Microfabricated nerve-electrode interfaces in neural prosthetics and neural engineering.

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Microfabricated nerve–electrode interfaces in neural prosthetics and neural engineering Yong-Ak Song

a b f

, Ahmed M.S. Ibrahim

Jongyoon Han & Samuel J. Lin

c d

e

, Amr N. Rabie ,

c d

a

Division of Engineering , New York University Abu Dhabi (NYUAD) , Abu Dhabi , UAE b

Department of Chemical and Biomolecular Engineering , Polytechnic Institute of New York University , Brooklyn , NY , USA c

Divisions of Plastic Surgery and Otolaryngology , Beth Israel Deaconess Medical Centre , Boston , MA , USA d

Harvard Medical School , Boston , MA , USA

e

Department of Otolaryngology , Ain Shams University , Cairo , Egypt f

Department of Electrical Engineering and Computer Sciences and Department of Biological Engineering , MIT , Cambridge , MA , USA Published online: 02 Aug 2013.

To cite this article: Yong-Ak Song , Ahmed M.S. Ibrahim , Amr N. Rabie , Jongyoon Han & Samuel J. Lin (2013) Microfabricated nerve–electrode interfaces in neural prosthetics and neural engineering, Biotechnology and Genetic Engineering Reviews, 29:2, 113-134, DOI: 10.1080/02648725.2013.801231 To link to this article: http://dx.doi.org/10.1080/02648725.2013.801231

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Biotechnology and Genetic Engineering Reviews, 2013 Vol. 29, No. 2, 113–134, http://dx.doi.org/10.1080/02648725.2013.801231

Microfabricated nerve–electrode interfaces in neural prosthetics and neural engineering Yong-Ak Songa,b, Ahmed M.S. Ibrahimc,d, Amr N. Rabiee, Jongyoon Hanf* and Samuel J. Linc,d* a

Division of Engineering, New York University Abu Dhabi (NYUAD), Abu Dhabi, UAE; Department of Chemical and Biomolecular Engineering, Polytechnic Institute of New York University, Brooklyn, NY, USA; cDivisions of Plastic Surgery and Otolaryngology, Beth Israel Deaconess Medical Centre, Boston, MA, USA; dHarvard Medical School, Boston, MA, USA; e Department of Otolaryngology, Ain Shams University, Cairo, Egypt; fDepartment of Electrical Engineering and Computer Sciences and Department of Biological Engineering, MIT, Cambridge, MA, USA


b

(Received 5 December 2012; accepted 27 February 2013) Neural interfaces and implants are finding more clinical applications and there are rapid technological advances for more efficient and safe design, fabrication and materials to establish high-fidelity neural interfaces. In this review paper, we highlight new developments of the microfabricated electrodes and substrates with regard to the design, materials, fabrication and their clinical applications. There is a noticeable trend towards integration of microfluidic modules on a single neural platform. In addition to the microelectrodes for neural recording and stimulation, microfluidic channels are integrated into a nerve–electrode interface to explore the rich neurochemistry present at the neural interface and exploit it for enhanced electrochemical stimulation and recording of the central and peripheral nervous system. Keywords: neural prosthetics; neural implant; neural engineering; microfabrication; MEMS; microelectrode array

1. Introduction Interfacing the central nervous system (CNS) and peripheral nervous system (PNS) with electrodes has become increasingly critical in treating patients living with different forms of neurological diseases and finding potential cures for them (Prochazka, Mushahwar, & McCreery, 2001; Sujith, 2008). The growing demand for high spatial selectivity in conjunction with anatomical constraints has increased the need for making smaller electrodes. Microfabrication with its capability to create electrodes with feature sizes in the range of few micrometers offers great potential for neural prostheses in terms of miniaturization and system integration (Stieglitz et al., 2009). It allows decreasing the size of the device drastically while increasing the number of the electrodes for a given space. From the very beginning, the material of choice for the microfabricated nerve–machine interfaces has been silicon since the microfabrication process was originally developed in the semiconductor industry (Cheung, 2007). However, since early 2000, polymers have gained popularity because of their flexibility, biocompatibility and material inertness. Among them, polyimide is the most *Corresponding authors. Email: [email protected]; [email protected] Ó 2013 Taylor & Francis


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widely used polymer because of its micromachinability in the cleanroom, with excellent electrical insulation properties (Stieglitz, Beutel, Schuettler, & Meyer, 2000). As for the electrode materials, there are tremendous research efforts going on to develop materials with long-term stability and less electrode polarization (Cogan, 2008). Microfabrication offers a potential to reduce the electrode polarization effect with its capability to modify the surface structures at the microscale precisely. Besides the material development towards more flexibility and biocompatibility, a research trend is clearly visible towards more integration of functionalities into the neuroprosthetic device for stimulation and recording of neural signals as well as for sampling or dispensing molecules such as neurotransmitters into the nerve. In this paper, we review recent advances in microfabricated neural interfaces with regard to design, materials and fabrication as well as translational and clinical applications. 2. Fundamental phenomena at nerve–electrode interface Electrical stimulation is based on the charge injection at the interface between a metal electrode and an electrolyte medium. There are two mechanisms of charge transfer at the electrode–electrolyte interface. The first mechanism known as the capacitive mechanism is based on charging and discharging of the electrical double layer, generating a redistribution of charge in the electrolyte, but no electron transfer from the electrode into the electrolyte (Merrill, Bikson, & Jefferys, 2005). The second mechanism is based on the Faradaic reaction, in which electrons are carried between the metal electrode and electrolyte. This electron transfer is causing a reduction or oxidation of chemical species in the electrolyte that may be reversible or irreversible. In Figure 1, the electrode–electrolyte interface with these two abovementioned mechanisms is shown schematically. The Faradaic charge injection may create products in the electrolyte that cannot be recovered by reversing the direction of current. This irreversible Faradaic reaction creates chemical byproducts that could potentially harm the nerve tissue or electrodes. For this reason, it is highly desirable to reduce the amount of electrical current to avoid such an irreversible Faradaic reaction if a small charge is sufficient to initiate action potentials. A more detailed description of the fundamental electrochemi-

Figure 1. Electrode and electrolyte interface showing faradaic charge injection and capacitive charge injection (a), and a two-element electrical circuit model of electrode/electrolyte interface (b). Note: Reprinted by permission of Merrill et al. (2005).

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cal aspects related to the electrical stimulation can be found in a review paper by Merrill et al. (2005). Attention has to be paid to the electrode polarization effect in the development of the electrodes for electrical stimulation and recording. The electrode polarization is due to the interaction of the charged electrode surface with free charges in the electrolyte solution, as shown in Figure 1. An electrical double layer (EDL) on the electrode surface with a typical thickness of only a few nanometers generates large capacitance that shields the electric field in the suspending medium such as physiological relevant phosphate buffered saline (PBS) at low frequencies. The predominant factors influencing the thickness of EDL are the electrolyte concentration and charge. It is inversely proportional to the electrolyte concentration, as shown in Equation (1). 1

j

rffiffiffiffiffiffiffiffiffiffiffiffiffi kB T ¼ 2z2 e2 c

ð1Þ

where к-1 is the Debye length, ɛ is the permittivity of the dielectric, kB is the Boltzman constant, T is the absolute temperature, z is charge, e is elementary charge, and c is the electrolyte concentration. A key consequence of the electrode polarization effect is that the sensitivity of the electrodes is lowered to changes in the electrical properties of the bulk (Malleo et al., 2010). The effect of the polarization effect on the impedance can be expressed as in Equations (2) and (3), provided that the polarization impedance Zp ¼ Rp þ 1=jw Cp is smaller than the sample impedance (Schwan, 1968). R ¼ Rs þ Rp þ Rs ðRx CÞ C ¼ Cs þ

1 x2 R2 Cp

2

ð2Þ ð3Þ

where R and C are the measured total resistance and capacitance, Rs is the sample resistance, Rp is the polarization resistance, Cs is the sample capacitance, Cp is the polarization capacitance. Based on Equations (2) and (3), the measured resistance is larger than the sample resistance of interest by Rp + Rs(RωC)2 and the measured capacitance by 1/(ω2 R2Cp). Often, errors in resistance R due the electrode polarization are small, while errors in capacitance C become large in the low-frequency range (Schwan, 1968). To minimize this bias of the measurement data through the electrode polarization, parameters such as electrode design and material have to be taken into consideration. One of the methods is four-electrode configurations by providing a non-current-carrying second pair of electrodes to measure the voltage across the sample (Schwan & Ferris, 1968). In the four-electrode system developed by Schwan, the polarization capacitance can be reduced by decreasing the effective contact area between the voltage pick-up electrodes and the sample. So, instead of using a solid gold layer, the gold electrode was patterned by photolithography while maintaining the same electrode dimensions (Padmaraj, Miller, Wosik, & Zagozdzon-Wosik, 2011). The mesh electrodes had square holes of 7.5 μm by 7.5 μm resulting in 38% less area than solid electrodes. By reducing the surface area of the electrode in the shape of the mesh electrodes, the impedance due to the polarization effect could be reduced by 89.6% at 100 Hz (Padmaraj et al., 2011).


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This micropatterned electrode was then applied to the low-frequency impedance spectroscopy of isolated mitochondria in buffer solution (Padmaraj et al., 2011). In two-terminal electrode systems, the effect of electrode polarization can be minimized by increasing the electrode–electrolyte interface area, either by mechanically roughening the surface through sand blasting, or by using electrochemical processes such as electrodepositon that generate a porous surface with a large effective surface area (Malleo et al., 2010). The most common method applied is the electrodeposition of platinum on the surface of a smooth platinum surface. However, the electrodeposited platinum (Pt) black film is not stable and has limited biocompatibility. Other electrode materials with large effective surface areas are titanium nitiride (TiN), iridium oxide (IrOx), and polypyrrole/polystyrenesulfonate (PPy/PSS). Especially, sputtered IrOx has been used for neural recording electrodes. However, its long-term characteristic is still under investigation. PPy/PSS is a conductive polymer used for neural recordings. It offers several advantages such as good aqueous solubility, easy deposition as thin films, high conductivity, and biocompatibility (Malleo et al., 2010). Impedance spectra of the abovementioned electrode materials are compared with each other in Figure 2. As shown in the impedance spectra, the impedance of the EDL dominates the impedance of the plain (smooth) Pt electrodes across the frequency spectrum up to 100 Hz; the impedance of the other materials with large surface area is dominated by resistance in the low-frequency range below 10 Hz. Capacitance values of the electrode materials measured are listed in Table 1. Capacitance values for plain Pt are in the range 0.2–0.3 F/m2, while the capacitance of the porous Pt black electrode is 66.87 F/m2 in 1x PBS, an increase of 230-fold compared with plain Pt with smooth surface. As this example shows, microfabrication with its ability to structure the geometry and surface at the microscale holds great promise to minimize the effect of electrode polarization. More works using microfabricated devices are expected in future to elucidate the basics of the electrode polarization effect and to explore ways to minimize it along with the new material development. In this context, the nanofabrication processes such as e-beam (Liu et al., 2002) or ion-beam lithography (Watt, Bettiol, Van Kan, Teo, & Breese, 2005) and scanning tunneling microscope (STM)- or atomic force microscope (AFM)-based nanomachining (Hla, 2005; Tseng, 2011; Tseng,

Figure 2. Impedance spectra of smooth platinum, Pt black, IrOx and PPy/PSS-coated electrodes in phosphate buffered saline (conductivity: 1.6 S/m), impedance magnitude (a), and impedance phase (b). Note: Reprinted by permission of Malleo et al. (2010).


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Table 1. Effective interfacial capacitance for each material in 1 M NaCl and 1x PBS (Malleo et al., 2010). Solution 103 M NaCl 102 M NaCl 101 M NaCl 1 M NaCl 1 PBS

Pt Plain (F/m2)

Pt black (F/m2)

IrOx (F/m2)

PPy/PSS (F/m2)

0.23 0.26 0.30 0.30 0.30

32.5 53.54 77.48 81.54 66.87

9.66 15.60 22.12 28.72 21.74

– 127.82 234.16 312.18 259.06


Note: NaCl (sodium chloride), PBS (phophate buffered saline), Pt (platinum), IrOx (Iridum oxide), PPy/PSS (polypyrrole/polystyrenesulfonate).

Notargiacomo, & Chen, 2005) could also be used to manipulate the surface structure at the nanoscale for less electrode polarization effect. 3.

Development of neuroprosthetics

3.1. Design concepts The microfabricated electrode design can be classified into two major groups. The first group comprises the planar electrode design which is based on the inherent 2D characteristics of microfabrication processes. This design allows constructing a massively parallel electrode array for stimulation or for recording of neural signals. A well-known application example is the 2D epicortical electrode array, which is particularly suitable for a spatiotemporal observation of brain signals (Csicsvari et al., 2003). The second group is the 3D micro electrode array, which can be further subdivided into an in-plane and an out-of-plane type. This 3D electrode array resembles freestanding high-aspectratio microneedle structures with the electrical contacts at the tip end (Bai, Wise, & Anderson, 2000; Normann, Maynard, Rousche, & Warren, 1999). These 3D electrodes are suited for intracortical recording or stimulating single neuron activity. One example for the out-of-plane electrode design is the Utah array (Badi, Kertesz, Gurgel, Shelton, & Normann, 2003; Kim et al., 2009), commercialized by Cyberkinetics under the brand name BrainGate™ (DiLorenzo & Bronzino, 2008). The Utah array is shown in Figure 3a. A similar electrode design has been developed at the University of Michigan at Ann Arbor, as shown in Figure 3b. Their electrodes are of the in-plane design and are parallel to the surface of shank. Their technologies have been licensed to NeuroNexus, a spin-off company from the University of Michigan, and it has successfully commercialized several neural probes. Based on this 3D in-plane neural probe design, the Kipke group has implemented a parylene-based open-architecture probe with subcellular dimensions (5 μm), as shown in Figure 4. Their major finding was that there was a relationship between the neural probe’s subcellular size and shape and its chronic reactive tissue response. After implanting the device in rat cerebral cortex for 4 weeks, encapsulation through cells within 25 μm of the thin lateral structure was reduced by almost one-third compared with the shank (Seymour & Kipke, 2007; Kipke et al., 2008). In addition, microglia reactivity and protein deposition were greatly reduced at the lateral edge. This result is particularly encouraging for microfabricated neural interface devices since it showed that microfabricated structures could reduce chronic tissue encapsulation with their geometry and size.


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Figure 3. Examples of 3D out-of-plane and in-plane electrode arrays. Note: (a) The Utah Electrode Array (UEA) for high-density neural recording and stimulation is a silicon-based device of 3D out-of-plane type. It has a 10 10 array of tapered silicon electrodes with a base width of 80 μm, a height of 1500 μm and an electrode pitch of 400 μm. Reprinted by permission of Badi et al. (2003). (b) Schematic view of a multielectrode stimulating neuroprobe and a scanning electron microscope (SEM) image of the tip of a multisite probe with a thickness of 15 μm and a distance of 200 μm between the electrode sites. Reprinted by permission of Tanghe, Najafi, and Wise (1990).

Figure 4. 3D neural probe consisting of a thick shank (48 68 μm) and a thin lateral platform with subcellular dimensions (5 100 μm). Note: Scale bar = 100 μm. Reprinted by permission of Seymour and Kipke (2007).

A significant advantage of using silicon material for the neural probe is that electric circuits for multiplexing and amplification as well as microfluidic channels for delivery of reagents or drugs can be integrated into the array monolithically (Stieglitz et al., 2009). For instance, the European NeuroProbes consortium has been formed to develop a highly integrated silicon probe with several functionalities (Ruther et al., 2008). The goal of this project is to develop multifunctional probe arrays containing a 3D stimulation and recording electrodes with microfluidic functionality and integrated biosensors to monitor glutamate and dopamine. With the help of this probe, it is aimed to investigate the complex interaction between electrical and chemical signals in the brain. To shorten the development time between the first conceptual design to a clinical trial, a


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modular concept was suggested by Stieglitz (Stieglitz, Schuettler, & Koch, 2005). Electrodes and standard modules can be fabricated separately from electronics such as amplifiers and multiplexers once the interface connections between the modules have been standardized. Such a modular design has been implemented in the multifunctional neural probe of the European consortium. 3.2.

Material selection

3.2.1.

Electrode materials


Another important issue when it comes to building a neural interface device is the choice of a proper material for the electrodes and substrate. The general requirements for an electrode material can be summarized as follows (Merrill et al., 2005): • • • •

biocompatibility long-term stability the amount of generated toxics through the Faradaic reactions should be minimal faradaic corrosion rates should be as low as possible to avoid a premature failure of the electrodes

Usually, the noble metals such as platinum (Pt), gold (Au), iridium (Ir), palladium (Pd) and rhodium (Rh) with high corrosion resistance have been used for electrical stimulation. Since platinum is a relatively soft material and therefore easily deformable during insertion, it is often alloyed with iridium to increase its mechanical strength. These metals inject charge by both faradaic reactions and double-layer charging. In most neural stimulation cases, the faradaic processes dominate doublelayer charging. To avoid irreversible processes of faradaic reactions that can cause electrode and tissue damage, a careful selection of charge injection waveform is necessary (Cogan, 2008). Titanium nitride (TiN) is a chemically stable metallic conductor with good biocompatibility and is one of the capacitive injection materials (Cogan, 2008). Large charge-injection capacities ( 1 mC/cm2) are achievable by creating electrodes with a high surface roughness through sputter deposition, as discussed in section 2. In fact, porous TiN is used as a coating for cardiac pacing electrodes because the electrode polarization during a pacing pulse is low so that the electrode can detect the cardiac contraction in the interpulse period (Norlin, Pan, & Leygraf, 2005). Even though TiN has a large charge storage capacity, access to all the available charge under the high rate and high current density conditions of a neural stimulation pulse, can be delayed by pore resistance (Goldberg, Bard, & Feldberg, 1972; Norlin et al., 2005; Posey & Morozumi, 1966). This behavior is shown schematically in Figure 5. The pore resistance R1-R3 in combination with capacitance C1-C3 creates a delay-line with a timeconstant dependent on the pore geometry, electrolyte resistivity and the interfacial double-layer capacitance (Cogan, 2008). Consequently, narrower and deeper pores lead to higher time constants, and their charge-injection capacity takes longer to access than shallow pores. One way to modify the pore geometry for faster access would be by varying the process parameters during the sputtering process. This approach, however, can be time-consuming. Alternatively, e-beam or ion-beam nanolithography as well as STM- and AFM-based machining with their capability to manipulate structures in the nanoscale could be a viable option to fabricate pore geometries so that the time constant can be positively influenced.

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(a)

(b)

Single pore

R1 C1

R2 C2


C3

Electrode

R3

Cross sectional view

Figure 5. Highly porous surface of sputtered TiN resulting in a large effective electrode surface area. Note: (a) SEM image of TiN surface. Reprinted by permission of Schaldach, Hubmann, Weikl, and Hardt (1990). (b) Schematic of a pore cross-section with pore resistance R and double-layer capacitance C. Modified from Cogan (2008).

New emerging materials are conducting polymers and carbon nanotubes as alternatives to Pt, Pt alloys, Ir oxide and TiN for chronic stimulation and recording (Cogan, 2008). Especially, the conductive polymer offers the possibility of surface modification with biomolecules to increase biocompatibility and create a biologically active interface between the electrode and tissue. A ~30-fold reduction in electrode impedance was achieved at very low frequencies ( 400 None 3

1420 5

1710 0.5

7.4

2.9

12

3.9

3 105

6 106 1016

3.1 20 103

0.002 8 106 1016

2 106 > 1016

(Akin, Najafi, Smoke, & Bradley, 1994) and used for neural implants (Richardson, Miller, & Reichert, 1993). The Stieglitz group used polyimide Pyralin PI 2611 for substrate and for insulation (Stieglitz et al., 2000). It has comparable insulation resistance and dielectric strength to those of silicon oxide and silicon nitride, at a lower density and higher flexibility, as shown in Table 2. This material has also been tested according to the international standard ‘Biological Evaluation of Medical Devices’ (ISO 10993) and USP-23. PI 2556 and PI 2611 showed excellent biocompatibility, and PI 2566 also had good biocompatibility (Stieglitz et al., 2000). Long-term study over 1 year showed an excellent biostability (Rubehn & Stieglitz, 2010; Stieglitz et al., 2005). However, polyimide is known to absorb moisture, leading to delamination at the interface between metal/polyimide (Weiland, Liu, & Humayun, 2005). Polyimide can be processed following the standard microfabrication procedure. Shortly, the polyimide resist is spincoated to the thickness required, prebaked, photolithographically patterned and then postbaked at 350°C. Once the flexible device has been fabricated, it can be connected to an external electronic circuitry via MicroFlex Interconnection (MFI) technique (Stieglitz et al., 2000; Stieglitz et al., 2005). It is similar to the common thermosonic ball wedge bonding, in which a gold ball is pressed again a flexible polyimide substrate onto a metal layer of an integrated circuit, thereby creating an interlayer mechanical and electrical connection. Parylene C is another flexible substrate material for neural interfaces. It has the USP class VI approval for chronic human implantations (Stieglitz et al., 2009). The cleanroom processing of parylene C is well established and is comparable to polyimide. However, handling of parylene C-based substrate seems to be a difficult task on account of its low stiffness, and adhesion of metal electrodes to parylene C surface is weak and prone to delamination (Stieglitz et al., 2009). Liquid crystal polymers (LCP) are new emerging materials for flexible neural prostheses (Hassler et al., 2011). These materials offer high mechanical strength at high temperatures, excellent chemical resistance, low moisture absorption and permeability. Compared with polyimide, the LCP offers the flexibility, chemical inertness and biocompatibility of polyimide combined with 100 times less absorption of moisture. LPC is available in sheet format with predefined thickness ranging from 25 μm to 3 mm.

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Meso scale

Precision machining

Single Nerve fiber

Single neurons

Micro scale

Laser machining (laser cutting, ablation)

Microfabrication via photolithography

Ion channel, Pores

Nano scale

Nanofabrication (E-beam lithography, STM-, AFM-, focused ion beam machining, electrochemical etching)

Available fabrication technologies

Figure 6. Classification of available fabrication technologies for neuroprosthetics based on the length scale of machinable structures. Modified from Stieglitz (2010).

The sheet material can be processed by laser machining or reactive ion etching. The Durand group has demonstrated the first neural prosthesis made out of this material (Wang, Liu, & Durand, 2009). A 2-μm-thin iridium oxide film was deposited on top of a 50-μm-thick LCP layer with a 50-nm-thick adhesion layer. The electrode allowed a charge injection limit at 4.5 mC/cm2. 3.3. Fabrication and packaging strategies At present three different manufacturing technologies for the fabrication of neuroprosthetic devices are available, as shown in Figure 6: traditional precision machining, laser machining, and micro/nanofabrication. Traditionally, the implant electrodes are precision machined-out metal sheets of several tens of micrometers up to 100 μm. These electrodes can withstand corrosion for several years before they break down. The complexity of the device, however, is significantly limited because of the manual assembly process. Using a laser beam, the structures can be cut more precisely and, therefore, a higher integration is possible. Feature sizes as small as 250um with a pitch size of at least 100um are possible (Stieglitz, 2010). To increase the integration density of the electrode array even further, either by decreasing the size of the electrodes or the gap between the electrodes, microfabrication via photolithography is more suitable than laser machining. To go further down in the length scale to the nanometer regime, e-beam (Liu et al., 2002; Malaquin et al., 2004; Ueno, Hayashida, Ye, & Misawa, 2005) or ion-beam lithography (Watt et al., 2005), focused ion-beam machining (Errachid et al., 2008) and STM- or AFM-based nanomachining (Hla, 2005; Tseng et al., 2005) could be used. The primary benefit of fabricating smaller electrodes such as nanoelectrodes would be the enhanced mass



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transport since diffusion becomes dominant at this scale (Arrigan, 2004). In principle, study of faster electrochemical and chemical reactions could be possible. Even though no application of these nanomanufacturing technologies to neuroprosthetics has been reported so far, the nanofabrication technologies offer a great potential for the field of electrochemical sensing on neural probes in conjunction with the use of nanomaterials as the next-generation electrode materials. Especially, for structuring the electrode surfaces to minimize the electrode polarization effect, the nanofabrication methods could open up a new exciting possibility. After assembling the modules to an integrated system, packaging is the next essential step. To protect electronic circuitry against body fluid, the implant can be hermetically or nonhermetically sealed (Stieglitz, 2010). Titanium or alumina are the most common materials for packaging. Nearly all commercially available neural implants use a hermetic sealing. However, there is a limitation of the system complexity because of the one-to-one connection from the electrode to the feed-through. A simpler way of sealing the device is to use nonhermetic sealing by polymers such as silicone rubber or parylene C (Stieglitz, 2010). Both materials are approved to be used in chronic implants for humans.

4.

Translational application examples of neural interfaces

4.1. ECoG-electrode array An electrocorticogram (ECoG)-electrode requires an array of electrodes for recording the signals from large areas of a cortex. At present, there is a commercially available array from Ad-Tech Medical Instrument Corporation, Racine, WI, USA. However, the electrode and pitch size are in the order of millimeters. One of the earliest developments towards microfabricated electrode arrays was reported by the D. Anderson group at the University of Michigan. They used silicon microfabrication to build a 3D high-aspectratio microelectrode array for recording and stimulation in the CNS. The array had 8 16 electrodes with a shank diameter of 200 μm to record cortical signals (Bai et al., 2000). Histological studies on cats showed that the tissue reaction to the arrays were minimal with healthy-looking neurons near the probe shanks. They also built a 4 4 microelectrode array with on-chip complementary metal–oxide–semiconductor (CMOS) preamplifier and buffers. Tsytsarev, Taketami, Schottler, Tanaka, and Hara (2006) built a rod-shaped multi-electrode array probe with 8 8 planar arrays of platinum recording electrodes. Each platinum electrode had a diameter of 50 μm. As a spacer, insulated copper wire was used between the platinum wires. The planar array had a contact surface area of about 0.6 mm2. This electrode array allowed recording of spatially localized field potentials as well as the spontaneous and stimulus-evoked action potentials of neurons in close proximity to the cortical surface. To increase the spatial resolution of the measurement even more, several groups have been working on the microfabricated planar electrode array. The Stieglitz group developed a micro-electro-mechanical systems (MEMS)-based flexible multichannel ECoG-electrode array consisting of a thin polyimide substrate with enclosed platinum electrode sites and conductor paths (Rubehn, Bosman, Oostenveld, Fries, & Stieglitz, 2009). They patterned electrodes with a diameter of 1 mm at a pitch size of 2, 2.5 and 3 mm. Even though the electrodes themselves were not on the micron scale, the conductor paths connecting the individual electrodes with the solder pads on the other end had a width of 15 μm at a pitch size of 30 μm. In this way, 252 electrodes covering an area of 60 35 mm could be


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Figure 7. Flexible electrode arrays. Note: (a) Polyimide-based 252-channel epicortical array for the brain hemisphere. Reprinted by permission of Hassler et al. (2011). (b) A highly flexible, high-density electrode array used on the visual cortex in an animal experiment. This flexible array van also be applied into the interhemispheric fissure, as the inset shows. Reprinted by permission of Viventi et al. (2011).

connected through a narrow ribbon cable to the connectors on the other end, as shown in Figure 7a. The impedance of the electrodes varied from 1.5 kΩ to 5 kΩ at 1 kH. The electrodes recorded on early visual areas to prefrontal regions for a period of 4.5 months after implantation without loss of signal quality. A similar approach was taken by Hollenberg and coworkers. They built flexible electrode arrays on a thin prefabricated 75 25-mm Kapton substrate with 64 gold electrodes in 8 8 arrays (Hollenberg, Richards, Richards, Bahr, & Rector, 2006). Each electrode was 150 μm in diameter and the electrodes covered an area of 5.25 5.25 mm. To insulate the electrode array, SU-8 was patterned on top of the electrodes. However, since an amplifier and multiplexer were not integrated in the Kapton array, a long-term in vivo study could not be done with the device. Long-term effects of SU-8 on the neural tissue need to be investigated. An exciting development with a more sophisticated flexible device of higher electrode array density has been reported by the Rogers group at the University of Illinois–Urbana Champaign. Using their flexible silicon nanomembrane transistor transfer printing technology, they built a massively parallel array of 720 silicon transistors and recorded on a large area of cat brain ( 80 80 mm), as shown in Figure 7b. This array enabled a 400-fold higher spatial resolution than the current ECoG electrode which led to an observation of complex spatial patterns such as spiral waves and clustering of spatiotemporal patterns (Viventi et al., 2011). In addition, the extremely flexible device can conform to uneven surface regions of the brain such as sulci and fissures where conventional flexible electrode array might have a problem accessing. This flexible electrode array is expected to be an efficacious tool for studying disorders such as epilepsy and dementia, affective disorders, movement disorders and schizophrenia, in which the brain activity has to be observed over a large area rather than at a small confined location (Viventi et al., 2011). To enhance the deformability of the electrode array, Rogers and his coworkers decreased the thickness of polyimide layer from 76 μm to 2.5 μm. In order to handle such a thin device with extremely low stiffness, they used a biologically dissolvable polymer – silk fibroin – with a thickness of 25 μm as a substrate enhancer (Kim et al., 2010). The advantage of using silk fibroin is that this material is completely dissolvable in biological fluids and biocompatible. So once the electrode array was placed on the brain cortex, it was dissolved and the remaining polyimide layer could lie down on the brain surface conforming itself to its nonuniform structure. By


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incorporating a mesh pattern into the polyimide substrate, its conformability could be further increased as tested on complex curvilinear surfaces such as feline brain. 4.2. Sieve and cuff electrodes Sieve electrodes on polyimide substrates are used to interface with regenerating peripheral nerves in the PNS and CNS. The Stieglitz group developed a sieve electrode for sciatic nerve of rats (Stieglitz et al., 2000). The diameter of the device was 2.1 mm and contained 570 holes with a diameter of 40 μm. The interconnect width was 1.0 mm. They also built cuff electrodes on a polyimide substrate for stimulation of motor fibers with a diameter of 1.6 mm. Twelve polar cuff electrodes implanted on the rat sciatic nerve enabled selective electrical stimulation of gastrocnemius and tibialis anterior muscle for alternating dorsiflexion and plantarflexion of the foot (Rodriguez et al., 2000). Another example for a flexible neural substrate developed by the Stieglitz group is a flexible retina stimulator for a spatiotemporal stimulation of ganglion cells of the retina. Their device contained 24 concentric, bipolar electrodes with integrated interconnects and connection pads (Stieglitz et al., 2005). So far the experiments with polyimides on peripheral nerves, retinal vision prostheses and epicortical electrode arrays have shown positive results with only mild immune reactions in chronic implantation and stable operation in vivo. The Durand group developed a flat interface nerve electrode (FINE) to flatten the nerve, thereby decreasing the distance between central axon populations and the electrode (Tyler & Durand, 2002). This new electrode design improved the more invasive intrafascicular approach, in which a thin metal or silicon probe penetrates through the membrane into the fascicles. The rectangular-shaped flat electrode shaped the nerve in a more elongated or flattened oval so the distance between the electrode and the fascicles became smaller. As an outcome of this flattening, the FINE electrode enabled higher selectivity than a standard electrode by individually addressing the fascicles. 4.3.

Neural probe

Most of the developed 3D neural probes are based on the in-plane design, as described in section 3.1. The Voldman group at MIT developed a microfabricated flexible neural probe for the CNS of the moth Manduca sexta. With the help of this flexible device in polyimide, they could stimulate the left and right abdomen of the insect as well as record neuronal activity in the ventral nerve cord of the moth (Tsang et al., 2012). By electroplating carbon nanotubes (CNT) on the gold electrodes, the interfacial impedance between the probe and the neural tissue could be reduced resulting in lower stimulation voltage. The same group also built a flexible split-ring electrode in polyimide for stimulation of the ventral nerve cord in insects (Tsang et al., 2010). Altuna et al. (2012) used SU-8 to build a 3D neural probe with integrated planar electrodes for neural depth recording. Using microfabrication, the recording capability has been improved by decreasing the impedance of the electrodes. Also, on-chip circuits as well as wireless systems have been integrated for chronic multisite neural systems. There have also been developments towards integrating microfluidic modules such as microchannels, shutters, valves and pumps into the 3D neural probe. The reason for the integration of fluidic modules is that the pure electrical stimulation is inherently causing electrochemical reactions at the electrode surface. These irreversible reactions can lead to damage of the electrode as well as of neurons. To alleviate this problem, there are two viable options available. Either the amount of required current can be


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reduced or a chemical stimulation with neurotransmitters delivered through the microchannel can be applied (John, Li, Zhang, Loeb, & Xu, 2011). In fact, neuronal action potentials are triggered chemically by the release of neurotransmitters from presynaptic neurons. Also, delivering drugs to the site of the inserted electrode could reduce tissue inflammation, biofouling, encapsulation or death of neurons. The parylene vapor-deposition technique was also applied to build a flexible 3D neural probe. Takeuchi, Ziegler, Yoshida, Mabuchi, and Suzuki (2005) developed a flexible neural probe out of parylene. They deposited 5-μm-thick parylene on top of a 10-μm-thick photoresist (AZP4620) and then dissolved the photoresist with the solvents. The dissolved area left behind a microchannel structure that could be used to infuse or draw fluids from the sample. In this device, six recording pads out of gold have been integrated around the microfluidic channel. To insert this highly flexible substrate into the tissue, they used polyethylene glycol (PEG) to fill the microchannel and enhance the stiffness of the probe. In this way, they could increase the stiffness of the probe from 1 mN to 12 mN and insert it into the tissue without bending. Once the probe had been inserted, PEG was dissolved by saline solution thereby decreasing the stiffness of the probe. Benzocyclobuthene (BCB) has also been used to build a flexible structure; however, it faced a similar stiffness problem (Lee et al., 2004). However, integrating microfluidic channels as well as other fluidic components such as reservoirs, valves and pumps remains a real engineering challenge and, even if one can manage to integrate all these modules into a single device, it would be quite challenging to maintain its stable operation in vivo. Therefore, it would be ideal to combine the advantages of electrical stimulation such as fast-acting stimulation with the advantages of chemical stimulation such as benign stimulation without generating

Figure 8. Electrochemical stimulation of sciatic nerve using ion-selective membrane. Note: With the help of an ion-selective membrane, the ion concentration around the nerve is modulated prior to the electrical stimulation. After depleting a specific ion such as Ca2+ ions using the membrane, an electrical current is injected to initiate the action potential. After depleting the Ca2+ions, the electrical threshold could be reduced by 40%.



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irreversible reaction processes at the electrode. In addition, it would be of great benefit with regard to device integration and operation if the movable microfluidic components can be omitted and the replenishment of neurotransmitters or other reagents in the fluidic reservoir is not required. Towards this goal, Song et al. (2011) have developed an electrochemical stimulation method that allows harvesting chemicals from the surrounding environment and utilizing ion concentration modulation to lower the electrical threshold. They have developed an electrochemical method to activate and inhibit a nerve by electrically modulating ion concentrations in situ along the nerve, as shown in Figure 8. Using an ion-selective membrane (ISM) to achieve different excitability states of the nerve, they observed either a reduction of electrical threshold by up to 40%, or reversible inhibition of nerve signal propagation. This low-threshold electrochemicalstimulation method is applicable in current implantable neuroprosthetic devices, whereas the on-demand nerve-blocking mechanism could offer an effective clinical intervention for chronic disease states caused by uncontrolled nerve activation, such as epilepsy and chronic pain syndromes. Kim, Abidian, and Martin (2004) investigated whether conducting polymers in connection with hydrogel can be a viable coating material on neural prosthetic devices. Hydrogel offers several advantages such as an additional mechanical buffer interface between the hard silicon-based probe and the soft brain tissue. Also, the reswelling of the cured gels allows better anchoring of the probe in the tissue. Since the hydrogel possesses intrinsic pores, these can be used as a storage for biomolecules such as nerve growth factors (e.g., neutrophin) or anti-inflammatory drugs that can delivered to a specific tissue in a controlled manner (Winter, Gokhale, Jensen, Cogan, & Rizzo, 2008). A neurotrophin-eluting hydrogel layer on top of an electrode has been tested and the result showed an improved neuron-electrode proximity, which could result in lower threshold value. Furthermore, the hydrogel matrix increases the effective surface area that can reduce the impedance of the electrode, as discussed in sections 2 and 3.2.1. The impedance of the electrode can be controlled by the thickness of the hydrogel, which in turn can be regulated by the polymerization time. The lowest impedance of the electrode achieved at 1 kHz was 7 Ωk compared with 100 kΩ of the polypyrrole film. However, the stability of a hydrogel–electrode interface remains a concern for its long-term usage and needs more experimental investigations. 5.

Clinical applications of nerve–electrode interfaces

Electricity is vital in sustaining the human body; it provides the current in which the human brain, nervous system and heart function. Electrical stimulation has been used in clinical practice since ancient times; texts from Egypt dating back to 2750 bc, describe the presence of electric fish used to treat patients suffering from conditions such as gout or headaches. In the present day, it is well known that a voltage applied to the human body creates an electric current through the tissues. Clinical applications of electrical stimulation are summarized in Table 3. If the current is sufficiently high, it causes muscle contraction, cardiac fibrillations and under extreme circumstances tissue burns (Baldridge, 1954). One of the oldest uses of electricity in medicine is electrocardiography (ECG); a simple, painless test that records the heart’s electrical activity, aiding in the diagnosis of various cardiac conditions. Also, millions of patients worldwide benefit from implanted cardiac pacemakers that use electrical pulses to help control abnormal heart rhythms, and implantable cardioverter-defibrillators (ICD) that save up to 100,000 American lives annually (Mackenzie, 2004).

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Table 3. Summary of the clinical applications of electrical stimulation. Electrical stimulation method

Disease state

Electrocardiography (ECG) Cardiac pacemaker Cardioverter-defibrillator Functional electrical stimulation (FES)

Monitor electrical activity of the heart Control abnormal heart rhythms Cardiac arrest Standing, ambulation, cycling, hand grasp and release, arm reaching, breathing, and bladder/bowel control, increase muscle bulk, improve cardiovascular performance, prevent and treat pressure ulcers, treat osteoporosis and joint contractures, control spasticity Paralysis, paresis, pain Refractory cluster headache Recovery of blink reflex Slow transit constipation, overactive bladder Fecal incontinence Facial nerve paralysis, synkinesis Trigeminal neuralgia, painful diabetic neuropathy

Spinal cord stimulation (SCS) Stimulation of occipital nerve Stimulation of palpebral nerve Stimulation of tibial nerve Stimulation of sacral nerve Stimulation of facial nerve Transcutaneous electric nerve stimulation (TENS) Electroconvulsive therapy (ECT) Deep brain stimulation Pulsed electromagnetic field (PEMF) therapy Electrical nerve stimulation Enhanced neural interface Electrotherapy

Depression, Parkinson’s disease, epilepsy Parkinson’s disease, Alzheimer’s disease Arthritis Healing of chronic diabetic ulcers Improved control of prosthetic limbs Bone fusion, rehabilitation of muscle

Various means of electrical stimulation have been recommended for the recovery of neurological functions and may be essential for the maintenance of neural circuitry should a cure be found. One of these – functional electrical stimulation (FES) – aims to restore certain functional motor activities, such as standing, ambulation, cycling, hand grasp and release, arm reaching, breathing, and bladder/bowel control (Ragnarsson, 2008). In addition, it has been used to increase muscle bulk, improve cardiovascular performance, prevent and treat pressure ulcers, treat osteoporosis and joint contractures, control spasticity, and improve general well-being (Ragnarsson, 2008). Spinal cord injuries (SCI) contribute to billions of dollars lost by the health care system each year. It has been reported that 5.6 million individuals in the current US population have forms of paresis or paralysis as a result of SCI (T. C. D. Reeve-foundation, 2010). Spinal cord stimulation (SCS) has been described to enhance the recovery of these patients by aiding in the restoration of motor function (Minassian, Hofstoetter, Tansey, & Mayr, 2012). Furthermore, SCS using high-frequency electrical pulses has been used for the relief of chronic intractable pain (Graybill, Conermann, Kabazie, & Chandy, 2011; Hegarty & Goroszeniuk, 2011; McAuley, Van Groningen, & Green, 2012). Patients with nerve paralysis experience serious psychological and functional problems. Treatment modalities have not seen significant development over time, with current management options having limited outcomes. Recent studies have shown that continuous electrical stimulation can promote peripheral nerve regeneration, producing positive effects in as little as 3 weeks (Shen & Zhu, 1995). As such, it has been used to stimulate the occipital nerve for the treatment of refractory cluster headache (Strand, Trentman, Vargas, & Dodick, 2011), the palpebral nerve for recovery of the blink reflex (Salerno, Bleicher, & Stromberg, 1990), the tibial nerve in slow-transit constipation and overactive bladder (Collins, Norton, & Maeda, 2011; Marchal et al., 2011), the sacral



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nerve for fecal incontinence (Boyle et al., 2011), but perhaps most impressively, as an implantable electrical prosthesis for facial reanimation alongside nerve grafting (Griffin & Kim, 2011; Song et al., 2011). Moreover, transcutaneous electric nerve stimulation (TENS) has emerged as a promising option for the management of patients with trigeminal neuralgia as well as painful diabetic neuropathy (Gossrau et al., 2011; Singla, Prabhaker, & Singla, 2011). Perhaps the greatest electrical mystery known to man is the human brain. Medical researchers are constantly finding new ways to use the electrical properties of the brain to treat diseases and injuries. Electroconvulsive therapy (ECT) has been used to treat psychiatric conditions, the most common of which is depression (Agarkar, Hurt, Lisanby, & Young, 2012; Allan & Ebmeier, 2011). While considered controversial by some, ECT has gained renewed acceptance in the treatment of Parkinson’s disease and is also now being touted as a means to treat epilepsy (Alon, Yungher, Shulman, & Rogers, 2012; Fisher, 2011; Shin et al., 2011). Similarly, high-frequency deep brain stimulation has been proven an effective therapy for the motor symptoms in Parkinson’s disease and, more strikingly, it has been shown to enhance memory for use in the management of patients with Alzheimer’s disease (Brocker et al., 2012; Suthana et al., 2012). Pulsed electromagnetic field (PEMF) therapy is now being used as a noninvasive means to ‘kick start’ the body’s natural inflammatory response, thus promoting rapid soft-tissue wound healing, relieving pain and improving other associated symptoms of various arthritic disorders (Ganesan, Gengadharan, Balachandran, Manohar, & Puvanakrishnan, 2009). Electrical nerve stimulation has also been deemed successful in improving the healing of chronic diabetic ulcers (Lundeberg, Eriksson, & Malm, 1992). More recently, studies have shown that scarless fetal skin wound repair requires neural stimulation for tissue regeneration (Stelnicki et al., 2000). In a time when military casualties and extremity injuries are on the rise, electrical stimulation has proven effectively to reduce muscle atrophy as well as improve muscle force in denervated muscles following limb amputation (Nemoto, Williams, Lough, & Chiu, 1988). Amputees have significant difficulties coordinating the separate functions of prosthetic limbs; the control of shoulder/hip-level disarticulation prostheses is significantly more difficult than that of prostheses for more distal amputations. Electrical stimulation could potentially improve neural interfaces in myoelectric prosthesis for better control of small muscle segments (Hijjawi et al., 2006). Electrotherapy devices are also being used to promote bone fusion as well as to rehabilitate muscles following injury or surgery (Hussman, 1969). Electrical stimulation has and will always serve as an important treatment modality in medicine. Ideally, these techniques should be clinically efficacious and well tolerated with the ability to operate in a low-power mode while maintaining optimal function. It will continue to be at the forefront in the management of nerve paresis/paralysis as well as chronic painful conditions, and will serve as a platform for the cure of nervous dysfunction syndromes that may result from civilian and military trauma. 6.

Summary and outlook

Microfabrication has found wide applications in neuroprosthetics and neural engineering. With its capability to fabricate electrode arrays at high intensity and integrate several functionalities into a single miniaturized platform, it offers unprecedented spatiotemporal resolution for stimulation, recording and sensing on the central and peripheral nervous systems. The progress in the material development has allowed more biocompatible and safer electrodes. With the help of micro- and nanofabrication, the


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geometries of electrodes as well as the surface roughness can be controlled at the micro- and nanoscale so that the electrode polarization effect is minimized for higher sensitivity in neural signal measurement. The substrate is becoming more flexible with integrated amplifiers and multiplexers so it can be implanted without damaging the tissue, and allows local stimulation of the nervous system and recording of bioelectrical signals for diagnostics. Numerous translational and clinical applications of the microfabricated devices have already been reported for the electrical stimulation of neuromuscular systems as well as for epicortical and intracortical recording on the CNS. With the rapid progress in design, materials and fabrication of the electrodes and substrates, which has been the focus of this review, we will see more of highly miniaturized, multifunctional and biocompatible neuroprosthetic implants and nerve–machine interfaces advancing into clinical applications and as research tools for neuroscientific investigations into neurodegenerative diseases soon. References Agarkar, S., Hurt, S., Lisanby, S., & Young, R. C. (2012). ECT use in unipolar and bipolar depression. Journal of Electroconvulsive Therapy, 28, e39–40. Akin, T., Najafi, K., Smoke, R. H., & Bradley, R. M. (1994). A micromachined silicon sieve electrode for nerve regeneration applications. IEEE Transactions on Biomedical Engineering, 41, 305–313. Allan, C. L., & Ebmeier, K. P. (2011). The use of ECT and MST in treating depression. International Review of Psychiatry, 23, 400–412. Alon, G., Yungher, D. A., Shulman, L. M., & Rogers, M. W. (2012). Safety and immediate effect of noninvasive transcranial pulsed current stimulation on gait and balance in Parkinson disease. Neurorehabilitation and Neural Repair, 26, 1089–1095. Altuna, A., de la Prida, L. M., Bellistri, E., Gabriel, G., Guimera, A., Berganzo, J., … Fernandez, L. J. (2012). SU-8 based microprobes with integrated planar electrodes for enhanced neural depth recording. Biosensors & Bioelectronics, 37, 1–5. Arrigan, D. W. M. (2004). Nanoelectrodes, nanoelectrode arrays and their applications. Analyst, 129, 1157–1165. Badi, A. N., Kertesz, T. R., Gurgel, R. K., Shelton, C., & Normann, R. A. (2003). Development of a novel eighth-nerve intraneural auditory neuroprosthesis. The Laryngoscope, 113, 833–842. Bai, Q., Wise, K. D., & Anderson, D. J. (2000). A high-yield microassembly structure for three-dimensional microelectrode arrays. IEEE Transactions on Biomedical Engineering, 47, 281–289. Baldridge, R. R. (1954). Electric burns: Report of a case. New England Journal of Medicine, 250, 46–49. Boyle, D. J., Murphy, J., Gooneratne, M. L., Grimmer, K., Allison, M. E., Chan, C. L., & Williams, N. S. (2011). Efficacy of sacral nerve stimulation for the treatment of fecal incontinence. Diseases of the Colon & Rectum, 54, 1271–1278. Brocker, D. T., Swan, B. D., Turner, D. A., Gross, R. E., Tatter, S. B., Koop, M. M., … Grill, W. M. (2012). Improved efficacy of temporally non-regular deep brain stimulation in Parkinson’s disease. Experimental Neurology, 239C, 60–67. Cheung, K. C. (2007). Implantable microscale neural interfaces. Biomedical Microdevices, 9, 923–938. Cogan, S. F. (2008). Neural stimulation and recording electrodes. Annual Review of Biomedical Engineering, 10, 275–309. Collins, B., Norton, C., & Maeda, Y. (2012). Percutaneous tibial nerve stimulation for slow transit constipation: A pilot study. Colorectal Disease, 14(e1), 65–70. Csicsvari, J., Henze, D. A., Jamieson, B., Harris, K. D., Sirota, A., Bartho, P., … Buzsaki, G. (2003). Massively parallel recording of unit and local field potentials with silicon-based electrodes. Journal of Neurophysiology, 90, 1314–1323. DiLorenzo, D. J., & Bronzino, J. D. (2008). Neuroengineering. Boca Raton, FL: CRC Press.



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Optical neural interfaces.

Review of Brain-Machine Interfaces Used in Neural Prosthetics with New Perspective on Somatosensory Feedback through Method of Signal Breakdown.

Neural Engineering for Rehabilitation.

Neural networks in enzyme engineering.

Combining decoder design and neural adaptation in brain-machine interfaces.

Corrigendum: Micro- and Nanotechnologies for Optical Neural Interfaces.

Elastomeric and soft conducting microwires for implantable neural interfaces.

Titania nanotube arrays as interfaces for neural prostheses.

MicroRNAs and their potential therapeutic applications in neural tissue engineering.

Corrigendum: Making brain-machine interfaces robust to future neural variability.

Spiking Neural Network Decoder for Brain-Machine Interfaces.

A Low Noise Amplifier for Neural Spike Recording Interfaces.

Neural interfaces for somatosensory feedback: bringing life to a prosthesis.

Invasive neural interfaces: the perspective of the surgeon.

Closing the Loop in Deep Brain Stimulation for Psychiatric Disorders: Lessons from Motor Neural Prosthetics.

Brains-Computers-Machines: Neural Engineering in Science Classrooms.

Microfluidic systems for stem cell-based neural tissue engineering.

Studies in RF power communication, SAR, and temperature elevation in wireless implantable neural interfaces.

Modulation depth estimation and variable selection in state-space models for neural interfaces.

Fabrication and electrochemical comparison of SIROF-AIROF-EIROF microelectrodes for neural interfaces.

Morphogenetic movements in the neural plate and neural tube: mouse.

Common neural correlates of real and imagined movements contributing to the performance of brain-machine interfaces.

Erratum: Optimal feedback control successfully explains changes in neural modulations during experiments with brain-machine interfaces.

In vivo Characterization of Amorphous Silicon Carbide As a Biomaterial for Chronic Neural Interfaces.