Biocompatible Materials for Optoelectronic Neural Probes: Challenges and Opportunities

The ability to understand and treat debilitating neurological conditions such as Parkinson’s disease, spinal cord injury, and chronic pain is limited largely by the lack of materials and devices that can seamlessly interface with neurons and restore or bypass the malfunctioning neural circuits (Cogan et al. 2008; Gilja et al. 2011; Normann 2007). But the technology involved in deep brain and spinal cord stimulation devices used to treat such conditions dates back to the 1950s (Hamani and Temel 2012; Kringelbach et al. 2007). Even cutting-edge experiments that enable tetraplegic patients to control robotic aids (Hatsopoulos and Donoghue 2009; Hochberg et al. 2012) depend on devices invented more than 20 years ago (Campbell et al. 1991). These devices do not take into account the fundamental materials properties of neural tissue, and so their reliability and long-term effectiveness are diminished (Lee et al. 2005; Polikov et al. 2005).

Flexible organic and hybrid electronics offers a compelling solution to the elastic and surface chemistry mismatch between neural probes and neural tissues, while enabling novel approaches for neural interrogation. Recent developments in materials chemistry and fabrication methods make flexible electronics ripe for tailored, biointegrated neuroprosthetics.

In this article I review challenges and opportunities in the materials selected for neural probes and the role of flexible electronics and optoelectronics at the frontier of neural engineering.


Methods of Neural Stimulation and Recording

Devices for neural recording and stimulation interact with neural tissues with different degrees of precision and invasiveness (Buzsáki et al. 2012). For example, electroencephalography (EEG) is performed noninvasively through the skull and thus offers a low-resolution map of smoothed field potentials associated mainly with the neural activity of the whole cortical surface. Electrocorticography (ECoG), involving devices placed directly on the cortical surface, yields higher temporal and spatial resolution and is routinely used to identify seizure loci in epilepsy patients.

Neural systems exchange information in the form of action potentials—voltage spikes that propagate along neuronal membranes—and fluctuations in local field potentials (LFPs) averaged across a neuronal subnetwork or even an entire structure in the nervous system. Detailed mapping of neural activity is clinically relevant not only in the cortex but also in deep brain regions (e.g., the subthalamic nucleus in Parkinson’s patients), the spinal cord, and peripheral nerves (e.g., in trauma patients or those in chronic pain). Moreover, many neurological disorders are associated with abnormal activity of specific types of neurons, and hence single-neuron resolution is essential to the development of effective therapies. I focus here on penetrating neural recording devices, designed to interface with individual cells in a particular region of the nervous system.

As with neural recording, neural stimulation offers varying degrees of precision and invasiveness. Noninvasive transcranial magnetic stimulation (TMS) allows for interrogation of cortical circuits via initiation of local flows of ions, which are hypothesized to cause changes in LFPs (Allen et al. 2007; Ridding and Rothwell 2007). However, there is currently no strategy for extending this approach to deep brain regions or targeting it to specific neuronal types because of the nonspecific nature and limited penetration depth of the low-frequency magnetic fields used in TMS.

In deep brain stimulation (DBS), an approved treatment for Parkinson’s and essential tremor patients, high-voltage pulses (1–10 V; as compared to membrane voltages, ~30–100 mV, or LFPs, ~1–5 mV) are used to stimulate the neural tissue surrounding the electrodes (Perlmutter and Mink 2006). But although the DBS therapeutic effect is well documented, its underlying mechanisms remain unclear; both electrically induced excitation and inhibition of neural activity have been proposed (Kringelbach et al. 2007). Furthermore, nonspecific interrogation of large tissue volume often yields undesirable side effects such as depression or compulsive behaviors (Frank et al. 2007; Temel et al. 2007).

Epidural electrical stimulation (in the spinal cord of chronic pain patients) is essentially equivalent to DBS, with the key difference that the electrode leads are placed on top of the dura (the thin barrier that isolates nerves from other tissues) rather than deep in the neural tissue.

Development of Optogenetics

With the development of optogenetics it became possible to excite or inhibit specific neuronal types with millisecond precision (Boyden et al. 2005; Zhang et al. 2007). This method uses genetic targeting of light-sensitive proteins, opsins (of algal, archaeal, and bacterial origin), to establish neuronal sensitivity to a variety of visible light wavelengths. Opsins can be roughly categorized as excitatory (used for evoking action potentials; e.g., cation channel channelrhodopsin 2, ChR2) or inhibitory (used for inhibiting action potential firing; e.g., modified chloride pump halorhodopsin, eNpHR3.0, and modified proton pump archaerhodopsin, eArch3.0) (Zhang et al. 2011).

Optogenetics is a powerful tool for scientific investigation of the behavioral correlates of neural dynamics, but its genetic and mechanical invasiveness impedes its clinical translation (Yizhar et al. 2011). As mammalian tissues are highly scattering and absorptive in the visible light range, implantation of optical waveguides or light-emitting devices is necessary for implementation of optogenetics. Thus, optical stimulation technologies face materials design and biocompatibility challenges similar to those of tissue-penetrating neural recording and stimulation electrodes.

Reliability Challenges of Implantable Neural Probes

Neural recording and stimulation devices have traditionally been fabricated out of hard materials with elastic moduli (Young’s modulus E~10s–100s GPa1) exceeding those of neural tissues (E~kPa–MPa) (Borschel et al. 2003; Green MA et al. 2008) by many orders of magnitude. For example, neural recording and electrical stimulation electrodes (Figure 1) are often based on silicon (silicon multielectrode or “Utah arrays”; Bhandari et al. 2008; Campbell et al. 1991), multitrode probes (Blanche et al. 2005; Kipke et al. 2003; Seymour et al. 2011), silica (cone electrodes; Bartels et al. 2008; Kennedy et al. 1992), or metals (individual microwires of tungsten, gold, platinum, or platinum-iridium alloys; tetrodes and stereotrodes of nickel-chromium alloys; Gray et al. 1995; Jog et al. 2002; McNaughton et al. 1983). Similarly, optical stimulation in optogenetic experiments is most often performed with standard commercially available silica optical fibers (E~50–90 GPa) implanted directly into neural tissue.

Figure 1

It is hypothesized that this mismatch in stiffness contributes to tissue damage and the resulting encapsulation of devices in dense scars composed of glial cells, leading to a decrease in recording quality (Lee et al. 2005; Polikov et al. 2005). It is reasonable to assume that the probe insertion itself produces a certain amount of initial damage as well (destruction or displacement of cells in the path of the implant), an assumption that is supported by the commonly observed improvement in recording quality approximately two weeks after implantation. However, the signal-to-noise ratio (SNR) and the total number of recorded neurons then decay steadily over the course of the implant lifespan.

Several mechanisms have been proposed to explain the neuronal death and glial scarring that compromise the probe’s effectiveness. One hypothesis is that, as neural probes are generally at least partially fixed to the skull/vertebrae, their motion is constrained, whereas the neural tissues may shift by tens to hundreds of micrometers due to movement, heartbeat, and respiration (Britt and Rossi 1982; Muthuswamy et al. 2003). This micromotion of soft neural tissues around the hard implants is thought to introduce additional tissue damage.

Another theory is that the disruption of glial networks by devices larger than an average cell (i.e., >10 µm) may increase astrocytic and astroglial responses that lead to thickening of the scar tissue around the device (Seymour and Kipke 2007). Furthermore, devices with particularly sharp edges have also been shown to be disruptive to the blood-brain barrier, inducing an inflammatory response that raises glial activity and the likelihood of scarring (Saxena et al. 2013).

Materials and Methods for Flexible Substrates

Flexible organic and hybrid electronics and optoelectronics offer opportunities to address the elastic, geometric, and chemical compatibility challenges of neural recording and stimulation devices.

Combining traditional metal and semiconductor technologies with flexible substrates provides a first transitional step toward stealthy bioinspired neural probes. Over the past decade polymer substrates have been used as a backing for metal and silicon-based neural recording electrodes. As illustrated in Figure 2, electrode arrays have been developed (using lithographic MEMS-inspired processing2) on silicone resins (poly(dimethylsulfoxane); PDMS), polyimide, and parylene C, to name a few (Kim BJ et al. 2013; Minev et al. 2012; Stieglitz et al. 2009; Viventi et al. 2011). Because these devices exhibit high flexibility and conformability to complex landscapes, they found immediate application in high-density microstructured cortical arrays (micro-ECoG) and nerve cuffs.

Figure 2

Contact printing methods developed by Rogers and colleagues have enabled highly innovative neural probes. This technology takes advantage of mature semiconductor-based (opto)electronics and combines it with flexible interconnects that enable transfer of circuit elements that are several microns thick onto polyimide and silk fibroin backing (Kim et al. 2010a, 2011). These flexible and foldable devices were recently introduced deep into the brain with the use of resorbable microneedles (Kim TI et al. 2013).

Meng and colleagues have taken an alternative approach by using a thermal molding process to produce soft cone electrodes based on parylene C, with active electrode pads facing inside the cone (Kim BJ et al. 2013; Tooker et al. 2004). This creative technology relies on earlier findings by Kennedy and colleagues (1992), who used silica capillaries seeded with nerve fragments to attract neuronal growth into the capillary containing an electrode, thus making a truly biointegrated device.

Yet there remain a number of challenges in the fabrication of neural probes on flexible substrates, such as relatively low resolution (dictated by contact printing methods), inability to scale to the high number of channels necessary for comprehensive mapping of brain activity, and inadequate capacity to interface with optical or drug delivery elements essential for neural interrogation (and potentially cell type identification). Robust reproducible manufacturing of probes suitable for use in human patients presents another challenge, as MEMS-style processing offers relatively low yield and is currently constrained to standard wafer sizes (several inches as compared to the several feet needed for a spinal cord).

Surface Modification and Encapsulation of Neural Probes

Because materials interfaces between devices and neural tissues play a critical role in both tissue response and the quality of neural recording, surface engineering is an important aspect of neural probe design. With their tunable chemical properties and low elastic moduli, organic materials offer a compelling toolbox for the engineering of intimate electrically and optically active interfaces between neurons and neural probes.

Surface Engineering

Polymers such as (poly(3,4-ethylenedioxythiophene); PEDOT) (Blau et al. 2011; Ludwig et al. 2011; Richardson-Burns et al. 2007), polylysine (Boehler et al. 2012; Hai et al. 2010), and polypyrrole (Abidian et al. 2010; George et al. 2005) have been shown to boost the reliability and SNR of neural recording electrodes by promoting cell adhesion and reducing the impedance of equivalent circuits between the devices and the neuronal membranes.

Hydrogels based on polymers and polymer blends of natural (agarose, alginate, xyloglucan, hyaluronan, methylcellulose, chitosan, and matrigel) and synthetic (methacrylate, polyethylene glycol (PEG), poly(vinyl alcohol), and poly(acrylic acid)) origins are used for most neural regeneration scaffolds (Frampton et al. 2011; Hanson Shepherd et al. 2011; Jhaveri et al. 2008; Nisbet et al. 2008; Seliktar et al. 2012; Shin et al. 2012) and have recently found application in surface modification of neural probes (Jun et al. 2008; Kim et al. 2010b; Lu et al. 2009). The advantages of hydrogels include elastic moduli comparable to those of the neural tissues as well as high permeability for nutrients and oxygen.

However, the electrical and optical properties of these soft gels have not yet been engineered for improved neural recording and stimulation. Consequently their application in neural probe engineering has been restricted to providing low-modulus biocompatible buffers, which may reduce the damage associated with micromotion.


Encapsulation is another form of surface modification routinely used during deep-tissue implantation of flexible neural probes. As mentioned above, flexible substrates make it possible to overcome the elastic modulus mismatch between an electronic or optoelectronic probe and the surrounding neural tissue. But it is difficult to target soft devices to a specific region of the nervous system as they are prone to buckling, which hampers straight-line penetration.

Dissolvable encapsulation temporarily stiffens the probe to permit targeted implantation. Organic and biopolymeric materials such as PEG, sugar, tyrosine-based polymers, and silk fibroin are often used because of their adjustable dissolution speed in aqueous environments as well as their versatile chemistry and biocompatibility. Silk fibroin (also used as a biocompatible adhesive) enables the introduction of PDMS-backed probes through silicon microneedles that are retracted shortly after implantation upon dissolution of the silk fibroin.

Opportunities with Polymer Optoelectronics

Two decades of advances in materials chemistry have propelled small-molecule organic optoelectronics into commercial applications in the display industry and beyond, but the sensitivity of these materials to environmental moisture and oxygen hinders their application in the human body. In contrast, environmentally stable polymers and polymer composites with versatile chemical and electronic properties and low elastic moduli present a promising materials system for the development of multifunctional tissue interfaces.

Despite their wide adoption throughout the medical community (orthopedic implants, encapsulation materials for stimulation electrodes, porous scaffolds for soft tissue regeneration), polymers have yet to be fully explored with respect to their applications in neural probes (Green RA et al. 2008). Pioneering studies by Martin and Kipke, among others, illustrate the potential of PEDOT, polypyrrole, and polymer-carbon composites (Abidian et al. 2010; Kozai et al. 2012) (Figure 3) to solve the elastic mismatch of neural recording devices while reducing the overall electrode impedance and thus increasing SNR. In parallel, Capadona and Tyler have applied biologically inspired design principles to create polymer composites with controllable elastic properties that mimic sea cucumber dermis (Capadona et al. 2008; Harris et al. 2011).

Figure 3

Despite growing evidence of the utility of polymers in neural probe design, various engineering challenges prevent widespread adoption of these materials systems by neuroscientists and clinicians. For example, polymer probes are primarily fabricated by electrospinning, chemical vapor deposition, thin film spin-casting, and lithography. The first two methods offer relatively low throughput and require painstaking postsynthesis assembly if multiple electrodes are desired, which is true for most neuroprosthetic applications. Furthermore, these methods currently do not allow for integration of optical elements, which are essential for neural stimulation applications. Although well-developed lithographic methods allow for integration of multiple functional elements, they are limited by the flat substrate geometry, which is not ideal for applications in deep brain regions.

In my laboratory we have recently explored a thermal drawing process (TDP) inspired by optical fiber production as a means of fabrication for multifunctional neural probes. During TDP a macroscale preform, which can be fabricated using low-end mechanical processing, is drawn into a fiber with microscale features (Abouraddy et al. 2007; Bayindir et al. 2004; Goff 2002; Varshneya 1994). The lateral dimensions are scaled by as much as 10,000-fold using, if necessary, multiple drawing steps, enabling the creation of structures on the nanometer scale without the need for high-resolution fabrication technology (Kaufman et al. 2011; Yaman et al. 2011). At the same time, the length is stretched by a factor of ~100, yielding hundreds of meters of fibrous devices with a conserved cross-sectional pattern.

Because TDP faithfully reproduces the cross-sectional geometry of the macroscopic preform, it enables the creation of sophisticated multifunctional structures on the microscale. In addition, it is compatible with a wide range of materials with varying optical and electrical properties, permitting, for example, the combination of waveguide core and cladding materials, conductive polymer composites, and low-melting-temperature metal microwires in a single device.

We have used TDP to produce a range of fiber-inspired neural probes (FINPs), from high-channel-count neural recording arrays of arbitrary lengths to multifunctional devices incorporating waveguides, drug delivery channels, and neural recording electrodes (Figure 4).

Figure 4

Our preliminary in vivo evaluation of FINPs suggests that TDP may provide a scalable fabrication tool for flexible optoelectronic devices compatible with implantation in a variety of regions of the nervous system. Furthermore, this process may complement recent materials discoveries by Martin, Capadona, Kipke, and others as it may not only enable the integration of these innovative polymer systems into multifunctional probes but also offer a pathway toward their high-throughput production.


High-fidelity recording and stimulation of neural activity are essential to the development of neuroprosthetic devices as well as to the mapping of neural circuits involved in neurological and neuromuscular disorders. While mature semiconductor technologies provided initial promise for neural probe design, recent tissue engineering studies illustrate the need for alternative biocompatible materials platforms. In this article I have reviewed the challenges of established neural probe technologies and the opportunities of flexible organic and hybrid materials platforms for improvements in the biocompatibility and longevity of these sensors. I have also emphasized the importance of integration of optical neural stimulation modules and discussed fabrication approaches that may enable flexible multifunctional neural prosthetics.


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1 GPa = gigapascals; kPa = kilopascals; MPa = megapascals.

2 MEMS = microelectromechanical systems.

About the Author: Polina Anikeeva is AMAX Assistant Professor in Materials Science and Engineering at the Massachusetts Institute of Technology.