Funct. Mater. 17 (2007) 1645–1653.

[9]

J.G. Hardy, R.C. Cornelison, R.C. Sukhavasi, R.J. Saballos, P. Vu, D.L Kaplan, C.E. Schmidt, Electroactive Tissue Scaffolds with Aligned Pores as Instructive Platforms for Biomimetic Tissue Engineering, Bioengineering. 2 (2015) 15–34.

[10]

D. Mawad, E. Stewart, D.L. Officer, T. Romeo, P. Wagner, K. Wagner, G.G. Wallace, A single component conducting polymer hydrogel as a scaffold for tissue engineering, Adv. Funct. Mater. 22 (2012) 2692–2699.

[11]

D. Mawad, C. Mansfield, A. Lauto, F. Perbellini, G.W. Nelson, J. Tonkin, S.O. Bello, D.J. Carrad, A.P. Micolich, M.M. Mahat, J. Furman, D.J. Payne, A.R. Lyon, J.J. Gooding, S.E. Harding, C.M. Terracciano, M.M. Stevens, A Conducting polymer with enhanced electronic stability applied in cardiac models, Sci. Adv. 2 (2016) e1601007.

[12]

M. Li, Y. Guo, Y. Wei, A.G. MacDiarmid, P.I. Lelkes, Electrospinning polyaniline-contained gelatin nanofibers for tissue engineering applications, Biomaterials. 27 (2006) 2705–2715.

[13]

P.M. George, A.W. Lyckman, D. A. LaVan, A. Hegde, Y. Leung, R. Avasare, C. Testa, P.M. Alexander, R. Langer, M. Sur, Fabrication and biocompatibility of polypyrrole implants suitable for neural prosthetics, Biomaterials. 26 (2005) 3511–3519.

[14]

B.J. Edgerton, W.F. House, W. Hitselberger, Hearing by cochlear nucleus stimulation in humans., Ann Otol Rhinol Laryngol Suppl. 91 (1982) 117–24.

[15]

T.F. Otero, J.M. Sansiñena, Bilayer dimensions and movement in artificial muscles, Bioelectrochemistry Bioenerg. 42 (1997) 117–122.

[16]

M.T. 2003 Otero, T. F. & Cortes, A sensing muscle., Sens. Actuat. B 96 (2003) 152 –156.

[17]

R.H.Baughman, Conducting polymer artificial muscles, Synth. Met. 78 (1996) 339–353.

[18]

W.G. Adeloju SB, Conducting polymers and the bioanalytical sciences: new tools for biomolecular communication. A review., Analyst. 121 (1996) 699–703.

[19]

C.E. Guimard, N. K., Gomezb, N. & Schmidt, Conducting polymers in biomedical engineering., Prog. Polym. Sci. 32 (2007) 876–921.

[20]

R. Ravichandran, S. Sundarrajan, J.R. Venugopal, S. Mukherjee, S. Ramakrishna, Applications of conducting polymers and their issues in biomedical engineering, J. R. Soc. Interface. 7 (2010) S559-S579.

[21]

D. Mawad, A. Lauto, G.G. Wallace, Polymeric Hydrogels as Smart Biomaterials, (2016) 19–45.

[22]

R. Balint, N.J. Cassidy, S.H. Cartmell, Conductive polymers: Towards a smart biomaterial for tissue engineering, Acta Biomater. 10 (2014) 2341–2353.

[23]

J.H. Min, M. Patel, W.G. Koh, Incorporation of conductive materials into hydrogels for tissue engineering applications, Polymers (Basel). 10 (2018) 1–36.

[24]

M. Izadifar, D. Chapman, P. Babyn, X. Chen, M.E. Kelly, UV-assisted 3D bioprinting of nano-reinforced hybrid cardiac patch for myocardial tissue engineering, Tissue Eng. Part C Methods. 24 (2017) 74-88.

[25]

E. Fantino, I. Roppolo, D. Zhang, J. Xiao, A. Chiappone, M. Castellino, Q. Guo, C.F. Pirri, J. Yang, 3D Printing/Interfacial Polymerization Coupling for the Fabrication of Conductive Hydrogel, Macromol. Mater. Eng. 303 (2018) 1–8.

[26]

S.J. Lee, W. Zhu, M. Nowicki, G. Lee, D.N. Heo, J. Kim, Y.Y. Zuo, L.G. Zhang, 3D printing nano conductive multi-walled carbon nanotube scaffolds for nerve regeneration, J. Neural Eng. 15 (2018) 016018.

[27]

B. Zhang, S. Li, H. Hingorani, A. Serjouei, L. Larush, A. a. Pawar, W.H. Goh, A.H. Sakhaei, M. Hashimoto, K. Kowsari, S. Magdassi, Q. Ge, Highly stretchable hydrogels for UV curing based high-resolution multimaterial 3D printing, J. Mater. Chem. B. 6 (2018) 3246–3253.

[28]

K. Tian, J. Bae, S.E. Bakarich, C. Yang, R.D. Gately, G.M. Spinks, M. in het Panhuis, Z. Suo, J.J. Vlassak, 3D Printing of Transparent and Conductive Heterogeneous Hydrogel–Elastomer Systems, Adv. Mater. 29 (2017) 1604827.

[29]

S.R. Shin, R. Farzad, A. Tamayol, V. Manoharan, P. Mostafalu, Y.S. Zhang, M. Akbari, S.M. Jung, D. Kim, M. Comotto, N. Annabi, F.E. Al-Hazmi, M.R. Dokmeci, A. Khademhosseini, A Bioactive Carbon Nanotube-Based Ink for Printing 2D and 3D Flexible Electronics, Adv. Mater. 28 (2016) 3280–3289.

[30]

M.A. Darabi, A. Khosrozadeh, R. Mbeleck, Y. Liu, Q. Chang, J. Jiang, J. Cai, Q. Wang, G. Luo, M. Xing, Correction to: Skin-Inspired Multifunctional Autonomic-Intrinsic Conductive Self-Healing Hydrogels with Pressure Sensitivity, Stretchability, and 3D Printability, Adv. Mater. 29 (2017) 1700533.

[31]

J. Odent, T.J. Wallin, W. Pan, K. Kruemplestaedter, R.F. Shepherd, E.P. Giannelis, Highly Elastic, Transparent, and Conductive 3D-Printed Ionic Composite Hydrogels, Adv. Funct. Mater. 27 (2017) 1701807.

[32]

S. Sayyar, S. Gambhir, J. Chung, D.L. Officer, G.G. Wallace, 3D printable conducting hydrogels containing chemically converted graphene, Nanoscale. 9 (2017) 2038–2050.

[33]

S.S. Robinson, K.W. O’Brien, H. Zhao, B.N. Peele, C.M. Larson, B.C. Mac Murray, I.M. Van Meerbeek, S.N. Dunham, R.F. Shepherd, Integrated soft sensors and elastomeric actuators for tactile machines with kinesthetic sense, Extrem. Mech. Lett. 5 (2015) 47–53.

[34]

Y. Zhao, B. Liu, L. Pan, G. Yu, 3D nanostructured conductive polymer hydrogels for high-performance electrochemical devices, Energy Environ. Sci. 6 (2013) 2856–2870.

[35]

F.L. Buchholz, A.T. Graham, Modern superabsorbent polymer technology, John Wiley & Sons Ltd., Chichester, 1997. pp. 1 – 282.

[36]

Q.A. Acton, Polyethylene Glycols—Advances in Research and Application, ScholarlyEditions, Atlanta, 2013. pp. 1 – 432.

[37]

L. Dai, Conjugated and fullerene-containing polymers for electronic and photonic applications: Advanced syntheses and microlithographic fabrications, J. Macromol. Sci. Polym. Rev. 39 (1999) 273–387.

[38]

G. Malliaras, R. Friend, An Organic Electronics Primer, Phys. Today. 58 (2005) 53–58.

[39]

H.K. Chitte, G.N. Shinde, N. V. Bhat, V.E. Walunj, Synthesis of Polypyrrole Using Ferric Chloride (FeCl3) as Oxidant Together with Some Dopants for Use in Gas Sensors, J. Sens. Technol. 01 (2011) 47–56.

[40]

T.K. Vishnuvardhan, V.R. Kulkarni, C. Basavaraja, S.C. Raghavendra, Synthesis, characterization and a.c. conductivity of polypyrrole/Y2O3 composites, Bull. Mater. Sci. 29 (2006) 77–83.

[41]

C. Cha, S.R. Shin, X. Gao, N. Annabi, X.S. Tang, A. Khademhosseini, Controlling mechanical properties of cell-laden hydrogels by covalent incorporation of graphene oxide, Small. 10 (2014) 514–23.

[42]

K.M.O. Håkansson, I.C. Henriksson, C. de la Peña Vázquez, V. Kuzmenko, K. Markstedt, P. Enoksson, P. Gatenholm, Solidification of 3D Printed Nanofibril Hydrogels into Functional 3D Cellulose Structures, Adv. Mater. Technol. 1 (2016) 1–9.

[43]

L. Pan, G. Yu, D. Zhai, H.R. Lee, W. Zhao, N. Liu, H. Wang, B.C.-K. Tee, Y. Shi, Y. Cui, Z. Bao, Hierarchical nanostructured conducting polymer hydrogel with high electrochemical activity, Proc. Natl. Acad. Sci. 109 (2012) 9287–9292.

[44]

S. Yang, L.K. Jang, S. Kim, J. Yang, K. Yang, S. Cho, J.Y. Lee, Polypyrrole / Alginate Hybrid Hydrogels : Electrically Conductive and Soft Biomaterials for Human Mesenchymal Stem Cell Culture and Potential Neural Tissue Engineering Applications, Macromol. Biosci. 16(2016) 1–9.

[45]

J. Yang, G. Choe, S. Yang, H. Jo, J.Y. Lee, Polypyrrole-incorporated conductive hyaluronic acid hydrogels, Biomater. Res. 20 (2016) 1–7.

[46]

Y. Wu, Y.X. Chen, J. Yan, S. Yang, P. Dong, P. Soman, Fabrication of conductive polyaniline hydrogel using porogen leaching and projection microstereolithography, J. Mater. Chem. B.3 (2015) 5352–5360.

[47]

Y. Wu, Y.X. Chen, J. Yan, D. Quinn, P. Dong, S.W. Sawyer, P. Soman, Fabrication of conductive gelatin methacrylate-polyaniline hydrogels, Acta Biomater. 33 (2016) 122–130.

[48]

J. Nilsson, A.H. Castro Neto, F. Guinea, N.M.R. Peres, Electronic properties of graphene multilayers, Phys. Rev. Lett. 97 (2006) 266801.

[49]

K.S. Novoselov, A.K. Geim, S.V. Morozov, Y. Jiang, S.V. Dubonos, I.V. Grigorieva, A.A. Firsov, Electric field effect in atomically thin carbon films, Science. 306 (2004) 666–669.

[50]

P.C. Ray, H. Yu, P.P. Fu, Toxicity and Environmental Risks of Nanomaterials: Challenges and Future Needs, J. Env. Sci Heal. C Environ. Carcinog. Ecotoxicol Rev. 27 (2009) 1–35.

[51]

P.H. Hoet, I. Brüske-Hohlfeld, O.V. Salata, Applications of nanoparticles in biology and medicine, J. Nanobiotechnol. 2 (2004) 12.

[52]

R.E. Gosselin, H.C. Hodge, R.P. Smith, M.N. Gleason, Clinical Toxicology of Commercial Products, 4th ed., Williams and Wilkins, Baltimore, 1976, pp. II-85.

[53]

H. Aral, A. Vecchio-Sadus, Toxicity of lithium to humans and the environment—A literature review, Ecotoxicol. Environ. Saf. 70 (2008) 349– 356.

[54]

J. Exon, A review of the toxicology of acrylamide, J. Toxicol. Env. Heal. B Crit. Rev. 9 (2006) 397–412.

[55]

K. Svennersten, M.H. Bolin, E.W.H. Jager, M. Berggren, A. Richter-Dahlfors, Electrochemical modulation of epithelia formation using conducting polymers, Biomaterials. 30 (2009) 6257–6264.

[56]

R. Saigal, E. Cimetta, N. Tandon, J. Zhou, R. Langer, M. Young, G. Vunjak-Novakovic, S. Redenti, Electrical Stimulation via a Biocompatible Conductive Polymer Directs Retinal Progenitor Cell Differentiation, 35th. Conf. Proc. IEEE Eng. Med. Biol. Soc. (2013) 1627–1631.

[57]

A.E. Jakus, E.B. Secor, A.L. Rutz, S.W. Jordan, M.C. Hersam, R.N. Shah, Three-dimensional printing of highcontent graphene scaffolds for electronic and biomedical applications, ACS Nano. 9 (2015) 4636–4648.

[58]

M.J. Yin, M. Yao, S. Gao, A.P. Zhang, H.Y. Tam, P.K. Wai, Rapid 3D Patterning of Poly(acrylic acid) Ionic Hydrogel for Miniature pH Sensors, Adv. Mater. 28 (2016) 1394–1399.

[59]

L.E. Engelmayr, G.C., Cheng, M., Bettinger, C.J., Borenstein, J.T., Langer, R., L.E. Freed, Accordion-like

honey- combs for tissue engineering of cardiac anisotropy., Nat. Mater. 7 (2008) 1003-1010. [60]

M. Brett Runge, M. Dadsetan, J. Baltrusaitis, T. Ruesink, L. Lu, A.J. Windebank, M.J. Yaszemski, Development of Electrically conductive Oligo(polyethylene Glycol) Fumarate-Polypyrrole Hydrogels for Nerve Regeneration, Biomacromolecules. 8 (2010) 2845–2853.

[61]

A.S. Rowlands, J.J. Cooper-White, Directing phenotype of vascular smooth muscle cells using electrically stimulated conducting polymer, Biomaterials. 29 (2008) 4510-4520.

[62]

M.R. Love, S. Palee, S.C. Chattipakorn, N. Chattipakorn, Effects of electrical stimulation on cell proliferation and apoptosis, J. Cell. Physiol. 233 (2018) 1860–1876.

[63]

M.L. Hernández-Bule, C.L. Paíno, M.Á. Trillo, A. Úbeda, Electric stimulation at 448 kHz promotes proliferation of human mesenchymal stem cells, Cell. Physiol. Biochem. 34 (2014) 1741–1755.

[64]

J. Du, G. Zhen, C. Huanwen, S. Zhang, L. Qing, X. Yang, G. Lee, H.Q. Mao, X. Jia, Optimal electrical stimulation boosts stem cell therapy in nerve regeneration, Biomaterials. 181 (2018) 347–359.

[65]

H.J. Kwon, G.S. Lee, H. Chun, Electrical stimulation drives chondrogenesis of mesenchymal stem cells in the absence of exogenous growth factors, Sci. Rep. 6 (2016) 1–13.

[66]

S. Snyder, C. DeJulius, R.K. Willits, Electrical Stimulation Increases Random Migration of Human Dermal Fibroblasts, Ann. Biomed. Eng. 45 (2017) 2049–2060.

[67]

M. Rouabhia, H. Park, S. Meng, H. Derbali, Z. Zhang, Electrical stimulation promotes wound healing by enhancing dermal fibroblast activity and promoting myofibroblast transdifferentiation, PLoS One. 8 (2013) e71660.

[68]

G.A. Rodan, L.A. Bourret, L.A. Norton, DNA synthesis in cartilage cells is stimulated by oscillating electric fields, Science 199 (1978) 690–692.

[69]

R.J. Fitzsimmons, S.L. Gordon, J. Kronberg, T. Ganey, A.A. Pilla, A pulsing electric field (PEF) increases human chondrocyte proliferation through a transduction pathway involving nitric oxide signaling, J. Orthop. Res. 26 (2008) 854–859.

[70]

Hwang J Y, Shin U S, Jang W C, Hyun J K, Wall I B and Kim H W. Biofunctionalized carbon nanotubes in neural regeneration: a mini-review, Nanoscale 5 (2013) 487–97.

[71]

F. Kunz, H. Rebl, A. Quade, C. Matschegewski, B. Finke, J.B. Nebe, Osteoblasts with impaired spreading capacity benefit from the positive charges of plasma polymerised allylamine. Eur. Cells Mater. 29 (2015) 177-89.

[72]

H. Rebl, B. Finke, J. Schmidt, H.S. Mohamad, R. Ihrke, C.A. Helm, J.B. Nebe, Accelerated cell-surface interlocking on plasma polymer-modified porous ceramics, Mater. Sci. Eng. C. Mater. Biol. Appl. 69 (2016) 1116–24.

[73]

D.N. Heo, S.J. Lee, R. Timsina, X. Qiu, N.J. Castro, L.G. Zhan, Development of 3D printable conductive hydrogel with crystallized PEDOT:PSS for neural tissue engineering, Mater. Sci. Eng. C. Mater. Biol. Appl. 99 (2019) 582-590.

[74]

X. Li, Y. Yang, B. Xie, M. Chu, H. Sun, S. Hao, Y. Chen, Y. Chen, 3D Printing of Flexible Liquid Sensor Based on Swelling Behavior of Hydrogel with Carbon Nanotubes, Adv. Mater. Technol. 4 (2019) 1800476.

[75]

Z. Deng, T. Hu, Q. Lei, J. He, P.X. Ma, B. Guo, Stimuli-Responsive Conductive Nanocomposite Hydrogels with High Stretchability, Self-Healing, Adhesiveness, and 3D Printability for Human Motion Sensing, ACS Appl. Mater. Interfaces. 11 (2019) 6796−6808.

[76]

X.Y. Yin, Y. Zhang, X. Cai, Q. Guo, J. Yang, Z.L. Wang, 3D printing of ionic conductors for high-sensitivity wearable sensors, Mater. Horiz. 6 (2019) 767-780.

[77]

R. Akthar, M. K. Sherratt, J. Kennedy Cruickshank, B. Derby, Characterizing the elastic properties of tissues, Material Today. 14 (2011) 96-105.

[78]

D. W. Evans, E. C. Moran, P.M. Baptista, S. Soker, J. L. Sparks, Scale-dependent mechanical properties of

native and decellularized liver tissue. Biomeh Model Mechnobiol. 12 (2013) 569-580.

Statement of Significance: Hydrogels are among the most frequently used biomaterials in tissue engineering (TE). A new class of hydrogels, namely, electrically conductive hydrogels (ECHs), has been introduced in recent years. Although ECHs have been comprehensively reviewed in the literature, the combination of ECHs with 3D printing technology has emerged only recently, representing a promising key development toward the

fabrication of functional 3D TE constructs. In this review, we compressively cover for the first time the current state of the art in the field of 3D printing of ECHs. Previous advances are presented, reviewing the 3D printing technologies utilised, spatial resolution and electrical conductivity values achieved, in addition to discussing the obtained mechanical properties and emerging applications of these materials.

Statement of Significance: Hydrogels are among the most frequently used biomaterials in tissue engineering (TE). A new class of hydrogels, namely electrically conductive hydrogels (ECHs), has been introduced in recent years. While ECHs have been comprehensively reviewed in the literature, the combination of ECHs with 3D printing technology has emerged only recently representing a promising key development towards the fabrication of functional 3D TE constructs. In this review, we compressively cover for the first time the current state of the art in the field of 3D printing of ECHs. Previous advances are presented, reviewing 3D printing technologies utilized, spatial resolution and electrical conductivity values achieved, in addition to discussing the obtained mechanical properties and emerging applications of these materials.