**3.2. Initial penetrating profile**

#### *3.2.1. Utah electrode array*

exhibited less inflammation-related biomarker distribution in tissues compared to the solid shanks (even though both have identical penetrating profiles). However, Section 3.4.2 will show the alternative theory of *mechanical property differences* [30] can also explain these results.

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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

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 [30]). 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 *mechanical property differences* theory (which is discussed further in Section 3.4.2).

**Figure 13.** Comparison of a solid shank vs. open-architecture designs with varying lattice sizes. Reproduced from [30].

provides the most compelling evidence for the biocompatibility [23].

**3.3. Reactive processes**

*3.3.1. Utah electrode array*

*3.3.2. Michigan neural probe*

**Figure 11.** Penetrating shafts of the Utah electrode array with probe tips coated with platinum, gold, or iridium.

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 [26]. This momentum transfer tool preserves uniformity and prevents dimpling of the cortex.

#### *3.2.2. Michigan neural probe*

**Figure 12.** Comparison of a solid shank vs. the open-architecture design with a lattice structure. Reproduced from [25].

Despite evidence that smaller implants can increase the survival of neurons [29], the research direction for Michigan probes has been primarily guided by the *critical surface area model* [25]. This theory is supported by results showing microelectrodes with lattice structures (**Figure 12**) exhibited less inflammation-related biomarker distribution in tissues compared to the solid shanks (even though both have identical penetrating profiles). However, Section 3.4.2 will show the alternative theory of *mechanical property differences* [30] can also explain these results.

#### **3.3. Reactive processes**

**3.2. Initial penetrating profile**

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 [26]. This momentum transfer tool preserves uniformity and prevents dimpling of the cortex.

**Figure 11.** Penetrating shafts of the Utah electrode array with probe tips coated with platinum, gold, or iridium.

Despite evidence that smaller implants can increase the survival of neurons [29], the research direction for Michigan probes has been primarily guided by the *critical surface area model* [25]. This theory is supported by results showing microelectrodes with lattice structures (**Figure 12**)

**Figure 12.** Comparison of a solid shank vs. the open-architecture design with a lattice structure. Reproduced from [25].

*3.2.1. Utah electrode array*

182 An Excursus into Hearing Loss

*3.2.2. Michigan neural probe*

#### *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 [23].

#### *3.3.2. Michigan neural probe*

**Figure 13.** Comparison of a solid shank vs. open-architecture designs with varying lattice sizes. Reproduced from [30].

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 [30]). 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 *mechanical property differences* theory (which is discussed further in Section 3.4.2).

#### **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 [29].

**3.5. Embedding of bioactive reagents**

The microneedle array can be fabricated with porous tips that can be loaded with drugs that have a high molecular weight [32]. 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 [33]. 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.

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**Figure 14.** Platinum coated tips of the Utah electrode. Scale bar = 0.5 mm. Reproduced from [32, 33].

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 [34]. 3D probe structures with microfluidic channels can also be fabricated via surface micromachining and deep reactive ion etching (DRIE) [35]. Neural probe designs (**Figure 15**) can also be bulk-microfabricated to include microchannels along the shanks for microscale and controlled

**Figure 15.** Probe design with microchannels along the shanks for controlled fluid delivery. Reproduced from [36].

*3.5.1. Utah electrode array*

*3.5.2. Michigan neural probe*

fluid delivery through the blood–brain barrier [36].

#### *3.4.2. Michigan neural probe*

In Sections 3.2.2 and 3.3.2, the *critical surface model* was used to explain the benefits of open-architecture designs [25, 30]. However, an alternative theory [30] was offered to explain how the open-architecture designs reduce tissue encapsulations compared to the solid shanks. The theory of *mechanical property differences* assumes that lattice structures reduce the induced strain from relative movements (e.g. micromotion) between the probe and the tissue (occurring from the continual pulsations of vascular and respiratory oscillations). In other words, the adjoining lattice structure is assumed to present a more flexible mechanical interface that can improve the mechanical surface properties (relative to the solid probe shank). This competing theory of *mechanical property differences* is supported by FEM results that demonstrate that the shape of the electrode substantially influences the local pattern and the magnitude of strain around the electrode (especially at the electrode tip where the shank has a spear-head shape or at regions that have discontinuous or sharp edges [31]).

The theory of *mechanical property differences* is also supported by electrophysiological data [28] showing the Michigan electrode's recording performance failed faster relative to a microwire electrode. Since persistent micromotions are known to trigger a complex cascade of events, it was hypothesized that the different material properties and design of the Michigan electrode was the cause for the diminished wound healing response, the increased inflammation, the enhanced blood–brain barrier permeability, and the infiltration of inflammatory myeloid cells. For instance, a genomic analysis revealed an upregulation of MMP-2 (for facilitating wound healing and promoting neuronal regeneration) and blood–brain barrier stabilizing proteins in microwire electrodes (indicating enhanced stability and reduced micromotions compared to Michigan electrodes). Despite these findings, the *critical surface area model* theory is preferred over the theory of *mechanical property differences* because of previous results that have shown negligible differences in glial encapsulation between the soft and flexible parylene/SU-8-based structures in [30] and the silicon-based arrays (that are orders of magnitude stiffer) in [25].

#### **3.5. Embedding of bioactive reagents**

#### *3.5.1. Utah electrode array*

**3.4. Anchoring of the probe's position**

closest to the tissue interface [29].

*3.4.2. Michigan neural probe*

edges [31]).

of magnitude stiffer) in [25].

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

In Sections 3.2.2 and 3.3.2, the *critical surface model* was used to explain the benefits of open-architecture designs [25, 30]. However, an alternative theory [30] was offered to explain how the open-architecture designs reduce tissue encapsulations compared to the solid shanks. The theory of *mechanical property differences* assumes that lattice structures reduce the induced strain from relative movements (e.g. micromotion) between the probe and the tissue (occurring from the continual pulsations of vascular and respiratory oscillations). In other words, the adjoining lattice structure is assumed to present a more flexible mechanical interface that can improve the mechanical surface properties (relative to the solid probe shank). This competing theory of *mechanical property differences* is supported by FEM results that demonstrate that the shape of the electrode substantially influences the local pattern and the magnitude of strain around the electrode (especially at the electrode tip where the shank has a spear-head shape or at regions that have discontinuous or sharp

The theory of *mechanical property differences* is also supported by electrophysiological data [28] showing the Michigan electrode's recording performance failed faster relative to a microwire electrode. Since persistent micromotions are known to trigger a complex cascade of events, it was hypothesized that the different material properties and design of the Michigan electrode was the cause for the diminished wound healing response, the increased inflammation, the enhanced blood–brain barrier permeability, and the infiltration of inflammatory myeloid cells. For instance, a genomic analysis revealed an upregulation of MMP-2 (for facilitating wound healing and promoting neuronal regeneration) and blood–brain barrier stabilizing proteins in microwire electrodes (indicating enhanced stability and reduced micromotions compared to Michigan electrodes). Despite these findings, the *critical surface area model* theory is preferred over the theory of *mechanical property differences* because of previous results that have shown negligible differences in glial encapsulation between the soft and flexible parylene/SU-8-based structures in [30] and the silicon-based arrays (that are orders

*3.4.1. Utah electrode array*

184 An Excursus into Hearing Loss

**Figure 14.** Platinum coated tips of the Utah electrode. Scale bar = 0.5 mm. Reproduced from [32, 33].

The microneedle array can be fabricated with porous tips that can be loaded with drugs that have a high molecular weight [32]. 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 [33]. 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 [34]. 3D probe structures with microfluidic channels can also be fabricated via surface micromachining and deep reactive ion etching (DRIE) [35]. 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 [36].

**Figure 15.** Probe design with microchannels along the shanks for controlled fluid delivery. Reproduced from [36].
