Abstract
Flexible deep brain neural interfaces, as an important research direction in the field of neural engineering, have broad application prospects in areas such as neural signal detection, treatment of neurological diseases, and intelligent control systems. However, chronic inflammatory responses caused by long-term implantation and the resulting electrode failure seriously hinder the clinical development of this technology. This review systematically explores the long-term stability issues of flexible deep brain neural interfaces, with a focus on analyzing the synergistic optimization of electrode geometric morphology and implantation strategies in regulating inflammatory responses. Additionally, this paper delves into innovative strategies, such as passive enhancement of biocompatibility through electrode surface functionalization and active inhibition of inflammation through drug-controlled release systems, offering new technical paths to extend electrode lifespan. By integrating and reviewing existing innovative methods for deep brain flexible electrodes, this study provides an important theoretical foundation and technical guidance for the development of high-stability neural interface devices.
Similar content being viewed by others
Introduction
In the new era of brain-machine interfaces, the long-term stability of deep brain flexible electrodes is a critical performance indicator for their application in treating neurological disorders and intelligent control1,2. Although flexible electrodes have a lower Young’s modulus compared to rigid electrodes, enhancing mechanical compatibility with brain tissue, they still inevitably trigger immune responses3,4. The immune response, a key factor affecting the long-term stability of electrodes, includes acute inflammatory reactions during implantation and chronic inflammation following implantation5. Acute inflammatory responses occur due to geometric and mechanical mismatches between the electrode and brain tissue during implantation6. This mismatch causes mechanical impact, tearing target neuronal tissue, damaging neurons and nerve fibers, and leading to tissue displacement and deformation. Injured tissue releases inflammatory factors that attract immune cells to the injured area to phagocytose cell debris and dead cells7,8. Additionally, when electrode implantation pierces blood vessels, it triggers clotting and thrombus formation, attracting immune cells that release additional inflammatory factors, worsening the acute inflammatory response9. Over time, the acute inflammatory response gradually diminishes, but the mechanical mismatch between the implant and brain tissue persists9. Macroscopic and microscopic movements lead to friction between the implant and brain tissue, causing ongoing damage. This foreign body injury rapidly activates microglia, which release inflammatory cytokines and reactive oxygen species to clear the foreign material10,11. Meanwhile, astrocytes proliferate and migrate toward the injury site, secreting extracellular matrix (ECM) components and cytokines12. ECM components accumulate around the foreign body, forming a dense physical barrier. Ultimately, glial cells proliferate and align at the injury site, forming a compact cell layer, which leads to persistent scar formation and eventually induces fibrosis13. Scar tissue forms an insulating sheath around the electrode, increasing the distance between neurons and electrode sites, causing rapid signal attenuation and a sharp rise in impedance10. The insulation effect reduces electrode transmission efficiency while damaging the electrode’s electrochemical, electrical/optical stimulation, and electrophysiological recording functions, ultimately leading to failure14. Additionally, micromovements of the electrode in brain tissue and surface conditions exacerbate the impact of chronic inflammatory responses15.
Since the implantation method of neural interfaces directly affects the degree of immune response, and electrode material and shape determine the implantation method, the substrate of implantable electrodes has gradually shifted from rigid to flexible materials to minimize immune responses9. Flexible materials reduce mechanical impact during implantation and relative motion with brain tissue post-implantation, making them more promising tools for neural interfaces3,16. However, the shape design, implantation method, and surface modification of flexible electrodes still cannot meet the requirement for years-long detection periods17. Therefore, summarizing strategies to further optimize flexible deep brain neural interfaces concerning immune response is urgently needed.
We believe that the shape of deep brain flexible neural interfaces determines the implantation method, which directly influences the extent of implantation-induced damage and the acute inflammatory response. The structural characteristics of the neural interface, as well as the geometric and mechanical mismatch between the implant and brain tissue, are the main factors that trigger chronic inflammatory responses. This chronic inflammation continues to affect the effective operational lifetime of the neural interface. Flexible electrodes have significantly mitigated the mechanical mismatch issues associated with traditional rigid interfaces, particularly in terms of Young’s modulus. The further goal of reducing immune responses is to make the electrodes “invisible” within neural tissue, passively “escaping” the immune system’s recognition, and thus aligning their shape and surface properties with the neural tissue. However, implantation inevitably causes some degree of brain tissue damage. Therefore, it is also crucial to enable the electrode to actively release anti-inflammatory substances to modulate the surrounding environment and promote tissue repair. Additionally, the compatibility between passive invisibility and active modulation strategies can maximize the electrode’s lifespan (Fig. 1).
The long-term stability of flexible brain-machine interfaces requires consideration of implantation methods and surface modifications of neural interfaces. Implantation methods that are compatible with electrode shapes can reduce acute inflammatory responses during implantation, such as tungsten wire-guided insertion for filamentous electrodes, carrier-guided insertion for mesh electrodes, and direct insertion for Utah-like electrodes. Passive surface modifications of neural interfaces help them blend into the microenvironment to ensure functional stability, while active modifications can adjust the microenvironment to maintain surrounding neural bioactivity. Balancing shape, implantation, and modification ultimately extends the detection effectiveness and lifespan of neural interfaces.
Stability Prerequisite: Coordinated Electrode Shape and Implantation Method
The mechanical properties of brain tissue, such as its low Young’s modulus (approximately 1–10 kPa) and high compliance, pose significant challenges for electrode implantation. To ensure long-term stability and minimize tissue damage, the mechanical properties of the electrode must be carefully matched to those of the brain. Flexible electrodes, with their low bending stiffness and Young’s modulus, are designed to mimic the softness of brain tissue, reducing the risk of chronic inflammation and mechanical mismatch. However, their inherent flexibility limits their ability to penetrate brain tissue directly, necessitating the use of rigid shuttles or surface stiffness enhancement techniques for precise implantation (formulas 1 and 2)18.
E represents Young’s modulus, and I is the moment of inertia. Formula (1) applies to circular cross-sections, with r as the cross-sectional radius. Formula (2) applies to rectangular cross-sections, with b as the width and h as the height. Flexible electrodes often utilize rigid shuttle guidance or surface stiffness enhancement techniques to enhance bending stiffness, ensuring precise implantation19. The choice of implantation method depends on the electrode’s shape. Simple-shaped electrodes, such as rod or filament-like electrodes, are compatible with various implantation schemes20. Complex-shaped electrodes, like mesh electrodes, require special attention to maintain stability during implantation21,22. The cross-sectional area of implantation directly affects the extent of acute injury. Therefore, selecting an implantation method that matches the electrode’s shape and optimizing the implantation cross-sectional area contribute to the long-term stability of electrode recordings.
Customized Coordination of Electrode Shape and Implantation Method
When using a rigid shuttle for electrode guidance, the electrode directly contacts brain tissue, allowing for simultaneous recording of neural signals23. During the implantation process, the shape and amplitude of spike signals assist in determining the implantation site. Common guided implantation methods include tungsten wire1, SU821, and injection24. Tungsten wire guidance is typically used for rod-like or filament-like electrodes due to its simplicity and speed1. In contrast, SU8 guidance and injection pump technology are more suitable for complex electrodes, such as mesh electrodes21. This is because the shape of complex electrodes needs to be preserved to the greatest extent possible to maintain the specific locations of functional sites.
Rod and Filament Electrodes: Coordinated Shape Design and Tungsten Wire Guidance
The classification of rod-like and filament-like electrodes is based on the maximum thickness and width of the implant shank1. Rod-like electrodes typically range from several hundred micrometers to millimeters, while filament-like electrodes range from submicron to tens of micrometers. A tungsten wire with a stepped tip passes through the guiding hole at the electrode tip and is fixed with a polyethylene glycol (PEG) coating (Fig. 2a)25. The electrode is guided into the target brain region. After the PEG melts, the tungsten wire is retracted.
a Stepped Tungsten Wire Tip-Guided Electrode Implantation Process33; b Single-Shank Integrated Electrode with 64 Detection Sites Distributed at Different Depths and Immunofluorescence Response of Tissue26. Microscopic images of consecutive tissue sections around the implanted micro-electrode array at 3 months after implantation. Left, schematic diagram showing three consecutive sections centered around the electrode array in the monkey brain. Right, bright-field microscopic images showing a piece of shank remained embedded in the brain tissue, with the insertion track(red circle) after needle removal. Immunostaining images of the three sections for the glial marker GFAP (green) and the neuron marker NeuN (red),together with DAPI (blue) staining of cell nuclei, showing no substantial difference in cellular organization between the tissues around the implantedelectrode and at a distance. Scare bar = 100 μm. Indicate that there is no significant difference between the electrode implantation site and the surrounding tissue; c 128-Channel Open Sheath Flexible Electrode and Immunofluorescence Response of GFAP and NeuN 2 Weeks Post-Implantation27. (i) Electrode and electrode tip; (ii) Immunofluorescence response; (iii) Enlarged view of the immunofluorescence response; d Filamentous Electrode and 3D Reconstruction of Blood Vessels (Green) Surrounding the Electrode (Red) 2 Months Post-Implantation33. (i) Filamentous electrode tip; (ii) Immunofluorescence response 0–400 μm below the brain surface; (iii) Immunofluorescence response 100–320 μm below the brain surface; e NeuroRoots Electrode, Electrode Transfer Method, and Immunofluorescence Response 90 Days Post-Implantation with 100 μm Diameter Microwire29. (i) Assembly method using capillary and surface tension to pull the electrode onto the microwire (the black arrow: the effect after electrode assembly); (ii) Tissue response around the implant tip; (iii) Horizontal cross-section of tissue response along the implant trajectory. The surrounding tissue of the implantation site is surrounded by a small number of astrocytes. Reproduced with permission33. Copyright 2017, Lan, L., Xiaoling, W., Zhengtuo, Z., et al., published by Science Advances. Reproduced with permission26. Copyright 2023, Advanced Science published by Wiley-VCH GmbH. Reproduced under terms of the CC-BY license27. Copyright 2024, Keundong, L., Angelique C., Yun Goo, R., et al., published by Springer Nature. Reproduced with permission33. Copyright 2017, Lan, L., Xiaoling, W., Zhengtuo, Z., et al., published by Science Advances. Reproduced with permission29. Copyright 2018, Marc D., F., Christopher M., P., Alexander, G., et al., published by bioRxiv.
Tungsten wire-guided implantation strategies are categorized into unified and distributed implantations. Unified implantation refers to the use of a single tungsten wire guidance system to deploy multiple electrodes simultaneously or in a single step. This method ensures that all electrodes are implanted in a coordinated manner, maintaining a predefined spatial arrangement. Unified implantation is suitable for high-throughput detection in a single brain area or at different depths along the same path. Signal recording from single-shank implants is commonly used in the cortex, where neural signals have been stably recorded for up to eight months and successfully used to train decoding algorithms to control a cursor on a computer screen26. One such single-shank electrode, with a 100 µm2 cross-sectional area and 64 detection channels, was used in the primary visual and motor cortex of macaques (Fig. 2b). Similarly, a polyimide(PI)-based 128-channel open-sleeve electrode, 15 µm thick and 1.2 mm wide, has increased thickness and width due to the open-sleeve design (Fig. 2c)27. Although this design reduces the risk of rigid shuttle detachment, it comes at the cost of increased acute injury. Two weeks post-implantation, a noticeable glial sheath formed around the electrode. The open-sleeve electrode’s U-shaped neck design adds extra length, making it one of the few flexible electrodes suitable for deep brain detection in non-human primates and the treatment of epilepsy27. Additionally, guiding multi-shank electrodes, such as folded electrodes, increases detection throughput with a single implantation but doubles the implantation thickness28. NeuroRoots innovatively separates all detection channels, with detection filaments 7 μm wide and 1.5 μm thick (Fig. 2d). Capillary action and surface tension transfer the filamentous electrodes to a single-shank guiding microwire with a diameter of 35 μm. After implantation, the microwire is retracted without causing additional injury, allowing for signal recording for up to 7 weeks29. Unified implantation is compatible with surgical catheters, where the rigid shuttle and flexible electrodes pass through the catheter to the cortical location, reducing the risk of surgical infection30. Unified implantation reduces the number of injuries while balancing the stiffness of the implant with long-term compatibility with brain tissue. Reducing mechanical mismatches helps to decrease the formation of glial scars and chronic inflammation, thereby enhancing the long-term stability of neural recordings.
Distributed implantation involves using multiple tungsten wire guidance systems to deploy electrodes sequentially or independently. Each electrode is implanted individually, allowing greater flexibility in placement and adaptation to tissue morphology. Distributed implantation places fewer functional sites on each electrode filament and narrows the electrode width. This minimizes the cross-sectional area of a single implantation, promoting wound healing with minimal scarring. Additionally, increasing the number of electrode filaments and distributing implantation sites enhances throughput and expands the detection range31. Two types of distributed filament electrodes are available: one is 50 μm wide and 1 μm thick with 8 detection throughputs, and the other is 10 μm wide and 1.5 μm thick with 4 detection throughputs. Both designs reduce the cross-sectional area to a subcellular level, matching single-cell traction (Fig. 2e)32. The carbon fiber or tungsten microwire guiding shuttle, with a diameter of 7 μm, allows for vascular recovery within a month post-surgery33. Subsequently, the nanowire cross-sectional area has been further reduced to 10 μm234. Robotic-assisted implantation technology has been introduced to improve surgical efficiency35,36. Neuralink has advanced high-density backend integration and high-throughput miniaturization36.
Unified implantation is better suited for deep brain detection, while distributed implantation requires the electrode and its implantation guide to be small enough, typically in the nano or submicron range. This helps reduce acute injury during implantation and prevents chronic inflammation caused by larger relative movements after implantation. However, the advantage of distributed implantation lies in its ability to capture a broader range of neural information, which is difficult to achieve with unified implantation.
Mesh Electrodes: Coordinated Complex Shape with SU-8 Guidance and Pump Injection
Submicron-thick mesh electrodes simulate the three-dimensional shape of the brain’s neural network9. The pores in the mesh electrodes facilitate the interweaving and growth of neuronal synapses, promoting nutrient flow and waste removal37,38. Mesh electrodes are primarily implanted using unified methods, including SU-8 guided arms and cylindrical injection with an injection pump.
SU-8 has good biocompatibility, and its thickness is controllable. A 40 μm thick SU-8 polymer shuttle serves as the support arm for a 2.5 μm thick flexible electrode (Fig. 3a)39. Before implantation, SU-8 is thermally treated to harden, causing molecular cross-linking and enhanced mechanical strength. However, SU-8’s high-temperature hardening is irreversible, and its electrostatic force makes it difficult to separate from the flexible electrode. Therefore, hardened SU-8 in long-term implants increases relative friction with the tissue. To address this, a 2 μm thick water-soluble glucose sacrificial layer is added between the SU-8 carrier shuttle and the electrode. After implantation, the glucose layer dissolves, allowing the SU-8 shuttle to be removed, leaving the planar mesh electrode in the tissue (Fig. 3a)40. During chronic inflammatory response, the pores in the mesh electrode facilitate neuronal recovery and support stable electrophysiological recordings for up to 24 months. Injection pump implantation of cylindrical mesh electrodes involves delivering the cylindrical mesh electrode into a glass capillary needle and implanting it using a microinjection pump and a stepper motor platform31,41. In a time-dependent immunohistochemical study of brain tissue with mesh electrodes, it was found that only slight accumulation of astrocyte and microglia signals occurred near the boundaries within two weeks after implantation, resulting in minimal inflammation. Three months later, the mesh electrodes were fully integrated with the brain tissue(Fig. 3b)41. Inspired by pyramidal neurons, the electrode is designed to mimic neuronal soma and axons with micrometer-scale mesh lines and thickness (Fig. 3c)24. The small diameter of the capillary needle limits the distribution of the electrode’s backend I/O ports, so a strategy of injecting first and then connecting is used. Automatic conductive ink printing independently addresses the I/O ports of the mesh electrode to the connector42. The direct contact connection method further enhances connection efficiency and avoids errors43. Subsequent suture-like injection achieves multiplexing of mesh flexible electrodes in different brain regions44. The cylindrical mesh electrode was linearly elongated without folds and inserted into the tissue at the desired location and depth. Furthermore, the cylindrical mesh electrode seamlessly integrates in a manner similar to a standard mesh structure, without damaging the brain tissue (Fig. 3d)44.
a SU-8 Polymer Shuttle-Guided Planar Mesh Electrode Implantation40. (i) Shape of the SU-8 polymer shuttle and mesh electrode; (ii) Representative 3D reconstructed confocal fluorescence imaging of mesh electronic device implanted after 6 weeks; The planar mesh electrode after implantation integrates well with the brain tissue, and the size of the detection sites is similar to that of the neurons; b Cylindrical Mesh Electrode Injection Implantation41. (i) Bright-field microscope image taken in wide-field transmission mode showing partially ejected mesh electronics through a glass needle with an inside diameter of 95 μm, exhibiting significant expansion and unfolding of mesh electronics in aqueous solution. (ii) Time-dependent histology of the mesh electronics/brain tissue interface. Confocal fluorescence microscopy images of 10-μm thick horizontal brain slices sectioned perpendicular to the long axis of mesh probes at 2, 4 and 12 weeks post-injection. Green, red, cyan and blue colors correspond to neuron nuclei (NeuN antibody), neuron axons (neurofilament antibody), astrocytes (glial fibrillary acidic protein, GFAP antibody) and mesh electronics, respectively; c Neuron-Inspired Mesh Electrode24. (i) Overall layout and schematic of the 16-channel electrode with different material layers, and close-up views; The local magnified view of the electrode is shown in the black frame, with an exploded view of the electrode highlighted within the blue dashed box, displaying the different material layers. SU-8, interconnects, and electrodes are represented in red, yellow, and green, respectively; below are the structural characterizations of NeuE, including SEM original images and optical images. Scale bar, 10 μm; (ii) 3D full-probe interface of the mesh electrode, neurons, astrocytes, and microglia 2 weeks post-implantation; From left to right, the images show 2 weeks post-implantation, the 3D interface between the Neuron-Inspired Mesh Electrode (red) and neurons (green), astrocytes (cyan), and microglia (magenta). Scale bar, 200 μm; d Glass Capillary Needle Stitching44. (i) A single cylindrical mesh electrode (green) was inserted into different brain regions of mice using a glass capillary needle and simulated in a hydrogel; (ii) 3D mapping of the cylindrical mesh electrode; On the left, a 3D rendering of the probe (red) and the interface with astrocytes (cyan) in the GFAP transgenic mouse brain two weeks after injection. Astrocytes express fluorescent proteins after genetic modification. (White dashed boxes indicate sites I and II); scale bar, 300 μm. On the upper right, high-resolution images of the highlighted regions at sites I and II. Scale bar, 100 μm. On the lower right, 3D mapping showing the interface between the probe (red) and neurons (green) in the YFP-H transgenic mouse brain two weeks after injection. (iii) Two weeks post-injection, the normalized fluorescence intensity of neurons and astrocytes as a function of the distance from probe surface sites I (black) and II (red). Reproduced with permission40. Copyright 2023, Siyuan, Z., Xin, T., Weiwen, T., et al., published by Springer Nature. Reproduced with permission41. Copyright 2018, American Chemical Society. Reproduced with permission24. Copyright 2019, Xiao, Y., Tao Z., Theodore J., Z., et al., published by Springer Nature. Reproduced with permission44. Copyright 2023, Jung Min, L., Dingchang, L., Young‐Woo, P., et al., Advanced Science published by Wiley-VCH GmbH.
The mechanical design of mesh electrodes is crucial for ensuring compatibility with the brain’s soft tissue. Their submicron thickness and porous structure not only lower the effective Young’s modulus but also enable these electrodes to mold to the brain’s three-dimensional curvature. This adaptability significantly reduces shear stress at the tissue-electrode interface, minimizing the risk of tissue damage and chronic inflammation. Moreover, the low bending stiffness of mesh electrodes allows them to accommodate the brain’s micro-motions caused by breathing and heartbeat, further boosting their long-term stability.
Utah-like Electrodes: Shape Adaptability and Surface Stiffness Coordination
Utah-like electrodes have 3D raised recording sites designed to capture signals from the brain’s superficial layers. They are classified as semi-flexible or fully flexible based on the stiffness of the base and 3D protrusions. Semi-flexible electrodes combine rigid protrusions with a soft base to maintain implantation shape and stability. Fully flexible electrodes conform to the brain tissue surface, ensuring close contact between the sites and neural tissue, thereby reducing the risk of displacement.
Semi-flexible electrode micro-protrusion arrays can directly penetrate the cerebral cortex, such as PI bases combined with SU-8 detection cones. The flexible base adheres to bodily fluids, conforming closely to the brain surface and maintaining conformity (Fig. 4a)45. The length of the cones can vary to meet the detection needs of different cortical functional layers. However, cones with diameters greater than 200 μm cause significant injury. Similarly, a PI-based 1024-channel silicon microneedle Utah-like electrode, with microneedle tips as small as 10 μm, covers the entire mouse brain laterally and has achieved stable recordings for up to six months (Fig. 4b)46. Due to its diameter being comparable to that of a 7 µm carbon fiber, minimal tissue response was observed 73 days post-implantation. Compared to the unexposed hemisphere of the same animal, there was no significant change in neuronal density. Additionally, the major advantage of Utah-like electrodes with external bases is the ability to construct microfluidic structures for controlled drug diffusion, enabling active anti-inflammatory treatment while minimizing tissue injury (Fig. 4c)47. In the same brain region, cell nuclei (appearing blue) and cells labeled with Texas Red-dextran (appearing red) were observed. The dextran-labeled cells were primarily located near the V-shaped incision formed after microneedle insertion, close to the interior of the incision. This suggests that Texas Red-dextran flowed along the electrode path and permeated into the surrounding brain tissue. It can be confirmed that the microfluidic neural interface successfully delivered chemicals to the electrode target site (Fig. 4c)47.
a Semi-flexible Utah-like Electrode with PI-based SU-8 Microneedles. (i) Photograph of the Utah-like Electrodes45; b High-throughput Semi-flexible Utah-like Electrode with PI-based Silicon Microneedles46. (i) Photograph of a 1024-channel, 32-channel silicon microneedle electrode and image of it implanted on the brain surface; (ii) Coronal tissue sections 73 days post-implantation, showing two-photon imaging and fluorescence microscopy photos; cell nuclei (DAPI, blue), neurons (NeuN antibody, green), astrocytes (GFAP antibody, red), and microglia (Iba1 antibody, magenta). The red arrow indicates tissue penetration by a single probe, and the yellow dashed box highlights the enlarged area on the right. c Microfluidic semi-flexible Utah-like electrodes47. (i) Images showing the electrode and cross-sectional view of the arched microfluidic channel; (ii) The microfluidic channel shown empty (left) and filled with stained water (right); (iii) SEM image of the electrode, featuring a pitted structure. The channel width is approximately 200 µm, and the height is 40 µm. Scale bar: 500 µm; (iv) Fluorescence image of rat brain slices containing the electrode implantation site. DAPI and Dextran Texas Red were used to label cell nuclei and visualize the delivery area, respectively. At the bottom, a magnified image of the region around the electrode. The white dashed line indicates the electrode trajectory. Scale bar: 500 µm. d Fully Flexible Origami Electrode49. (i) Full-view photograph of the electrode. At the top, a rotational view (top right) of the 3D neural mapping device of the Fully Flexible Origami Electrode, consisting of 9 ECoG electrodes and 4 probes, with each probe containing 6 electrodes. The recording area of each electrode is 20 × 20 µm², with a 500 µm spacing between the ECoG electrodes and a 200 µm spacing between the probe electrodes. Each probe has a height of approximately 1.5 mm and a thickness of <20 µm, including both sides of the folded flexible substrate. Scale bar, 250 µm. On the bottom left, the Fully Flexible Origami Electrode is wrapped around a glass rod with a radius of 5 mm. Scale bar, 500 μm. On the bottom right, a photo of the PEG-coated 3D neural mapping device vertically inserted into the anterior part of the mouse brain, with part of it inserted perpendicularly into the brain. Scale bar, 250 μm; (ii) Immune response around the implanted electrode, showing neurons; fluorescence images of activated astrocytes (GFAP, green fluorescence) and microglia (IBA1, red fluorescence) in the tissue surrounding the electrode array, and an expanded image (yellow box). (White arrow: electrode implantation site, white dotted line: electrode position). Reproduced with permission45. Copyright 2016, Zhuolin, X., Jingquan, L., Chengkuo, L., et al. Reproduced with permission46. Copyright 2022, Wiley‐VCH GmbH. Reproduced under terms of the CC-BY license47. Copyright 2021, Yoo Na, K., Namsun, C., Jae-Won, J., et al. Reproduced under terms of the CC-BY license49. Copyright 2022, Ju Young, L., Sang Hoon, P., Yujin, K., et al.
Fully flexible electrodes offer higher mechanical compliance, allowing them to move with the micro-movements of brain tissue caused by breathing, heartbeat, and head motion. Their soft implant shanks typically undergo surface stiffness enhancement to increase rigidity. For example, planar comb-shaped floating electrodes made entirely of PI material are coated with poly(lactic-co-glycolic acid) (PLGA) to enhance stiffness, allowing for up to one year of detection in the visual cortex of monkeys48. Using a pop-up origami fabrication technique, the base can transform from 2D to 3D, featuring four shanks and 20 detection throughputs (Fig. 4d)49. In vivo histological results of the glial response observed 2 days after the device was inserted into the brain show that due to the flexibility of the device, there is minimal immune response and glial scarring around the implantation site (Fig. 4d)49. Coating with PEG completes the implantation. Compared to PLGA, PEG dissolves quickly; combining the two allows for adjustable dissolution rates50. Semi-flexible Utah-like electrodes focus on expanding the lateral detection range. Combined with 3D silicon microneedles, which have small electrode radii and close spacing, extends brain surface distribution and increases recording throughput. The detection throughput of fully flexible 3D electrodes is limited by spacing, however multi-channel detection at different depths on a single shank can be further developed.
Utah-like electrodes are specifically engineered to overcome the mechanical mismatches between rigid electrodes and the soft brain tissue. These electrodes feature semi-flexible designs that incorporate rigid protrusions for effective penetration paired with a soft base that contours to the brain’s surface, effectively reducing shear stress and mechanical mismatch. Fully flexible electrodes, characterized by their high mechanical compliance, seamlessly adapt to the brain’s micro-motions, further diminishing the risk of tissue damage. By aligning the mechanical properties of the electrodes with those of the brain, Utah-like electrodes strike an optimal balance between stable implantation and enduring biocompatibility.
Intravascular Electrodes: Shape Expansion and Self-expanding Stent Coordination
Intravascular electrodes utilize blood vessels as conductive pathways, providing a minimally deep brain approach to reach target brain regions51. Arterial walls have a thickness of 10–20 μm, and neural electrode spike detection ranges within 130 μm, allowing vascular electrodes to cover spike detection.
Six to twelve platinum electrodes are attached to a self-expanding stent, which is inserted via the jugular vein in the neck into the superior sagittal sinus of sheep using a stainless-steel needle. Recording can last up to 6 months51,52. However, the stiffness and bulkiness of metal catheters and stents limit detection in most brain areas, and navigating the micrometer-scale tortuous cerebral vasculature can cause tissue injury and inflammation. Magnetically actuated flexible guides enable dynamic steering at branch points, facilitating intravascular navigation53. Mesh electrodes paired with soft probes enhance precision in intravascular spike recording. Saline injection is used to select 100 μm branching vessels, successfully recording spike signals in the cortex and olfactory bulb54. The main challenges of vascular electrodes include maintaining an extended state while smoothly advancing, effectively selecting branches, and adhering to the vessel wall at the target location while in the extended state.
Broad Electrode Shape Adaptability with Enhanced Surface Stiffness
Brain tissue features low Young’s modulus and high compliance, necessitating electrodes with adjustable stiffness: high stiffness for penetration during implantation and reduced stiffness post-implantation to minimize mechanical mismatch. Surface stiffness enhancement techniques, such as biodegradable coatings (e.g., PEG, silk fibroin, and PLGA) and cryogenic freezing, allow for temporary hardening during implantation while maintaining flexibility afterwards55. However, each method has its pros and cons. Biodegradable coatings reduce the risk of chronic inflammation and enhance long-term stability of neural recordings but inevitably increase the electrode’s cross-sectional area. Cryogenic freezing temporarily increases stiffness without altering the electrode’s shape, though precise temperature control during implantation is required. Additionally, because it is difficult and unnecessary to separate the electrode after surface stiffness enhancement, a general approach is to use overall enhancement and integrated implantation. Table 1 summarizes the advantages and disadvantages of the various electrodes and their implantation methods mentioned, while Table 2 summarizes the parameter comparison of all the technical solutions presented in this paper.
PEG and PLGA Coatings: Balancing Stiffness and Implantation Efficiency
Based on the biocompatibility, melting point, and hydrophilic/hydrophobic characteristics of synthetic polymers, an appropriate ratio of synthetic polymers can be drop-coated or dip-coated to wrap the outermost layer of flexible electrodes. The polymer coating not only strengthens the electrodes and simplifies implantation but also preserves the integrity and functionality of the modified sites after degradation8.
PEG solidifies into a wax-like state at room temperature, promoting neuron enrichment and protecting electrode functionality. After PEG coating, electrophysiological detection sites modified with platinum nanoparticles (PtNPs) and PtNPs/ Poly(3,4-ethylenedioxythiophene) (PEDOT)23,56, as well as dopamine electrochemical detection sites modified with Nafion57, remain effective. The functional protection formed enables the electrode surface to maintain continuous detection under long-term animal movement49. PEG is also suitable for the implantation of multi-shank electrodes48,49. Despite the quick dissolution of the coating post-implantation, the increased electrode thickness due to coating still causes tissue injury49. Using PDMS molds for single-sided coating can reduce the thickness of the synthetic polymer58.
The rapid curing properties of PEG and PLGA are utilized to achieve the integration of multifunctional electrodes in situ and enhance electrode flux. PLGA copolymers, with their lower hydrophilicity and slower degradation rate, maintain biocompatibility while allowing control over the dissolution rate when appropriately combined with PEG50. Silk fibroin optical fibers combined with 128-channel flexible electrodes59 and optical fibers combined with 1024-channel flexible electrodes, under the influence of elasticity and capillary forces, create dual-functional electrodes for synchronous optogenetic stimulation and electrophysiological detection (Fig. 5a)60.
a Neural Fringe Electrode (i) Neural fringe/PEG assembly; (ii) Immunohistochemical staining and bright-field image of horizontal brain slices 5 weeks post-implantation; b Bilayer Silk-Parylene Electrode61. (i) SEM image of the bilayer silk-parylene shank; (ii) Enlarged SEM image of (i); (iii) Schematic of the bilayer silk-parylene electrode with feature dimensions; (iv) Enzymatic degradation of silk in the bilayer silk-parylene probe; c 3D Nanomesh Electrode66. (i) Shape of the 3D nanomesh electrode; (ii) 3D reconstructed confocal micrograph projection of cross-sectional slices immunochemically labeled 5 weeks post-implantation. Blue: nuclei (Hoechst); Green: NeuN (neuronal nuclei); White: SU-8; Red: GFAP (reactive astrocytes). The right panel shows the fluorescence intensity distribution of astrocytes (red channel) along the long axes of the two dashed box regions; d Liquid Metal Electrode67. (i) Diagram of the liquid metal electrode and its filling pressure; (ii) Liquid metal electrode in rigid state and integrated drug delivery state (The scanning electron microscope image of the electrode tip in the box shows the drug delivery channel). Reproduced under terms of the CC-BY license60. Copyright 2019, Shouliang, G., Jinfen, W., Xiaowei, G., et al., published by Science Advances. Reproduced with permission61. Copyright 2022, Clement, C., Adrian, L., Lionel G., N., et al. Reproduced with permission66. Copyright 2015, Chong, X., Jia, L., Tian-Ming, F., et al., published by Springer Nature. Reproduced with permission67. Copyright 2019, Elsevier B.V.
The 128-channel dual-functional electrode, implanted in the mouse brain for two months, showed less microglial growth and neuronal apoptosis compared to silicon-based electrodes59. The 1024-channel Neurotassel electrode can record stable neuronal activity in the prefrontal cortex for up to six weeks60. Therefore, in situ stimulation and synchronous electrochemical detection of multiple neurotransmitters can be achieved using PEG to integrate multifunctional electrodes. PEG functional integration expands the functional capacity and reduces the risk of interference between stimulation and detection processes by optimizing the arrangement and combination of electrodes.
Silk Fibroin Coatings: Controllable Degradation and Long-term Stability
The thickness of the silk fibroin layer can be precisely controlled by adjusting the polymer percentage in the solution or the volume of the fixed area. Thus, the degradation time of silk fibroin can be adjusted by varying the layer thickness, ranging from a few hours to a week, allowing for more flexible applications (Fig. 5b)61. In the fabrication process, the electrode substrate is first treated with cellulose acetate, followed by silk fibroin deposition, which is compatible with conventional Microelectromechanical Systems (MEMS) processes. The electrode thickness mainly depends on the 30 μm thick silk fibroin, resulting in a final implantation thickness that increases but remains within acceptable limits.
The biodegradable properties of the materials simplify electrode implantation and eliminate the need for secondary extraction. However, it is crucial to maintain the original thickness of the electrode to avoid increasing acute injury. Each surface hardening material has its suitable application scenarios. For instance, PEG dissolves quickly, providing a small implantation time window and higher operational difficulty, making it suitable for short-term, shallow brain region implantation. In contrast, silk fibroin has a longer and more controllable degradation time, making it more suitable for long-term, deep brain region implantation. The method of one-sided deposition of silk fibroin can simultaneously achieve information recording61. The biodegradable materials also facilitate the active delivery of anti-inflammatory drugs to the implanted brain area. The combination of PEG and silk fibroin can meet implantation needs while exploring different timescales for anti-inflammatory drug delivery62. Besides PEG and silk fibroin63, other biodegradable polymers such as polylactic acid and chitosan64,65 are also excellent choices18. These materials can be used as additives in polymer implants to regulate the degradation rate and achieve drug deliver. The application of drug delivery will be described in detail in Part 3.2.1.
Cryogenic Freezing and Liquid Metal: Shape and Stiffness Coordination
Liquid nitrogen freezing maximally preserves the electrode’s morphology, making it suitable for complex-shaped electrodes such as 3D nano-mesh electrodes. When transitioning from air to a water environment, the electrode naturally expands into a mesh shape. Upon withdrawing from the liquid back into the air, the liquid surface tension increases the overall curvature, forming a hollow central three-dimensional cone with a diameter of approximately 100–200 μm (Fig. 5c)66. The hollow center and mesh gaps promote symbiotic growth and communication with neural tissue without triggering an immune response. Therefore, the key to implantation is to avoid deforming the mesh shape and cone gaps. Liquid nitrogen freezing directly hardens the electrode while maintaining its original state and allows controlled implantation. Confocal microscope images taken 5 weeks post-implantation show that the tissue cross-section initially has a hollow center. Five weeks later, neuronal synapses have moved, interpenetrated, and partially filled the central void created during the initial implantation. The neurons and the electrode closely integrate and fuse into a single unit.
Liquid metal gallium, with a melting point of approximately 29 °C, is introduced into PDMS channels at lower temperatures, increasing electrode stiffness upon cooling. A few minutes after implantation, body temperature causes the Ga to return to a liquid state, reducing friction with brain tissue (Fig. 5d)67. However, the microchannels in PDMS increase the electrode thickness by several tens of micrometers, necessitating the integration of controlled anti-inflammatory drug delivery channels.
Cryogenic freezing and liquid metal technologies further enhance the mechanical compatibility of electrodes with brain tissue. Cryogenic freezing temporarily increases the stiffness of complex-shaped electrodes, such as 3D nanomesh electrodes, without altering their form. This allows for precise implantation while maintaining post-implantation softness, reducing mechanical mismatches with brain tissue. Similarly, liquid gallium provides adjustable stiffness, transitioning from a rigid state during implantation to a flexible state afterwards. These technologies ensure that the mechanical properties of the electrodes dynamically match those of the soft brain tissue, minimizing tissue damage and enhancing long-term stability.
Key to Long-Term Stability: Surface Modification to Reduce Chronic Inflammation
Based on the interaction mechanisms between the electrode surface modifications and brain tissue, modification strategies are classified into passive and active types. Passive modifications involve the use of natural or synthetic materials to alter the electrode surface, making its physical or chemical properties more similar to brain tissue and thereby reducing inflammatory responses. Active modifications build upon passive modifications by incorporating drugs or bioactive molecules68. As the carrier material dissolves or degrades, drugs or molecules are released, actively intervening to regulate the biochemical environment of the tissue68. Both passive and active modifications require a clear definition of the modification scope, including modifications of the entire implanted surface and specific electrode sites. The scope of modification determines the conductivity requirements of the target modifier, its ability to adhere to the electrode, the retention time on the implanted part or electrode sites, and the thickness of the modification. Additionally, passive modification and active drug delivery can be carried out simultaneously without conflict. Through different functions, materials, and design approaches, the two complement each other on the electrode surface, enhancing the electrode’s performance. This ensures both long-term stability and biocompatibility, while also enabling therapeutic drug release, maximizing the therapeutic effect.
Passive Modification to Evade Foreign Body Recognition for Functional Stability
Passive modifications are categorized into natural material modifications and synthetic material modifications based on the type of material used. Modifications using natural materials can exhibit inherent biological functionality and low immunogenicity, which may support cell adhesion and growth while facilitating the integration of the modified electrode with the implantation environment69,70. Synthetic material modifications offer precise control over chemical and physical properties, enabling the creation of electrode surfaces with specific functions. Both types of modifications can be applied to the same electrode to complement each other.
Natural Material Coatings for Biocompatibility
Natural materials, due to their biocompatibility and structural similarity to endogenous components, can be engineered to reduce immune recognition and mitigate foreign body reactions69,70. ECM is a critical component of the tissue environment, maintaining structural stability and regulating cell behavior in brain tissue71. Therefore, ECM or its extracts can directly mimic the natural neural environment, providing necessary biological support and structural foundation. ECM includes various proteins, polysaccharides, and growth factors. Proteins such as fibrin, collagen, and laminin; polysaccharides such as hyaluronic acid72. Additionally, proteins are extracted from natural sources like silk fibroin59 and gelatin; and polysaccharides such as alginate, chitosan, and bacterial cellulose73. These extracts can be mixed in certain proportions to prepare hydrogels for electrode modification.
Flexible electrodes were fully coated with collagen membranes and collagen/laminin membranes. One month after implantation, the collagen membranes were only partially degraded, and the electrodes’ recording and noise characteristics were unaffected by the biomembranes. The modified biomembranes reduced immune cell accumulation at the electrode-tissue interface, similar to brain regions without electrode implantation74. The slow hydrolysis of protein membranes allows the incorporation of anti-inflammatory drugs, achieving both passive and active anti-inflammatory effects75. Unlike collagen, laminin modification requires pretreatment of the electrodes, followed by layer-by-layer self-assembly through continuous cation-anion adsorption76. Four weeks post-implantation, immune cell expression was halved compared to uncoated electrodes77. ECM derivatives and neural adhesive L1 coatings also attracted neurons to the coated electrode sites78. Although full-surface modifications align the electrodes’ physicochemical properties with the tissue environment, proteins as insulating media may introduce parallel impedance into the detection circuit, and its impact on spike signal detection is unknown74. Additionally, the effectiveness of drug delivery by polysaccharide biomembranes has been proven79. Chitosan combined with hyaluronic acid constructs a polyelectrolyte with sustained biodegradability, incorporating nanocomposites to achieve biofunctionalization80. As the biomembrane degrades, nanocomposites are absorbed by cells. Natural protein or derived hydrogel matrices coated on electrodes serve as mechanical and biological buffer layers, with neurotrophic factors incorporated to modulate immunity, enhancing the unmodified Inflammatory effect for over three months75.
Natural materials can be directly used to construct electrode substrates. Bacterial cellulose has high biostability and is resistant to degradation by the body’s enzyme systems. Its Young’s modulus, approximately 7 kPa, is similar to that of brain tissue81. The three-dimensional microfibril network shape serves as a growth scaffold, promoting neuronal attachment and growth82. Four weeks post-implantation, the bacterial cellulose-based electrodes induced inflammation reduced to nearly one-third compared to PI-based electrodes (Fig. 6a)83.
a Bacterial Cellulose-Based Electrodes83. (i) Photograph of the bacterial cellulose-based electrode; (ii) Immunofluorescence staining 2 weeks and 4 weeks post-implantation of Au-bacterial cellulose-based (left) and Au-PI (right) electrodes, showing neurons (NeuN, red), cell nuclei (DAPI, blue), and astrocytes (GFAP, green). The white dashed line indicates the cross-section of the electrode (location); (iii) Magnified image of astrocytes; (iv-v) Relative fluorescence intensity of neurons and astrocytes connected to the Au-BC and Au-PI electrodes in the brain 2 weeks and 4 weeks post-implantation. Scale bar: 500 μm. b Hydrogel Hybrid Electrodes88. (i) Concept illustration and application of the hydrogel hybrid electrode; (ii) Average spike waveform over time; (iii) Average spike amplitude (red) and noise level (blue) over time, with average spike interval per unit; c Microporous Drug-Releasing Electrodes103. (i) Preparation process for microporous drug-releasing electrodes; (ii) Optical image of the electrode with CsA-loaded PLGA, MS, and PEG hydrogel; (iii) Another optical image of the same electrode; (iv) SEM cross-section showing CsA-loaded MS distribution in the hydrogel layer; (v) Another SEM cross-section showing CsA-loaded MS distribution; d Biomolecule-Modified Electrodes110. (i) Conceptual illustration of the biomolecule-modified electrode. Scale bar: 500 μm; (ii) Detailed view of the microelectrode, scale bar: 35 μm. Inset: SEM image of a microelectrode filament, 12 μm in width, featuring a 10 μm recording site, scale bar: 7 μm; (iii) Schematic diagram of EK-IKVAV modified electrode; (iv) QCM-D characterization of EK-IKVAV modification. Reproduced with permission83. Copyright 2018, American Chemical Society. Reproduced with permission88. Copyright 2021, Seongjun, P., Hyunwoo, Y., Ruike, Z., et al., published by Springer Nature. Reproduced with permission103. Copyright 2016, published by Elsevier Ltd. Reproduced with permission110. Copyright 2021, published by Elsevier B.V.
Natural materials used for modification are mostly in the form of hydrogels. Hydrogels, as a polymeric three-dimensional network shape, can absorb large amounts of water without dissolving. They have good biocompatibility; however, the swelling ratio needs to be controlled to avoid excessive implantation injury.
Synthetic Material Films for Mechanical, Chemical, and Electrical Stability
Synthetic materials generally offer superior biological and mechanical stability compared to natural materials. Conductive polymers can be used for targeted modification of electrode sites8,84,85,86, increasing surface roughness, enhancing specific surface area, reducing impedance, and improving neuronal adhesion87. Non-conductive synthetic polymer hydrogels, such as polyvinyl alcohol (PVA), polyacrylic acid (PAA), and PEG, are commonly used for full-surface modifications58.
These polymers can form various mixed hydrogels, adjusting their physical properties according to specific needs. For example, PVA, alginate, and poly(ethylene glycol) diacrylate-alginate (PEGDA-Alg) hydrogels have low fracture toughness, while poly(acrylamide)-alginate (PAAm-Alg) hydrogels offer softness and mechanical robustness similar to brain tissue. Combining the properties of these materials, a composite PVA/alginate/PEGDA-Alg/PAAm-Alg hydrogel can be applied to the full surface of electrode implants. The hybrid hydrogel allows for neural information detection for up to 6 months while integrating optical, electrical, and fluid functionalities into the electrodes. (Fig. 6b)88.
Conductive polymers can be combined with natural or non-conductive polymers to improve mechanical stability, biostability, and conductivity. For example, the biocompatibility of alginate gel combined with the conductivity of PEDOT/PSS forms a soft interpenetrating network at neural electrode sites through electrochemical growth, enhancing charge storage capacity and reducing electrochemical impedance89. Adding conductive polymers, such as P3HT6S, to polyacrylamide and poly(N-isopropylacrylamide-co-N-isopropylmethacrylamide) hydrogels enhance their electrical performance90. Using perfluoropolyether as an electrode surface lubricant reduces protein and immune cell adhesion91, thereby improving biocompatibility; however, its effect on neuron adhesion requires further study.
Active Intervention to Regulate Microenvironment for Bioactivity
Enhancing the biocompatibility of electrode coatings typically only delays inflammation. To actively reduce inflammation, strategies involving anti-inflammatory agents and carrier coatings have been proposed. Anti-inflammatory substances, such as the broad-spectrum drug dexamethasone (Dex), act by regulating the biochemical environment of tissues. When combined with carriers like hydrogels and electrospun materials, the degradation rate and diffusivity of the carriers determine the amount and duration of drug release. The choice of carrier must consider its degradation behavior within the tissue. Additionally, passive modification materials can also serve as carriers for active modifications. Active anti-inflammatory approaches can be categorized into drug delivery systems and biomolecular modifications.
Drug Release Systems for Immunomodulation
Hydrogels achieve high spatial and temporal resolution drug delivery through their lattice-like internal structure and interconnected pores92. These pores can incorporate various materials to enhance specific functions, such as anti-inflammatory drugs to reduce Inflammatory responses, conductive polymers to improve conductivity, polypyrrole (PPy) and oligomers to reduce swelling93, and materials like polycaprolactone (PCL), polyacrylamide, and polyvinyl alcohol to enhance mechanical properties and control degradation rates. Common drug release mechanisms include electrically stimulated release and dissolution/degradation-based release94.
Electrically controlled drug release allows precise regulation of the release location and amount. Carriers like PPy, PEDOT:PSS nanotubes, and aniline oligomers are used for delivering drugs such as Dex. PPy mixed with Dex enables precise control of release through electrical stimulation95. Besides step potential stimulation, cyclic voltammetry using alternating positive and negative potentials can also promote the release of anionic Dex, enhancing neural synapses96. Vertically growing PPy on an alginate hydrogel layer can reduce impedance by an order of magnitude97. Coating the outermost layer with ECM to form a PPy/Dex/ECM three-dimensional nanonetwork structure adds passive protection to active anti-inflammatory effects98. For carrier material selection, biodegradable electrospun poly(L-lactic acid) and PLGA with high surface areas help increase drug loading99. Additionally, coating the outermost layer with an alginate hydrogel extends the drug release time100. Both techniques use electroactive materials to control drug release. Electrospinning technology is a method that uses a high-voltage electric field to produce nanofibers, commonly employed to fabricate high-surface-area carriers for drug delivery systems. It utilizes biodegradable fibers and conductive layers, making it ideal for precise, large-dose releases99. The PPy/Dex system relies on the electroactive properties of PPy, making it ideal for applications requiring electrically stimulated responses.
In degradation/dissolution-controlled release, drugs and carriers can be functionally interchangeable. Synthetic hydrogels support cell adhesion, growth, and proliferation, and their slow degradation/dissolution helps to re-regulate drug release73. For instance, alginate forms a three-dimensional network hydrogel matrix encapsulating PLGA nanoparticles. Alginate stabilizes the nanoparticles and adjusts cross-linking density to control degradation time and release rate. The biodegradable nature of PLGA allows encapsulated dexamethasone to be released slowly and steadily101. Chitosan, as a carrier mixed with aniline oligomers, transforms the microstructure of the sites from honeycomb to aligned patterns, enhancing conductivity, mechanical properties, and surrounding cell activity102. Agarose, a natural gel-forming agent, is sometimes added to improve the mechanical strength and stability of hydrogels102. Drugs in drug delivery systems can also be substituted. For example, cyclosporine A can replace dexamethasone in PLGA microspheres to control the release rate103. Additionally, introducing microporous structures on electrode surfaces provides drug-loading spaces, and aminopropyl methacrylamide chemical treatment can improve coating adhesion (Fig. 6c)103. However, for 25 μm PI-based electrodes, the 200 μm thick hydrogel increases the final implantation thickness of the electrode by nearly ten times. To address this, incorporating aniline oligomers into the hydrogel can reduce the swelling ratio and slow down the degradation rate102.
Biomolecular Functionalization for Neuro-regeneration
The bioactivity of biomolecules on the surface of neural electrodes can actively promote cell adhesion, growth, and signal transmission. The biocompatibility enhancement improves the integration between the electrode and surrounding neural tissue, promoting nerve regeneration and repair104.
The L1 cell adhesion molecule is an immunoglobulin that, when modified on the surface of neural electrodes, directly promotes cell adhesion, growth, signal transmission, nerve regeneration, and repair105. Covalently fixing L1 on the surface of silicon-based neural electrodes significantly increases axon density within a 100 μm area and significantly reduces the number of immune cells106. Nerve growth factor has a similar effect to L1 and can be combined with dexamethasone for dual active anti-inflammatory action75.
Compared to proteins, peptide fragments are more stable, less prone to rapid degradation, and less likely to become inactive. Specific bioactive fragments are often designed from laminin, fibronectin, and collagen proteins, such as Ile-Lys-Val-Ala-Val (ILVAV), Tyr-Ile-Gly-Ser-Arg (YIGSR), and Arg-Gly-Asp (RGD)107. Besides extraction from proteins, chemical synthesis allows precise control over peptide composition, structure, and purity, producing complex sequences with specific functions. For example, laminin-derived peptides covalently linked to vinyl and amino-functionalized polyimide sheets form biocompatible coatings that support neuron adhesion and neurite outgrowth while reducing fibroblast contamination108. Non-active, non-adhesive peptides fixed on dextran coatings promote low levels of adhesion and spreading, enhancing neuron cell adhesion and minimizing fibroblast and glial cell adhesion109. Anti-fouling zwitterionic peptides used to modify ultra-flexible neural probes enable consistent neuronal activity monitoring for up to 16 weeks and reduce neural injury (Fig. 6d)110.
Peptide chains combined with conductive polymers can enhance electrode transmission efficiency. When integrated with hydrogels, electrode sites provide a sufficient physical structure for biomolecule adhesion and control the sustained release of biomolecules111,112. Modifying carboxylic acid-functionalized PEDOT copolymer films with a nonapeptide from the basement membrane protein laminin leads to an average neurite length of 90 μm113. Similarly, peptide chains can be interspersed within hydrogel gaps to modify electrodes, promoting nerve regeneration80. Peptoid foldamers and supramolecular β-peptides maintain proteolytic stability. Peptoid foldamers and supramolecular β-peptides are prepared as non-biodegradable hydrogels with carbon nanotubes added to form a three-dimensional network structure, making the impedance of the modifications controllable and maintaining signal transmission stability. β-peptide hydrogels seamlessly adhere to surrounding neural tissue, showing no significant inflammatory response on the cortical surface or within brain tissue over several months of observation114.
Outlook
The long-term stability of flexible deep brain neural interfaces is driven by two key trends: first, the miniaturization of the deep brain interface and the adaptation of minimally invasive implantation techniques; and second, the biocompatibility of the neural interface115. Achieving these two goals requires a three-step strategy: first, optimizing the shape and surface morphology of the implant; second, developing minimally invasive implantation techniques that are compatible with the neural interface; and third, ensuring that the microenvironment of the long-term implant closely resembles the brain tissue environment or neuronal distribution. By incorporating these designs, electrodes can evade immune system recognition and actively reduce immune responses. Thus, the future optimization of flexible electrodes for long-term stability is likely to follow the technological paths outlined below:
Miniaturization and High Surface Area Structures
Nano- or submicron-scale biomimetic structures, designed to resemble neurons, can better match the microstructure of neural tissue, promoting natural integration between the neural interface and the tissue116. High surface area designs not only ensure the effective detection range of neural signals but also minimize rigid contact during implantation, reducing mechanical stimulation of surrounding brain tissue and consequently lowering immune rejection57,117,118. Examples of such electrode types include nanowire electrodes and three-dimensional mesh electrodes.
Minimally Invasive Implantation Methods
Flexible, miniaturized electrodes can be paired with implantation guides, or electrodes with a hard base that soften post-implantation, such as shape-memory electrodes. These electrodes adjust their stiffness by controlling the temperature difference between the natural and implanted environments, thus minimizing excessive pressure on the brain tissue during implantation119. However, the relatively large size of shape-memory electrodes suggests the potential for a shift in approach towards developing simpler, more compact implantation guides to reduce implantation-induced damage.
Electrode Interface Materials to Minimize Immune Responses
To mitigate immune responses, one approach involves using biodegradable materials that degrade over time after implantation, preventing chronic immune reactions caused by long-term foreign material presence120. For example, natural silk proteins can degrade and be absorbed by brain tissue after a certain period. A challenge, however, lies in precisely controlling the degradation timeline and ensuring complete absorption by the brain tissue. Another strategy involves applying nanocoatings, such as collagen or hydrogels, which can reduce the immune system’s recognition of the electrode by mimicking the brain’s biological environment. Hydrogels, in particular, offer excellent biocompatibility and flexibility, helping to reduce friction and damage between the electrode and neural tissue.
Additionally, several future research directions are proposed:
Standardized Metrics for Stability Assessment
There is currently no unified evaluation framework to scientifically and objectively measure the long-term stability of flexible deep brain neural interfaces. Future efforts could focus on establishing standardized assessment criteria based on parameters such as electrode impedance, signal-to-noise ratio, and interface condition to provide reliable benchmarks for stability.
Integration of In Vivo and In Vitro Models for Rapid Prototyping
Most current research is focused on either in vivo or in vitro experimental platforms. The future challenge lies in integrating both approaches to accelerate the development and validation of new neural interface devices.
In summary, optimizing the immune response of flexible deep brain electrodes is an interdisciplinary field that spans neuroscience, electronics, materials science, and physics. It serves as a crucial driver for advancing future intelligent control systems and long-term therapeutic solutions.
Data availability
The graphical data and other findings in this paper are available from the corresponding author upon reasonable request.
Code availability
All custom codes regarded as central to the conclusions are available from the corresponding authors upon reasonable request.
References
Zhao, Z. et al. Ultraflexible electrode arrays for months-long high-density electrophysiological mapping of thousands of neurons in rodents. Nat. Biomed. Eng. 7, 520–532 (2022).
Oganesian, L. L. & Shanechi, M. M. Brain–computer interfaces for neuropsychiatric disorders. Nat. Rev. Bioeng. 2, 653–670 (2024).
Tang, X., Shen, H., Zhao, S., Li, N. & Liu, J. Flexible brain–computer interfaces. Nat. Electron 6, 109–118 (2023).
Wang, J. et al. Flexible electrodes for brain–computer interface system. Adv. Mater. 35, 2211012 (2023).
Rivnay, J., Wang, H., Fenno, L., Deisseroth, K. & Malliaras, G. G. Next-generation probes, particles, and proteins for neural interfacing. Sci. Adv. 3, e1601649 (2017).
Salatino, J. W., Ludwig, K. A., Kozai, T. D. Y. & Purcell, E. K. Glial responses to implanted electrodes in the brain. Nat. Biomed. Eng. 1, 862–877 (2017).
Skaper, S. D., Facci, L., Zusso, M. & Giusti, P. An inflammation-centric view of neurological disease: beyond the neuron. Front. Cell. Neurosci. 12, 72 (2018).
Wu, N. et al. Electrode materials for brain–machine interface: a review. InfoMat 3, 1174–1194 (2021).
Shen, K., Chen, O., Edmunds, J. L., Piech, D. K. & Maharbiz, M. M. Translational opportunities and challenges of invasive electrodes for neural interfaces. Nat. Biomed. Eng. 7, 424–442 (2023).
Wang, Y., Yang, X., Zhang, X., Wang, Y. & Pei, W. Implantable intracortical microelectrodes: reviewing the present with a focus on the future. Microsyst. Nanoeng. 9, 7 (2023).
Lyu, J. et al. Microglial responses to brain injury and disease: functional diversity and new opportunities. Transl. Stroke Res. 12, 474–495 (2021).
Oakes, R. S., Polei, M. D., Skousen, J. L. & Tresco, P. A. An astrocyte derived extracellular matrix coating reduces astrogliosis surrounding chronically implanted microelectrode arrays in rat cortex. Biomaterials 154, 1–11 (2018).
He, Y., Liu, X. & Chen, Z. Glial scar—a promising target for improving outcomes after CNS injury. J. Mol. Neurosci. 70, 340–352 (2020).
Song, E., Li, J., Won, S. M., Bai, W. & Rogers, J. A. Materials for flexible bioelectronic systems as chronic neural interfaces. Nat. Mater. 19, 590–603 (2020).
Gulino, M., Kim, D., Pané, S., Santos, S. D. & Pêgo, A. P. Tissue response to neural implants: the use of model systems toward new design solutions of implantable microelectrodes. Front. Neurosci. 13, 689 (2019).
Chen, Y. et al. How is flexible electronics advancing neuroscience research? Biomaterials 268, 120559 (2021).
Servais, B., Mahmoudi, N., Gautam, V. & Tong, W. Engineering brain-on-a-chip platforms. Nat. Rev. Bioeng. 2, 691–709 (2024).
Lecomte, A., Descamps, E. & Bergaud, C. A review on mechanical considerations for chronically-implanted neural probes. J. Neural Eng. 15, 031001 (2018).
Seo, K. J. et al. A soft, high-density neuroelectronic array. npj Flex. Electron 7, 40 (2023).
Hong, G. & Lieber, C. M. Novel electrode technologies for neural recordings. Nat. Rev. Neurosci. 20, 330–345 (2019).
Huang, Y. et al. Flexible electronic-photonic 3D integration from ultrathin polymer chiplets. npj Flex. Electron 8, 61 (2024).
Hong, G., Yang, X., Zhou, T. & Lieber, C. M. Mesh electronics: a new paradigm for tissue-like brain probes. Curr. Opin. Neurobiol. 50, 33–41 (2018).
Mo, F. et al. Single-neuron detection of place cells remapping in short-term memory using motion microelectrode arrays. Biosens. Bioelectron. 217, 114726 (2022).
Yang, X. et al. Bioinspired neuron-like electronics. Nat. Mater. 18, 510–517 (2019).
Du, Z. J. et al. Ultrasoft microwire neural electrodes improve chronic tissue integration. Acta Biomater. 53, 46–58 (2017).
Tian, Y. et al. An ultraflexible electrode array for large‐scale chronic recording in the nonhuman primate brain. Adv. Sci. 10, 2302333 (2023).
Lee, K. et al. Flexible, scalable, high channel count stereo-electrode for recording in the human brain. Nat. Commun. 15, 218 (2024).
Kim, J. H. et al. Flexible deep brain neural probe for localized stimulation and detection with metal guide. Biosens. Bioelectron. 117, 436–443 (2018).
Ferro, M. D. et al. NeuroRoots, a bio-inspired, seamless Brain Machine Interface device for long-term recording. AIP Adv. 14, 085109 (2024).
Wang, Y. et al. Flexible multichannel electrodes for acute recording in nonhuman primates. Microsyst. Nanoeng. 9, 93 (2023).
Fu, T.-M. et al. Stable long-term chronic brain mapping at the single-neuron level. Nat. Methods 13, 875–882 (2016).
Du Roure, O. et al. Force mapping in epithelial cell migration. Proc. Natl. Acad. Sci. USA 102, 2390–2395 (2005).
Luan, L. et al. Ultraflexible nanoelectronic probes form reliable, glial scar–free neural integration. Sci. Adv. 3, e1601966 (2017).
Wei, X. et al. Nanofabricated ultraflexible electrode arrays for high‐density intracortical recording. Adv. Sci. 5, 1700625 (2018).
Wei, C. et al. Distributed implantation of a flexible microelectrode array for neural recording. Microsyst. Nanoeng. 8, 50 (2022).
Musk, E. & Neuralink. An integrated brain-machine interface platform with thousands of channels. J. Med. Internet Res. 21, e16194 (2019).
Lee, J. M. et al. Scalable three-dimensional recording electrodes for probing biological tissues. Nano Lett. 22, 4552–4559 (2022).
Lee, J. M. The ultra-thin, minimally invasive surface electrode array NeuroWeb for probing neural activity. Nature Communications (2023).
Chik, G. K. K. et al. Flexible multichannel neural probe developed by electropolymerization for localized stimulation and sensing. Adv. Mater. Technol. 7, 2200143 (2022).
Zhao, S. et al. Tracking neural activity from the same cells during the entire adult life of mice. Nat. Neurosci. 26, 696–710 (2023).
Hong, G. et al. Tissue-like neural probes for understanding and modulating the brain. Biochemistry 57, 3995–4004 (2018).
Hong, G. et al. Syringe injectable electronics: precise targeted delivery with quantitative input/output connectivity. Nano Lett. 15, 6979–6984 (2015).
Lee, J. M. et al. Nanoenabled direct contact interfacing of syringe-injectable mesh electronics. Nano Lett. 19, 5818–5826 (2019).
Lee, J. M. et al. Stitching flexible electronics into the brain. Adv. Sci. 10, 2300220 (2023).
Xiang, Z., Liu, J. & Lee, C. A flexible three-dimensional electrode mesh: An enabling technology for wireless brain–computer interface prostheses. Microsyst. Nanoeng. 2, 16012 (2016).
Lee, S. H. et al. Scalable thousand channel penetrating microneedle arrays on flex for multimodal and large area coverage brainmachine interfaces. Adv. Funct. Mater. 32, 2112045 (2022).
Kang, Y. N., Chou, N., Jang, J.-W., Choe, H. K. & Kim, S. A 3D flexible neural interface based on a microfluidic interconnection cable capable of chemical delivery. Microsyst. Nanoeng. 7, 66 (2021).
Merken, L., Schelles, M., Ceyssens, F., Kraft, M. & Janssen, P. Thin flexible arrays for long-term multi-electrode recordings in macaque primary visual cortex. J. Neural Eng. 19, 066039 (2022).
Lee, J. Y. et al. Foldable three dimensional neural electrode arrays for simultaneous brain interfacing of cortical surface and intracortical multilayers. npj Flex. Electron 6, 86 (2022).
Yang, Y. et al. Ultraflexible neural probes for multidirectional neuronal activity recordings over large spatial and temporal scales. Nano Lett. 23, 8568–8575 (2023).
Opie, N. L. et al. Focal stimulation of the sheep motor cortex with a chronically implanted minimally invasive electrode array mounted on an endovascular stent. Nat. Biomed. Eng. 2, 907–914 (2018).
Oxley, T. J. et al. Minimally invasive endovascular stent-electrode array for high-fidelity, chronic recordings of cortical neural activity. Nat. Biotechnol. 34, 320–327 (2016).
Pancaldi, L. et al. Flow driven robotic navigation of microengineered endovascular probes. Nat. Commun. 11, 6356 (2020).
Zhang, A. et al. Ultraflexible endovascular probes for brain recording through micrometer-scale vasculature. Science 381, 306–312 (2023).
Li, H., Wang, J. & Fang, Y. Bioinspired flexible electronics for seamless neural interfacing and chronic recording. Nanoscale Adv. 2, 3095–3102 (2020).
Lv, S. et al. Tentacle microelectrode arrays uncover soft boundary neurons in hippocampal CA1. Adv. Sci 2401670 (2024).
Fan, P. et al. Flexible microelectrode array probe for simultaneous detection of neural discharge and dopamine in striatum of mice aversion system. Sens. Actuators B Chem. 390, 133990 (2023).
Pas, J. et al. A bilayered PVA/PLGA-bioresorbable shuttle to improve the implantation of flexible neural probes. J. Neural Eng. 15, 065001 (2018).
Zhou, Y. et al. A silk-based self-adaptive flexible opto-electro neural probe. Microsyst. Nanoeng. 8, 118 (2022).
Guan, S. et al. Elastocapillary self-assembled neurotassels for stable neural activity recordings. Sci. Adv. 5, eaav2842 (2019).
Cointe, C. et al. Scalable batch fabrication of ultrathin flexible neural probes using a bioresorbable silk layer. Microsyst. Nanoeng. 8, 21 (2022).
Lecomte, A. et al. Silk and PEG as means to stiffen a parylene probe for insertion in the brain: toward a double time-scale tool for local drug delivery. J. Micromech. Microeng. 25, 125003 (2015).
Wen, D.-L. et al. Recent progress in silk fibroin-based flexible electronics. Microsyst. Nanoeng. 7, 35 (2021).
Rauhala, O. J. et al. Chitosan‐based, biocompatible, solution processable films for in vivo localization of neural interface devices. Adv. Mater. Technol. 5, 1900663 (2020).
Ojeda-Hernández, D. D., Canales-Aguirre, A. A., Matias-Guiu, J., Gomez-Pinedo, U. & Mateos-Díaz, J. C. Potential of chitosan and its derivatives for biomedical applications in the central nervous system. Front. Bioeng. Biotechnol. 8, 389 (2020).
Xie, C. et al. Three-dimensional macroporous nanoelectronic networks as minimally invasive brain probes. Nat. Mater. 14, 1286–1292 (2015).
Wen, X. et al. Flexible, multifunctional neural probe with liquid metal enabled, ultra-large tunable stiffness for deep-brain chemical sensing and agent delivery. Biosens. Bioelectron. 131, 37–45 (2019).
Villa, J., Cury, J., Kessler, L., Tan, X. & Richter, C.-P. Enhancing biocompatibility of the brain-machine interface: a review. Bioact. Mater. 42, 531–549 (2024).
Sayyad, P. W., Park, S.-J. & Ha, T.-J. Bioinspired nanoplatforms for human-machine interfaces: recent progress in materials and device applications. Biotechnol. Adv. 70, 108297 (2024).
Kämmerling, L. et al. Mitigating the foreign body response through ‘immune-instructive’ biomaterials. J. Immunol. Regenerative Med. 12, 100040 (2021).
Song, I. & Dityatev, A. Crosstalk between glia, extracellular matrix and neurons. Brain Res. Bull. 136, 101–108 (2018).
Chelyshev, Y. A., Kabdesh, I. M. & Mukhamedshina, Y. O. Extracellular matrix in neural plasticity and regeneration. Cell Mol. Neurobiol. 42, 647–664 (2022).
Amini, S., Salehi, H., Setayeshmehr, M. & Ghorbani, M. Natural and synthetic polymeric scaffolds used in peripheral nerve tissue engineering: Advantages and disadvantages. Polym. Adv. Technol. 32, 2267–2289 (2021).
Vitale, F. et al. Biomimetic extracellular matrix coatings improve the chronic biocompatibility of microfabricated subdural microelectrode arrays. PLoS ONE 13, e0206137 (2018).
Lee, C. D. et al. Matrigel coatings for Parylene sheath neural probes. J. Biomed. Mater. Res 104, 357–368 (2016).
Qi, Y., Kang, S.-K., Fang, H. & Editors, Guest Advanced materials for implantable neuroelectronics. MRS Bull. 48, 475–483 (2023).
He, W., McConnell, G. C. & Bellamkonda, R. V. Nanoscale laminin coating modulates cortical scarring response around implanted silicon microelectrode arrays. J. Neural Eng. 3, 316–326 (2006).
Ghane-Motlagh, B. et al. Physicochemical properties of peptide-coated microelectrode arrays and their in vitro effects on neuroblast cells. Mat. Sci. Eng. C 68, 642–650 (2016).
Tarus, D. et al. Design of hyaluronic acid hydrogels to promote neurite outgrowth in three dimensions. ACS Appl. Mater. Interfaces 8, 25051–25059 (2016).
Hu, Y. et al. Resilient and self-healing hyaluronic acid/chitosan hydrogel with ion conductivity, low water loss, and freeze-tolerance for flexible and wearable strain sensor. Front. Bioeng. Biotechnol. 10, 837750 (2022).
Swingler, S. et al. Recent advances and applications of bacterial cellulose in biomedicine. Polymers 13, 412 (2021).
Robbins, M. et al. Biofunctionalised bacterial cellulose scaffold supports the patterning and expansion of human embryonic stem cell-derived dopaminergic progenitor cells. Stem Cell Res Ther. 12, 574 (2021).
Yang, J. et al. Bacterial cellulose as a supersoft neural interfacing substrate. ACS Appl. Mater. Interfaces 10, 33049–33059 (2018).
Kozai, T. D. Y. et al. Chronic in vivo evaluation of PEDOT/CNT for stable neural recordings. IEEE Trans. Biomed. Eng. 63, 111–119 (2016).
Li, X. et al. Flexible electrocorticography electrode array for epileptiform electrical activity recording under glutamate and GABA modulation on the primary somatosensory cortex of rats. Micromachines 11, 732 (2020).
Li, X. et al. PDMS–parylene hybrid, flexible micro-ECoG electrode array for spatiotemporal mapping of epileptic electrophysiological activity from multicortical brain regions. ACS Appl. Bio Mater. 4, 8013–8022 (2021).
Lim, J. et al. Hybrid graphene electrode for the diagnosis and treatment of epilepsy in free-moving animal models. NPG Asia Mater. 15, 7 (2023).
Park, S. et al. Adaptive and multifunctional hydrogel hybrid probes for long-term sensing and modulation of neural activity. Nat. Commun. 12, 3435 (2021).
Wang, K. et al. Electrodeposition of alginate with PEDOT/PSS coated MWCNTs to make an interpenetrating conducting hydrogel for neural interface. Compos 26, 27–40 (2019).
Rinoldi, C. et al. In vivo chronic brain cortex signal recording based on a soft conductive hydrogel biointerface. ACS Appl. Mater. Interfaces 15, 6283–6296 (2023).
Lee, Y. et al. A lubricated nonimmunogenic neural probe for acute insertion trauma minimization and long‐term signal recording. Adv. Sci. 8, 2100231 (2021).
Sung, C. et al. Multimaterial and multifunctional neural interfaces: from surface-type and implantable electrodes to fiber-based devices. J. Mater. Chem. B 8, 6624–6666 (2020).
Ziai, Y. et al. Conducting polymer‐based nanostructured materials for brain–machine interfaces. WIRES Nanomed. Nanobi 15, e1895 (2023).
Qazi, R. et al. Wireless optofluidic brain probes for chronic neuropharmacology and photostimulation. Nat. Biomed. Eng. 3, 655–669 (2019).
Seyfoddin, A. et al. Electro-responsive macroporous polypyrrole scaffolds for triggered dexamethasone delivery. Eur. J. Pharm. Biopharm. 94, 419–426 (2015).
Wadhwa, R., Lagenaur, C. F. & Cui, X. T. Electrochemically controlled release of dexamethasone from conducting polymer polypyrrole coated electrode. J. Control Release 110, 531–541 (2006).
Kim, D., Abidian, M. & Martin, D. C. Conducting polymers grown in hydrogel scaffolds coated on neural prosthetic devices. J. Biomed. Mater. Res. 71A, 577–585 (2004).
Wu, C. et al. Electrochemical deposition of Ppy/Dex/ECM coatings and their regulation on cellular responses through electrical controlled drug release. Colloids Surf. B 222, 113016 (2023).
Abidian, M. R., Kim, D. ‐H. & Martin, D. C. Conducting‐polymer nanotubes for controlled drug release. Adv. Mater. 18, 405–409 (2006).
Abidian, M. R. & Martin, D. C. Multifunctional nanobiomaterials for neural interfaces. Adv. Funct. Mater. 19, 573–585 (2009).
Kim, D.-H. & Martin, D. C. Sustained release of dexamethasone from hydrophilic matrices using PLGA nanoparticles for neural drug delivery. Biomaterials 27, 3031–3037 (2006).
Bagheri, B. et al. Self-gelling electroactive hydrogels based on chitosan–aniline oligomers/agarose for neural tissue engineering with on-demand drug release. Colloids Surf. B 184, 110549 (2019).
Heo, D. N. et al. Multifunctional hydrogel coatings on the surface of neural cuff electrode for improving electrode-nerve tissue interfaces. Acta Biomater. 39, 25–33 (2016).
Zhong, Y. & Bellamkonda, R. V. Controlled release of anti-inflammatory agent α-MSH from neural implants. J. Control Release 106, 309–318 (2005).
Eles, J. R. et al. Neuroadhesive L1 coating attenuates acute microglial attachment to neural electrodes as revealed by live two-photon microscopy. Biomaterials 113, 279–292 (2017).
Azemi, E., Lagenaur, C. F. & Cui, X. T. The surface immobilization of the neural adhesion molecule L1 on neural probes and its effect on neuronal density and gliosis at the probe/tissue interface. Biomaterials 32, 681–692 (2011).
Cui, X. In vivo studies of polypyrrole/peptide coated neural probes. Biomaterials 24, 777–787 (2003).
Righi, M. et al. Peptide-based coatings for flexible implantable neural interfaces. Sci. Rep. 8, 502 (2018).
Kil, D. et al. Dextran as a resorbable coating material for flexible neural probes. Micromachines 10, 61 (2019).
Zou, Y. et al. Anti-fouling peptide functionalization of ultraflexible neural probes for long-term neural activity recordings in the brain. Biosens. Bioelectron. 192, 113477 (2021).
Povlich, L. K. et al. Synthesis, copolymerization and peptide-modification of carboxylic acid-functionalized 3,4-ethylenedioxythiophene (EDOTacid) for neural electrode interfaces. BBA-Gen. Subj. 1830, 4288–4293 (2013).
Aregueta-Robles, U. A., Woolley, A. J., Poole-Warren, L. A., Lovell, N. H. & Green, R. A. Organic electrode coatings for next-generation neural interfaces. Front. Neuroeng. 7, (2014).
Bhagwat, N., Murray, R. E., Shah, S. I., Kiick, K. L. & Martin, D. C. Biofunctionalization of PEDOT films with laminin-derived peptides. Acta Biomater. 41, 235–246 (2016).
Nam, J. et al. Supramolecular peptide hydrogel-based soft neural interface augments brain signals through a three-dimensional electrical network. ACS Nano 14, 664–675 (2020).
Kozai, T. D. Y., Jaquins-Gerstl, A. S., Vazquez, A. L., Michael, A. C. & Cui, X. T. Brain tissue responses to neural implants impact signal sensitivity and intervention strategies. ACS Chem. Neurosci. 6, 48–67 (2015).
Steinmetz, N. A. et al. Neuropixels 2.0: A miniaturized high-density probe for stable, long-term brain recordings. Science 372, eabf4588 (2021).
Liu, Y. et al. A high-density 1,024-channel probe for brain-wide recordings in non-human primates. Nat. Neurosci. 27, 1620–1631 (2024).
Zhou, Y. et al. A mosquito mouthpart-like bionic neural probe. Microsyst. Nanoeng. 9, 88 (2023).
Wei, S. et al. Shape-changing electrode array for minimally invasive large-scale intracranial brain activity mapping. Nat. Commun. 15, 715 (2024).
Cho, M. et al. Fully bioresorbable hybrid opto-electronic neural implant system for simultaneous electrophysiological recording and optogenetic stimulation. Nat. Commun. 15, 2000 (2024).
Acknowledgements
Research supported by the National Key Research and Development Program of China (2022YFC2402501, 2022YFB3205602), the National Natural Science Foundation of China (Nos. 62121003, T2293730, T2293731, 62333020, 62171434, and 62471291), the Major Program of Scientific and Technical Innovation 2030 (2021ZD02016030), the Joint Foundation Program of the Chinese Academy of Sciences (No. 8091A170201), and the Scientific Instrument Developing Project of the Chinese Academy of Sciences (No. PTYQ2024BJ0009), the National Natural Science Foundation of Beijing (F252069).
Author information
Authors and Affiliations
Contributions
S.Y.L.: Formal analysis, Writing - Original Draft. Z.J.X., F.M. & Y.W.: Investigation. Y.M.D., Y.Y.L., L.Y.J., & J.S.: Validation. Q.L.J., M.C.W., S.Y.Z., & Y.L.: Investigation. J.T.L. & J.P.L.: Formal analysis. Y.R.W., Y.L.S., M.X.W. & X.X.C.: Funding acquisition, Project administration, Supervision. All authors read and approved the final manuscript.
Corresponding authors
Ethics declarations
Competing interests
The authors declare no competing interests.
Additional information
Publisher’s note Springer Nature remains neutral with regard to jurisdictional claims in published maps and institutional affiliations.
Rights and permissions
Open Access This article is licensed under a Creative Commons Attribution-NonCommercial-NoDerivatives 4.0 International License, which permits any non-commercial use, sharing, distribution and reproduction in any medium or format, as long as you give appropriate credit to the original author(s) and the source, provide a link to the Creative Commons licence, and indicate if you modified the licensed material. You do not have permission under this licence to share adapted material derived from this article or parts of it. The images or other third party material in this article are included in the article’s Creative Commons licence, unless indicated otherwise in a credit line to the material. If material is not included in the article’s Creative Commons licence and your intended use is not permitted by statutory regulation or exceeds the permitted use, you will need to obtain permission directly from the copyright holder. To view a copy of this licence, visit http://creativecommons.org/licenses/by-nc-nd/4.0/.
About this article
Cite this article
Lv, S., Xu, Z., Mo, F. et al. Long-term stability strategies of deep brain flexible neural interface. npj Flex Electron 9, 40 (2025). https://doi.org/10.1038/s41528-025-00410-x
Received:
Accepted:
Published:
DOI: https://doi.org/10.1038/s41528-025-00410-x
This article is cited by
-
Integrated implantable bioelectronic system based on intrinsically stretchable and conductive nanocomposites
Health Nanotechnology (2025)
-
Innovative Approaches in Regenerative Medicine for the Development of Sustainable and Bioactive Implants
Biomedical Materials & Devices (2025)
