
==== Front
Bioact Mater
Bioact Mater
Bioactive Materials
2452-199X
KeAi Publishing

S2452-199X(24)00372-4
10.1016/j.bioactmat.2024.08.034
Review Article
Enhancing biocompatibility of the brain-machine interface: A review
Villa Jordan Jordan.villa@northwestern.edu
a
Cury Joaquin Joaquin.cury@northwestern.edu
a
Kessler Lexie Lexie.kessler@northwestern.edu
a
Tan Xiaodong Xiaodong.tan@northwestern.edu
ab
Richter Claus-Peter Cri529@northwestern.edu
abcd⁎
a Northwestern University-Feinberg School of Medicine, Department of Otolaryngology, USA
b The Hugh Knowles Center, Department of Communication Sciences and Disorders, Northwestern University, USA
c Department of Communication Sciences and Disorders, Northwestern University, USA
d Department of Biomedical Engineering, Northwestern University, USA
⁎ Corresponding author. E Superior St 13-564, Chicago, IL, 60611, USA. Cri529@northwestern.edu
11 9 2024
12 2024
11 9 2024
42 531549
9 4 2024
5 8 2024
27 8 2024
© 2024 The Authors
2024
https://creativecommons.org/licenses/by-nc-nd/4.0/ This is an open access article under the CC BY-NC-ND license (http://creativecommons.org/licenses/by-nc-nd/4.0/).
In vivo implantation of microelectrodes opens the door to studying neural circuits and restoring damaged neural pathways through direct electrical stimulation and recording. Although some neuroprostheses have achieved clinical success, electrode material properties, inflammatory response, and glial scar formation at the electrode-tissue interfaces affect performance and sustainability. Those challenges can be addressed by improving some of the materials' mechanical, physical, chemical, and electrical properties. This paper reviews materials and designs of current microelectrodes and discusses perspectives to advance neuroprosthetics performance.

Graphical abstract

Image 1

Highlights

• Biocompatible coatings can promote long-term electrode coatings.

• Pharmaceutical, peptide, and polymer coatings reduce inflammatory responses in implantable electrodes.

• Mechanical, thermal, and electrical properties affect chronic microelectrode array stability.

• Implantable electrodes activate immune response and promote glial scar formation.

Keywords

Brain-machine interface
Neuroprostheses
Implantable devices
Electrode coatings
Immune reaction
==== Body
pmc1 Introduction

Injuries to the nervous system and neurodegenerative diseases often result in severe functional and sensory impairments. The typical treatment includes neural prostheses facilitating direct communication between the neural tissues and electronic devices. These devices typically consist of electrodes connected to electrical circuits (processors) to deliver electrical current and record neural responses to form the brain-machine interface (BMI). In the remainder of the manuscript, “Brain” in BMI describes brain tissue and all other neural tissue. The BMI offers a potential pathway not only to alleviate debilitating conditions but also to restore lost functions.

Developing a functional long-term BMI requires neural stimulation and recording of responses at a high spatial resolution over the patient's lifetime. At the same time, tissue damage and failure of the BMI should be minimized. It is important to recognize that failure of a BMI, often defined as a lack of any recording or stimulatory activity, is multi-faceted. The failure can include idiopathic causes, material failures due to broken probes/sensors following implantation [1], mechanical failure due to displacement of the electrode from the intended target [2], or biological failure from the body rejecting a foreign object [3]. Failure may occur within a few weeks or several years. In case of BMI failure, it typically happens during the first year after implantation [4].

In this Review, we discuss the advancements in electrode design and materials of BMIs. We explore the integration of new biocompatible coatings that aim to reduce the body's inflammatory reaction and use cutting-edge materials designed to mirror the mechanical properties of neural tissue, thereby minimizing tissue damage and rejection.

2 Brain-machine interface: design considerations

2.1 Mechanical properties

Mechanical properties describing a BMI may include hardness, softness, flexibility, and Young's modulus. For better communication, these terms are defined as follows. Hardness or softness is the resistance of a material to local plastic deformation by an indenter. It is the material's property to withstand localized scratch, rebound, and indentation. Flexibility is the ability of a material to deform elastically. After removing the applied stress, the material will return to its original shape during elastic deformation. The term flexibility also characterizes Young's modulus; flexible materials have a low Young's modulus, and stiff materials have a high Young's modulus. Young's modulus describes a solid material's mechanical properties in the material's linear elastic region. It measures the tensile and compressive stiffness of the material to a force along an axial axis. It is calculated as the ratio of the force applied per area (stress) and the resulting displacement or deformation (strain). Hardness and Young's modulus differ as hardness measures the resistance to surface deformation. Young's modulus quantifies the degree to which a material can resist elastic deformation under loading conditions.

Today, the mechanical properties of neural devices differ by order of magnitudes from those of neural tissue (Fig. 1, Table 1) [5]. For example, Young's modulus of brain tissue is roughly 1 kPa, whereas the BMIs' Young's modulus of commonly used materials is up to 100 GPa (Fig. 1, Table 1).Fig. 1 Relative Young's modulus values for commonly used materials in producing a BMI. A. Biocompatible, non-toxic materials that may be used when designing electrodes for the BMI. B. Young's modulus values of neurotoxic materials; the materials should be avoided when designing a BMI.

Fig. 1

Table 1 Material properties of diverse materials used in neural interfaces.

Table 1Material	Young's Modulus (GPa)	Thermal conductance (W/m*K)	Electrical conductance (S/m) (20 °C)	Neurotoxicity	Relative
Allergenic History	
Metals, Metalloids, and Alloys used as Conductive Tracks	
Graphene	950–1000	140	1.0x 105	Nontoxic [52]	Rare [42,64]	
Tungsten	411	170	2.0x 107	Variable results [48,49]	Rare [42,64]	
Gold	79	320	4.5x 107	Nontoxic [47]	Common [42,64]	
Silicon	47	150	1000	Nontoxic [54]	Rare [42,64]	
Platinum	168	71	9.4x 106	Minimal toxicity [47]	Rare [42,64]	
Iridium	528	140	2.1x 107	Nontoxic [47]	Rare [42,64]	
Tin	50	67	9.1x 106	Minimal toxicity [47]	Rare [42,64]	
Titanium	116	22	2.5x 106	Nontoxic [47]	Rare [42,64]	
Copper	130	400	5.9x 107	Toxic [51]	Common [42,64]	
Nickel	200	91	1.4x 107	Toxic [51]	Most common [42,64]	
Cobalt	209	100	1.7x 107	Toxic [51]	Common [42,64]	
Silver	83	430	6.2x 107	Toxic [51]	Common [42,64]	
Zinc	108	120	1.7x 107	Toxic [51]	Common [42,64]	
Chromium	279	94	7.9x 106	Toxic [51]	Common [42,64]	
Biocompatible Polymers for Electrode Coating	
PEDOT	2.6 [65]	0.17–0.37 [66]	2.8x 105 [67]	Nontoxic	–	
PVA	2.75 [68]	8.49 [58]	1.05x 10−9 [69]	Nontoxic	–	
PLGA	1.57–2.43 [70]	–	9.8x 10−9 [71]	Nontoxic	–	
PEG	0.32 [70]	0.297 [72]	–	Nontoxic	–	
PDMS	0.00036–0.000870	0.15	1.0x 105	Nontoxic	Rare	
Encapsulation Materials used as Structural Scaffolding	
Parylene-C	3.2	0.084	–	Nontoxic	Rare	
Silk	0.515	0.083	–	Nontoxic	Common	
Silicon Dioxide	66.3–74.8	1.3–1.5	1.0 x 10−12	Nontoxic	Rare	
Teflon	0.575	0.24	–	Nontoxic	Rare	
Bioactive Glass	55–126	0.5–1.38	–	Nontoxic	Rare	
Sapphire	400	25–46	–	Variable Toxicity	Rare	

Table 2 Electrode coatings for long-term fidelity.

Table 2Electrode Coating	Material used	Neuron Adhesion	Foreign Body Response	Length of Study	Reference(s)	
Polypeptide and ECM Coatings	
Poly-D-Lysine	Electrode materials: platinum, iridium, gold	Improved neurite attachment and growth compared to uncoated controls	Not Discussed in Referenced Articles	14 days	[59,165]	
7 days	
PEI-LN	Oxide-covered silicon substrate	Improved neurite growth and attachment with neuron differentiation	Not Discussed in Referenced Articles	5 days	[163,165]	
No change in impedance	
eNL-1	Gold-deposited titanium nitride electrodes with indium tin oxide track lines	Selectively helps in the formation of excitatory synapses at the brain-machine interface	Not Discussed in Referenced Articles	12–14 days	[166,167]	
L1 Protein	Electrodes: iridium oxide
Substrate: Silicon dioxide, glass	There was no change in impedance compared to uncoated controls at 4 weeks.	Lower Iba-1 markers within 60 μm of the implant compared to control.
Significantly lower GFAP staining 240 μm from the implantation site	2–4 months	[[169], [170], [171]]	
Significantly higher neuron density in L1-coated probes.	
Increased NF-100 staining showing mature neurons.	
Electrically Conductive Polymers	
PEDOT (Coating)	Electrode: gold	Enhanced surface contact at the brain-machine interface and lower impedance levels	The gliosis process matched that of uncoated controls (starting at week 3)	3 weeks	[56,57]	
Substrate: Silicon dioxide and silicon nitride	
PEDOT:Poly (Structural scaffold)	Electrode: Platinum, gold, and indium tin oxide	Enhanced neurite attachment and growth with decreased impedance	Diminished GFAP markers over 4-week in vivo study	4 weeks	[174]	
(SS-4VP)
(Structural scaffold)	Substrate: PEDOT:PSS	
PVA with PLGA	In vitro: silica substrate	Enhanced surface area of brain-machine interface	Not Discussed in Referenced Articles	4 weeks	[15]	
PEG	Electrode: Gold	Not Discussed in Referenced Articles	Diminished mechanical stress in multiple directions	8 weeks	[[176], [177], [178], [179]]	
Increased BBB leakage and implantation trauma from increased size resulting in astrocyte activation	
LIPS	Electrode: Platinum	Neuron attachment supported through signal: noise ratio 2x that of uncoated probes	>95 % reduction in protein adhesion within in vitro activation of microglial cells	Not Discussed in Referenced Articles	[181]	
Substrate: Silicon	
Antioxidants	
DFX	Substrate: Parylene-coated silicon electrodes	Not Discussed in Referenced Articles	Stabilized BBB leakage within 7 days post-implantation	1 week	[182]	
iSOD	Electrode: glassy carbon, platinum, and silver/silver chloride	Significant reduction in neuronal apoptosis 1-week post-implantation	Reduced BBB capillary damage.	1 week	[183]	
No effects on Iba-1 staining compared to controls.	
Reduced NAPDH oxidase activation.	
PVT	Substrate: Silicon	Not Discussed in Referenced Articles	Reduced cell/neuron death following immune reaction.	Not Discussed in Referenced Articles	[184]	
Non-cytotoxic effects on astrocytes.	
Diminished NADPH oxidase-driven damage.	
Nanotube Coating	
Graphene	In vitro characterization	Not Discussed in Referenced Articles	Inhibition of TLR4 innate signaling pathway	Not Discussed in Referenced Articles	[187]	
Gold	Polymer-based probe of PBMS and PEG	Not Discussed in Referenced Articles	Significant reduction in GFAP and Iba1 levels at 2 and 6 weeks.	4 weeks	[188]	
Pharmaceutical Agents	
Corticosteroids	Electrode: Platinum	Reduced distance in brain-machine interface	Reduced microglial and astrocyte activation in acute and chronic settings.	12 weeks	[[189], [190], [191],193,194]	
Substrate: Polyimide, PLGA	
Limited glial scar formation.	
Minocycline	Electrode: Gold	Not Discussed in Referenced Articles	Reduced pro-inflammatory cytokine release and Iba-1 staining.	4 weeks	[149]	
Substrate: Parylene-coated silicon electrodes	
GFAP levels unchanged.	
α-MSH	Nitrocellulose-coated silicon wafer	Not Discussed in Referenced Articles	Reduced pro-inflammatory cytokine and NO production	21 weeks	[198]	
Melatonin	Michigan-style [185]silicon probe	–	Reduced pro-inflammatory cytokine and NO production in chronic implantation	4 weeks	[199]	
HOE-642	Michigan Silicon electrode	–	Reduced microglial activation and encapsulation compared to controls up to 72 h.	72 h	[196]	
Reduced radius of microglial activity 72 h post-implantation.	

Following the implantation of the BMI, micromotions between the implanted devices and the neural tissue may occur from differences in Young's modulus and lead to tissue reactions, limiting biocompatibility and affecting performance [[6], [7], [8]]. Changing the size and shape of the electrode can alter the bending stiffness; however, the stiffness of a device can never be entirely eliminated [9].

Rigid neural interfaces, like those made of tungsten, platinum, or carbon, result in more severe tissue damage than those made of a more compliant material, such as gold. The effects are independent of whether the BMIs are tethered to the skull or not [[10], [11], [12]].

Neural interfaces with material stiffness close to that of tissue encounter different challenges. The thin, flexible electrodes are not stable enough to be inserted into neural tissue. It has been addressed by incising the pia mater [13], treating it with collagenase [14], and stiff transient coatings with dextrose [[15], [16], [17]], maltose [18], poly-glycolic acid [19,20], poly(lactic-co-glycolic acid) [21,22], gelatin [23,24], silk [25], and carboxymethylcellulose [[26], [27], [28], [29]]. The various coatings have succeeded in implantation by providing a stiff vector and the ability to dissolve, allowing the formation of the BMI through direct contact with the neurons and the electrode. The mechanical properties play a prominent role in the overall stability and biocompatibility of the BMI. Hence, the material selection, the size of the electrodes, and how they exist within biological environments are crucial for the sustainability of BMIs. In an ideal world, a BMI would be as small as possible and have properties like hardness, stiffness, and Young's modulus, close to the tissue it is designed to exist within. Fig. 1 below compares Young's modulus of commonly used metals and polymers (left column) and those that should be avoided due to neurotoxic effects (right column). It is important to understand that both biocompatibility and biotoxicity are highlighted in Fig. 1. Biocompatibility in this context refers to the materials' relative similarity to actual biological tissue. In other words, the closer the material is to Young's modulus of brain tissue, the more biocompatible. Biotoxicity, however, refers to potential neural damage following tissue contact with the listed materials. Damage frequently originates from tissue reactions, resulting in excessive reactive oxygen species (ROS) formation.

2.2 Thermal properties

BMIs convert only a fraction of the supplied electrical energy into electrical currents stimulating neurons. Most of the remaining energy is converted into heat with a subsequent temperature increase, which must be dissipated by the surrounding tissue [30]. While the heat production at the electrode is typically neglectable, the heat production by the processor must be considered. Consequently, materials with higher thermal conductance should be used cautiously in the implants.

Seese et al. (1998) studied how temperature changes and chronic heating affect tissues. They showed that a maximum temperature elevation of 1°Kelvin (K) can be observed before tissue necrosis and subsequent fibrotic encapsulation occur [31,32]. Other studies have found cortical necrosis at temperature elevations of 2 K [33] and 3.4 K [34]. Kim et al. (2006) simulated the net change in cortical temperature after implanting a three-dimensional (3D) Utah Electrode Array (UEA). The results showed that the electrode elevated the tissue temperature by 1.3 K, exceeding the limit for damage described previously [35].

The Food and Drug Administration allows a temperature increase of less than 2°Celsius (2 °C or 2 K) over 15 min if the device is adjacent to thermally sensitive tissue, such as neural tissue, eyes, testes, or ovaries. The increase in temperature can be less than 4 °C (4 K) if the device is adjacent to tissue that is not thermally sensitive.

2.3 Chemical properties

The most important chemical property to consider when designing a BMI is how prone the material is to developing excessive ROS [36]. Metal generation of ROS is well documented within the literature and has even been categorized in terms of their relative ability to form such harmful molecules. Nanoparticles of silver, copper, iron, manganese, and cobalt have been shown to generate the most ROS in physiologic environments. Meanwhile, gold, silicon, and platinum have been categorized into less corrosive and damaging materials for a BMI [37]. Like the electrode material, it is important to ensure that any BMI coatings used do not increase ROS production. This is especially true as coatings are often used for enhancing the implantation of flexible structures and cause inevitable implantation damage. Poly(3,4-ethylenedioxythiophene)-poly(styrenesulfonate) (PEDOT:PSS) is a conductive polymer that has been shown to enhance implantation of microwire BMIs, as well as reduce intracellular ROS compared to non-coated controls [38]. Similar polymers like poly-vinyl alcohol (PVA) have shown the ability to scavenge ROS in wounds, which supports its potential use as an electrode coating for a BMI [39]. The significance of generating ROS in biological tissue is thoroughly described later in this review in the “Typical Reactions in Electrode Implantation” subsection. In short, the generation of excessive ROS damages the tissue required for recording and stimulating with the BMI. The damage will accumulate over time until the once structurally normal tissue is infiltrated with inflammatory responders and the BMI is no longer functional.

2.4 Electrical properties

A BMI delivers electrical current to stimulate neurons and measures voltages to characterize responses of the neural system. For each measurement, the voltage and current amplitude ratio provides the impedance. The impedance of the electrodes refers to the resistance and reactance these electrodes offer to the flow of electrical current. This property is crucial because it influences the efficiency of electrical signal transmission (stimulation and recording) between the electrodes and neural tissues. The impedance depends on the stimulus frequency. High impedances of the BMI require higher voltages to pass sufficient current levels to stimulate neurons. Consequently, devices will consume more power, leading to faster battery depletion in portable or implantable BMIs [40].

A derived measure is the current density, the current level divided by the electrode contact's cross-sectional area. Current density values provide boundaries for safe neural stimulation. Large current densities facilitate the formation of gas bubbles at the electrode contact, free neurotoxic compounds from the electrode, and may change the pH at the electrode-to-fluid interface. Bubbles lead to mechanical tissue damage, and toxic compounds directly damage the cells and their neural transmission, impacting the safety and long-term viability of the device [41]. The fluid environments and the electrodes' contact material also determine the BMI's electrical properties. While the ion content of the extracellular space cannot be altered, the electrode materials can. As shown in Table 1, the most promising materials include iridium, gold, and tungsten [42].

A successful BMI delivers electrical currents at levels sufficient to stimulate the neurons but small enough to avoid tissue damage. The BMI should also be able to simultaneously record neural responses from the target tissue, requiring a good signal-to-noise ratio (SNR). The first microelectrodes approved for chronic deep brain stimulation (DBS) had current densities below 30 μC/ cm−2, the tissue and neuron damage threshold determined by in vivo animal studies [43,44]. Other experiments on the maximum current density for safe electrical stimulation found current densities of 12 μC/ cm−2 and 60 μC/ cm−2 [44,45]. In particular, implants for spinal cord stimulation required larger charge densities before any damage was reported [45]. Overall, the electrical properties of a particular BMI may depend on its intended use. Thus, to maximize the BMIs' ability to record and stimulate activity, the impedance, current density, and materials used should match that intended use. Further research must outline appropriate values for such variables in different applications. Guidelines on these values will provide a more streamlined use of BMIs for researchers with similar intentions to further evaluate their use.

2.5 Biocompatibility

A BMI is considered biocompatible if it is mechanically and chemically matched to the tissue it will exist within [46]. Besides optimizing the various discussed physical, thermal, and electrical properties, it is important to consider other factors, including avoiding toxic materials and prominent allergenic responses.

2.6 Neurotoxicity of materials

Neurotoxic materials lead to cellular necrosis and loss of neurons. Selvakumaran et al. (2002) examined the stability and cytotoxic effects of implants made with gold (Au), platinum (Pt), iridium (Ir), indium tin oxide (ITO), and titanium (Ti) in direct contact with the tissue. Compared to a polystyrene tissue culture control, Au, ITO, and Ir showed no decrease in neuronal activity or growth following 72 h of direct contact. Despite Pt and Ti samples decreasing the number of living cells, they were not labeled cytotoxic as over 75 % of cells remained [47].

Neural interfaces using Tungsten have been used for several decades. However, in the presence of oxygen or other oxidizing species, tungstic ions have toxic effects on the neural tissue near the implantation site. Tungsten's cytotoxic effects, high Young's modulus, and hardness challenge its use in BMIs [48,49]. Other metals, including iron, silver, copper, cadmium, manganese, lead, chromium, and nickel, show elevated neurotoxic effects, most likely through free radical species formation [50,51].

In the form of graphene and carbon nanotubes (CNTs), carbon is a promising material for neuro-prosthetic devices [52]. Graphene and CNTs have a high Young's modulus, are not cytotoxic, and can be used to record and stimulate neural activity. They also promote neuronal proliferation and adhesion at the implanted site [53]. The mechanical properties allow flexible, biocompatible BMI designs. Similar characteristics for silicon microelectrode implants have been found, supporting their use in BMIs like the UEA [54].

Besides metals and metalloids, recent research efforts explored electrically conductive polymer-based microelectrode arrays [55]. These polymers, which will be discussed further in this review as potential coatings to enhance biocompatibility, include polypyrrolen [56], poly (3,4-ethylenedioxythiophene, (PEDOTor PEDT) [57], Poly-vinyl alcohol (PVA) [22,58], poly-(lactic-co-glycolic acid) (PLGA) [22], Poly-D-Lysine [59], and poly(ethylene glycol) (PEG) [60]. All polymers were tested in vivo and showed no evidence of tissue damage or neurodegeneration [22,[61], [62], [63]].

2.7 Allergenic considerations

Some materials trigger tissue reactions, manifested as acute and chronic tissue inflammation, with subsequent failure of the BMI. An allergenic response to a material in a BMI necessitates device removal, even if the response is minimal [42]. Since allergic reactions have been described for certain materials, each patient receiving an implant should undergo testing to identify sensitized materials.

The most common allergenic metals and alloys include nickel, cobalt, chromium, zinc, beryllium, mercury, copper, silver, and gold [42,64]. However, gold is commonly used in BMIs due to its electrical properties. When designing a BMI, all components of the BMI, including polymer coatings or materials used to support the electrode (encapsulation materials), must be biocompatible, not just the electrodes themselves. A detailed list of used neural interface materials and their potential for immune reaction is listed in Table 1.

2.8 Contemporary electrode designs

The ideal neural interface has as many contacts for recording or stimulation as required by the neural system, and it does not elicit any tissue reaction acutely or chronically after implantation. Despite many efforts to build the perfect neural interface, this currently does not exist. The following subsections review and discuss the major developments toward achieving this goal.

2.9 Utah Electrode Array

The Utah Electrode Array (UEA) has become a standard neural interface in basic science research. It has a high signal-to-noise ratio (SNR) and well-documented safety [73]. The interface has been used in humans since 2004 with no reported severe adverse events, such as death, hospitalization, congenital anomalies, or interventions to prevent permanent impairment [73]. The device contains up to 100 electrically conductive channels, with the potential to customize the device to add new channels. Each silicon needle on the array has a metal tip, frequently made of iridium oxide or platinum [74]. Standard lengths of the electrode needles for neuroscience research and clinical use are 0.5–1.5 mm and 1.0–1.5 mm, respectively. A standard UEA design can be seen in Fig. 2 [75].Fig. 2 The Utah Electrode Array contains 100, 1.5 mm long penetrating microelectrodes (Reproduced and modified with permission from Fernandez et al., 2016; [75]).

Fig. 2

The number of needles the UEA has, would intuitively enhance recording and stimulation at the BMI. A study by Kelly et al. (2007) compared a 10x10 UEA to a single electrode implant. They found that the single electrode provided a higher SNR than the UEA. Combining the recordings obtained with the UEA on the separate needles compares with the recordings with the single probe. The study also found that the UEA can record from neuron pairs and allows for the study of neural connections that the single electrode cannot achieve [76]. In vivo recordings have shown mean SNR values of 23 dB [77] within the cortex of rats for six months. In addition to recording neural activity with the UEA, Branner et al., 2001 showed that it also allows stimulation of the sciatic nerve of cats at low current intensities [78].

Although commonly used, the UEA promotes immune reactions, neural death, and glial scarring. The tissue damage is related to the device's 100 needles. A study by Patel et al. (2023) showed that Utah Slanted Electrode Arrays (USEAs) implanted for 590 and 848 days reduced the neuron population by 63 % [73]. Other studies using the USEAs over 502 days found that the SNR did not significantly change during the experiment. Still, functional recording sites decreased in number within two months of insertion [79]. A meta-analysis completed by Sponheim et al. (2021) analyzed over 6000 datasets and found that UEAs have an average functional lifespan of 622 days. Some devices lasted over nine years [80].

The studies mentioned above showed that over time, the UEA lost function. However, it is important also to point out their relative sustainability of months to years, resulting in the FDA approval of chronic human studies (designated as over 30 days) of neuron recording and stimulation.

2.10 Michigan probes

The Michigan probe (NeuroNexus) is a single silicon shank often insulated with silicon dioxide, consisting of 16 conductive sites for recording [81]. The individual electrodes in the Michigan probes are 2–15 mm long compared to Utah arrays of 0.5–1.5 mm. The difference in design allows the Michigan to be inserted deeper into the brain tissue [81,82]. A NeuroNexus probe can be seen below in Fig. 3.Fig. 3 The 16-channel silicon-based microelectrode array is an example from Neuronexus Technologies.

Fig. 3

Although not FDA-approved for human implantation, many animal studies support the in vivo use of the Michigan probe. A work by Vetter et al. (2004) used Michigan-like planar silicon electrodes to record neural activity from the cortex of 10 rodents. They found that 90 % of their probes could record neural activity for up to 127 days with sufficient SNRs, 13.1 dB–24.4 dB. The corresponding average noise levels were between 9.7 μV and 16.6 μV. The histology showed fibrous tissue formation near the implant but did not show significant glial fibrillary acidic protein-positive (GFAP+) staining. This indicates little glial scarring from the tissue reaction [83]. Other in vivo trials used these electrodes to record and stimulate neural activity following a severe spinal cord injury in rats [84].

2.11 Microwire electrodes

Designs vary from singular to a few (tetrad array) wires. Most materials employed include stainless steel [85], silicon [86], gold [87,88], tungsten [89], carbon [90], platinum [91], silver/silver chloride [92], and Parylene-C [92]. The major advantage of wire electrodes is their flexibility. However, their flexibility imposes difficulty in crossing dural structures without bending or breaking.

The impedance values and the SNR of BMIs using microwires are similar to the UEA devices [22,93]. For instance, Suner et al., 2005 compared the UEA and microwires, obtaining mean SNRs of 13.6 dB and 14.8 dB, respectively [94].

2.12 Neuropixel probes

Neuropixel probes are fully integrated silicon digital neural interfaces with on-chip circuitry for signal conditioning and digitization Fig. 4 [5,95]. These devices have nearly 1000 recording sites along the singular probe. When implanted simultaneously, two probes could record responses from over 700 individual neurons [95]. Neuropixel devices provide benefits in studying neural systems by monitoring neighboring areas and their connections. These small electrodes, with a wide recording surface area, have furthered the research of neural networks and have already been used in humans. Unlike the UEA, which often requires a latent period following insertion to start recording, these probes recorded neural activity within 1 min after implantation. The electrodes were only left implanted to record for about 30 min in humans. Their SNRs ranged from 8.6 dB to 16.7 dB [96].Fig. 4 Neuropixel Electrode. A) Labeled schematic of Neuropixel electrode B) Image of Multi shank Neuropixel electrode 2.0 with silicon cap. D, (Schematic image was reproduced and modified with permission from Jun et al., 2017) [95].

Fig. 4

3 Typical Reactions in Electrode Implantation

3.1 Foreign body response to implanted electrodes

Foreign body responses are triggered regardless of the material used in the BMI. Astrocytes, microglia, and oligodendrocytes are major cellular contributors to this multifaceted response. They are known to be present in both acute and chronic responses. For a better understanding, we provide a short overview of the cascades of events that might trigger cell death and tissue damage after BMI implantation.

3.2 Mitochondrial-independent cell damage and apoptosis

Inflammatory signaling pathways, most commonly the nuclear-factor kappa B (NF-κB), mitogen-activated protein kinases (MAPKs), and JAK-STAT pathways, are activated by BMI implantation. NF-κB is a transcription factor mediating inflammatory response by regulating the expression of various pro-inflammatory genes such as tumor necrosis factor-α (TNF-α). NF-κB translocation can occur much earlier (1–2 h) than maximal reactive oxygen species (ROS) generation (1 day), suggesting an early involvement of the inflammatory pathways or even an upstream event of ROS formation [[97], [98], [99], [100], [101]]. TLR4 is a transmembrane protein that plays a fundamental role in pathogen recognition and activation of innate immunity. TLR4 activates proinflammatory cytokines such as interleukin-6 (IL-6), IL-8, and NF-κB, etc. [102]. Translocation of NF-κB to the nucleus, mediated by TNF-a and IL-6, further induces the de novo synthesis of TNF-a, IL-6, IL-1b, and iNOS as well as activates caspase-3 in a mitochondrial-independent manner [98,103]. Expression of TNF-α can also further activate NF-κB [102,104] and the extrinsic apoptosis pathway by binding to its receptor, TNFR. It leads to caspase-8 activation, which activates caspase-3, leading to mitochondrial-independent apoptosis (Fig. 6) [105,106]. This is also known as the death receptor pathway [107] or extrinsic apoptosis [108].

The activation of MAPK/ERK also facilitates the secretion of the pre-existing TNF-a, IL-1b, and IL-6 [109] (Fig. 5), which in turn activates the translocation of NF-kB to the nucleus [100].Fig. 5 Schematic of multi-faceted inflammatory response that promotes tissue damage and neuronal death following BMI implantation.

Fig. 5

Fig. 6 Mechanism of mitochondrial-dependent and mitochondrial-independent apoptosis after cellular stress pathways are initiated due to BMI implantation.

Fig. 6

STAT1 [110] and NF-kB [111] induce increased expression of iNOS, which produces nitric oxide (NO) [for reviews see Refs.[98,103]. Superoxide (O2−*) can react with NO to form peroxynitrite (ONOO−), a highly reactive oxidizing molecule that damages proteins [112]. Peroxynitrite induces protein peroxidation via nitration of tyrosine residues (nitrotyrosine) [112,113], which alters protein configuration and function [for reviews, see Refs.[97,114,115]. Downstream effects of superoxide can also cause lipid peroxidation. Catalyzed by superoxide dismutase (SOD), O2−* is converted to hydrogen peroxide, which is catalyzed by iron to form hydroxyl free radicals [97]. These highly reactive ROS react with polyunsaturated fatty acids in cellular membranes, producing the highly toxic aldehyde, 4-hydroxynonenal (4-HNE), and malondialdehyde (MDA) [114]. Antioxidant enzyme depletion has been linked to increased levels of MDA, indicating high lipid peroxidation [116]. Lipid peroxidation, especially the production of 4-HNE, induces calcium influx into the cell [97,117]. ROS further enhances calcium influx. According to a review article, ROS can open the endoplasmic reticulum (ER) calcium channel, RyR, L-type, T-type, and the TRPV1 plasma membrane calcium channels [98]. A study by Yoshida et al. found that intracellular NO and H2O2 activate TRPV1 channel opening and subsequent calcium influx (Yoshida 2006 nitric oxide activates TRP channels).

3.3 Mitochondrial-dependent cell damage and apoptosis

Mitochondrial-dependent cell and tissue damage, including apoptosis, originates in releasing pro-apoptotic factors from the mitochondria to the cytosol. This process can be triggered by BMI implantation and is associated with different factors/pathways, including lipid peroxidation, DNA damage, signal transducer and activator of transcription 1 (STAT1), p53 activation, B-cell lymphoma 2 gene (Bcl2)/Bcl2 Associated X (Bax) activation, cyclin/cyclin-dependent kinase (CDK) activation, Cytochrome C, Caspases, etc. (Fig. 6).

DNA damage activates p53, initiating the intrinsic mitochondrial apoptosis pathway [118]. P53 increases the expression of the pro-apoptotic molecule Bax [103], which translocates to the mitochondria, where it permeates the outer mitochondrial membrane. Mitochondrial membrane permeabilization leads to the loss of the mitochondrial membrane potential, ROS production, and release of Cytochrome C and ROS from the mitochondria into the cytoplasm [119,120]. NOX3-dependent ROS production [98,110,117,121,122] and ROS-mediated activation of extracellular signal-regulated kinases 1 (ERK1) [99,123] activates STAT1 signaling. Cyclin A is another critical player for mtROS production, which can be upregulated after BMI implantation and activates CDK2 kinase, consequently facilitating mtROS production [124].

DNA damage activates p53, initiating the intrinsic mitochondrial apoptosis pathway [118]. P53 increases the expression of the pro-apoptotic molecule Bax [103], which is translocated to the mitochondria, where it permeates the outer mitochondrial membrane. Mitochondrial membrane permeabilization leads to the loss of the mitochondrial membrane potential, ROS production, and release of Cytochrome C and ROS from the mitochondria into the cytoplasm [119,120]. NOX3-dependent ROS production [98,110,117,121,122] and ROS-mediated activation of extracellular signal-regulated kinases 1 (ERK1) [99,123] activates STAT1 signaling.

As an essential component of the mitochondrial electron transportation chain, cytochrome C is an apoptotic protease activating factor that activates caspase 9. Caspase 9 activation subsequently activates caspase 3, causing the fragmentation of chromosomal DNA through the cleavage of its substrates [125]. Cytochrome C release and activating caspase-9 and caspase-3 mark the final stage of mitochondrial-dependent apoptosis [114] (Fig. 5).

Apoptosis can also be induced in the cytoplasm through ROS overload, inflammation, lipid peroxidation [98], and endoplasmic reticulum (ER) stress [126]. ER stress is induced by oxidative damage, intracellular calcium imbalances, and protein damage [127]. Caspase-12 then activates caspase-9, which activates caspase-3, leading to apoptosis in an extrinsic- and intrinsic-independent mitochondrial-independent manner [121,128,129]. This ER-specific apoptosis pathway is also closely linked to the mitochondrial-dependent path due to the simultaneous activation of C/EBP homologous protein (CHOP) [126]. CHOP plays an important role in ER stress-induced apoptosis while regulating the Bcl2 family expression. Decreased Bcl2 enables increased Bax activity, releasing apoptotic active substances from mitochondria to the cytoplasm [130] (Fig. 6).

3.4 Other cell death pathways

TNF-a induces cell death via necroptosis pathway activation [131,132] [for reviews, see Refs.[99,133]. Apoptosis occurred at lower doses, while necroptosis occurred at higher doses. As opposed to the organized breakdown of cells during apoptotic cell death, necroptosis results from cellular, as well as organelle, membrane permeation, leading to the release of intracellular substances and exacerbating inflammation [108]. Activation of caspase-8 inactivates RIPK1 (receptor-interacting serine/threonine-protein kinase 1) and RIPK3 (receptor-interacting serine/threonine-protein kinase 3), which induces activation of the extrinsic mitochondrial-independent apoptosis pathway as described above. The inactivation of caspase-8 leads to the activation of RIPK3, shifting death receptor-mediated cell death from apoptotic pathways to necroptotic pathways [133]. In necroptosis, TNF-a binds its receptor (TNFR1) and, without caspase-8, leads to the formation of the RIPK1/RIPK3 complex, also known as necrosome [131,133]. Downstream effects of necrosome involve activation of MLKL, which induces calcium influx via TRPM7 channels and creates pores in the plasma membrane, leading to leakage of substances, cell lysis, and cell death [134].

Autophagy differs from apoptosis because it is caspase-independent and characterized by protein, mitochondrial, ER degradation, cytoplasmic vacuolization, and eventual cellular component digestion [135,136]. Although oxidative-, nitrosative-, and ER stress can stimulate both autophagy and apoptosis, autophagy does not lead to apoptosis, suggesting separate cell death pathways [98,99,135].

Like necroptosis, autophagy occurs during excessive cellular stress without caspase activation [136]. Under normal conditions, autophagy is cytoprotective by removing damaged cellular components. Under stress conditions, autophagy can lead to cell death [135]. The cascade of damage starts with intracellular organelle or protein damage, resulting in AMPK activation, which inhibits mTOR, leading to autophagy induction [137]. This leads to the activation of PI3K, inducing Beclin-1 formation of the phagosome [136]. Beclin-1 is the initial autophagy promoter [137]. This activates the conversion of unlipidated LC3-I to lipidated LC3-II. Increased LC3-II levels correlate with increased autophagosome amount [137]. LC3-II is required for complete autophagosome formation and autophagosome-lysosome fusion [135,136].

Ferroptosis is another form of non-apoptotic programmed cell death. Cellular iron content promotes ferroptosis, while autophagy generates the necessary iron [138]. Ferroptosis also depends on the chain of reactions leading to lipid degradation (lipid peroxidation) [139]. In lipid peroxidation, ROS attack -CH2- of polyunsaturated fatty acids deposited in the cell membrane and causes cell injury [140]. Ferroptosis takes place when GSH-dependent antioxidant capability is impaired [141]. Erastin and RSL3 are known to cause ferroptosis. Erastin inhibits glutamate-to-cysteine conversion, while RSL3 inhibits GPX4, a lipid repair enzyme. Differentiation of ferroptosis from necroptosis is due to cell death without RIPK1/RIPK3, and its differentiation from apoptosis is due to cell death without caspases and BAX [139].

Pyroptosis is a programmed, pro-inflammatory cell death. DFNA5, NLRP3, and Gasdermin family members, such as GSDMD, produce “inflammasome”, resulting in pyroptosis [142]. Caspase-8 can cleave GSDMD to produce IL-1β and IL-18, while DFNA5 can be cleaved by caspase-3 to cause pyroptotic cell death [143,144]. Pyroptosis causes flattening of cells, unlike apoptosis and necroptosis [145].

3.5 Acute glial cellular responses: Microglia, astrocytes, and oligodendrocytes

Microglial cells are a group of specialized phagocytic macrophages in the central nervous system that act within 30 min in the acute response to pathogens and cell injury, with the eventual release of pro-inflammatory cytokines, including IL-1β, IL-6, IL-12, TNF-a, IFN-y, and ROS [[146], [147], [148], [149]].

Astrocytes, without an inflammatory response, provide structural support for neuronal circuits and regulate neuronal synapses, the extracellular environment, and the fluid composition within the central nervous system (CNS) [150,151]. Pro-inflammatory states activate astrocytes, resulting in their hypertrophy and proliferation. The process is identifiable through elevated GFAP expression [150]. Growth of the activated astrocytic processes causes attachment to BMIs, which can eventually encapsulate it, forming a glial scar. In case the inflammation is chronic and not acute, astrocytes will form a scar [152].

Oligodendrocytes have a primary role separate from inflammatory responses in promoting myelination within the CNS [3]. Although recent research addresses the role of the cells in reactions to BMIs, their exact role has yet to be determined [153].

Mild to moderate reactive astrogliosis comprises variable changes in molecular expression and functional activity together with variable degrees of cellular hypertrophy. Such changes occur after mild trauma or at sites distant from a more severe injury or after moderate metabolic or molecular insults, milder infections, or inflammatory activation.

Severe reactive astrogliosis with persisting scar formation generally occurs along borders to areas of overt cell and tissue damage and inflammation. Fig. 7 provides a scheme for glial scar formation, including activated microglia in a matter of hours following implantation (orange), neovascularization of prominent vessels within the tissue (red), and newly proliferated astrocytes (green).Fig. 7 Illustration of the glial encapsulation response (A) before implantation, (B) 12 h post-implantation, (C) 1 week post-implantation, (D) 4 weeks post-implantation, and (E) 12 weeks post-implantation. Panels (F), (G), and (H) represent cross-sectional views of (C), (D), and (E), respectively (Reproduced with permission from Cambell and Wu (2018) [46].

Fig. 7

Astrocytes no longer occupy discrete domains in the mature glial scar but have overlapping processes. Mature glial scars persist for long periods and act as barriers to axon regeneration and inflammatory cells in a manner that protects healthy tissue from nearby areas of intense inflammation. A simplified schematic of the inflammatory cascade following BMI implantation can be seen below in Fig. 8.Fig. 8 Schematic of the inflammatory cascade following BMI implantation.

Fig. 8

3.6 Implantation trauma and blood-brain barrier (BBB) leakage

The neural interface can directly damage the neuronal and vascular structures needed for functionality within the nervous tissue. Markers for vascular damage include proteins, such as albumin, immunoglobulins, fibrinogen, thrombin, complement, and products of red blood cells (RBCs). They leak into the extracellular space and trigger inflammatory pathways [46,146,154,155]. Specifically, albumin can activate both microglia and astrocytes through the MAPK pathway, which results in downstream production of IL-1β and ROS. The production of the pyrogen IL-1β promotes more leakage in the BBB and entry of other cells mediating inflammation, including T-cells and lymphocytes promoting astrogliosis [11,154,156]. Furthermore, IL-1β is part of the apoptosis pathway, and its activation leads to the loss of neurons.

3.7 The complement cascade

The complement cascade is part of the innate immune system. Its activation by pathogens or foreign bodies triggers inflammatory pathways with the production and release of anaphylatoxin. Three of the four activation pathways are intrinsic; one is extrinsic, following fibrinolysis and blood clotting activation. For example, implantation of the UEA activated the complement cascade, including C1q, and promoted the recruitment of other inflammatory mediators [146]. The complement proteins, C1q, C3, and C5, activated the alternative and classical complement pathways on the BMI as a foreign body reaction [146]. The source of acute inflammatory infiltration around the BMI likely originated in the production of acute inflammatory markers, like IL-1β and anaphylatoxins from the complement system. However, the role of complement proteins on the tissue reaction beyond four weeks after implantation has yet to be studied [146].

Fibrinogen persists at the neural interface for more than eight weeks and most likely has a role in maintaining the foreign response, resulting in neural interface failure [154].

3.8 Chronic inflammation and electrode failure

A primary concern following the electrode insertion is glial scar formation, as inflammation persists via reactive gliosis. As prefaced, microglial activation is quick following implantation, creating an inflammatory environment that persists within the CNS and promotes astrocyte hypertrophy and proliferation [46,157,158]. This can be seen through the study by Ravikumar et al. (2014), which found a positive relationship between microglial infiltration and neurodegeneration, with the lowest densities of neurons 16 weeks after implantation [159]. As resident macrophages of the CNS, these cells have direct negative impacts on neuron survival through inflammatory damage and promotion of astrogliosis and scar formation over time, as seen in Fig. 7 [46].

Microglia activation immediately leads to downstream astrocyte hypertrophy. However, glial scar formation does not occur right away. It is the presence of a chronic inflammatory response or severe trauma to the tissue that triggers scar formation. Scars increase electrode impedance, affecting neural recordings and stimulation [152,159,160]. Budoff et al. described a continuous progression in glial scar formation between day three and seven months after implantation [161]. Furthermore, Thelin et al. (2011) chronically implanted electrodes with a diameter of 200 μm. Thelin's study showed that the device's size and implantation length are directly linked with inflammation and glial scarring. Twelve weeks after implantation, the tissue around the electrodes showed increased GFAP and decreased NeuN staining, which was used for identifying mature neurons, compared to 6 weeks after implantation [162].

3.9 Electrode coatings

Strategies that attract, attach, or preserve neurons at the BMI increase the ability to record and stimulate neural activity [160]. Activating the foreign body response to the implant cannot be avoided entirely. Efforts to minimize damage by the BMI have focused on coating electrodes with bioactive molecules. The following sections describe electrode coatings with polypeptide and extracellular matrix (ECM), electrically conductive polymers, anti-oxidative, and pharmaceutical coatings.

3.10 Polypeptide and ECM protein coatings

Synthetic polypeptide coatings have been used to facilitate neural growth and neural adhesion. For example, poly-d-lysine (PDL), a synthetic polypeptide, enhances neural-cell adhesion and proliferation. Silicon electrodes coated with multi-layers of polyethyleneimine and laminin (PEI-LN) significantly promoted neuron growth and differentiation onto the electrode surface [59,163,164]. In contrast to those promising results, studies have also found that PDL and PEI are toxic at certain concentrations, and further studies are required to determine the safe amount of the synthetic peptides that can be used in coatings [165].

Neuroligin-1 is an endogenous protein that plays a role in neuronal synapse formation and stability. Bioengineered synthetic neuroligin-1 (eNG-1) coated electrodes have established a neuronal connection at the BMI, which needs to be maintained for the long-term function of the electrodes [166,167]. Its success has initiated further exploration of synthetically produced proteins like neurexins and leucine-rich repeat transmembrane proteins [168].

Moreover, coating the electrodes with surface proteins may promote the continual use of the electrode by introducing resident peptides to diminish glial expansion [[169], [170], [171]]. Covalent modification of the probes, involving the brain-derived neuronal specific cell adhesion molecule L1, can reduce the foreign body reaction to the probe and enhance the recordings' durability and quality [169,170]. It has been shown that L1 on a silicon probe had diminished activation of pro-inflammatory microglia cells from 8 to 16 weeks post-implantation in the cortex of mice [169,170]. L1 on neural probes, in combination with thiol-functionalized silica nanoparticles (TNP), offered significant benefits in reducing oxidative stress and inflammatory cytokines over eight weeks in rats [171]. Inflammatory cytokines were reduced, and anti-inflammatory proteins increased. The antioxidant effects counteracted microglia activation using the L1 protein [171].

3.11 Electrically conductive coatings

Electrically conductive polymers of polypyrrole and poly (3,4-ethylenedioxythiophene) (PEDOT) are also able to enhance the contact between the interface and adjacent neuronal tissue [56,57]. By incorporating the YIGSR peptide fragment from laminin into the polypyrrole coating, the recording sites supported more neuron and astrocyte attachment in vitro and in vivo than peptide-free sites. However, the peptide-incorporated and peptide-free electrodes produced similar recordings and exhibited similar glial scar reactions [172]. Incorporating other peptides like sericin and gelatin into PEDOT/Poly-vinyl-alcohol (PVA) conductive hydrogels (CH) provided neuro-adhesive effects with decreased impedance. Adding sericin to the CH provided a greater net negative charge to the BMI, which is thought to be the mechanism behind its greater efficiency compared to CH with gelatin. Sericin could also significantly increase neurite density while maintaining charge transfer, validating its use alongside electrically conductive polymers and hydrogels like PEDOT/PVA [173]. Similarly, the long chain and conducting polymer (poly (styrene sulfonate-co-4-vinyl pyridine) (Poly (SS-4VP)) and PEDOT have been chemically cross-linked to fabricate a biocompatible platinum wire with a PEDOT:Poly(SS-4VP) hydrogel coating. The coating was compared to PEDOT:PSS and bare platinum wire controls and showed decreased GFAP markers after implantation, enhanced neurite viability and attachment, and lowered impedance in recordings over a four-week in vivo study in the hippocampus of rats. It is important to highlight that PEDOT-coated electrodes and electrodes made of PEDOT:PSS materials differ. PEDOT coating is used to help with implanting microwires and reducing inflammation after that. Making electrodes with PEDOT:PSS uses conductive polymers. Note that the material is also crucial as a supporting substrate for the electrodes. In this study, PEDOT:PSS forms the BMI, not its coating. PEDOT was used as a coating. Regardless, this study showed the potential for coatings to open gateways using conventional metal electrodes like platinum and tungsten while providing a biocompatible device [174].

PVA and poly-(lactic-co-glycolic acid) (PLGA) are two FDA-approved biodegradable polymers, with the former used as a transient adhesive to neuronal tissues and the latter utilized to stiffen the microelectrode for implantation [15]. Using PVA and PLGA may be an approach for implanting a perfect device and promoting neurogenesis near the surface. Similarly, a study by Capeletti et al., 2016 found the use of a silica-based polymer coating prepared via a sol-gel process was able to both augment neuron growth and diminish astrocyte activation through the addition of amino groups with (3-aminopropyl) triethoxysilane [175]. It was concluded that the positive charge on the amino groups was the main contributor to the beneficial effects [175].

A hydrophilic polymer poly(ethylene glycol) (PEG) surface coating has been utilized to provide resistance to nonspecific protein adsorption and cell adhesion [[176], [177], [178]]. The PEG block copolymer coating showed a 95 % reduction in serum protein adsorption compared to unmodified control surfaces [60].

Further research into the benefits of PEG hydrogels compared uncoated vs coated electrodes and their effect at two different diameters. PEG hydrogel-coated electrodes with diameters of 400 μm showed significantly decreased tissue strain with lateral movements compared to 200 μm coated and uncoated controls. The 200 μm PEG-coated electrodes had significantly decreased tissue strain compared to other samples when undergoing in and out motions, meaning the PEG hydrogel may provide benefits in inhibiting glial cell adhesion triggered by micromotions in chronically implanted probes [179]. Similar effects have been seen in vitro, using dextran DLC-poly-lysine coatings, but these findings have yet to be investigated in vivo [180]. Potential limitations in recording and stimulation originate from an increased distance in the neuron-electrode interface by any coating, such as PEG and DLC poly-lysine.

Among other advancements, using lubricated immune-stealthy probe surface (LIPS) coated electrodes has been studied in microelectrode development. LIPS-coated electrodes utilize a hydrogel or polymer coating that is both lubricious (to reduce friction) and bioinert (to minimize protein adsorption and cell adhesion). The LIPS electrode demonstrated an SNR double that of the platinum control electrode over 16 weeks. LIPS-coated and bare platinum electrodes were placed in samples of blood and albumin, resulting in microglial activation and encapsulation of the bare electrode. The LIPS coating showed >95 % less adhesions, and the necessary insertion force for the implantation of the LIPS probe was seven times less than that of the control [181].

3.12 Anti-oxidative coatings

Blood-brain barrier damage leads to a larger presence of iron within the CNS, producing reactive oxygen and nitrogen species (RONS) following microglial activation. The production of RONS is directly involved in the downstream diminished function of microelectrodes through their pro-inflammatory effects. Simple iron chelator coatings like deferoxamine (DFX) have seldom been used but have found success in chronic implantation (>7 days) [182]. Besides iron-chelating drugs, the ability to coat the implanted probe with immobilized superoxide dismutase (iSODm) provides intrinsic antioxidant properties that reduce the oxidative stress from RONS within 25 μm of the implanted electrode. iSODm-coated electrodes can minimize damage to the intracranial capillaries in acute implantation models and thus dampen glial cell activation by decreasing leakage across the blood-brain barrier. Further, the iSODm coating may potentially promote neuronal growth and adhesion through its positively charged amine groups in a similar method as poly-lysine previously discussed. The use of iSODm, though, requires further investigation as the use of an iSODm monolayer in chronic implantation beyond one week and its effects greater than 25 μm from the probe has yet to be seen [183].

With the known benefits of antioxidants for protecting neural probes seen in literature, producing biocompatible and noncytotoxic compounds may have a considerable impact. Synthetic production of such a compound with antioxidative properties is seen in the fabrication of poly(triethylene glycol methyl acrylate-co-α-tocopheryl acrylate) (PVT). PVT has shown cytoprotective effects against hydrogen peroxide for astrocytes in in vitro studies. It gives hope for future research with PVT to show its potential neuroprotective effects from radical damage in vivo [184].

3.13 Nanotube metal coatings

Some of the most innovative advances towards a biocompatible implant include nanotube coatings. These nanotube coatings came to the forefront as they were found to increase the available surface area of a BMI. Carbon nanotubes, in particular, are the most popular nanosized coatings used in BMIs. Such nanotubes have been nicely characterized [185]. In vivo studies have already been completed using standard tungsten wires with nanotube coatings. These studies have shown that the impedance values compared to non-coated controls decreased 23 times at 1 kHz while charge transfer increased 45 times [186]. Small pure carbon, or graphene, nanotube sheets have also been shown to have intrinsic anti-inflammatory properties, which are thought to be through inhibiting TLR-4 stimulation and its downstream inflammatory pathway that was previously described [187]. Another frequently used nanoparticle is gold. Won et al. (2022) used a polymer-based shank embedded in gold nanoparticles for chronic in vivo brain recordings. The probe recorded spontaneous and evoked neural signals for up to four months. It was also found to have significantly lower GFAP staining around the insertion point than a stainless steel electrode at 2 and 6 weeks post-implantation. Iba1 levels were also significantly decreased, most likely due to the reduced flexibility of the stainless steel wire [188]. Although nanoparticle-based probes have been around for a while, their newfound abilities to reduce inflammation and enhance neural signals have brought them to the forefront of research in the BMI.

3.14 Pharmaceutical coatings

Pharmaceutical drug coatings and physical and mechanical modifications have been studied to minimize protein and effector cell activation that causes glial scar formation. Anti-inflammatory drug-eluting coatings feature wound healing suppression and immunosuppression, which can mitigate the initial immune response and perhaps even the glial scar formation. Acute and chronic inflammatory reactions can be drastically reduced by dipping the neural interface in dexamethasone before implantation or by coating the interface with materials containing and releasing dexamethasone [189,190]. Steroids like dexamethasone are standardly used to diminish microglial pro-inflammatory effects and attenuate the formation of potential fibrous capsules [[189], [190], [191]]. Nitrocellulose coatings loaded with dexamethasone on Michigan probes reduced reactive astrocyte and microglia numbers [192]. Apart from the slow-releasing dexamethasone, this coating also contained neuro-adhesive 10.13039/100007084 PEI and laminin to support neuronal and glial attachment and growth. PEDOT coatings, which actively deliver dexamethasone, might have neuron survival close to the microelectrode [189].

Systemic administration of corticosteroids reduced the wound healing response for CNS implants [193,194]. Dexamethasone doses given to cats 12 h before and after implantation surgery for cortical edema management showed considerable improvements in implant performance. Peripheral dexamethasone injection during electrode implantation also resulted in less glial scar formation [195]. However, the use of systemic and peripheral dexamethasone injection requires further investigation due to the large dose needed.

Other drugs, such as antibiotics, hormone therapies, and ion exchange inhibitors, have been beneficial [149,192,196]. Pharmaceuticals may provide the extra anti-inflammatory effects needed for chronic electrode use and brain tissue preservation if combined with physical and mechanical modifications. A Teflon sheet might contribute to BMI success as an alternative to pharmaceuticals because it prevents adhesions between the underlying dura and the neural probe following implantation [197].

Minocycline is a broad-spectrum protein synthesis inhibiting antibiotic. It also has well-known effects on downregulating the release of pro-inflammatory cytokines like IBA1 produced following microglial activation up to four weeks after the procedure. Minocycline has been used to reduce the inflammatory response to electrodes in the brain. Interestingly, studies have shown that the drug advantageously minimizes the effects of microglial cells, but GFAP levels remain elevated, suggesting that it does not affect astrocyte response [149].

The α-melanocyte stimulating hormone (α-MSH) is a neuropeptide that inhibits proinflammatory cytokine production and related mediators of inflammation, allowing it to reduce reactive gliosis [198]. The α-MSH has been incorporated into nitrocellulose coatings, released for over 21 days, and inhibited nitric oxide production in microglia in cell culture. Investigations using α-MSH alongside standard silicon electrodes have yet to be studied entirely in vivo [198].

Another hormone, melatonin, has been used in intraperitoneal injections in chronic microelectrode implantation, which has shown similar beneficial effects in antioxidant and anti-inflammatory effects compared to other compounds like corticosteroids and minocycline. The use of melatonin, however, has been found only to have these effects with sustained, chronic injection as the effects of melatonin were diminished in acute treatment (4 weeks) groups compared to those with continued therapy after four weeks [199].

Pro-inflammatory cells like microglia depend on NHE-1, a 1:1 Na+/H+ exchanger, to maintain an optimal environmental pH for producing oxygen radicals in its primary role as a defense mechanism in the CNS. HOE-642, also known as cariporide, is an NHE-1 inhibitor with significant inhibitory effects on microglia activation and neural probe encapsulation when used post-implantation. Immune cell expansion over the electrode was significantly reduced 5 h after implantation using HOE-642. Also, microglial activation was significantly reduced up to 72 h after implantation. The benefits of HOE-642 were seen in the initial and chronic neural probe insertion, making it a novel approach deserving of future investigation [196]. The various coatings used to prolong the fidelity of the implanted electrodes are summarized in Table 2. A visual display of the mechanisms of the different coatings that were discussed can be seen below in Fig. 9. Some coatings act in several mechanisms, while others simply limit implantation trauma. Regardless of the mechanism, all aim to limit device failure.Fig. 9 Mechanisms of electrode coatings in promoting chronic implantation of the BMI. These mechanisms can be further visualized in Table 2 below.

Fig. 9

3.15 Emerging technologies

The emerging technologies in BMI design are progressing toward interfaces that closely resemble biological tissues. This includes the integration of electronics and probes created to be flexible, stretchable, and comparably soft to the brain's tissue. Such innovations allow seamless integration of the BMI with the brain, decreasing disruption and enhancing long-term compatibility. Key to this progress is thin, flexible electronics in several configurations (thin films, fiber probes, and 3D mesh), an advancement that has shown promising results in reducing chronic mechanical damage and immune responses.

For instance, Khodagholy et al. created an ultrathin organic biocompatible film substrate, denoted as “Neurogrid”, capable of detecting single-cell action potentials from the brain's surface [200]. The BMI comprises a ‘soft’ substrate with a high mechanical compatibility with the brain tissue. In contrast to ‘hard’ electronics, soft substrates are minimally invasive, reducing the risk of trauma and immune responses. Following this, Guo et al. proposed a micro-electro-mechanical systems (MEMS) BMI with an ultrathin polyimide film developed for low invasive implantation. The device showed a diminished bending stiffness to suppress inflammatory responses and caused little trauma [201]. Immunohistochemical analysis showed that the thinner substrate reduced neuronal cell loss and reduced immune reaction. Similar approaches have been conducted using other flexible substrates, such as parylene and the photoresist SU-8 [202,203].

Fiber-like BMIs have also been developed. In this context, a recent study proposed graphene microwires engineered with a platinum coating [204]. These microwires showed biocompatibility along with unique mechanical and electrochemical characteristics. The platinum coating reduces the impedance of the microwires, and with its porosity, this strategy allows the detection of neural activity over a smaller region in the tissue [205]. Ferro et al. proposed an alternative design for flexible microwires called "NeuroRoots". This device comprises individual roots made from a polymer-metal/polymer structure. These roots have a different design compared to traditional BMIs. Each root is independent in its movement, providing flexibility with respect to its adjacencies [206]. This enhances long-term stability while decreasing the immune response. Similar designs were proposed to mimic biological structures. For example, a design for neural probes, known as neuron-like electronics (NeuE) that resembles the subcellular structural features and mechanical properties of neurons showed negligible immune response and revealed continuous interpenetration interfaces with the brain after implantation. Furthermore, these fiber-like probes showed a conducive substrate for the migration of newborn neurons, indicating its potential to guide these cells toward areas of injury for tissue regeneration [207].

In line with this advancement in BMI development, Liu et al. proposed mesh-like electronics in which sensors are linked with flexible connectors and coated in sub-micrometer-thick encapsulation layers [208]. This BMI resembles the architecture of a tissue scaffold, which reduces the bending stiffness and permits the unobstructed movement of chemicals and biological agents throughout the implanted BMI, improving its integration and functionality in biological systems while lowering the chronic immune response.

To design stretchable mesh BMIs that can adapt to the volume variation of neural tissue during development, growth, aging, and disease, Li et al. proposed a cyborg brain organoid platform. This work created an ultrathin mesh-like tissue with nanoelectronics from the photoresist SU-8 and integrated it with human-induced pluripotent stem cell-derived progenitors and neurons through organogenesis [209]. Later works have been conducted by the same team to characterize the electrophysiological properties of these brain organoids [210]. These results showed a combined stretchable mesh nanoelectronic without interfering with the development of brain organoids. This novel BMI adjusted to the changes in size and shape that occur during the formation of brain organoids and maintained stable electrical connections with the neurons inside the brain organoids throughout their development.

New horizons in BMIs are also being considered. For example, Green et al. introduced the concept of living electrodes [211]. In their work, they proposed a soft, cell-integrated electrode coating to address the problem of scar tissue encapsulation in neuroprosthetics. Aligned with this, Adewole et al. [212] introduced an implantable living electrode comprised of cortical neurons and axonal tracts encapsulated within soft hydrogel cylinders designed for optical modulation and monitoring of brain activity.

Further improvements in BMI developments emphasize multifunctional interfaces for multimodal recording and stimulation of different cell types [213]. These interfaces can include stimulation and monitoring electrodes, optical probes, microfluidics channels for drug delivery, and cell gene expression for connectivity information with single-cell electrophysiology. For instance, Canales et al. developed micrometric polymer-based neural probes with multimodal capabilities, including simultaneous optogenetic stimulation, neural recording, and drug delivery in freely moving mice [214]. These probes consisted of two designs. One design incorporated a cylindrical waveguide with two microfluidic channels and two electrodes. The second included one microfluidic channel, four electrodes, and a surrounding waveguide. Results showed a stable interface for at least two months. Keeping with this trend, comparable devices have been reported. For example, Park et al. [215], Jeon et al. [216], and Qazi et al. [217] exhibited optogenetics probes for stimulation, including microfluidics channels and recording electrodes. These microchannels were used for viral vectors carrying opsin genes and pharmacological drugs. Other modalities have also been explored. Yang et al. manufactured a closed-loop BMI for optogenetic stimulation and recording in the mouse brain [218]. Liu et al. [219] designed a stretchable optical array, denoted as a “skin display” that exhibited mechanical features similar to the human skin. Li et al. [220] developed a mesh electronic integrated into human induced pluripotent stem cell-derived cardiomyocyte and a hydrogel for continuous monitoring and in-situ sequencing.

Challenges in BMIs remain on the frontline. Safe, reliable, and effective surgical implantation protocols are still required. One of the main concerns is the surgical placement of these emergent BMIs into the brain without causing them to buckle. This scenario highlights the requirement for a guiding system that allows control during the entire implantation procedure. Recent innovations in surgical procedures include the use of a robot inspired by sewing machines for the insertion of neural threads [221], the implementation of syringe and injection tools [208,222], and anti-blocking systems [223]. The discussed emerging technologies can be seen in Fig. 10.Fig. 10 Emerging technologies for BMIs. These new technologies mimic tissues' mechanical properties and size, reducing mechanical mismatch to enhance long-term biocompatibility and prevent inflammation. A-D] Thin-film flexible BMIs. A-C) Neurogrid interface. A) The interface adapts to the surface of an orchid petal. Scale bar: 5 mm. B) The interface adapts to the surface of the rat somatosensory cortex. Scale bar: 1 mm. C) Action potential signals recorded over a time of 10 days. Blue: hippocampus signals. Black: Cortex signals. Red: Spike detection threshold. Reproduced with permission of Khodagholy et al., 2015. D) Neural implant with ultrathin film substrate. Reproduced with permission of Guo et al., 2022. [E-K] Fiber-like BMIs. E-F) Neuroroots interface. E-F) Microscopic images of the interface. This device consists of arrays of individual electrodes, ≈7 μm wide, ≈1.5 μm thick, organized in axon-like tendrils. G) To insert these roots into the brain, an electrode self-assembly method was developed. This approach uses capillarity to organize the arrays onto a microwire as small as 20 μm in diameter. Reproduced with permission of Ferro et al., 2018. H-K) Neuron-like electronics (NeuE). H) Representation of the NeuE compared to a neuron in green. Electrodes and interconnects are in yellow, and polymer layers are in red. I) Fluorescence image of a neuron (I) and false-colored scanning electron microscope image of two NeuE designs (II and III). Scale bars: 10 μm. J) 3D histology of the NeuE interface and neurons. The figure shows the interface NeuE (red) and neurons (green) after two days post-implantation. K) Bending stiffness of the axons, the NeuE, and previous designs reported in the literature. Reproduced with permission of Yang, Zhou, and Zwang et al., 2019. [L-O] Mesh-like BMI. L-M) The mesh is introduced into the mouse's brain using a syringe. N) Bright-field microscopy image of the mesh inside a glass pipette. The red arrow indicates the ending of the mesh. O) Image showing the structure of the mesh. The red box shows the location of the sensors. The green box indicates the interconnections, and the black box the input and output connectors. P) Representation of the setup for electrophysiology using the mesh-like BMI in cardiac organoids during organogenesis. Q) Images of the organoid growth during 48 h while connected to the recording setup. Reproduced with permission of Liu et al., 2015 and Li, Nan, and Lin et al., 2019 [219,220].

Fig. 10

3.16 Perspectives

Currently, BMIs in everyday clinical practice are used in certain applications, with the most popular being the cochlear implant (CI). Other typical applications include deep brain stimulation (DBS) for diseases like Parkinson's disease and essential tremor, retinal implants for stimulation of photoreceptors, or restoration of motor control in conditions like Amyotrophic Lateral Sclerosis (ALS) or post-stroke debilitations. Although these applications are no short of amazing, the success of some BMI applications brings aspirations to use these same devices in more complex applications like restoring sensation in prosthetic devices, linking cortical planning of movements to actual movement in those with brain-stem lesions, or translating thought into words. Additionally, having access to a BMI that is known to provide information on neuronal connections in any situation would be invaluable. Still, there are many issues to overcome.

Looking at successful applications like the CI, it becomes evident that having the entire system that makes up the BMI, including the sensor, processor, and contact with neural tissue that is both practical and portable, is necessary to be used daily. Today, one of the largest issues lies in maintaining implant-neuron connections within that system. Another issue is that the CI causes intra-cochlear inflammation over time. The implantation process often results in residual hair-cell loss, damage to spiral ganglion neurons, and fibrosis around the effective area of the implant, resulting in increased impedance and poor outcomes. Implantation of a new cochlear implant is often necessary but not plausible due to excessive intracochlear fibrosis and osteoneogenesis [224]. Research towards understanding the processes within the cochlea and how to minimize them would advance the usage of the world's most popularly used BMI.

The retinal implant is another popular BMI application for conditions like retinitis pigmentosa or macular degeneration. Recent research toward better understanding effective pulse durations to minimize unintended stimulation has been conducted. The study found that extending pulse durations of the electrical stimulus by almost double that of what was normal resulted in better spatial resolution for visual stimulation [225]. Although the principle of these BMIs is successful, there are still many more things to consider. Future research on all BMI applications should look further into the following:1. Matching mechanical properties of electrodes to the tissue to be implanted to minimize tissue damage, immune responses, and electrodes shifting away from the intended site.

2. Matching the charge density of the BMI with that needed for its intended use. Smaller electrodes may not allow for this. Other applications of stimulation, such as optical stimulation, ought to be further studied.

3. Safe and efficient methods of implanting flexible neural interfaces.

4. Achieving high spatiotemporal resolution to avoid stimulation of unintended areas.

5. Achieving high throughput in BMIs to handle large datasets that will be decoded in real-time stimulation.

6. Allowing the device to interact with the largest area of neurons.

7. Issues relating to safety and privacy of BMIs.

This review provides insight into modern advancements in fabrication techniques of microelectrodes to promote efficient implantation, recording and stimulation, and long-term stability. Current designs to maximize efficiency often result in a device that is too thin to prevent buckling of the probe upon insertion or too large to prevent a substantial inflammatory response. The biocompatibility and implementation of these microelectrodes in future projects ought to combine the known strategies to enhance neuronal cell growth, minimize the foreign body response, and allow for simplified implantation to maximize the full potential of the microelectrodes.

Microelectrode implantations' impact on many neurological conditions has driven the search for better means of implantation and long-term stabilization through minimizing astrocyte and microglial responses. Still, the challenges lie in the lack of a perfect electrode in size, stiffness, conductivity, and biocompatibility, which enhances increased glial scar formation. As a result, mechanisms of enhancing neuronal attachment to the probe's tip to enhance the reading have become more popular, improving stability and viability. Further, anti-inflammatory drugs that decrease the neuroinflammatory response have been incorporated into such coatings to allow for successful chronic implantation. Modifications of the probes involving physical coatings and covalent or morphological/topographic changes have shown promising results in neuronal attachment and growth and the potential for selectively inhibiting glial encapsulation. Although physical, mechanical, and chemical properties alongside electrode modification can enhance their efficiency, a disadvantage of probes like those of microwire electrodes is their inability to deposit within the tissue following implantation; however, if they were able to do so, they would provide the most biocompatible electrode.

Ethics approval and consent to participate

IUCUC approval and consent to participate was unnecessary as no experiments were conducted in writing this review.

CRediT authorship contribution statement

Jordan Villa: Writing – review & editing, Writing – original draft, Methodology, Investigation, Conceptualization. Joaquin Cury: Writing – review & editing, Writing – original draft, Supervision, Methodology, Investigation, Funding acquisition, Conceptualization. Lexie Kessler: Writing – review & editing, Investigation. Xiaodong Tan: Writing – review & editing. Claus-Peter Richter: Writing – review & editing, Writing – original draft, Resources, Methodology, Investigation, Funding acquisition, Formal analysis, Conceptualization.

Declaration of competing interest

The authors declare no competing interests.

Acknowledgments

This work was funded through the NIH by grant R01DC018666 at Northwestern University.

Peer review under responsibility of KeAi Communications Co., Ltd.
==== Refs
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