Implantable microelectrodes are useful for monitoring neural response patterns in the auditory cortex, however chronic neural recordings can often deteriorate with time (e.g. impedance measures across electrode arrays generally increase monotonically over the first 7 days post-implant). This problem is caused by the increasing spatial distribution of reactive tissue responses (corresponding to changes in impedance spectra along the electrode-tissue-interface). Therefore, the design of microelectrode probes must ensure that the neuronal ensembles lie within a cylindrical radius of the recording electrodes. In this chapter, chronic neural recording failure is examined via cortical spike patterns, histological analyses, indentation experiments, and finite element models. Next, the microfabrication of the “Utah” electrode array and the “Michigan” probe is compared to determine how their size, shape, and geometry address: (1) the spatial distribution of neurons (as related to recording quality); (2) the initial penetrating profile (as related to insertion killzones); (3) the reactive cell responses (as related to glial encapsulation); (4) the anchoring of the probe’s position in the tissue (as related to micromotions) and (5) the embedding of various bioactive reagents (ex: growth factors, anti-inflammatory drugs, etc.). Finally, a novel hydrogel “Dropping Method” is proposed for controlling the biocompatibility and mechanical properties at the electrode-tissue-interface.
- auditory cortex
- implantable microelectrode array
- chronic neural recording
- neural probe
- brain machine interface
- inflammatory response
- insertion killzone
- glial encapsulation
- finite element model
This chapter will focus on the
Multichannel microelectrodes (Figure 1B) can be used to monitor the activity in the
In Section 2, chronic neural recording failure is examined by reviewing cortical spike patterns, histological analyses, indentation experiments, and
In Section 3, the microfabrication of the
In Section 4, a
2. Chronic neural recording failure
2.1. Deterioration of neural recordings over time
Silicon-substrate micromachined probes can provide high-quality multichannel spike activity recordings and local field potentials in the auditory cortex. For acute neural recordings, these measurements are generally stable. However, chronic long-term recordings are limited by the foreign body responses that occur during the wound healing process . Previous research showed signal-to-noise-ratios (SNR) decreased as a function of postoperative time (Figure 2A) . Impedance measures across electrode arrays were also shown to increase monotonically over the first 2 weeks (Figure 2B). This problem (of electrical signals gradually decreasing over time post-implant) must be solved to improve the chronic recording of neural activity.
Histological analyses have revealed tissue reactions to the implanted probe shanks during the assessment periods. For instance, confocal microscopy and immunohistochemistry protocols were used to measure the cellular density around the implanted electrodes and confirmed that a region of extensive reactive responses occurs around each individual electrode track . By differentiating the spatial distribution of reactive tissue responses corresponding to changes in impedance spectra along the electrode-tissue interface, these findings confirm that reactive tissue responses directly influence the impedance spectra over time.
The recorded spike amplitudes were also measured as a function of the distance between the neuron and the electrode (by making comparisons from a multisite tetrode wire ). Figure 3 shows that the extracellularly recorded spike amplitudes decreased rapidly with distance (a finding that indicates that the neuronal ensembles must lie within a cylindrical radius of ~140 μm of the recording electrode to effectively monitor the extracellular spike amplitudes). These “
2.2. Insertion killzone
When inserting a microelectrode, the microshaft must penetrate through arteries, veins, tight junction cells, collagen, and smaller vessels in the pia. The shaft must then penetrate through tissues composed of neurons, glial cells, capillaries, arterioles, and venules. The tip of the shaft must be sharp enough to puncture through the microtubules and the fibrous structures within the neurofilaments. This electrode insertion process inevitably causes death and degeneration of its neighboring neurons and capillaries during the surgical implantation . The
Penetrating shafts should be designed to minimize the initial mechanical trauma, to preserve the neuronal density in a local area around the implanted array, and to minimize the early reactive responses caused by the pathway of tissue damage. Indentation experiments have shown that immediate vascular and brain damage results in the recruitment of cells from the peripheral immune system and the activation of resident astrocytes or microglia. The reactive responses can be compared by immunochemically labelling for glial fibrillary acidic protein (GFAP), vimentin (for astrocytes), or ED1 (for microglia). When devices of different size were inserted, it was found that the volume of the reactive tissue responses was proportional to the cross-sectional area . For instance, the smallest device (smallest cross-sectional area) with smooth surfaces and rounded corners caused less damage to the tissue, produced a smaller volume of reactive tissue, and left a smaller hole when removed from the tissue 1 week later.
2.3. Prolonged injury response
An immunohistochemistry experiment  showed that the early reactive response (within 1 week (Figure 5A)) was associated with the amount of damage generated during the insertion (which depends on the device size or shape). However, this same study also showed that the prolonged reactive responses (after 4 weeks (Figure 5B)) eventually became similar (ultimately resulting in a compact cellular sheath containing astrocytes and microglia). While it is always better to minimize the insertion killzone, these results clearly indicate that a second sustained response occurs that is related to the tissue-device interactions. For instance, a long-term study  showed that the persistent ED1 up-regulation and neuronal loss (associated with the foreign body responses) were not observed in microelectrode stab controls (which indicates that the phenotype did not result from the initial mechanical trauma of the electrode implantation). Moreover, chronically-implanted electrodes were also covered in ED1/MAC-1 immunoreactive cells and released the pro-inflammatory cytokines MCP-1 and TNF-α (which are only characteristics of chronic inflammatory reactions). Figure 6 shows that the zones of astrocytosis and connective tissue vary in proportion to an implant’s reactivity . Although the normal long-term response to an indwelling microelectrode is to develop a reactive glial tissue (that eventually forms into a fibrotic encapsulation layer or a glial scar), this sustained injury response has also been associated with chronic recording failure since the encapsulation (from persistent inflammatory reactions) isolates the electrodes from the surrounding neurons and decreases the stability and quality of the neural recordings.
The mechanical properties of the brain can be researched by mathematical simulation models that faithfully represent the geometry, material properties, and boundary or load conditions. For instance, an in-vivo indentation experiment  described the soft tissue deformation with a three-dimensional non-linear finite element model (FEM) of geometric information from the brain (obtained via magnetic resonance imaging techniques). Another FEM study measured the effects of tethering forces, probe-tissue adhesion, and the stiffness of the probe substrate on the interfacial strains induced around the implant site . The results indicated that the interfacial strains were created by
3. Comparing “Utah” electrode arrays versus “Michigan” neural probes
Implantable probe designs have primarily focused on modifying the size, shape, or geometry to minimize reactive cell responses. For the
3.1. Spatial distribution of neurons
3.1.1. Utah electrode array
The Utah array was designed from the ground up (with new manufacturing techniques ) to meet the specific needs of a multichannel neural interface . Psychophysical experiments usually require evoking discriminable patterned percepts from many electrodes. Therefore, the Utah array (Figure 8A) was designed to possess a large number (~100) of large electrodes (~1.5 mm) that are typically in a square grid that projects out from a thin (~0.2 mm) substrate. Figure 8B shows the tapered electrodes suspended in a “sea of glass” substrate (that isolates each of the individual electrodes in the array from each other (~0.4 mm separation) to form a very effective dielectric insulating layer between the adjacent electrodes). The electrode probe tips are coated with platinum, gold, or iridium to facilitate the electronic to ionic transduction. The array’s substrate was designed to be thick enough to prevent breaking upon insertion, but thin enough to rest on the cortical surface without producing a constant downward force on the array that could push it further into the cortex.
3.1.2. Michigan neural probe
Unlike the Utah electrode array , the planar Michigan probe  was only designed to take advantage of photolithographic manufacturing techniques from the semiconductor industry (Figure 9A). The extensibility of this platform technology enables these micromachined probes to be built with batch fabrication, easy customization of recording site placements or substrate shape, high reproducibility of geometrical/electrical characteristics, and the ability to integrate with ribbon cables or to incorporate on-chip electronics for signal conditioning (Figure 9B).
Unfortunately, the Michigan probes are known to induce a chronic breach of the blood–brain barrier  which leads to more chronic inflammation and culminates in neurodegeneration and ultimately to electrode failure (as described in Section 2). For instance, Michigan probes showed a significantly higher breach and worse wound-healing in comparison to microwires.
Since there is a trade-off between the size (spatial selectivity) and quality of signal recordings (sensitivity) in a neural microelectrode, previous research have altered the electrical properties by synthesizing biocompatible conducting polymers (Figure 10A) such as polypyrrole (PPy) and poly(3,4-ethylenedioxythiophene) (PEDOT) directly onto the electrode. Electrochemical deposition (Figure 10B) allows polymer films to be formed in a one-step process with a high degree of control over the film thickness and surface properties (e.g. with nanotubes in ).
3.2. Initial penetrating profile
3.2.1. Utah electrode array
The penetrating electrodes of the Utah array (Figure 11) were designed to be slender enough to retain sufficient strength for withstanding the implantation procedure, yet to compromise as little cortical volume as possible (only 80 μm in diameter at the base). The needles were intentionally designed to have a cylindrical/conical geometry (rather than a planar geometry) to displace the tissues they are inserted (rather than cutting their way through). In addition, an “Impact Insertion” technique was designed to inject the array into the cortex at a high velocity . This momentum transfer tool preserves uniformity and prevents dimpling of the cortex.
3.2.2. Michigan neural probe
Despite evidence that smaller implants can increase the survival of neurons , the research direction for Michigan probes has been primarily guided by the
3.3. Reactive processes
3.3.1. Utah electrode array
The electrodes are built from biocompatible materials such as silicon, silicon nitride, silicon dioxide, platinum, titanium, tungsten, and silicone. Histological analyses show the neuronal cell bodies remain in close apposition to the electrode tracks (especially in sections where the tracks are similar in diameter to blood vessels). The benign tissue response showed that a thin capsule (~2–5 microns) forms around each electrode track. Histological samples also revealed gliosis, fibrotic-tissue buildup between the array and the meninges, bleeding in some tracks, and some array displacement through the cortex. Despite these reactive responses, electrodes were shown to be able to record single- and multi-unit responses in the cortex for over 3 years (the longest intervals studied). Therefore, the stability of these units over periods of months provides the most compelling evidence for the biocompatibility .
3.3.2. Michigan neural probe
Glial sheaths (consisting of activated microglia and hypertrophied astrocytes, meningeal cells, and oligodendrocyte precursors that produce extracellular proteins that hinder local nerve regeneration) have been shown to form around the probe tract. As mentioned in Section 2, this tissue encapsulation is concomitant with a decrease in signal quality for neural recordings. Therefore, probe geometry has been investigated as a parameter for reducing chronic tissue encapsulation. Open-architectures (Figure 13) can reduce tissue encapsulation by presenting a narrow edge of surface area to prevent the prototypical attachment and spreading of foreign body responses (as cells would theoretically be unable to attach or create cytoskeleton tension below a certain dimension ). The tissue reactivity around the smallest lattice structure was expected to induce the least encapsulation. However, the comparison between probe designs indicated that the differences in encapsulation and neuronal loss was insignificant. Qualitative histology also showed similar responses inside the lattice region and around the shank. This has led to the
3.4. Anchoring of the probe’s position
3.4.1. Utah electrode array
The large number of penetrating electrodes presents a very large surface area to the cortex and the implanted array tends to self-anchor to the cortical tissues. More importantly, the “sea of glass” allows the array to float in the cortical tissues as the cortex moves due to respiration, blood pumping, and skeletal displacements. This design feature produces an extremely stable interface with the surrounding neurons as it moves with the cortex and thereby produces very little micromotion between the electrode tips and the neurons near its active tips. In addition, GFAP staining showed that a fixation mode that un-tethers the implant from the skull elicits a smaller tissue reaction and results in the survival of a larger number of neurons in the region closest to the tissue interface .
3.4.2. Michigan neural probe
In Sections 3.2.2 and 3.3.2, the
The theory of
3.5. Embedding of bioactive reagents
3.5.1. Utah electrode array
The microneedle array can be fabricated with porous tips that can be loaded with drugs that have a high molecular weight . These porous tips are nano-structured silicon (Figure 14A), a material that is biocompatible, bioactive, biodegradable, and appropriate for the cultivation of adherent cells in-vivo without noticeable toxicity for biological applications. Drug delivery is also possible in the 3D floating silicon array by using a silicon platform that has microfluidic cables with adapters used for liquid supply . Selected shafts can be equipped with fluidic integration (Figure 14B) with either one independent fluidic channel (to maximize the volume coverage) or two independent microchannels (to infuse both a drug and a buffer reference at the same location) when combining drug delivery within the array of recording electrodes.
3.5.2. Michigan neural probe
Planar microelectrodes can directly utilize innovations from micro-fabrication technology and lab-on-a-chip delivery devices. For instance, aerosol jet printing technology can be utilized to directly construct a patterned neural interface with anti-inflammatory nanocarriers . 3D probe structures with microfluidic channels can also be fabricated via surface micromachining and deep reactive ion etching (DRIE) . Neural probe designs (Figure 15) can also be bulk-microfabricated to include microchannels along the shanks for microscale and controlled fluid delivery through the blood–brain barrier .
4. Hydrogel improves biocompatibility and mechanical properties
4.1. Improved biocompatibility
Alginate hydrogels have been successfully applied as a coating on the Michigan probes. One study  showed that biodegradable neurotrophin-eluting hydrogels can be applied on the microelectrode arrays to attract neurites to the surface of the electrodes (for improved neuron-electrode proximity). Hydrogels can also be loaded with nanoparticles of anti-inflammatory agent dexamethasone (DEX) for drug delivery . In-vitro drug release kinetics has revealed that 90% of the DEX can be successfully released from entrapped nanoparticles over 2 weeks. In addition, the impedance of the nanoparticle-loaded hydrogel coatings on microfabricated neural probes were equivalent to the unmodified/uncoated probes controls (indicating that the loaded-nanoparticles do not hinder electrical transport). Most importantly, the chronically implanted electrodes loaded with DEX were shown to maintain impedances in-vivo (unlike the control electrodes that showed increases after 2 weeks of implantation). This improvement in chronic neural recordings indicates the DEX-modified neural probes reduced the amount of glial inflammation via local administration of therapeutic agents. Hydrogels can also be used as a scaffold to encapsulate drug-incorporated biodegradable nanofibers for a more sustained and slower release (to reduce the burst effect ).
4.2. Improved mechanical properties
Hydrogels have also been proposed to improve the mechanical properties of the implantable microelectrodes. First, hydrogel coatings can be dehydrated to minimize the initial penetrating profile . Second, the degree of reswelling can be controlled by using different cross-linking molecules to better anchor the position of the probe in the tissue after the implantation . Third, the shear moduli of a hydrogel can be tightly regulated by controlling the cross-linking densities to reduce the stiffness mismatch between the hard silicon-based probes (~100 GPa) and the soft tissue (~10 kPa). Fourth, the uniformity of the hydrogel coating can be controlled by changing the concentration of the solution. Lastly, the hydrogel thickness can be controlled by quickly drying the coated surfaces in air and then overlaying additional layers on top.
Despite the many advantages of hydrogels, the spatial distribution of neurons (as it relates to recording quality) has been a barrier to the practical implementation of hydrogel coatings on microelectrodes. For instance, neural recordings in the auditory cortex showed a significant loss in functionality (as determined by the number of clearly detectable units and the average signal-to-noise ratios ). In fact, in-vivo experiments showed that only 30% of the electrode sites (with 80 μm thick hydrogel coatings) could record detectable signals from surrounding neurons and that the average signal-to-noise ratio dramatically decreased (even for 5 μm thick coatings). The loss in neural recording quality is hypothesized to be caused by the reswelling properties of the dehydrated hydrogel coatings (which can affect the spatial distribution of the target neurons around the implanted electrode). In other words, the low signal detection is hypothesized to be caused by the post-implantation water absorption which pushes the target neurons away from the electrode surfaces as they rehydrate (and beyond the
To solve the hydrogel swelling problem,  proposed to modify the electrode design to place the recording sites in closer proximity to the hydrogel’s surface (Figure 17). The deposition of conducting polymer PEDOT has also been shown to restore some of the lost functionality of electrode sites with thicker hydrogel coatings . Alternatively, fibrin can be selected (rather than Alginate) to be reabsorbed in the surrounding tissues 7 days after the implantation .
4.3. The “Dipping Method”
While previous solutions to the hydrogel swelling problem have mostly proposed physical changes to the actual microelectrode or the hydrogel [39, 40, 41, 42, 43, 44, 45], most attempts still rely on the “
In practice, the Dipping Method has several problems. First, lateral movements of the neural electrode effect the reproducibility of the alginate coating around the shank of the probe . Second, it is impossible to control the average thickness by just approximating the number of dipping cycles (as verified by ellipsometry data in Figure 4B of ). Third, it is impossible to control the maximum radial thickness by just altering the number of dips (as demonstrated in Figure 2C of ). Fourth, the speed of dipping effects the uniformity of the alginate coating around the shank of the probe . These limitations of the Dipping Method are confounding and counter-productive when considering that the entire rationale for using hydrogel coatings was to control the mechanical properties of the electrodes in . Furthermore, the Dipping Method requires a human observer to subjectively monitor the entire dipping process with an optical microscope . Therefore, this proposed Dipping Method essentially eliminates one of the primary advantage of the Michigan probes: exploiting semiconductor manufacturing techniques to allow batch fabrication for the high reproducibility of geometrical/electrical characteristics (as highlighted in Section 3.1.2).
4.4. The “Dropping Method”
Techniques for coating hydrogels onto microelectrodes include spray coating, brush coating, Dipping Methods [39, 40, 41, 42, 43, 44, 45], and micropipetting . Unfortunately, none of these methods can produce uniform coatings in a reproducible manner . Therefore, novel strategies such as the Molding Method (which involves the photopolymerization of hydrogels in a polyethylene tube) have been proposed to provide hydrogel coatings that are more reproducible and have a means to control coating uniformity and thickness (e.g. by varying the polyethylene tubing).
Figure 19 proposes a novel procedure called the “Dropping Method” to address the reported limitations of the Dipping Method. First, the microelectrode is permanently attached in one location to eliminate lateral movements (to improve reproducibility ). Second, a stage that contains solutions is automatically revolved to eliminate the requirement of approximating movements (to improve the average thickness ). Third, the stage is dropped downward to allow gravity to align the droplets and to evenly pull down the edges (to improve the radial thickness ). Fourth, the speed of dropping is automated by a digital micromanipulator (to improve the uniformity of the alginate coating around the shank of the probe ). Finally, the optimal contact angle is automated on a Dropmeter stage via programmable software (instead of relying on subjective interpretations of a human observer via optical microscopy as in ).
FEM and histological analyses have shown that elevated local strains correspond to increases in gliosis (with a three-fold increase in gliosis at the probe tips compared to other areas of the explanted arrays ). In addition, FEM pressure profiles predict that the strain at the tip can be reduced by utilizing more flexible or softer substrates, reducing the opening angle of the probe face, and by promoting tissue integration to reduce excessive adhesion. Therefore, the Dropping Method is recommended for coating hydrogel at the tips of the microelectrodes to optimize the mechanical effects for improved chronic recording stability. Applying hydrogel coatings on the probe tips should stabilize the mechanical interface and lower the interfacial tension with the surrounding biological environment. Constraining hydrogel coating depths to just the tips will also avoid the electrode coverage issue . Finally, the initial penetrating profile should be reduced since only the dehydrated probe tips will reswell in thickness upon water absorption (unlike previous studies that coated the hydrogel layers along the entire length of the neural electrode shanks in efforts to reduce the sustained injury responses).
This chapter reviewed how controlling the biocompatibility and mechanical properties of a microelectrode is critical when implanting deeper within the auditory system. To effectively monitor extracellular spike amplitudes in the cortex, several design considerations should be factored to ensure that the neuronal ensembles lie within a cylindrical radius of the recording electrode. First, penetrating shafts should reduce the initial mechanical trauma during surgical implantation (to minimize the reactive responses caused by the pathway of tissue damage). Second, the indwelling microelectrodes should embed bioactive reagents since chronic neural recording failure is associated with sustained injury responses from persistent inflammatory reactions. Third, the mechanical properties should be controlled since micromotions can cause shearing and compression to the surrounding tissues (due to the large stiffness mismatch of the implanted substrates). This chapter also describes how hydrogels can minimize the cross-sectional area while ensuring the neuronal density is maintained in local areas surrounding the implant. A novel hydrogel coating “Dropping Method” is proposed instead of the Dipping Method to eliminate the problems of lateral movements, approximated cycles, and uniformity.
In this chapter, some parts of this work was performed when the first author was with the Centre for Biosystems, Neuroscience, and Nanotechnology, City University of Hong Kong, Hong Kong, with Stella Pang. The author is now working at the Center for Information Technology Innovation (CITI), Academia Sinica, Taipei, Taiwan with Yu Tsao.
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