Neural interfaces engineered via micro- and nanostructured coatings

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Review

Neural interfaces engineered via micro- and nanostructured coatings Nuan Chen a,b , Lingling Tian b , Anoop C. Patil b,c , Shengjie Peng a,∗ , In Hong Yang b , Nitish V. Thakor b,c,d,∗∗ , Seeram Ramakrishna a,e,∗ a

Center for Nanofibers and Nanotechnology, Department of Mechanical Engineering, National University of Singapore, Singapore 117576, Singapore Singapore Institute for Neurotechnology (SINAPSE), National University of Singapore, Singapore 117456, Singapore Department of Electrical & Computer Engineering, National University of Singapore, Singapore, 117576, Singapore d Department of Biomedical Engineering, Johns Hopkins University, Baltimore, MD 21218, USA e Guangdong-Hongkong-Macau Institute of CNS Regeneration (GHMICR), Jinan University, Guangzhou 510632, PR China b c

a r t i c l e

i n f o

Article history: Received 14 January 2017 Received in revised form 19 March 2017 Accepted 24 April 2017 Available online xxx Keywords: Neural interface Electrode coating Carbon Conducting polymer Nanotechnology Functionalization

a b s t r a c t During the past decade, neural interfaces have attracted great interest due to their potential of body-machine communication for disease diagnosis and therapy. Considering the significant material mismatch between the electrode implants and the native tissue, a thin coating is employed on the electrode sites as an intermediate layer to bridge the material differences and has been proved to play an important role in the promotion of neural cell attachment and signal transmission. Micro- and nanostructured coating materials, together with surface functionalization such as biological cues, provide not only high surface areas for signal transduction but also a biomimetic platform for the cells. In this paper, we review the performance of different kinds of micro- and nanostructured coating materials including metallic materials, carbon materials, conducting polymeric materials and composite materials and then complement the discussion with the influence of the fabrication process on the performance of the coatings. The coating could be functionalized by various advanced techniques, which are also reviewed. The existing challenge and future research directions of electrode coatings are briefly discussed at the end of the review. © 2017 Elsevier Ltd. All rights reserved.

Contents Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 The transmission of neural signal . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Transmission of neural signals between neurons . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Charge transfer at the electrode-tissue interface . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Advanced coating materials for neural interfaces . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Metallic materials and their derivatives . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Platinum and gold . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Iridium oxide . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Titanium nitride . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Carbon materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Carbon nanotube . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Graphene . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Conducting polymeric materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Composite materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Metal-carbon composite . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Carbon-polymer composite . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00

∗ Corresponding authors at: Center for Nanofibers and Nanotechnology, Department of Mechanical Engineering, National University of Singapore, Singapore, 117576, Singapore. ∗∗ Corresponding author at: Singapore Institute for Neurotechnology (SINAPSE), National University of Singapore, Singapore, 117456, Singapore. E-mail addresses: [email protected] (S. Peng), [email protected] (N.V. Thakor), [email protected] (S. Ramakrishna). http://dx.doi.org/10.1016/j.nantod.2017.04.007 1748-0132/© 2017 Elsevier Ltd. All rights reserved.

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Challenges and opportunities. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .00 Acknowledgements . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00

Introduction Nervous system diseases such as Amyotrophic Lateral Sclerosis [1], Parkinson’s [2] and seizures [3] are among the thorniest problems in clinical medicine. Due to the serious consequences and especially hopeless self-recovery from damages to the central nervous system caused by these diseases [4], it is crucial to employ sensing and/or modulating devices that can “read” and “write” neural signals assisting diagnosis and therapy of nervous disorders. During the past decades, neural electrodes have been developed as promising interface technology for direct communication with the neural tissues [5]. Due to the superior electrical conductivity, chemical stability and biocompatibility to a certain extent, noble metals, such as gold [6], platinum [7] and iridium [8], are commonly used materials for neural electrodes. However, traditional metallic electrodes have their limitation due to the mismatch with resident tissue in electrical, mechanical and biological properties, which is a huge hurdle to achieve a neural electrode with a satisfactory signal delivery [9], signal detection [10] and least tissue response [11]. The electrical difference comes from two different charge transport carriers, ions in neural tissue and electrons in electrodes [12]. The large disparity in mechanical properties causes deformation difference and tiny movement in micrometer or even millimeter ranges at the interface [13]. From a biological perspective, the implanted electrodes are recognized as foreign objects and are rejected by the immune system [14]. To date, designing an ideal electrode using analogous organic materials closely mimicking the soft material properties of the resident tissues remains a significant challenge. Biomimetic and smart device with organic materials to model neural signal transduction in vivo should be an excellent design, which require the use of combination of neuroscience, materials science and engineering. However, it has not been realized. Compared to metallic and carbon materials, organic materials are not good candidate for signal transduction inside the electrode due to their low electrical conductivity [15]. The softness and flexibility of organic materials make them difficult to penetrate the body and be immobilized in specific position. Although new fabrication methods for the organic chip like optical lithography have been gradually developed [16], traditional fabrication methods for inorganic electrodes and chips still show their superiority on both quality and efficiency. Thus, traditional metallic neural electrodes still play an important role, but the significant materials mismatch between tissues and them is inevitable at the present. As a result, an intermediate layer should serve as a charge transfer platform and mechanical buffer with good biocompatibility for the neural interface. Coating, as a convenient and straightforward method to create a layer at the electronic device/tissue interface and change the properties of electrode sites locally, become an important consideration. This review focuses on the application of micro- and nanoscale coatings for neural interfacing. The junction between neural interfacing and electrode coatings is worthy of research interests for several reasons: (i) Neural tissues are electrically active. Understanding the functioning of the underlying neural circuitry is possible through reliable electrode interfaces. (ii) The material properties of the electrode coatings are critical for reliable integration of the neural electrodes with the neural tissues as chronic implants. (iii) The development of an ideal neural interface is limited by the availability of specialized materials and coatings that can render the electrode interfaces friendlier to the surrounding tis-

sues. It is thus quintessential to assess the current state of the neural electrode coatings and the recent developments in coating procedures/technologies that have influenced neural interfacing. As an intermediate layer between the neural electrode and the target tissues, the electrode coating plays an important role in determining the efficiency of the abiotic/biotic interface. For example, a coating layer with poor electroactivity (a material property determining difficulty of charge transfer between ions and electrons) and electrical conductivity will significantly limit the effectiveness of signal transduction at the interface [10,17]. Failure of providing an interface with good chemical and mechanical properties may cause severe degradation to the surrounding tissues, affecting the signal acquisition and/or modulation [7,18]. The combination of these factors motivates us to undertake the review of micro- and nanoscale coatings to neural electrode surfaces, as well as to highlight the successes of various research efforts in this field and the challenges faced in achieving an ideal neural interface. It is essential to note that both the factors: the selection of materials and the related processing (including fabrication and surface functionalization) should be carefully considered while defining a coating layer for the neural electrode. Strategies for neural electrode coating design is shown in Fig. 1. The scope of the article is limited to the coating for the electrode sites that are in direct contact with the tissues. The coating applied to the neural electrodes affecting the electrical modality of the interface is discussed. The modification of the neural electrode interface with regard to chemical and optical modalities is beyond the scope of this article. In the following sections, the mechanism of signal transmission in the nervous system and the signal transduction at the electrode/tissue interface is discussed first. We then present a holistic view of the materials employed for the neural electrode coating (including metallic materials, carbon materials, conducting polymeric materials and other composite materials) along with key process related parameters and results for each of the coating materials. We use this as a template to further discuss and review in detail, the electrical, mechanical and biological properties of each of the coating materials and the performance of the resulting coating layers in recording and stimulating of neural signals. Furthermore, advanced techniques for surface functionalization of the electrode coating layers are reviewed. As a conclusion to the review, the existing challenges that face the field and future research directions regarding neural electrode coatings are discussed. The transmission of neural signal Transmission of neural signals between neurons Signals transmission in the nervous system depends on the electrical carrier (ions) and chemical carrier (neurotransmitters) [19]. Electrical signals are transported along the axon and dendrite while chemical signals are transported at the synapses. When an action potential reaches a neuron through the dendrites, the signals are processed and sent out to another neuron or effector through the axon and the synapse. The depolarized state of the cell membrane is transported along the unmyelinated axon while ions exchanged across the cell membrane occur only at the Ranvier nodes to refresh the signals in the myelinated axon [20]. When the action potential reaches the chemical synapse (the most common one), the voltage-gated Ca2+ channels are activated to allow Ca2+ flow into the presynaptic terminal, which induces the release of neurotrans-

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Fig. 1. Strategies for neural electrode coating design.

mitters into the synapse [21,22]. The released neurotransmitters molecules bind to the postsynaptic receptors, followed by the activation of ligand-gated Na+ channel and the entry of Na+ into the postsynaptic cell. Finally, the postsynaptic potential initiated by the increased Na+ concentration is delivered to the neuron cell body. The movement of ions like Na+ and K+ through the cell membrane changes the concentration of ions in the extracellular fluid (ECF) locally. Charge transfer at the electrode-tissue interface As discussed in the last section, the signal transport in the nervous system induces the flux of ions through the neural cell membrane and changes the concentration of the ions in the ECF. Since ion is the charge carrier in ECF, its movement would modulate the electric field. A neural electrode either detects the change of the electric field in the ECF (recording) or alters the electric field for excitation or inhibition of neurons (stimulating) [23,24]. However, charges are carried by electrons in the conductive body of the neural electrode, implying transduction from ionic conduction as observed in the ECF to electronic conduction in the neural electrode. Thus, electroactive materials which enable charge transfer between ions and electrons are necessary at the neural interface [7,12]. A schematic illustration of the signal transmission between the neuron and the neural electrode is shown in Fig. 2. According to the difference in the charge transfer mechanism, electroactive materials can be classified as capacitive and faradaic (pseudocapacitive) materials [7,23,25]. Comparisons between faradaic and capacitive materials are shown in Table 1. It should be noted that the highly reversible charge transfer is conducted within the potential window of the materials. If the applied voltage on the materials exceeds the potential window, the irreversible electrolysis of the water occurs at the interface [7,23,26]. The electrolysis of the water has been shown to be harmful to the surrounding tissues as it would generate gas and cause localized changes in the pH of the ECF [7,23]. Also, the electroac-

tivity and conductivity of a material are not consistent as they represent the difficulty of charge transfer at electrode/electrolyte interface and charge transport through a conductive material, respectively. It should be noted that a highly conductive material may not be a highly electroactive material. For example, platinum has higher conductivity but lower electroactivity than conducting polymers (CPs) at room temperature [32–35]. However, since high conductivity allows fast and efficient electron transport during charge transfer at the interface, the increase in conductivity of the material may contribute to the increase in its electroactivity. In addition to the electroactive nature of the interface materials, the electroactivity of the neural interface is also significantly influenced by the surface morphology of the interface [7,23]. Since the charge transfer occurs at the electrode/ECF interface, larger surface area of the interface would provide more reaction sites for charge exchange and consequently increase the effective capacitance of the interface [24,28,31,36–39]. Besides, the high effective surface area is significantly necessary for small-sized electrode aiming to locally record and/or stimulate as it would provide an intimate contact with the surrounding tissues while maintaining the small geometry, which achieves both selectivity and sensitivity. The nanoscale structure shows marked superiority at the neural interface due to its ultrahigh surface-to-volume ratio. The electroactivity of a neural interface is shown by some electrochemical quantities such as impedance, charge storage capacity (CSC) and charge injection limit (CIL). Impedance is the evaluation parameter of the opposition that a circuit presents to a current when a voltage is applied. During stimulation, a lower impedance at the neural interface allows smaller applied voltages to achieve similar current amplitudes, which would reduce the likelihood of harmful effects including abnormal electroporation of cells [23,40], polarization of electrode [41–43] and electrolysis of water [7,23]. At the same time, it is probable to achieve localized high current stimulation safely at the electrode-tissue interface with lower impedance [41]. As for recording, the reduction of the impedance will increase signal intensity and reduce the thermal

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Fig. 2. Schematic illustration of the signal transmission process neural interface.

noise [7,44–46], which enables clearer detection of the neural signals. Thus, a neural interface with low impedance is favorable for both stimulating and signal recording. CSC is the measure of the sum of passed charges at the interface during the cyclic voltammetry (CV) which is just within the potential window [7,23,47]. Larger CSC of a neural interface would mean more delivered charge during stimulation and higher signal intensity during recording [7]. CIL is a very important parameter for electrical stimulation as it sets up the threshold of charge injection density without electrolysis of water [42]. High CIL will enable large charge delivery under the safe conditions, which permits effective stimulation of the neurons. Also, it allows the safe usage of small-sized electrode for local stimu-

lation whose charge density would be high [42]. In the following sections, the electrochemical performances of different coatings for neural interfaces will be reviewed and compared mainly based on the aforementioned three electrochemical quantities. Advanced coating materials for neural interfaces In the following section, different kinds of coating materials for neural interface including metallic materials and their derivatives, carbon materials, conducting polymeric materials and composite materials are reviewed. An overview of the key properties and biological studies that have been done for different coating materials

Table 1 Comparison between two categories of electroactive materials.

Examples Charge transfer mechanism Chemical reaction of common materials at the interface

Main factor to influence the electroactivity Advantages

Faradaic materials

Capacitive materials

gold (Au), platinum (Pt), iridium oxide (IrOx ) and conducting polymer (CP) [7] electrochemical reaction between the surface of the electrode and ECF [7,23,26] Au+3H2 O ↔ Au2 O3 +6H+ +6e− (2.2.1) (2.2.2a) Pt+H2 O ↔ PtO+2H+ +2e− (2.2.2b) PtO+H2 O ↔ PtO2 +2H+ +2e− (2.2.2c) Pt-H ↔ Pt+H+ +e− (2.2.3) Ir2 O3 +H2 O ↔ 2IrO2 +2H+ +2e− − x+ − − [CP]+xA ↔ [(CP) (A )x ]+xe (small  dopant)  + (2.2.4a) x+ − + [ CP A− x (C )x ] ↔ [(CP) (A )x ]+xC +xe− (big

titanium nitride (TiN), carbon nanotube (CNT) and graphene [23] the redistribution of ions/electrons at the electrode/ECF interface [7,23,26] –

dopant) (2.2.4b) [23,27–29] the efficiency of the redox reaction of the materials [7] high electroactivity unit surface area [7,28,30]

the adsorption of the ions to the materials and the formation of the electric double layer [23] high electrochemical stability [7,23,31]

Representative cyclic voltammetry curve [7,24,30,31]

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Charge Transfer Mechanism

Coating Thickness (nm)

Conductivity (S/m)

Impedance Reduction (impedance of bare site/impedance of coated site) at 1 KHz

Charge Storage Capacity increase

Charge Injection Limit (mC/cm2)

Water Window(V)

Young’s modulus

Cells for in vitro biocompatibility test

Cell and animal type in electrophysiological and electrical stimulation tests

Metallic materials and their derivatives

Pt

Faradaic

103 [48]

106 –107 [49]

∼2–800 times [18,41,42,48,50–56]

∼40 times [52]

∼0.05–24 [18,42,52–54]

−0.6–0.8 [48]

∼140 GPa [57]

SH-SY5Y cell [52]

rat cortical neural cell [58], mouse hindbrain and spinal cord [41], hamster substantia gelatinosa neuron [56] (In vitro); rat [50,54,55], crucian carp [53] (In vivo);

Au

Faradaic

102 [59–62]

106 –107 [60]

∼4–50 times [59,60,62–65]

∼10 times [64]

∼1 [64]

−0.5–0.8 [60,64]

∼70 GPa [57]

NG108-15 cell [60], rat primary cortical cell [61]

rat primary neuronal cell [63,64], rat organotypic hippocampal slice [62] (In vitro);

IrOx

Faradaic

102 –103 [9,66–71]

105 –106 [72]

∼4–150 times [44,71,73–78]

∼3–24 times [68,73,75,77–79]

∼1–4 [30,71,73,77]

−0.6–0.8 [80]

–

rat primary cerebral cortex neurons [66], human embryonic stem cell derived neuronal cell [74], rat cortical neuron [76,81]

rat cortical neuron [76], rat brain slices [82], human embryonic stem cell derived neuronal cell [74] (In vitro); cat [79], rat [83] (In vivo);

TiN

Capacitive

103 [30]

105 [84]

∼1.2times [85]

∼40times [85]

∼0.8–4.5[30,86,87]

−1–1.2 [86]

∼80–640 GPa [88]

CNT

Capacitive

102 –105 [89–96]

104 –105 [97–99]

∼4–70 times [45,91,92,95,97,100–104]

∼3–140 times [92,100,105]

∼1–4 [90,100]

−1.6–1 [90,92,100,106]

∼1GPa-1.8TPa [107]

mouse dissociated frontal cortex [105], rat cortical cells [108], crayfish lateral giant neuron [109], rat hippocampus neuron [90,102,104], rat ventricular myocytes [91], PC12 cell [97,110,111], mouse hindbrain-spinal cord and hippocampal cells [106], 42MG-BA cell [97],3T3 cell [97], 293T cell [97], NG108-15 cell [98,99], L-929 cell [112]

NG108-15 cell [99], rat hippocampal neuron [90,104], rat cortical cells [89], crayfish lateral giant neuron [45,94,101,102,109], rat hippocampus slide [113], mouse hindbrain-spinal cord and hippocampal cells [106], rabbit retina [97], mouse retina [96] (In vitro); rat [92,100,103,105], monkey [93,105] (In vivo);

Graphene

Capacitive

100 –102 [11,114,115]

103 [116]

∼1.5–10 times [115,117,118]

∼6 times [115]

∼0.02 [117]

–

∼1TPa [119,120]

rat cortical neuron [117], rat hippocampus neuron [116], PC12 cell [11,121], mice hippocampus neuron [115]

rat hippocampal neuron [116], mouse hippocampal neuron [115], PC12 cell [121] (In vitro); rat [11] (In vivo)

PPy

Faradaic

102 –105 [122–125]

102 –103 [15,126,127]

∼1–40 times [122,128–131]

∼2400 times [122]

–

−0.9–0.5 [122,128]

∼2 MPa–3 GPa [12]

rat cortical cells [132], rat primary dorsal root ganglia (DRG) explants [122,133], rat and mouse neural stem cell [134], rat spinal ganglion neuron (SGN) explant [123], PC12 cell [124,131,135,136]

rat SGN explant [123], PC12 cell [124] (In vitro); rat [128], pig [130] (In vivo);

PEDOT

Faradaic

102 –105 [122,137–144]

102 –104 [126,145]

∼2–300 times [10,122,128,137– 143,146–157]

∼15–5000 times [122,141,144,146, 147,150,151]

∼1.2–3 [146,147,158]

−0.9–0.5 [122,151]

∼2 MPa–3 GPa [12]

rat cortical cell [138], rat primary DRG explants [122], murine fibroblastic cell L929 [35], PC12 [35,131,136,144,154, 157–160], SH-SY5Y cell [142,152,153], rat primary neurons and astrocyte [147], rat primary motor neurons [161], rat hippocampus neuron [162]

rat cortical cell [138] (In vitro); rat [10,128,139,147, 149,155,163,164], pig [165] (In vivo);

Carbon materials

Conducting Polymeric Materials

rat [9] (In vivo);

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Table 2 Overview of properties of different coating materials for neural interface.

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is shown in Table 2. In the following sections, detailed discussion on the electrical, mechanical and biological properties and the performance of resulting coatings in stimulating and recording will be given, followed by the effect of fabrication processes on the coatings and advanced functionalization techniques. All characterizations on electrical and mechanical performance of the materials were done without living things unless otherwise specified. Metallic materials and their derivatives Due to the superior electrical conductivity, biostability and corrosion resistance [48], noble metals, such as gold, platinum, and iridium, are commonly used materials for neural electrodes. However, the poor electroactive nature due to the high chemical stability, limits their applications at the neural interface [7,23,28,29,38]. Thus, some surface modifications are needed to improve the electrode-tissue interface. From an engineering perspective, it is always preferable to use as few different materials as possible to avoid parasitic effects like bimetallic corrosion or delamination [48,52,166] so metallic materials and their derivatives such as platinum black and iridium oxide are chosen as the coating materials for metallic electrode sites. Besides, metallic materials and their derivatives exhibit higher mechanical strength than other materials [52], which is essential for preventing the failure of the electrodes. Commonly metallic materials and their derivatives for neural electrode coating include platinum, gold, iridium oxide and titanium nitride. Platinum and gold As platinum and gold are the most widely used neural electrode materials [10,91,100,102,167], coatings made of these materials are considered for application on the surface of the electrode sites. To improve the electrochemical property of the coating, the effective surface area is enlarged by employing micro- and nanoparticles (NPs) or structures with high porosity. Platinum black coating is a layer of fine platinum powder that is commonly formed by electrochemical deposition (ED). The ED of platinum black, which is also called platinization, can be dated back to 1894 [168] and it has become a mature technique currently [58]. The high surface-to-volume ratio of these micro-sized platinum powders is quite beneficial to charge transfer at the neural interface [41,48,50,53,58]. With the rough platinum black layer, the impedance of the electrode at biologically relevant frequency (1 kHz) is reduced at least by a factor of 4 [51,53,55,56] and the impedance did not show significant increase after a 3-week postimplantation [50], indicating good electrochemical stability of the coating. Also, the CIL of the electrode increased by a factor of 6 after the employment of the platinum black coating, which contributed to a more efficient and safer charge delivery during stimulation [51,53]. However, studies showed that the purely plated platinum black coatings exhibited low mechanical durability, which greatly limited their long-term application [50]. Ultrasonic agitation was found to be an effective method to improve the stability of the plated platinum black coating [50,169]. The mechanical durability of the platinum black coating was significantly improved by the application of ultrasonic agitation (Fig. 3a & b). After ultrasound cleaning test, only 2.5% of the coating was lost on sonicoplated coating while more than 80% of the coating was lost in case of non-ultrasonically coated electrode sites [169]. Also, the platinum black coating displayed a more uniform but rougher morphology when coupled with the ultrasonic agitation, with as much as 6 times larger electrochemical areas than their non-ultrasonically plated counterparts [50,169]. Owing to the high surface area, the impedance of the sonicoplated coating was lower than those of pulsed plated one and direct-current plated one during in vivo operation [50]. Compared

to the bare electrode, sonicoplated electrode showed 43%, 98% and 83% reduction on the stimulation voltage, duration, and amplitude of stimulation artifacts respectively, indicating a more efficient and low-noise neural interface [50]. Besides, the sonicoplated coating displayed excellent chronic stability with no evident of severe impairment of electrochemical property or failure during several months of implantation [50,53,169]. Although the mechanical stability of the platinum black coating can be improved by using electroplating of platinum black under sonication, the cytotoxicity of the platinum black remains a serious issue. A study showed that the DNA synthesis of the rat oligodendrocytes in vitro was inhibited when exposed to a platinum black extract, possibly due to the release of toxic lead, an ingredient of the conventional electroplating electrolyte [41,48,52,170]. Thus, new methods for forming porous platinum coating were developed to improve the neural electrode interface. Chemical and electrochemical depositions without cytotoxic components like lead were employed to form three-dimensional (3D) nanofibrous structure of Pt. The tissue-friendly performances of these two methods have been confirmed in vitro by the nontoxic effect of the elution products of the resulting coatings on cells [52]. Chemical deposition allows the formation of nanowire structures which have a higher effective surface area while ED could produce clearly defined pattern of the coating efficiently without post-cleaning. Due to the high effective surface area of the nanostructures, the electrochemical property of the coating was significantly improved, with around 45-fold reduction of impedance (at 1 kHz) and 40-fold increase of area under the CV-curve compared to the bare electrode, as shown in Fig. 3c. The highly porous Pt coating showed two times charge delivery capacity as large as those of sputtered IrOx and electrodeposited Poly(3,4-ethylenedioxythiophene) (PEDOT) coatings. By these highly electroactive coatings, the size of the electrode sites can be reduced by 90% without the impairment of the performance [52]. In addition to the electrochemical property, the coatings also exhibited good mechanical and electrochemical stability with the evidence of negligible change on surface morphology and electrochemical performance after cleaning procedures and pulse testing [52]. To increase the surface area of the coating for a more intimate interface, surface roughening methods were usually included in the fabrication process to form fuzzy Pt coating on electrode sites. The ED of platinum-copper alloys and subsequent removal of copper was shown to fabricate a cauliflower-like micro-sized Pt coating with good mechanical and electrochemical stability. It also exhibited a better electrochemical property than sputtered Pt due to higher effective surface area. Although there are still ∼9% copper residues which may be toxic to the cells, it might be cleaned ultimately by an additional etching step [48]. Templating is also a useful technique to create a porous structure. By adopting lyotropic liquid crystal and silica microbeads, a well-ordered, micro-structured and mechanically robust Pt coating was formed on the neural electrode. Although the impedance reduction and CSC increase of the coating are not as effective as those of platinum black, it still enabled low-noise and clear detection of low-amplitude spikes [41]. The formation of Pt coating also can be achieved by electrochemical roughening via repeated oxidation and reduction cycling and the morphology of the resulting coating is shown in Fig. 3g. With highly rough surface, the coating displayed comparable charge injection capacity (CIC) to that of CNT. Good mechanical stability and electrochemical stability of the coating were confirmed under harsh ultrasonication and 14-day biphasic stimulation, respectively [42]. In addition to incorporating surface roughening methods during deposition, pretreatment of the substrate facilitate the roughening of the coating as well. Using reactive-ion etching before sputtering of Pt, the effective surface area of the coating was significantly

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Fig. 3. (a) Scanning electron microscope (SEM) images of sonicoplated Pt black coating before and after 60 min ultrasonic durability test. (b) Mean impedance traces of the coatings by different plating modes and percentage increase in the impedance values after being subjected to 60 min of sonication in saline.**p = 0.001 compared to control; *p < 0.05 compared to pulsed plated and direct current plated electrodes. (a)(b) Reproduced with permission from Ref. [50], Copyright 2010, The Authors. (c) SEM image of chemically and electrochemically deposited nanostructured Pt coating and the CV enlargement factor of these two kinds of coating with respect to bare Pt. Reproduced with permission from Ref. [52], Copyright 2015, The Authors. (d) SEM images of nanostructured Au coating by employing NPs. Reproduced with permission from Ref. [60], Copyright 2012, American Chemical Society. (e) SEM images of nanostructured Au coating by employing nanowires. Reproduced with permission from Ref. [65], Copyright 2007, IOP Publishing. (f) SEM images of nanostructured Au coating by alloy corrosion process. Reproduced with permission from Ref. [61], Copyright 2015, American Chemical Society. (g) AFM images of nanostructured Pt coating by electrochemical roughening. Reproduced with permission from Ref. [42], Copyright 2015, American Chemical Society. (h) SEM images of nanostructured Au coating by employing roughened polysilicon substrate. Reproduced with permission from Ref. [59], Copyright 2003, IOP Publishing. (i)SEM image of the Pt-elastomer mesocoposite coating. (j) Optical micrographs and (k) CV curve (PBS pH 7.4) of the Pt-elastomer mesocoposite coating (with stretchable gold film interconnect visible) at rest at 0% and at 45% uniaxial tensile strain. a.r. denotes aspect ratio. 0% rec denote upon returning the electrode to its original length. (i)(j)(k) Reproduced with permission from Ref. [54], Copyright 2015, The Authors.

increased, which resulted in a dramatic decrease in impedance (∼850 times) and excellent CIC (24 mC/cm2 ) [18]. To improve the compatibility between the neural electrode and the resident tissues, soft materials were incorporated into the Pt coating. A platinum-elastomer composite coating was deposited on the electrode sites by mixing and screen-printing, as shown in Fig. 3i. The combination of the electrochemical properties of platinum with the mechanical compliance of silicone enables a soft mechanical buffer with comparable CIL to that of Pt at the electrode-tissue interface. The efferent and afferent neural pathways were successfully activated in vivo over the entire range of possible muscle activation. Besides, this coating also displayed stable electrochemical and electromechanical responses [54], as shown in Fig. 3j & k. In addition to Pt, gold (Au) is also a promising material for neural electrode coating. ED [61,63–65,171], evaporation [59,62] and

sputtering [59] are commonly used method for Au coating formation. With the electrodepositing Au coating, the modified electrode displayed four-time lower impedance and ∼40% reduction of noise level [63]. To further improve the electrochemical property of the electrode, deposition associated with surface roughening techniques were used to enlarge the effective surface area of the coating. By employing nanoscale structures as templates (Fig. 3e), the impedance at the interface and the noise level were shown to be reduced owing to increased active surface area [65,83,171]. Higher porosity and pore interconnectivity of the coating (Fig. 3f) can be obtained by co-deposition of Ag-Au alloy and then chemical de-alloying of Ag [61,62,64]. The impedance value of the modified electrode was observed to decrease sharply by more than 25 times, which is comparable to Pt black coating by ultrasonic electroplating [62,64]. Besides, this coating exhibited comparable CSC and CIL to TiN and CNT [64]. The biocompatibility of the nanoporous

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Au-modified electrode was also confirmed [61,64]. Results even showed that the nanoporous Au coating selectively reduced astrocyte attachment by about 56% in an in vitro neuron−glia co-culture model [61]. With the nanoporous Au coating, the sensitivity of the electrode was increased owing to reduced noise and more active channels, which enabled sufficient temporal resolution for single spike analysis and resolved single spikes from the burst in vitro [62,64]. Low voltage stimulation was also achieved and more stimulus-induced spikes were observed in vitro for the nanoporous Au-modified electrode than the electrodeposited Au-modified one [64]. Moreover, the nanoporous Au coating displayed long-term electrochemical and mechanical stability [62,64]. In addition to templating and de-alloying, other techniques such as direct deposition of Au NPs (Fig. 3d) [60] and substrate roughening (Fig. 3h) [59] were adapted to improve the effective surface area of the coating. Layer-by-layer assembly coating of Au NP was shown to have superior performance on electrical conductivity and charge transfer than CNT coating. Its biocompatibility was evidenced by good adherence, viability and differentiation of cultured neurons as well [60]. Surface roughening of the electrode sites enabled a uniform and highly rough Au coating, which resulted in approximately fiftytime impedance reduction on coated electrodes than the bare ones [59].

Iridium oxide Although the electrochemical property of Pt and Au can be improved by enlarging the active surface area, their limited charge delivery capacity remained a great challenge especially when the area of the electrode sites decreases due to their faradaic nature. With weak or no faradaic reaction charge transduction, the performance of the double-layer capacitive coupling in charge delivery is often limited [75]. Thus, materials with higher electroactivity and ability to support reversible faradaic reactions are needed to promote charge transfer at the neural interface. Owing to its four oxidation states and ease of valence change between Ir3+ and Ir4+ states which is shown in Eq. (2.2.3) [66,68,70,71,73,75,76], iridium oxide (IrOx ) exhibits high electroactivity among noble metalcontaining coatings by employing reversible faradaic reaction. The good hydrophilicity and formation of hydrated oxide surface layers also contribute to an increased electrochemically active surface and thus a higher charge delivery capacity [70,172,173]. With the high charge delivery capacity, IrOx shows a significantly high CIL up to 4mC/cm2 [7,80,174], which is more than 25 folds larger than that of Pt [175]. The high CIL is a highly desirable behavior for safe electrical stimulation. Moreover, IrOx displays good biocompatibility and high corrosion resistance [66,68,70,73,75], which supported the absorption of polycations like polylysine and neural cell behaviors in both in vitro and in vivo study [66,74,76,172], as shown in Fig. 4f. With the good electrochemical property and biocompatibility, both successful stimulation of individual neurons and recording multiple single-unit spike activity with good signal-to-noise ratio (SNR) were achieved in vitro on IrOx −coated neural electrode [76,82]. Electrode with the IrOx coating was also shown to detect enhanced spontaneous activity in vitro than that with Pt coating due to lower thermal noise [77,82]. The fabrication of the IrOx coating is shown to play an important role in its structure and properties [66]. Common kinds of IrOx coating classified by the preparation methods includes activated iridium oxide film (AIROF) [30,44,67,68,75,78,79,177], sputtered iridium oxide film (SIROF) [70,71,76–78,82,176–180], electrodeposited iridium oxide film (EIROF) [66,73,78,83], atomic layer deposited iridium oxide film [74] and physical vapor deposited iridium oxide film [69]. A schematic of the fabrication process of the SIROF, AIROF, and EIROF coating and the morphologies of resulting coatings are shown in Fig. 4a & b, respectively.

Due to the difference in the fabrication, AIROF, SIROF and EIROF behave differently both in structure and properties [70,71,77,78,177,178]. EIROF and AIROF showed open and porous iridium oxide layer while SIROF exhibited a dendritic surface structure with much higher density (porosity/openness: EIROF > AIROF > SIROF) [66,70,71,77,78,178]. Owing to the open structure, EIROF exhibited higher CSC (∼88% and 70% higher, respectively) and lower impedance (15% and 7% lower, respectively) than SIROF and AIROF at similar thickness [78] and AIROF behaved slightly better than SIROF [71,78]. As for CIL, AIROF seemed to perform better (∼4 mC/cm2 ) [30] compared to SIROF (∼2–3.5 mC/cm2 ) [71,77] and EIROF (∼1.4–2.6 mC/cm2 ) [73]. All of them showed significant improvement on electrochemical property of bare electrode [44,66,71,73,75,77,78] and some of them performed better than other coatings like Pt, Au and TiN [30,71,76,77], which are summaried in Table 2. Regarding electrochemical and chemical stability, SIROF performed better than AIROF. SIROF was found to be more durable than AIROF under continuous high-charge-density stimulation [177] and maintained its electrical property in ambient condition for a month [82]. Although AIROF exhibits good electrochemical property, it was found that the activated state was not robust enough to last for a few days even when reactivation was employed [44]. The electrochemical property of the IrOx coating can be improved by optimization of fabrication parameters such as the type of plasma excitation, substrate heating, sputtering pressure and oxygen flow (Fig. 4c) [179,180] or employing electrochemical activation [70,79,179]. A trade-off between electroactivity and mechanical durability of the coating exists in the case of AIROF when a higher degree of activation leads to more porous structure and better electrochemical property but poorer mechanical adhesion [67,79]. In vivo activation of iridium is not preferred as the presence of iridium-containing deposits in the tissue near the charge-injection tips has been observed in the histology studies [79]. The mechanical stability make a difference on the long-term function of thick and porous coating which would have high CSC. The EIROF coating with not-so-reliable structure was only able to maintain mechanical and electrochemical stability with a low CSCC value of no more than 45 mC/cm2 [73] while AIROF and SIROF can achieve a stable coating with CSCC of 52 mC/cm2 [79] and 180 mC/cm2 [71], respectively. A recent study showed that CSC of more than 300 mC/cm2 could be achieved on template-induced porous AIROF without the impairment of mechanical stability, as shown in Fig. 4d & e [68]. The combination of IrOx with other metallic materials was also shown to either improve the property of the coating [83,181] or reduce the cost [81]. EIROF formed on the surface of Au nanowire coating significantly decreased the impedance of the electrode by almost three orders compared to plain Au electrode, which enabled the clear measurement of spontaneous neural spikes from the hippocampus in vivo [83]. The incorporation of ruthenium oxide (RuOx ) in the IrOx coating by thermal decomposition led to an enlarged apparent electrochemically active surface area which would facilitate charge transfer. The IrOx /RuOx mixture exhibited excellent mechanical and electrochemical stability verified by a prolonged extreme electrochemical cycling within a 5 V potential window [181]. Although IrOx coating is a good candidate for neural electrode coating, the cost of such a noble metallic coating is quite high. Thus, relatively cheap metallic materials such as titanium oxide (TiO2 ) were considered to incorporate into the pure IrOx coating and the IrOx /TiO2 mixture had been shown to preserve the good property of IrOx [81]. Titanium nitride Titanium nitride is also one of the metallic compounds which have been commonly used in medical device design such as

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Fig. 4. (a) Schematic of fabrication process of the SIROF, AIROF and EIROF coating. SEM image (top) and impedance spectroscopy (bottom) of SIROF, AIROF and EIROF coating (from left to right). (a)(b) Reproduced with permission from Ref. [78], Copyright 2014, IEEE. (c) Charge delivery capacity after 400 cycles of activation and film pumping speed of SIROF coating with respect to oxygen flow to the deposition chamber. Reproduced with permission from Ref. [176], Copyright 2009, IOP publishing. (d) SEM image of nanoporous AIROF coating by employing anodized aluminum oxide template. (e) CV scans in phosphate-buffered saline of flat (solid line) and nanoporous AIROFs by five seconds (dot line) and ten seconds (dashed line) Ir(acac)3 exposure. (d)(e) Reproduced with permission from Ref. [68], Copyright 2010, Elsevier. (f) Neuron survival on iridium oxohydroxide thin coatings after 4 days in vitro (DIV) in the different stages of development of hippocampal neurons in culture with respect to control (borosilicate-H glass). Reproduced with permission from Ref. [66], Copyright 2012, American Chemical Society. (g) SEM image of porous TiN coating. (h) Porous/smooth ratios of some electrochemical quantities of the TiN coating including electrode potential(Ep ), Estimated marginal means (EMM ± SE) of the CV and pulse derived capacitance (C0.05 , C1.0 and Cpulse ) and the impedance magnitude (|Z (1 kHz)| and |Z (100 mHz)|) as a function of time, including both in vitro and in vivo measurements. (g)(h) Reproduced with permission from Ref. [85], Copyright 2016, IOP publishing.

cardiac pacemakers and dental implants due to its good corrosion resistance [70,77,182,183]. Owing to its desirable capacitive charge-injection process and high effective surface area, it was found to be a promising material for the coating of neural electrodes [9,30,70,77,86]. With the columnar and porous TiN coating (Fig. 4g), the modified neural electrode exhibited high CIL up to 4.45 mC/cm2 , which is comparable to that of IrOx [9,87]. Thanks to the high CIC, the size of the electrode site could be reduced by more than 50% for effective in vivo stimulation within the safe potential window [9]. Moreover, TiN showed excellent electrochemical stability after prolonged pulsing [86,87]. However, the biocompatibility of TiN for neurons is still not clear [30,77] and more studies are needed. The electrochemical property of the porous TiN coating was shown to be markedly impaired by protein adhesion in vivo [85], as shown in Fig. 4h.

the effective surface area or improving the electroactive nature of the interface, these efforts to create an intimate and stable neural interface are unsatisfying due to the significant mismatch with the surrounding tissues in mechanical and biological properties [184,185]. Thus, carbon material is emerged as promising candidates. As a chemically stable element [105,186] and one of the fundamental compositions in the body, the toxicity of carbon is minimal compared to metallic materials which may release harmful metallic deposits in the body [44,79,187]. Furthermore, recently discovered forms of carbon such as CNT and graphene exhibit superior electrical and mechanical properties [188–192], which make them most promising carbon materials for microelectronic medical devices. CNTs have been proved to be a superior coating material at the neural interface by numerous studies while the study of graphene has emerged as a new trend in recent years [13,105], which will be discussed in the following sections.

Carbon materials Although metallic materials and their derivatives were shown to improve the performance of the neural electrodes by increasing

Carbon nanotube Thanks to its great electrical conductivity [188,189] and mechanical strength [190–192], CNT finds its role in neural elec-

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Fig. 5. (a) Idealized representation of defect-free (n,m) SWNTs with open ends. Reproduced with permission from Ref. [194], Copyright 2002, John Wiley and Sons. (b) CV of the CNT coated and bare electrodes. (The inset) the CV for Pt electrode at a different scale. Reproduced with permission from Ref. [92], Copyright 2015, IOP publishing. (c) Impedance spectra of CNT coating withstanding rat dura mater penetration. Coated electrodes (red), after a mechanical pretreatment (green), and after recording sessions in rat brain (blue) versus an uncoated one (black). Reproduced with permission from Ref. [100], Copyright 2011, American Chemical Society. (d) Electrochemical stability of CNT-, PEDOT-, and IrOx - coated electrode under repeated CV scanning cycles and SEM images of coatings after 300 CV cycles. Reproduced with permission from Ref. [95], Copyright 2009, American Chemical Society. (e) Picture of the whole embryonic mouse hindbrain-spinal cord preparation, layout of the 4 × 15 3D-nanostructured CNT-coated electrode, five minutes recording of the rhythmic activity detected in the rostral region (from left to right), local field potential recorded at the electrode marked with a triangle in the electrode layout (top), burst of spikes detected on the marked with a diamond in the layout (medium) and bursts of spikes recorded on electrode marked with a circle in B layout and triggered when applying a 1-ms biphasic (cathodic first) current of 10 ␮A on one microelectrode located in the upper medulla (electrode marked with a square in the layout) (bottom). Reproduced with permission from Ref. [106], Copyright 2015, The Authors.

trode design. Here, we focus on its application in electrode coating. The biggest advantage of CNTs as neural electrode coating is their high surface-to-volume ratio [108,189,193,194] (Fig. 5a), which provides an extremely large area for charge transfer and cell attachment, thus facilitating high sensitivity for recording and high efficiency for stimulating. Another critical advantage of CNTs is their capacitive interfacial behavior during charge transfer [25,195,196]. This type of charge-injection mechanism allows reversible and reliable charge exchange without degradation and significantly reduces the possible toxic products formed in the redox reaction. Also, the great electrical conductivity of CNTs enables a better electron transport from the interface to the electrode body. Besides, the high mechanical strength and flexibility of CNTs may enhance the adhesion between the coating layer and the electrode sites [95,110,111]. The biocompatibility and cytotoxicity of CNTs with neuronal cells has been widely studied both in vivo and in vitro [108,197–206]. Although these studies have not reached an agreement on the biocompatibility and cytotoxicity of CNTs, the majority of them showed promising results including negligible effect on cell electrophysiological function, formation of stable

cell network, boosting of neuronal electrical signaling and limited in vivo tissue response [90,198,202,206,207]. CNT coated electrode showed superior performance on characterization and in vivo work. Compared to the bare electrode, the impedance of CNT-coated electrode decreased significantly, ranging from 10 to 60 times (Fig. 5c) [92,95,100,106], which performed better than the titanium nitride (TiN) coating [106]. CNT coating made by layer-by-layer assembly (LBL) was found to be more effective for reducing the interfacial impedance value than PEDOT and IrOx by 30% and 60%, respectively [95]. In addition, the CSC was improved by applying CNT coating (∼3–140 times), as shown in Fig. 5b and Table 2 [92,100,105], which was more effective than PEDOT and IrOx with the same thickness [95]. Besides, the coating of CNT achieved more than 10 times CIL than bare platinum (Pt) electrode. Although its CIL is still not as high as that of iridium oxide, it is sufficient for safe in vitro stimulation [90]. The wider potential window of CNT coating than that of Pt and IrOx would reduce the possibility of harmful water electrolysis [90]. Furthermore, CNT coating exhibited good electrochemical stability with negligible loss of electrical performance after hundreds of pulse and CV scan-

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ning cycles while CP and IrOx coating showed different degrees of degradation (Fig. 5d) [95,100]. In additional to the electrochemical stability, CNT coating possesses the advantage of chemical and mechanical stability. Less than 2% variation of impedance magnitude at 1 kHz was shown after two-month storage for CNT coating while the impedance of poly(pyrrole)(PPy)-CNT and Au-CNT coatings increases 220% and 60% respectively [100]. The CNT coating was shown to be more robust than IrOx and PEDOT coating where cracking occurred after 300 CV scanning cycles [95]. The attractive properties of CNT coating as discussed above benefit the performance of neural electrode dramatically on both stimulating and recording. Compared to Pt electrode, only 14% of threshold voltage was needed to initiate the spike of the neurons in vivo by employing CNT coating [92]. At the same stimulation current, CNT-coated electrode was found to elicit stronger cell response than tungsten wire and Pt electrode in vitro [113]. For recording, the employment of CNT coating significantly reduced the noise in vitro [96,106] and increased the SNR by more than 6 times in vivo [92]. With the high sensitivity of CNT coating, significant neural activities with higher amplitudes were collected in vitro and in vivo [93,96]. Noticeably, low frequency signals like local field potential and small amplitude spiking signals which are less than 20 ␮V could be clearly detected in vitro [106], as shown in Fig. 5e. A long-term in vivo recording that lasted for a year proves the great potential of CNT coating for chronic implantable application [93]. The fabrication of CNTs deposition determines the properties of the coating. There are several methods which have been used to deposit CNTs on the electrode sites, including chemical vapor deposition (CVD) [45,89–94,96,100–102,106,108,110,111,113], ED [100,103,105], layer-by-layer assembly (LBL) [95,98,99], solution casting [97] and covalent attachment [105]. Among these methods, CVD is the most widely used technique. By employing metal-oxidesemiconductor as catalyst, CNTs can be grown either to a fluffy mat (Fig. 6a) or to a highly ordered and vertically aligned pillar bundle (Fig. 6b). The size, geometry, and location of the CNT pillars are precisely controlled by lithographic patterning of the catalyst [90]. The formation of highly porous structure creates large effective surface area for charge transfer, which significantly improved the electrochemical performance of the coating [100]. In addition to the good electrical property, CNTs coating made by CVD displays excellent mechanical durability and adhesion. Vertically aligned CNTs can hold the cell bodies with their ultrasmall contact areas (several square micrometers) and sustain high degree bending (16◦ ) exerted by the cell bodies without breakage or delamination from the electrode sites during in vitro studies, thus showing superior flexibility and strength [110,111]. After dozens of blending and later implantation into the animal’s body, the CNT bundles still remained firm on the electrode sites without distinct variations on the interface impedance [92]. The proximity between the electrode and the target is also improved owing to the CVD-formed 3D protruding structure, which enabled tighter integration with the neural tissue [90]. The protruding structure may be a promising design to improve the cell-electrode coupling through the activation of endocytotic-like mechanism [208].Compared to randomly oriented CNTs, vertically aligned CNTs were shown to display faster electron transfer kinetics as results of their similar structure to highly oriented pyrolytic graphite [91,209,210]. However, aligned CNTs bundle with extremely high density may lose their electrochemical superiority over random CNTs due to reduced porosity [90,100]. Vertically aligned CNTs were able to penetrate through the cell membrane while randomly oriented CNTs only contact the cell surface during in vitro studies [91,110]. This intimate contact may allow not only extracellular but also intracellular recording and stimulating. However, the biggest disadvantage of CVD is their high temperature manufacturing process, which constrains the choice of electrode and substrate materials. Higher temperature applied

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on CNT growth was shown to obtain longer and denser CNT with higher degree of graphitization and thus better electrical conductivity and higher SNR during in vitro and in vivo signal recording [92,94,101]. Some methods such as plasma-enhanced CVD [211] and secondary transfer [92,94] were tried to solve this problem and vertically aligned CNTs integrated on parylene substrate have been achieved [92,94]. Compared to CVD, the biggest advantage of ED is that the manufacturing process could be realized in ambient condition and mild solution. Other than simple manufacturing process, ED exhibits high flexibility with the controllable pattern and electrochemical properties of the coating by manipulating parameters such as CNT deposition charge and choice of additives [103,105]. Usually, ED results in coatings with rice-like morphology formed by bundle of nanotube, as shown in Fig. 6c [105]. Nevertheless, the coatings made by ED are always fragile, which limits their chronic performance [95]. Another simple method to form coating is the solution casting [97]. However, it faces the same problem of poor contact with the substrate as ED as it only relies on physical absorption to form the coating [95]. What’s more, it only enables planar morphology of the coating (Fig. 6d), which significantly limits its electrochemical property. In contrast, covalent attachment provides much more coherent and stable coating films by utilizing covalent bonds. Results showed that the covalently attached CNTs (Fig. 6e) remained intact on the electrode sites even though the parylene insulation had been peeled back during penetrating [105]. The coating made by covalent attachment was shown to perform better on interface impedance reduction than those made by ED [105], probably due to its rough and bubble-like morphology and increased adhesion with the electrode sites. LBL is a recently developed method to form multilayered coating with the help of electrostatic force. Polyelectrolytes are usually employed to form this multilayered structure with CNTs. This method enables highly precise control on coating thickness which is engineered at the nanoscale level [98,99,212]. The coating made by LBL displays exceptional uniformity with exclusive dispersion of individual CNTs rather than bundles (Fig. 6f) [98,212]. CNTs can be oriented to the same direction by air-water interface force [212]. The incorporation of polyelectrolyte introduces ionic conductivity to the coating, which aided the charge transfer at the neural interface [98]. Furthermore, LBL fabricated coating showed unique material properties such as high mechanical strength and good adhesion to the substrate due to molecular species acting at the nanometer length scale in the layered architecture [95]. Although CNTs show attractive performance as neural electrode coating, it still has great potential for improvement of properties due to the materials nature and fabrication. Some surface modification strategies were studied to boost the performance of the CNT-based coating. Amorphous carbon (Fig. 7a) is often formed on the surface of CNTs during CVD process, resulting in a barrier to the electrolyte solution [213] and the leakage of current during electrode working [101]. Methods such as plasma treatment [101] and heat treatment [213] were adopted to remove amorphous carbon. After treatment, structures with larger porosity were obtained with 2-fold increase in capacitance compared to the CNTs coating without treatment [213]. However, the removal of amorphous carbon introduced some defects in the structure and reduced the mechanical stability [101,213]. Another challenge associated with CNTs coating is its high hydrophobicity, which significantly lowers the accessibility of electrolyte and limits charge injection at the interface [90]. Thus, surface treatments like plasma treatment [45,91,101], ultraviolet (UV)-ozone treatment [104,109], aminofunctionalization [102] and binding of amphiphilic layer [90] were utilized to solve the problem. The formation of hydrophilic chemi-

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Fig. 6. SEM images of CNT coating by (a) chemical vapor deposition (random nanofibers), (b) chemical vapor deposition (vertically aligned nanofibers), (c) electrodeposition, (d) drop casting, (e) covalent attachment and (f) layer-by-layer assembly. (a) Reproduced with permission from Ref. [104], Copyright 2015, The Royal Society of Chemistry. (b) Reproduced with permission from Ref. [113], Copyright 2009, The Authors. (c)(e) Reproduced with permission from Ref. [105], Copyright 2008, Macmillan Publishers Ltd. (d) Reproduced, with permission, from Ref. [97], Copyright 2009, Elsevier. (f) Reproduced with permission from Ref. [95], Copyright 2009, American Chemical Society.

cal bonds such as C OH, C O and OH C O after plasma treatment and UV-ozone treatment dramatically decreased the water contact angle of the surface(∼145◦ ) [45,101,104,109]. With this greatly increased effective surface area, the interfacial capacitance of the CNTs coating was shown to increase up to 80 folds [101], leading to substantial decrease of the impedance (Fig. 7b & d) [45,91,101,104]. The changes of hydrophobicity also promoted the attachment and differentiation of the neural cells in vitro (Fig. 7e) [104,109]. Thanks to the improved contact and reaction between the cell and the electrode surface, high fidelity neural recording was facilitated in vitro [101,104]. Both the plasma-treated and UV-ozone- treated CNT coating exhibit good chemical stability and mechanical stability, with maintained hydrophilic and electroactive state for at least a month [45,109] and good adhesion with electrode sites after ultrasonication or surgery [109]. Although these two kind of treatments slightly damage the graphitization of CNT coating [45,109], they benefit the CNT coating more by enhancement of hydrophilicity. Similarly, the attachment of amino groups on the surface of the CNT coating by amino-functionalization increased its hydrophilicity and electrochemical properties (Fig. 7c). Besides, amino group was shown to promote the adhesion and growth of neuronal cells in vitro. The CNT-coated electrode with amino-functionalization exhibited 7 times SNR than Au electrode in vitro [102]. In addition to binding of functional groups, the incorporation of amphiphilic materials such as poly (ethylene glycol)-lipid conjugates in CNT coating is a promising method to improve surface hydrophilicity. Effective surface area and thus interfacial double-layer capacitance increase extremely after the surface treatment with the poly (ethylene glycol)-lipid conjugates layer, about 300 times than that one without treatment. A more than 50-fold improvement of the CIL was also demonstrated, which significantly benefited the employment of electrical stimulation [90]. Mechanical properties of the CNT coating such as adhesion [101,214] and strength [112] also can be improved by surface modification. With microwave treatment, almost 100% of CNT remained on the substrate after sonication while that without treatment was nearly completely removed (Fig. 7f & g) [214]. The incorporation of chitosan and DNA in solution-casting CNT film was shown to increase the mechanical strength of the film [112]. The employment of bacterial cellulose membrane as template for CNT coating

was found to change the brittle CNT film to flexible realizations [215]. To facilitate the interaction between neural cells and electrode surfaces, the biocompatibility of the CNT coating was adjusted by incorporation of highly biocompatible materials. A very thin collagen layer on the CNT coating was shown to notably increase PC12 cells adhesion and growth in vitro [110,111]. Biocompatible boron-doped diamond was used to encapsulate the CNT coating for promotion of favor of the cells and reduction of the potential release of CNTs into tissues, as shown in Fig. 7h [106]. Graphene Graphene (Fig. 8a) is a particularly attractive candidate for bioelectronic applications, due to its superior physical and chemical properties such as extremely high charge carrier mobility [216–218], good chemical stability [219], and high mechanical strength [220]. Besides, its thin-film form factors exhibit high transmittance [221,222]. In recent years, the study of graphene as neural electrode coatings has emerged as a new trend of research. Graphene coating of the neural electrode can be prepared by either pre-growing on a substrate by approaches like graphite exfoliation [114,116] and CVD [117] followed by transferring to the target electrode or growing directly on the electrode by CVD (Fig. 8b) [11,115]. Direct growth of graphene coating allows seamless contact with the substrate, but it may not be compatible to all electrode materials [11]. Reduced graphene oxide, another kind of graphene-based materials, was also utilized as the coating of the electrode due to its good solution-processability and ease of functionalization [118,223]. Electrophoresis was employed to simultaneously reduce graphene oxide and firmly attach it to the metal electrode by electrostatic interaction (Fig. 8c) [223]. Due to the superior electrical property of the graphene, the coating exhibited high conductivity of 2000 S/m and low impedance [11,115–118]. However, the charge storage capacity of graphene seemed not as high as CNT, probably due to its smoother morphology [116,121]. In vitro studies showed that neural cells on graphene-based coating exhibited healthy viability (more than 90%), welldeveloped network and unaltered neuronal signaling properties [11,115–117,121]. Neuronal density on the graphene was even higher than glass substrate during in vitro study (Fig. 8d) [116].

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Fig. 7. (a) High-resolution transmission electron microscope images of multi-walled CNT with amorphous carbon before and after O2 plasma treatment. The insets show the contact angle of these two samples. (b) CV curve of the electrode without CNTs, with as-grown CNTs, and with plasma-treated CNTs in PBS solution. The scan rate was 100 mVs−1 . (a)(b) Reproduced with permission from Ref. [101], Copyright 2010, Elsevier. (c) Process flow of multi-walled CNT amino-functionalization. Reproduced with permission from Ref. [102], Copyright 2011, Elsevier. (d) Interfacial impedance and capacitance of CNT-coated electrodes per unit area as a function of UV-ozone exposure time (top) and ZC and ZR of as-grown and 40-min UV-ozone-modified CNTs (bottom). (e) Fluorescent images of neuron cells cultured on control glass, as-grown CNTs, and UV-ozone-modified CNTs for 10 min and 40 min and SEM images of Hippocampal neuron cells cultured on the 40 min-UV-ozone-modified CNTs. (d)(e) Reproduced with permission from Ref. [109], Copyright 2010, John Wiley and Sons. (f) Schematic of how microwave treatment enhanced the mechanical adhesion of CNT coating. (g) The percentages of CNTs remaining after 5 min sonication versus microwave treatment time at various powers. (f)(g) Reproduced with permission from Ref. [214], Copyright 2010, Elsevier. (h) Fluorescent images of hippocampal cell culture on CNT coating with boron-doped diamond top layer at 8 DIV. Reproduced with permission from Ref. [106], Copyright 2015, The Authors.

Additionally, the membrane of the neuron was found to securely engulf the coating in vitro, which indicated an intimate and robust coating/cell interface and demonstrated the cell affinity of the graphene-based coating [115]. Furthermore, the biocompatibility of the coating was supported by reduced tissue response in vivo. Microglia and astrocytes showed scattered distribution around graphene-coated electrode while they formed a tight sheath around Pt electrode, which may indicate the antifouling property of graphene (Fig. 8e) [11]. Similar results were obtained in another study by in vivo impedance measurement and analysis based on equivalent circuit model [114]. The graphene-based coating allowed in vitro electrical stimulation with a high current density at same applied voltage and strong cell response [121]. Amphiphilic functionalization of the coating was found to boost the cell response further due to improved biocompatibility and charge injection ability (Fig. 8f) [121]. The good recording ability of graphene-coated electrode was demonstrated by both in vitro and in vivo study. With the graphene coating, around 38% more spikes were detected with one-fold increase in

SNR in vitro (Fig. 8g) [115]. Single-unit recording was realized in vivo with a high SNR (∼4–7) [11]. In addition to superior electrical and biocompatible property for neural stimulating and recording, the ultrathin form factors, close magnetic susceptibility to tissue and great transparency of graphene find a good opportunity for simultaneous imaging [11,224,225]. The graphene-coated electrode showed high compatibility to magnetic resonance imaging (MRI), where no image artifacts were found around the implanted electrode in the scanner (Fig. 8h) [11]. Simultaneous electrophysiological recording and optical imaging was also enabled in vitro and in vivo by utilizing graphene as the bulk materials of the electrode [224,225]. Conducting polymeric materials Polymeric materials, with low Young’s modulus and good biocompatibility, have been considered as coating materials for neural interface. As a coating material for device-tissue interfaces in bioelectronics, a polymer should be electroactive. CPs are suitable

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Fig. 8. (a) Schematic of graphene as the mother of all graphitic forms. Reproduced with permission from Ref. [216], Copyright 2007, Macmillan Publishers Ltd. (b) SEM image of graphene coating on the electrode. Reproduced with permission from Ref. [11], Copyright 2016, American Chemical Society. (c) SEM image of reduced graphene oxide coating. Reproduced with permission from Ref. [118], Copyright 2016, American Chemical Society. (d) Immunofluorescence staining of cultures developed on control and graphene substrates, marked for neurons (NeuN, green) and nuclei (DAPI, blue) and plot summaries of neuronal density in the two substrates. Reproduced with permission from Ref. [116], Copyright 2016, American Chemical Society. (e) Histological studies of tissue response to chronically implanted microelectrodes. Tissue was labeled for astrocytes (purple), microglia (red), neurons (green), and nuclei (blue). Scale bar, 300 ␮m. Reproduced with permission from Ref. [11], Copyright 2016, American Chemical Society. (f) Fluorescence imaging of the cells on mPEG–rGO (top) and rGO (bottom) films before (left) and after (right) electrical stimulation. Reproduced with permission from Ref. [121], Copyright 2014, Royal Society of Chemistry. (g) In vitro extracellular recording of action potentials from primary hippocampal neurons cultured on top of electrodes without (top) and with (bottom) graphene-based coating. Reproduced with permission from Ref. [115], Copyright 2014, Elsevier. (h) Schematic diagram and MRI image of implanted electrode in the brain. Reproduced with permission from Ref. [11], Copyright 2016, American Chemical Society.

candidate for coating layer in neural interface and they have been widely investigated. The discovery of polyacetylene by Alan J. Heeger, Alan G. MacDiarmid and Hideki Shirakawa broke the traditional view that polymers are insulating materials. The inherent molecular structure with the presence of conjugated double bond in polymer chain and the assistance of dopant as charge carriers allow charge to migrate both along the backbone and laterally between molecules in the polymer matrix, which provide CPs electrical conductivity [12,226]. The discovery of CP bagged the Nobel Prize in Chemistry in 2000. The intrinsically conductive properties of CPs make them a useful class of materials for a wide range of electronic and biomedical applications, such as supercapacitors [227], actuators [228], organic solar cell [229], biosensors [230] and tissue engineering [231–233]. CP is a faradaic material whose redox reaction happens with repeated intercalation and depletion of the charge balance ions [7,27,28,32,36]. The swelling nature of the polymer matrix in

an aqueous environment provides high ionic conductivity and a large contact area between the electroactive material and the electrolyte [138]. Additionally, CPs could be chemically modified and biologically functionalized with ease and high flexibility. As there have a few comprehensive reviews on the applications of CPs in the design and engineering of biointerfaces [12,13,126,234–239], we focus more on the improvement of CP coating with recent progresses. Common examples of CPs which have been applied on biomedical engineering include polythiophene (PT) [240], Polyaniline (PANI) [241] and poly (p-phenylene vinylene) (PPV) [242]. For neural interface, most of focuses have been placed initially on PPy and recently on PEDOT for their high conductivity and biocompability. Due to different functional groups and molecular structures (Fig. 9a) [238], PEDOT and PPy have a few difference in electrical properties, mechanical properties and biocompatibility. Compared to PPy, PEDOT has higher conductivity, electroactivity and elec-

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Fig. 9. (a) Chemical structures of PPy and PEDOT. Reproduced with permission from Ref. [238], Copyright 2014, IOP publishing. (b) CIL of PEDOT-coated electrode and bare Pt electrode. Reproduced with permission from Ref. [158], Copyright 2012, Elsevier. (c) The effect of charge density on the morphology, electrical and biological property of the coating. Reproduced with permission from Ref. [144], Copyright 2014, Elsevier. (d) SEM image of PEDOT nanotube. The insert is an image with higher magnification. (e) Percentage of high-quality recording sites units with SNR > 4 with respect to number of days after surgery. (d)(e) Reproduced with permission from Ref. [139], Copyright 2009, John Wiley and Sons. (f) In vivo polymerization of PEDOT cloud in the living rat hippocampus. (g) Confocal image of glial scar formation both on the insertion track and near the PEDOT deposition area (the dashed shape indicates the position of the original electrode). Green, red and blue staining indicated the existence of positive granule cells, astrocytes and neuronal nuclei, respectively. (f)(g) Reproduced with permission from Ref. [143], Copyright 2014, IOP Publishing. (h) Schematic of pretreatment of electrode site using EDOT-acid and images of PEDOT films on bare substrate and EDOT-acid modified substrate before and after ultrasonication adhesion test. Reproduced with permission from Ref. [249], Copyright 2015, American Chemical Society. (i) Procedure of multilayer poly-l-lysine and laminin attachment on PPy coating by employment of polyglutamic acid as the connector. Reproduced with permission from Ref. [133], Copyright 2006, Elsevier. (j) Images of cochlear neural explants grown on PPy/pTS coating with and without neurotrophin and electrical stimulation. Reproduced with permission from Ref. [252], Copyright 2010, Elsevier. (k) Schematic of the release of dexamethasone from the PEDOT nanotubes controlled by electrical stimulation. Reproduced with permission from Ref. [151], Copyright 2006, John Wiley and Sons.

trochemical stability owing to its dioxyethylene bridging group [155] and the lower band-gap that facilitates the charge transport and transfer in the polymer chain [157]. PEDOT is also more electrochemically stable than PPy (89% vs 5% of the original electrochemical activity after 16 h polarization) [243]. The good electrochemical stability of PEDOT comes from more capacitive charge transfer at the electrode/electrolyte interface, which is supported by a nearly featureless CV curve [128,130,141,146,154]. In addition, PEDOT shows better biocompatibility than PPy. Less protein absorption was observed on the PEDOT coating than on PPy coating in vivo [128]. The release of potentially cytotoxic dopant and degradation products was prevented in PEDOT film due to higher chemical stability in aqueous solutions [134,144,243] while PPy is prone to be reduced by biologically relevant agents in tissue [140,157,160]. However, PEDOT performs less satisfyingly on mechanical adhesion to the electrode site due to relatively brittle structure. Larger delamination area was observed on the edges of PEDOT film than on PPy film [122]. Another drawback of PEDOT is the low solubility of EDOT in water, which slows down the ED process [244]. As either PEDOT or PPy alone exhibits limitations as neural electrode coating, strategies such as functionalization and composition with other materials should be developed to improve their properties. PEDOT-PPy layered composite film has been considered to optimize the design but only a trade-off can be obtained [136].

Generally, PPy and PEDOT are directly polymerized on the surface of the electrode by ED. Advantages of the electrochemical synthesis of CPs include: (1) formation of coherent and stable coating based on chemical redox reactions; (2) nontoxic and simple deposition process without high temperature and subsequent purification steps; (3) controllable process at the coating sites, and tailorable structure and thickness; (4) ease of incorporating polyelectrolytes and biomolecules [126,153,234–236]. The coating layer on neural interface was proved to improve the properties of the electrode by a number of studies which will be discussed below. CP coated electrode showed superior electrical performance to their bare counterparts inclusive of low impedance, large CSC, good response to sub-millisecond current pulses and high CIL [137,138,146]. The improvement in electrical properties made the ultrasmall electrode recording (electrode site area <200 um2 ) realizable [10]. Compared to IrOx coated electrode, PEDOT coated one was more stable after 24 h voltage cycling with charge densities up to 3 mC/cm2 [146]. Through implantation, the ability of PEDOT coated electrode to record neural signals are demonstrated well supported by increased SNR than that achieved with bare electrodes in vivo [146,163,245–247]. The favorable reactive cellular responses of CP coating have been shown by both in vitro test and in vivo test [132,134]. PEDOT/PSS was found to support the neural network with down regulation of glial reaction [162]. A long-term animal study up to 7 months further demonstrated the biocompat-

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ibility and chronic stability of CP coating, where consistently larger signal amplitudes were obtained throughout the study without causing any toxic effects [163]. The selection of dopant was shown to have a great impact on the structural, chemical, and physical properties of electrodeposited CP film. The size of dopant was shown to be a dominant factor of the structure of CP coating during the formation of the electrodeposited film. Among the commonly used anions as dopant, the size increases in order [35,144,148] across sodium chloride (NaCl) [15,134,148], Lithium perchlo(LiClO4 ) [35,122,137,139,140,144,148,158,161,163,248], rate Sodium benzenesulfonate (BS) [35], Sodium p-toluenesulfonate (pTS) [15,35,123,128,144,158], Tetrabutylammonium perchlorate (TBAP) [10,155,159,249], Sodium dodecylbenzene sulfonate (DBS) [35,134], heparin [124,148], Poly(sodium 4-styrenesulfonate) (PSS) [15,35,127,131,132,138,142–144,146–148,153,156,158,164] and Hyaluronic acid(HA) [129]. CP polymerized with smaller dopants such as LiClO4 showed high structural regularity, more compact internal structure and thus higher conductivity and flexibility compared to those with larger dopants like PSS and heparin [35,148,158]. Employment of dopant with smaller size, higher deposition charge density and higher dopant concentration in the electropolymerization solution was shown to result in rougher surface morphology of the coating generally [15,35,144,148,250]. The uniformity of the coating on electrode site also varies with the size of dopant. PSS doped PEDOT displayed aggregation on the edge of the electrode site while ClO4 and pTS doped ones showed more even distribution [158]. The structure of the CP film including inner structure and surface morphology directly influences their electrical, mechanical and biological properties. Rougher surface morphology brings larger effective surface area for charge transfer at the electrode/electrolyte interface [35,144,158]. The mobility of the dopant mainly determined by their sizes also affects the interfacial charge transfer [35,144]. Small dopants move in and out easily when CP is electrically inactive while immobilized large dopants need to employ cations in the electrolyte as the charge transfer carriers. Thus, CP coatings doped with smaller anions usually exhibit better electroactivity shown by lower impedance, higher CSC and higher CIL [35,144,149,158,164,251], as shown in Fig. 9b. Although large dopants seem not as favorable as small dopants on electroactivity, they perform better regarding electrochemical stability with a limited loss of CSC after hundreds of CV cycles [144]. In addition to electrical properties, the selection of the dopants also influences the mechanical property of the CP coating. The elastic modulus of the CP coating was found to increase with the size of the dopant [35]. However, the modulus of CP coating, even that doped with large dopant PSS, was still two orders magnitude lower than that of traditional electrode materials Pt [35], which provides a mechanical buffer at the electrode-tissue interface. Smaller dopants were shown to have better performance on mechanical stability because of their compliance [126,149,158,164,251]. Compared to electrical and mechanical properties which have relatively clear relations with the dopant induced structure, the biological property of the CP coating appeared more complex, with a combined dependence on many factors such as surface morphology and material chemistry (Fig. 9c) [35,144]. More uniform surface morphology with patterns seemed to be more favourable to cells, which was supported by the excellent cell attachment and neurite outgrowth on the coating with this kind of surface like CP/ClO4 in vitro [35]. It is worth mentioning that the biocompatibility and nucleophilicity of the dopant needs to be considered especially for those small mobile anions [35,126,144]. Various strategies on structural design and surface functionalization were utilized to improve the performance of the conventional electrodeposited CP coating. An increased effective

surface area without change in the geometric surface area is desirable as it allows better electrical properties, more intimate interactions with cells and softer buffer layer [122,132]. Template technique is a useful way to create porous structure on CP coating during electroplating. The electrode with the 3D porous coating formed by polystyrene (PS) latex spheres showed significant impedance reduction compared to bare electrode over the whole range of frequencies (1Hz–100 kHz) and the maximum reduction at 1 kHz (∼80 times) was achieved when using PS sphere with a diameter of 300 nm [31,248]. Although the surface area could be increased by decreasing the particle size of PS sphere, rough structures with much smaller pore size (diameter ≤ 100 nm) did not show comparable electrical properties as that with medium size (diameter: 100–300 nm) due to their irregularity and poor connectivity [31]. Similar results were got in another study which obtained optimized structure with a nanofibrils diameter of 400 nm [140]. These indicate that an optimized feature size of nanostructure of the coating for electrical performance may be obtained in the medium size (a few hundred nanometers) instead of small size (below one hundred nanometers) [31,140]. However, further studies are needed to confirm it. Also, flexible electrospun nanofibers were employed as the template of CP coating (Fig. 9d) and the nanotube structure they induced not only extremely improved the electrical properties but also benefited the mechanical adhesion and cell attachment [24,122,139]. To the best of our knowledge, this nanotube structure performed best on the reduction of impedance for CP coating, roughly 200 fold in magnitude. Owing to CPs nanotube coating, the number of highly effective recording sites was doubled for both in vitro and in vivo recording tests (Fig. 9e). Neural cells were also utilized as template for CP coating in vitro. Despite the limited performance in impedance compared with electrospun nanofibers, the cell-shaped feature encouraged cells to re-populate, which would serve as a more intimate neural interface [142]. Further studies even directly polymerized CP in vivo with tissue employed as scaffolds (Fig. 9f) to bypass the glial scar near the electrode sites. However, secondary scarring was found around this conductive network, as shown in Fig. 9g [143,156]. Templating provides a simple and ordered method to generate controlled nanostructure on coating layer and improve electrochemical property, but it suffers from additional steps to remove templates and residual of harmful organic solvents. Thus, template-free techniques were utilized [10,147,148,152,253–257]. Spherically symmetric defect structure and improved electrical properties were attained with the addition of surfactant but the biological response was not favorable [10,152]. The PEDOT coating synthesized in ionic liquid (IL) shows rougher surface morphology, lower electrical impedance and higher SNR than the one synthesized in water. Biocompatibility of ILs was also demonstrated by high neuron viability and much reduced astrocyte fouling in vitro [147]. Nanostructured CP hydrogel showed good electrochemical property and mechanical softness, which allowed more clearly detectable units in vivo [253–257]. Recently developed techniques for patterning of CP micro- and nanostructures by methods like laser interference and photolithography show promise for the design of biomimetic topographical features targeting cell manipulation, but the aforementioned techniques have yet to be utilized for neural electrode coating [258–261]. Based on the improvement of electrical properties, more attention has been paid on promotion of mechanical durability, hydrophilicity and biocompatibility, where chemical strategies such as conjugation with other polymers and incorporation of functional groups [124,129,141,153,159] play an important role. The change of polymer properties by chemical synthesis exhibits superiority regarding good stability, and reproducibility. Due to the intrinsic chemistry and preparation process, delamination or cracking may occur with high possibility during mechanical

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deformation. PEDOT co-polymerized with EPh clearly changed the coating into more uniform and mechanically robust films with reduced extent of buckling without significant impairment of electrical property and biocompatibility [159]. A more recent study employed covalent bonded 2,3-dihydrothieno(3,4-b)(1,4)dioxine2-carboxylic acid (EDOT-acid) to promote the adhesion between PEDOT coating and electrode sites without compromise of electrical property (Fig. 9h) [249]. Wettability of the coating film can be tuned by functionalization with carboxylic acid group while electrical performances including both the CSC and the magnitude of impedance were still similar to those of pure PEDOT [141]. To improve the biocompatibility of the coating, biological polyelectrolytes like hyaluronic acid [129] and heparin [124] was co-synthesized with CPs monomer for better neuron attachment and less tissue response without severe impairment of electrical properties. Favourable reactive tissue responses through ∼60% attenuation of glial fibrillary acidic protein expression was exhibited during three-week in vivo test of poly(PyHA)-coated iridium microwires [129]. Moreover, strategies by introducing biological cues are needed to direct behavior of neurons and glial cells and serve as a mediator to promote the interaction with targeted neurons or reduce inflammatory response [262]. Although CPs have favorable reactive tissue responses through attenuation of glial responses and enhanced integration and signaling of neuronal processes, the overall tissue response after implantation of the CP coated electrode remains a serious problem. An acute inflammatory response and then a chronic wound-healing response occur upon several weeks after surgery. The nonconductive fibril encapsulation generated in the tissue response has a notably negative effect on the signal transduction at the electrode/tissue interface [139]. Different kinds of bioactive molecules can be easily incorporated in the CP coating during electrochemical synthesis of polymers. Depending on their acting mechanisms, commonly used bioactive molecules can be classified into three groups, ECM-based protein [130,133,154,161], neurotropic factor [123,131,135,154,160,252,263,264] and antiinflammatory drug [125,150,265]. ECM-based proteins like laminin increase cell attachment and adhesion through binding of cells on the surface of the electrode sites. Cell adhesion on electrode sites multiplied after incorporation of ECM- based proteins-coated CP. In a study of polypyrrole/CDPGYIGSR coated neural probe, 67% of the coated electrode sites showed positive immunostaining for neurofilaments, while only 6% positive immunostaining was shown for the uncoated gold electrode sites in vivo [130]. To increase the adhesion between the ECM-based protein and the CP, some functional groups like carboxylic acid [161] were utilized to covalently bind ECM-based proteins. Cells cultured on PEDOT-PEDOTacid-RGD film showed 3-fold cell density than on common RGD-entrapped PEDOT film and high elongation of neurites [161]. The enhanced adhesion and growth of neurons were attributed to the increased incorporation of RGD. In another study, two ECM-based proteins poly-l-lysine and laminin were attached to the coating with polyglutamic acid in sequence to form a multilayer structure (Fig. 9i) [133]. Neural cells had a more outgrowth on mixed multi-layered substrate than on the substrate with poly-l-lysine or laminin alone in vitro. In addition, the multilayer structure increased the elasticity and compliance of the coating, which benefited the interaction with neurons and attenuation of tissue response on electrode sites. Neurotropic factors also show great attraction to cells as they promote neuronal survival and growth towards the electrode sites [263]. The functions of different neurotrophins and their combination have been studied and compared [123,131,135,160,252,264]. For example, Adenosine 5 -triphosphate, an important trophic factor in synapses, facilitated double percentage of cell adhesion for CP coating in vitro [160]. Brain derived neurotrophic factor (BDNF)

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presented a more positive effect on promoting growth of neurites than NT-3. Summation effect happened on neurite growth during dual incorporation of BDNF and NT-3 in vitro (Fig. 9j) [252]. Unlike ECM-based proteins and neurotropic factors which exert attraction to neurons to form intimate electrode-tissue interface, anti-inflammatory drugs like Dexamethasone (DEX) promote the electrode-tissue interaction through attenuation of barrier induced by tissue response. DEX appeared effective inhibition of reactive astrocytes around neural electrode in vivo [265]. Highly controlled and efficient delivery of DEX was realized with electrical stimulation [125]. The rate of drug release was significantly slowed down by more than 50% through a CP nanotube structure (Fig. 9k), which provides possibility for a long-term treatment of inflammation [150]. With the bioactive molecules, the biocompatibility of neural electrodes was improved, but the electrical properties and mechanical adhesion were compromised. The effects on the electrical properties and mechanical adhesion depend on the type of bioactive molecules and their interaction with polymer chain. The impedance of PEDOT/DEDEDYFQRYLI/NGF film showed negligible difference with PEDOT/pTS film while the impedance of PEDOT-MeOH/CDPGYIGSR film was four folds greater than that of PEDOT-MeOH/PSS film [153,154]. Minor loss of charge transfer capacity was observed with the attachment of poly-l-lysine while more than 70% of charge transfer capacity was lost with dual incorporation of DEDEDYFQRYLI and NGF [133,154]. The interaction mechanism between bioactive molecules and CP also has an impact on the electrical property of neural electrode. For example, direct covalent attachment of biomolecules has a more negative influence on the electrical property compared to incorporation as a dopant [161]. The loss of electroactivity with the addition of biomolecules may be attributed to the incorporation of non-electroactive component, the inhibition on CP formation and covalent attachment of the binding agent [154,161]. Besides, the addition of biomolecules negatively affected the mechanical adherence. Loss of PEDOT coating dramatically increased from 0.5% to 10% with NGF, and further increased to 30% with peptide DEDEDYFQRYLI as the dopant [154]. Considerable types of biomolecules are associated with the central and peripheral nerve system. Choosing one for neural interface design is mainly based on their applications. Some specific types of peptide domain like RGD [161] and biomolecules with positive charge like poly-l-lysine [266] gained more favour from neural cells in vitro. In addition to selection of biomolecules, the incorporating method also plays an important role. The common practice to incorporate dual or multi biomolecules is realized by the same method. However, the incorporating efficiency might not be satisfactory as the different biomolecules might respond differently to the same attaching method. The neurite length per adhered cell on PEDOT/DEDEDYFQRYLI film was showed to be twice as great as that on PEDOT/pTS film while that on PEDOT/DEDEDYFQRYLI/NGF film was 50% as that on PEDOT/pTS/NGF film in vitro [154]. Hence, dual or multiple methods such as doping, covalent tethering and entrapment are suggested to be employed simultaneously for incorporating different biomolecules. Composite materials Although different materials have been studied as coating materials for neural electrode, the design of the coating layer is yet to be perfected. To date, there is no ideal candidate material which performs well on all properties such as electrical, mechanical and biological properties. Thus, composites are being recently considered to make best use of the advantages of several materials and also bypass their disadvantages. The composites reviewed here are those which include at least two different kinds of electroactive materials. Those composites which only include one kind of

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electroactive materials have been reviewed in the aforementioned sections based on the classification of electroactive materials. Two main categories are metal-carbon composite and carbon-polymer composite. Metal-carbon composite Metallic materials and their derivatives have excellent electrical properties while carbon materials like CNT and graphene exhibit high effective surface area and good mechanical property. The combination of these two kinds of materials as coating has the potential to notably improve the electroactivity and mechanical property of the neural electrodes. The Au-CNT composite was generated on the electrode sites by ED. The addition of CNT in gold coating was shown to act as an electroplating inhibitor, leading to continuous deposition of new crystallites and improved mechanical adhesion [103]. The formation of the Au-CNT coating with high porosity significantly increased the surface area of the coating (Fig. 10a) [46,103,105,167,267,268]. With the combination of electrical property of Au and CNT, the impedance of the electrode was decreased by more than 10 folds than bare electrode and 3 folds than Pt (Fig. 10b) [46,103,105,167,268]. The CSCc of the Au NP-CNT coated electrode was showed to be higher than that of AIROF [268]. However, the CIL of Au-CNT coated electrodes was less than 1mC/cm2 , probably due to limited faradaic charge transfer at the interface [100]. The mechanical durability of the coating was confirmed by capability of withstanding dura mater penetration with only 20% delamination after twice implantation and not significant change in impedance [100,267]. Neurons cultured on the coating for nearly three months in vitro still maintained their viability and activity [105]. The Au-CNT modified electrode showed good performance during recording. Fast single neuron spikes recording with larger amplitudes and shorter latency were achieved both in vitro and in vivo on the coating (Fig. 10c) [103,105]. The noise level during in vitro recording was shown to be 65% lower than that of gold-coated electrodes in average [105]. SNR was significantly improved, which allowed the detection of low-amplitude signals in vivo [46,267]. The coating was able to sustain ∼106 pulses [100] and mimetic physiological environment for at least 12 months [46], which indicates good electrochemical and chemical stability. The combination of IrOx and carbon were also considered to improve the property of the neural interface. The scaffold-type structure was utilized to display the respective advantages of two components [269,270]. As highly conductive and mechanically robust material, carbons form an interconnected and stiff scaffold for the composite, which would facilitate the charge transport within the coating and improve the mechanical durability. Furthermore, the large surface areas of some carbon materials like CNT and graphene would significantly enlarge the porosity and roughness of the coating and allowing the exposure of more active centers at the electrode-tissue interface [269–271]. Compared to carbon, IrOx has higher electroactivity owing to the faradaic charge transfer and hence it serves as an encapsulating layer outside the carbon scaffold. Not only would more charge exchange be promoted at the interface, but also the release of carbon would be prevented [270]. Different carbon materials such as CNT [270], graphene oxide [271] and graphene [269] have been studied to generate composite coating together with IrOx for neural electrode. A tridimensional IrOx -CNT composite was deposited on the electrode sites by ED (Fig. 10d). IrOx was attached to CNT by the carboxylate groups present in the nanotubes instead of simple mixing process, which would notably improve the connection between two components. The homogeneously dispersed CNT bundles were formed surrounded by IrOx NPs. Due to the combination of the high electroactivity of IrOx and the large effective surface area of CNT, IrOx -CNT composite coating exhibited improved CSC (Fig. 10e). An

excellent CSCc of 101.2 mC/cm2 was shown at 10 mV/s, which was three folds and eighteen folds greater than that of IrOx and Pt, respectively. No evidence of polarization was observed in the CV curve and less than 30% decrease of capacity was shown after ∼103 cycles, which indicates a good electrochemical stability. Additionally, the mechanical stability of the coating was confirmed by good maintaining of surface roughness after cleaning and in vitro cell culturing. The same cell viability and expression of postsynaptic receptors as those of IrOx and borosilicate in vitro serve as evidence for good biocompatibility of the IrOx -CNT hybrids (Fig. 10f) [270]. Graphene oxide [271] and graphene [269] were combined with IrOx to form a composite coating for neural electrode as well. Graphene oxide was connected with IrOx by chemical covalent interaction while graphene flakes was electrostatically attracted by IrOx . Despite the difference in connecting method between IrOx -graphene oxide and IrOx -graphene, the resulting multilayered structures with IrOx exposed at the surface were similar after deposition of the coatings (Fig. 10g). Among the three IrOx −carbon composite, IrOx -graphene oxide showed slightly larger CSCc than IrOx -graphene and IrOx -CNT [269,271]. Although IrOx -graphene does not have a surface morphology as rough as IrOx -graphene oxide which may reduce its CSCc, it has a CSCc per real surface ares 8 times larger than IrOx -graphene oxide owing to its better electrical and electronic properties than that of graphene oxide [269]. Both of IrOx -graphene oxide and IrOx -graphene were shown to maintain their reversibility for at least 103 cycles (Fig. 10h) [269,271]. IrOx -graphene oxide even showed more than 10% higher CSC retained than bare IrOx and IrOx -CNT after 1000 cycles [271]. However, cracking of the coating was observed when high vacuum conditions were applied. Comparable neuronal cell proliferation and the release of neurotransmitters were obtained in vitro on IrOx graphene oxide and IrOx -graphene to those on IrOx -CNT, which indicates their good biocompatibility (Fig. 10i) [269,271]. With the durable IrOx -graphene coating, in vitro electrical stimulation by neural electrode was shown to be safer with more than 10 times direct current field application than bare IrOx [269].

Carbon-polymer composite CPs are highly electroacitve and soft while carbons are mechanically robust and highly conductive. The combination of the two materials may provide a highly electroactive, mechanically strong and biocompatible coating. By electrochemical co-deposition, nanostructured carbon materials (NC) like CNTs and graphene oxide were incorporated into CP coatings as sole dopants [47,272–277] or co-dopants with other counter ions like PSS [127,278–280] or passive fillers in the polymeric matrix [281]. The morphologies of co-deposited PEDOT-CNT and PEDOT-graphene oxide coating are shown in Fig. 11a & i, respectively. Due to the good electrical property of the two components, the good electroactivity of CP and the high effective surface area of the NC, the CP-NC coating showed dramatically improved electrical property, with two orders of magnitude more CSC [43,280] and one-order of impedance reduction than that of bare electrode [43,47,272,273,275–279]. The CNT ratio in the composite and the way it is attached with CP influenced the charge transfer at the interface, which further affected the electroactivity of the composite [275]. Regarding electrochemical property, 6% reduction in impedance was observed for CNT doped CP compared with PSS doped CP, which indicated that CNT could be a better candidate as dopant for CP [43,275]. In addition, the incorporation of nanostructured carbon was shown to improve the electrochemical stability by more than 50% as they act as highly conductive pathways for efficient charge transport and reduced polarization (Fig. 11b) [43,127,275,276]. Besides, the poor mechanical durability of CP coating was well addressed by the addition of NC as reinforced

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Fig. 10. (a) SEM image of CNT-Au NP composite coating. Reproduced, with permission, from Ref. [267], Copyright 2015, IEEE. (b) Impedance spectra of CNT-Au NP composite coated electrode in comparison with Au coated electrode. Reproduced with permission from Ref. [167], Copyright 2016, IEEE. (c) Data recorded from a bare tungsten (red trace) and CNT/gold-coated (black trace) stereotrode tip over 150 ms (left top), power spectra calculated from 60 s of neural activity (left bottom) and spectrograms (1–2000 Hz) of bare (right top) and CNT-coated stereotrodes over 4 s (right bottom). Reproduced with permission from Ref. [105], Copyright 2008, Macmillan Publishers Ltd. (d) SEM images of IrOx -CNT coatings. (e) CV of bare platinum, IrOx and IrOx –CNT coated electrode in PBS (pH = 7.4) electrolyte at sweep rate of 20 mV s−1 . (f) Fluorescent microphotographs of cortical neurons on IrOx (left top) and IrOx –CNT(left bottom) after 5 DIV culture and the neuronal counting(right). (d)(e)(f) Reproduced with permission from Ref. [270], Copyright 2014, Elsevier. (g) SEM images of the IrOx -graphene hybrid coatings in PBS (pH = 7.4) at 10 mV s−1 . (h) Stability test of CSCc values of IrOx -graphene hybrid and bare IrOx coatings during 1000 consecutive cycles. (i) GABA and glutamate release of neurons grown on Borosilicate, IrOx and IrOx -graphene hybrid. Cells were exposed to a basal HBSS solution (5 mM K+ ) and to a depolarizing condition (90 mM K+ ) to induce the release of neurotransmitters. **P b 0.01, ***P b 0.001 against 5 mM K+ condition on material after One-way ANOVA statistical analysis. (g)(h)(i) Reproduced with permission from Ref. [269], Copyright 2015, Elsevier.

material, by which crack, delamination, and noticeable change in either appearance or impedance in CP-NC coating were eliminated for both in vitro and in vivo test [43,47,272,273,275,278]. Moreover, the interdigitated network of the CP-NC coating and reduced anion movement by employing immobile NC as part of change balancing dopants markedly reduced the volumetric change of CP during redox reaction, leading to a more intimate adhesion to the electrode sites [43,47,274]. Furthermore, the incorporation of reduced graphene oxide was shown to improve the hydrophilicity of the coating, enabling a more intimate contact with cells and ECF in vitro [276]. Another kind of CP-NC composite was achieved by first deposition of carbon nanostructures using chemical vapor deposition [110,111,113], electrophoresis [282,283] or drop casting [127] followed by ED of CP. Compared to co-deposited one, layered CP-NC composite enable more controllable deposition of carbon scaffold rather than random dispersion [283]. The morphologies of layered PPy-CNT coating by electrophoresis and chemical vapor deposition are shown in Fig. 11e and g, respectively. The introduction of CP on the surface of carbon nanostructure induces an improvement on the electrical property through faradaic charge transfer

and mechanical property through increased mechanical strength. The resulting 3D free-standing core-shell structure of the coating provided a more electroactive and fuzzy surface for charge transfer [110,111,113,282]. More than 20 times reduction in impedance and nearly two times increase in charge transfer storage [111,282,283] were obtained for vertically aligned carbon nanofibers (CNFs) by CVD-growth with CP shell than as-grown CNFs (Fig. 11f). The deposition of CP on the surface of carbon nanostructure was also shown to induce enhanced electrochemical stability due to effective inhibition of the migration of carbon [282]. The electrochemical stability of drop-casting CP-CNT coating was shown to be slightly lower compared to the co-deposited one [127]. However, it performed better in conductivity owing to less interference on polymer formation during ED [127]. The improved mechanical stability of the coating after the deposition of CP was confirmed by the good maintaining of free-standing structure in vitro while collapse of CNF usually happened on as-grown CNF coating, as shown in Fig. 11g [111]. The better mechanical contact with cells with minimized local mechanical stress and penetration into cell membranes was observed in vitro, which indicates a more intimate neural interface [110]. Among the depositing method of carbon, electrophoresis was

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Fig. 11. (a) SEM image of co-deposited PEDOT/CNT-(COOH)x coating. Reproduced with permission from Ref. [275], Copyright 2016, IEEE. (b) Plots of capacitance shift against charge–discharge cycle number. Potential scan rate: 100 mV/s for PANI-CNT, PANI-CNTb and PPy-CNT, and 200 mV/s for PEDOT-CNT. Deposition charge: 60, 60, 50 and 6.5 mC, respectively. Reproduced with permission from Ref. [274], Copyright 2007, Elsevier. (c) Fluorescent (top) and SEM (bottom) images of neurons cultured on PEDOT/CNT surfaces. Reproduced with permission from Ref. [47], Copyright 2011, Elsevier. (d) Chronic spike recording performance of PEDOT/CNT-(COOH)x (black) and PEDOT/PSS (blue) electrode sites. Reproduced with permission from Ref. [275], Copyright 2016, IEEE. (e) SEM image of layered PPy/CNT coating. Reproduced with permission from Ref. [283], Copyright 2015, Elsevier. (f) CV of an as-grown CNF coating (top), and two CNF coatings with 24 nm (dashed line) and 127 nm (dotted line) PPy covering layer (bottom), respectively, in comparison with the original as-grown sample (solid line). The scan rate was 100 mVs−1 and the measurements were carried out in 1.0 M KCl. Reproduced with permission from Ref. [111], Copyright 2006, John Wiley and Sons. (g) SEM images of the PC12 neural network interpenetrating with freestanding VACNFs with a PPy coating. Reproduced with permission from Ref. [110], Copyright 2007, IEEE. (h) Viability of PC12 cells cultured on different surfaces. Data are shown in terms of mean ± SD (***p < 0.01 versus cells cultured in complete media). Reproduced with permission from Ref. [282], Copyright 2016, Elsevier. (i) SEM image of PEDOT/N-doped reduced graphene oxide coating. Reproduced with permission from Ref. [276], Copyright 2016, The Royal Society of Chemistry. (j) Neuron growth on PEDOT surfaces doped with GO and PSS at 3 d. Reproduced with permission from Ref. [277], Copyright 2013, The Royal Society of Chemistry. (k) Fluorescence microscopy images of tracks for non-encapsulated (left panels) and pHEMA-encapsulated (right panels) PEDOT-PSS-CNT-coated electrode 4 weeks after implant. Reproduced with permission from Ref. [284], Copyright 2016, The Authors. (l) The comparison of mean maximal amplitude between the polypeptide coated and bare PEDOT/PSS coating. Error bars indicate the standard error of the mean; * indicate two sample t-tests, unequal variance statistical significance, p < 0.05. Reproduced with permission from Ref. [285], Copyright 2015, The Royal Society of Chemistry.

shown to allow an enhanced adherent coating to electrode sites than drop casting due to the bonding of surface hydroxides [283]. As carbon nanomaterials are embedded or encapsulated in the CP matrix for both carbon-polymer composites, their exposure to the tissue is significantly reduced, which would improve the in vitro biocompatibility of the coating (Fig. 11c, h & j) [47,127,282]. Additionally, the slight positive charges of the CP may benefit the contact between tissue and electrode [113]. Better neural cell affinity and differentiation were showed in vitro on CP-NC coated electrode when compared to bare one [43,47,127,275–277,282,283]. Neurons cultured on CP-NC microgroove coating displayed directional growth along the features [282]. Successful in vitro stimulation of neural cells including hippocampous pathway and retinal interneurons and strong cell response with short latency were achieved at low voltage amplitude for CP-NC coating because of its excellent electrical, mechanical and biological properties [47,113,281]. Also, the

recording performance of the electrode was dramatically enhanced where more than two fold reduction of SNR and clear detection of single unit with long-term stability were obtained in vivo (Fig. 11d) [275,279,280]. To further improve the biocompatibility of the coating, soft hydrogels, biomolecules and drugs were incorporated into the coating. A buffer hydrogel layer on the surface of the coating was shown to markedly reduce the mechanical difference between the tissue and the electrode sites and avoid the direct contact with the nanocomposite without noticeable impairment of electrochemical property [278,284]. The reduction of the in vivo tissue response at the resulting interface is shown in Fig. 11k. Based on the ease of surface functionalization of CP, biomolecules like peptides were bound to the CP-NC coating [285]. Although slightly negative effect of the peptides existed on the electrochemical performance of the electrode [277,285], the improved biocompatibility was shown to benefit in vivo recording supported by more active channels and

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higher signal power than that without peptide bonding, as shown in Fig. 11l [285]. Anti-inflammatory drug dexamethasone was also incorporated in the CP-CNT coating for inflammation alleviation [272,273]. Benefiting from the porous structure and deformation of CP during redox reaction, dexamethasone exhibited a sustained and controllable release, which reduced tissue response and cell death in vivo [273].

Challenges and opportunities The electrical, mechanical and biological properties of the neural electrodes have been significantly improved by the employment of a coating layer on the electrode sites, but the design of the coating layer is yet to be perfected. The nanostructured manipulation of the coating plays an important role on the properties and functioning of the neural interface. Although the highly porous structures like hollow tubes and sponge-like structures [64,122,150] have been employed to form the coating layer, the overall performance is limited by the high hydrophobicity and poor biological affinity of most coating materials. The resulting poor contact limits the signal transduction at the interface. Firstly, to address the problem of hydrophobicity, some surface treatments [45,91,102] have been used but it may not work for inner surface. Functionalization such as chemical modification should be considered to improve the hydrophilicity of bulk materials [141]. Secondly, the limited biological affinity of the coating causes the poor contact between the electrode site and target neural cells. Studies showed that cells freely stood on the coating surface without further penetration or neurite ingrowth [110]. To address this problem, the manipulation of the coating structure is required to focus more on mimicking of native cell living environment in the nervous system. Specific topological cues like aligned nanostructures inspired from nerve tissue engineering [286,287] may be promising structures for more intimate neuron contact and neurite guidance. Additionally, the surface functionalization of biomolecules and controlled delivery of growth factor would contribute to the good guidance and contact [262]. More importantly, great challenges still exist regarding the long-term performance of the coating in vivo. Firstly, the degradation of the coating in vivo limits the long-term performance of the coating. The electroactivity of the materials will decrease after repeated stimulation due to the degradation of the materials [60,136,144,154,158]. It is also possible that the coating reacts with some reductive agents in vivo and loses its electroactivity [140,157,160]. Moreover, mechanical delamination may happen with high possibility due to the volumetric change of the materials and friction with tissue and this problem is especially serious for polymeric materials which have markedly distinct intrinsic properties from common electrode materials like metal [53,58,69,122,154,158,213]. The poor adhesion between the coating and electrode sites may result from lack of strong binding force, which significantly limits the chronic performance of the neural interface. The enhanced mechanical binding forces by preroughening of the electrode sites have been studied and proved to provide a strong interlocking effect [36,58,158]. The introduction of chemical bonding between the coating and the electrode site should also be a promising way to improve the adhesion. Secondly, the induced tissue response at the neural interface remains a serious problem. Although good attachment and viability of neural cells was observed in vitro, different degree of tissue responses occur in vivo [129,130,143,284], probably due to the large differences between the simple in vitro environment and complicated in vivo environment. The acute protein fouling and chronic scar formation set a big nonconductive barrier between the electrode and resident tissue, which dramatically impair the function of

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the electrode. Besides, this barrier increases the distance between the electrode and neural cells and thus attenuates the signal transmission and increase signal loss along the pathway. The tissue response notably impairs the improved electrical properties of the interface by the coating in vivo. The impedance of the electrode was increased by nearly 45 fold after 49-day implantation [139]. The SNR of in vitro electrophysiological tests were shown to be considerably (∼100) [94,102] while those of in vivo test were ∼4 [128,147,155]. The noise levels of coated electrode sites were also shown to have no significant difference with those of bare electrode sites [50,55,149]. Due to the limited biocompatibility of the coating, most of coating materials are still under research and only Pt coating applied on the Utah electrode has been approved by food and drug administration (FDA) [288]. Thus, there is still a long way to go before more coating materials can be approved by FDA for clinical applications. Considering the biocompatibility, naturally occurring conductive materials like melanin and biomimetic synthetic polymers may be promising candidates. Although naturally occurring CPs have been studied for tissue engineering [289,290], the research studies on their potential applications as the neural sensing materials is limited. A recent study has successfully deposited polydopamine melanin on indium tin oxide substrate but the performance of the coating is limited to a CIL of 0.1 mC/cm2 [291]. The fabrication process of melanin coating may need to be further optimized and the mechanical and biological performances need to be evaluated. The low conductivity and high biodegradability of naturally occurring CPs may need to be addressed for chronic implantation. A synthetic cell membrane-mimicking CP has been shown to recognize target cell and significantly enhance neurite outgrowth [292]. Another recent biomimetic idea for neural electrode coating would be the combination of tissue engineering and neural interface. A local self-inspired tissue-electrode connection formed by body’s own healing process would be highly promising for stable and biocompatible neural interface. Neural cell embedded conductive layered hydrogel, an innovative concept, has been recently studied for living neural interfaces [293]. Results showed that it had good electrical performance but cell manipulation may need to be further improved [294,295]. Although considerable researches have been devoted to the fabrication of nanostructured coating for neural interface, the reliability-yield is yet to be perfected. Future research efforts should target a more cell-favorable interface. Combination of different materials, manipulation of the functional nanostructure and employment of biological cues should be considered to provide a desirable abiotic/biotic plane for neural signal transduction. A deeper understanding of tissue-electrode interaction in vivo would be beneficial for a more robust and efficient coating layer. Acknowledgements This work was supported by the National Research Foundation [No. NRF-CRP10-2012-01]. The authors would like to thank Annkathrin Swanson for her assistance on the figure drawing and thank Baiwen Luo, Agata Blasiak, Chaolemeng Bao, Tristan Lee and Jing Wang for their useful discussion. References [1] M.C. Kiernan, S. Vucic, B.C. Cheah, M.R. Turner, A. Eisen, O. Hardiman, J.R. Burrell, M.C. Zoing, Lancet 377 (2011) 942–955. [2] J. Simon-Sanchez, C. Schulte, J.M. Bras, M. Sharma, J.R. Gibbs, D. Berg, C. Paisan-Ruiz, P. Lichtner, S.W. Scholz, D.G. Hernandez, R. Kruger, M. Federoff, C. Klein, A. Goate, J. Perlmutter, M. Bonin, M.A. Nalls, T. Illig, C. Gieger, H. Houlden, M. Steffens, M.S. Okun, B.A. Racette, M.R. Cookson, K.D. Foote, H.H. Fernandez, B.J. Traynor, S. Schreiber, S. Arepalli, R. Zonozi, K. Gwinn, M. van der Brug, G. Lopez, S.J. Chanock, A. Schatzkin, Y. Park, A. Hollenbeck, J.J. Gao, X.M. Huang, N.W. Wood, D. Lorenz, G. Deuschl, H.L. Chen, O. Riess, J.A. Hardy, A.B. Singleton, T. Gasser, Nat. Genet. 41 (2009) 1308–1312.

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[288] E. Katz, Implantable Bioelectronics, John Wiley & Sons, Hoboken, 2014. [289] D. Kai, M.P. Prabhakaran, G.R. Jin, S. Ramakrishna, J. Mater. Chem. B 1 (2013) 2305–2314. [290] C.J. Bettinger, P.P. Bruggeman, A. Misra, J.T. Borenstein, R. Langer, Biomaterials 30 (2009) 3050–3057. [291] I.S. Kwon, Y.J. Kim, L. Klosterman, M. Forssell, G.K. Fedder, C.J. Bettinger, J. Mater. Chem. B 4 (2016) 3031–3036. [292] B. Zhu, S.C. Luo, H.C. Zhao, H.A. Lin, J. Sekine, A. Nakao, C. Chen, Y. Yamashita, H.H. Yu, Nat. Commun. 5 (2014) 9. [293] R.A. Green, K.S. Lim, W.C. Henderson, R.T. Hassarati, P.J. Martens, N.H. Lovell, L.A. Poole-Warren, 35th Annual International Conference of the Ieee Engineering in Medicine and Biology Society, IEEE (2013) 6957–6960. [294] U.A. Aregueta-Robles, A.J. Woolley, L.A. Poole-Warren, N.H. Lovell, R.A. Green, Front. Neuroeng. 8 (2015) 57. [295] U.A. Aregueta-Robles, K.S. Lim, P.J. Martens, N.H. Lovell, L.A. Poole-Warren, R. Green, 37th Annual International Conference of the Ieee Engineering in Medicine and Biology Society, IEEE (2015) 2600–2603. Nuan Chen received her B.E. degree from Sun Yat-sen University in 2015. She is currently a PhD candidate under the supervision of Prof. Seeram Ramakrishna and Prof. Nitish V. Thakor at the Department of Mechanical Engineering and Singapore Institute for Neurotechnology (SINAPSE), National University of Singapore. Her research involves micro- and nanostructured materials and their applications in neural interface.

Dr. Lingling Tian was a research fellow working on the project ‘Electrospun Biomimetic Nanofibres as Regenerative Medicines − Application for Nerve Tissue Engineering’ at the Center for Nanofibers & Nanotechnolgy in NUS (NUSCNN) before she joined Singapore Institute for Neurotechnology (SINAPSE) in February of 2016. She has already co-authored over 20 journal papers/book chapters/book with a citation of 276 and a h-index of 9 as of March 2017 (Google Scholar). Her research interests include fabrication of electrospun nanofibers, design of neural electrodes and the applications in sensing neural signals and tissue regeneration. Anoop C. Patil received his M.Tech. (VLSI Design) and B.E. (ECE) degrees from National Institute of Technology, Karnataka and Visvesvaraya Technological University, Karnataka (India) respectively. He has pursued his doctoral research under Prof. Nitish Thakor, at the Department of ECE, NUS, Singapore. He submitted his thesis (January 2017) and is currently with SINAPSE, NUS driving silk-based technology for implantable devices. During his masters, Anoop researched on Slow Light at RINCE, DCU, Ireland. He was a Digital Design Engineer at Cypress Semiconductors, Bangalore (2011–2012). His current research interests include development of soft and conformal electrical sensor implants including neural electrodes, electromagnetic sensors and flexible circuits for biomedical applications.

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Dr. Shengjie Peng received his PhD degree in Nankai University (P.R. China) in 2010. He is now working as a senior research fellow in Prof. Seeram’s group in National University of Singapore. He has been working on functional nanomaterials in energy and environment for more than ten years. He has co-authored over 70 peer-reviewed publications. His current research interests include the design and development of nanomaterials and their applications in energy.

Dr. In Hong Yang is a Head of the Neuro-Chip group and principle Investigator in the Singapore Institute for Neurotechnology (SINAPSE) at the National University of Singapore. He is also a faculty of Biomedical Engineering at Johns Hopkins University. He received a B.S. and M.S. degrees from Kyoung Hee University and Ph.D. degree from Texas A&M University. Dr. Yang’s technical expertise is in the areas of bio-electronic medicine (electroceuticals), neural microsystems and drug development. Dr. Yang was a recipient of IDEA and CONCEPT award from Department of Defense in USA. His research focuses on the development of innovative tools and therapeutics for nerve regenerations. Professor Nitish V. Thakor is the Director the Singapore Institute for Neurotechnology (SINAPSE) at the National University of Singapore. Dr. Thakor’s technical expertise is in the field of Neuroengineering, where he has pioneered many technologies for brain monitoring to prosthetic arms and neuroprosthesis. He is the author of 340 refereed journal, more than a dozen patents, and co-founder of 3 companies. Dr. Thakor is a recipient of awards from NIH, NSF and IEEE. He is the Fellow of IEEE, AIMBE, IFMBE, and BMES.

Professor Seeram Ramakrishna, FREng, FBSE is the Director of Center for Nanofibers & Nanotechnology at the National University of Singapore (NUS). He is a Highly Cited Researcher in Materials Science. Thomson Reuters lists him among the World’s Most Influential Scientific Minds. He co-authored ∼1000 articles with ∼62,000 citations and ∼118 H-index. He is a Fellow of UK Royal Academy of Engineering (FREng); International Union of Biomaterials Science and Engineering (FBSE); and American Institute for Medical & Biological Engineering (AIMBE). He authored a book Medical Devices: Standards, Regulations and Practices. He is an Editor of Current Opinion in Biomedical Engineering.

Please cite this article in press as: N. Chen, et al., Neural interfaces engineered via micro- and nanostructured coatings, Nano Today (2017), http://dx.doi.org/10.1016/j.nantod.2017.04.007