**2. Chronic neural recording failure**

**1. Introduction**

174 An Excursus into Hearing Loss

*Brain machine interfaces* connect the human brain to electronic devices and computer software [1]. Machine intelligence can offer the inherent advantages of greater processing speeds, computing power, memory capabilities, and even unrestricted sensory perception (e.g. infrared, ultra violet, X-ray, and ultrasonic spectra). With the emergence of deep brain stimulation, neuroprosthesis, neurofeedback, and exoskeleton technologies, science fiction is being bridged into a modern reality. However, the ultimate realization of a brain machine interface is to have a computer system that can chronically interface with the neural tissues. This neurotechnology will require neuroscientists and engineers to work together to address the technical challenges of accessing neural communication channels (for data routing and transmission), preserving the *biocompatibility* (to interface electronic components within the biological neural tissues), and

maintaining the bio-signal processing (for selecting the appropriate control signals).

This chapter will focus on the *microelectrode*: a neural probe used in neurophysiology for either recording neural representations in the brain or for electrically stimulating the nervous tissue. Micromachined electrodes and microwires are often used to monitor the neuronal activities by characterizing extracellular field potentials of multiple active neurons. These neural interfaces are artificially-engineered extensions of the nervous system that must coexist in the precise connections of supporting glial cells, oligodendrocytes, astrocytes, and microglia (**Figure 1A**).

Multichannel microelectrodes (**Figure 1B**) can be used to monitor the activity in the *auditory cortex* and to investigate the functional organization of the auditory system [2–4]. For example, a chronic microelectrode investigation of the human auditory cortex [5] revealed a "tonotopic" pattern where the sound-driven units had excitatory receptive fields with sharply tuned best-frequency response over a range of frequencies. These micro-machined devices are currently being developed to control brain machine interfaces [6–8] and can benefit many applications in medicine, communication, entertainment, military, and education. However, there has been one limiting factor that has obstructed their reliability as a fully-implantable neural prosthesis:

**Figure 1.** (A) Depiction of cellular changes induced by an implanted electrode and (B) multichannel neural recordings.

*chronic neural recordings* have been shown to deteriorate [9, 10] with time (**Figure 2**).

#### **2.1. Deterioration of neural recordings over time**

**Figure 2.** (A) Mean SNR across implanted probes over time and (B) mean impedance values. Reproduced from [10].

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 [9]. Previous research showed signal-to-noise-ratios (SNR) decreased as a function of postoperative time (**Figure 2A**) [10]. 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.

**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 [13]. The *insertion killzone* is defined as the region around the shaft where the local neuron density is lower than the expected neuron density by a 90% confidence interval (**Figure 4A**). The tearing of neurons will cause the temporary loss of action potential capability and cell death due to the influx of calcium. The cutting of vessels and capillaries will cause microhemorrhages that can displace the neural tissue. If the neural tissue is compressed or stretched, the leakage of ions can lead to the loss of cellular homeostasis. Therefore, the penetrating shafts should be carefully fixed to a manipulator to prevent

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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 [14]. 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.

**Figure 4.** (A) Damage during insertion of a sharp microelectrode (where the shaft stretches the tissue beyond its elastic limits while passing through) and (B) significant axial misalignment of a shaft creates a swath of damage during insertion.

significant misalignment and to minimize the killzone (**Figure 4B**).

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 [11]. 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.

**Figure 3.** Unit isolation quality varies as a function of distance from the electrode. The spike amplitudes of the neuronal ensembles must lie within a cylindrical radius (n = 140 μm) to be detected. Reproduced from [12].

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 [12]). **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 "*cylindrical radius"* distances will be important to consider since previous studies have shown that the inflammatory responses can extend to ~100 μm.

#### **2.2. Insertion killzone**

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 [9]. Previous research showed signal-to-noise-ratios (SNR) decreased as a function of postoperative time (**Figure 2A**) [10]. 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.

176 An Excursus into Hearing Loss

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 [11]. By differentiating the spatial distribution of reactive tissue responses corresponding to changes in impedance spectra along the electrode-tissue interface, these findings confirm

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 [12]). **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 "*cylindrical radius"* distances will be important to consider since previous studies have

**Figure 3.** Unit isolation quality varies as a function of distance from the electrode. The spike amplitudes of the neuronal

ensembles must lie within a cylindrical radius (n = 140 μm) to be detected. Reproduced from [12].

shown that the inflammatory responses can extend to ~100 μm.

that reactive tissue responses directly influence the impedance spectra over time.

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 [13]. The *insertion killzone* is defined as the region around the shaft where the local neuron density is lower than the expected neuron density by a 90% confidence interval (**Figure 4A**). The tearing of neurons will cause the temporary loss of action potential capability and cell death due to the influx of calcium. The cutting of vessels and capillaries will cause microhemorrhages that can displace the neural tissue. If the neural tissue is compressed or stretched, the leakage of ions can lead to the loss of cellular homeostasis. Therefore, the penetrating shafts should be carefully fixed to a manipulator to prevent significant misalignment and to minimize the killzone (**Figure 4B**).

**Figure 4.** (A) Damage during insertion of a sharp microelectrode (where the shaft stretches the tissue beyond its elastic limits while passing through) and (B) significant axial misalignment of a shaft creates a swath of damage during insertion.

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 [14]. 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**

**2.4. Micromotion**

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 [17] 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 [18]. The results indicated that the interfacial strains were created by *micromotions* of the chronically implanted electrode, and that these mechanical strains around the implant site are likely responsible for the sustained tissue responses in chronic implants. In addition, the elevated strains at the probe tip were shown to cause poor probe-tissue adhesion and delamination of the tissue from the probe. The simulated probes also induced strain fields that displayed high radial tethering forces, with pressure profiles revealing extensive tension and maximum frictional shear stress at the tips of the arrays (**Figure 7**). Therefore, these findings indicate that softer substrates should be used to reduce the strain at the probe-tissue interface (to reduce tissue responses in chronic implants). Specifically, the magnitude of the micromotions should be reduced at the microelectrode's tip position to ensure the stability of the recordings at the cortical surface since gliosis is typically observed at the probe tips [13] (with up to a three-fold increase in the size of the surrounding glial sheath compared to at other areas of the arrays [19]). Micromotions and cortical surface displacements can also result from respiratory pulsations [20], behavioral sources [21], and the translational movements of the electrode's tethering lead

**Figure 7.** (A) Finite element model (FEM) strain profile of the radial tethering forces in the brain tissue that results from a 1 μm displacement of a silicon probe extended to 100 μm from the interface (s = surface, m = midpoint, t = tip). (B) Normalized strain values decreased exponentially as a function of distance in the brain tissue. Reproduced from [18].

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wire [22] (which can occur due to the rotational acceleration of the head [23]).

**Figure 5.** (A) Early reactive responses and (B) prolonged reactive responses to a chronically implant. Reproduced from [14].

An immunohistochemistry experiment [14] 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 [15] 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 [16]. 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.

**Figure 6.** Histopathological changes that occur around a non-reactive implant, a reactive implant, and toxic implants.

#### **2.4. Micromotion**

**2.3. Prolonged injury response**

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neural recordings.

An immunohistochemistry experiment [14] 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 [15] 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 [16]. 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

**Figure 5.** (A) Early reactive responses and (B) prolonged reactive responses to a chronically implant. Reproduced from [14].

**Figure 6.** Histopathological changes that occur around a non-reactive implant, a reactive implant, and toxic implants.

**Figure 7.** (A) Finite element model (FEM) strain profile of the radial tethering forces in the brain tissue that results from a 1 μm displacement of a silicon probe extended to 100 μm from the interface (s = surface, m = midpoint, t = tip). (B) Normalized strain values decreased exponentially as a function of distance in the brain tissue. Reproduced from [18].

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 [17] 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 [18]. The results indicated that the interfacial strains were created by *micromotions* of the chronically implanted electrode, and that these mechanical strains around the implant site are likely responsible for the sustained tissue responses in chronic implants. In addition, the elevated strains at the probe tip were shown to cause poor probe-tissue adhesion and delamination of the tissue from the probe. The simulated probes also induced strain fields that displayed high radial tethering forces, with pressure profiles revealing extensive tension and maximum frictional shear stress at the tips of the arrays (**Figure 7**). Therefore, these findings indicate that softer substrates should be used to reduce the strain at the probe-tissue interface (to reduce tissue responses in chronic implants). Specifically, the magnitude of the micromotions should be reduced at the microelectrode's tip position to ensure the stability of the recordings at the cortical surface since gliosis is typically observed at the probe tips [13] (with up to a three-fold increase in the size of the surrounding glial sheath compared to at other areas of the arrays [19]). Micromotions and cortical surface displacements can also result from respiratory pulsations [20], behavioral sources [21], and the translational movements of the electrode's tethering lead wire [22] (which can occur due to the rotational acceleration of the head [23]).
