The presence of different types of cells and their correlated functions including ECM synthesis, remodeling, and degradation (mainly fibroblasts) makes a tissue (connective tissue) unique within an organ. So far, biochemical properties of tissues have provided a large amount of information about the presence of tissue or cell specific biomarkers. These biomarkers reveal the distinction between the healthy and diseased state of a tissue, which may help in synthesizing specifically targeted drugs. Cell mechanics has now become a potential biomarker to discriminate between the different physiological and pathological states of cells (Rianna and Radmacher, 2016). Similarly, investigating tissue mechanics opens up a new platform in the biomedical field to diagnose pathological states of different tissues.

Generally speaking, brain tissue has three distinct parts: the cerebrum, cerebellum, and the brainstem. Each part has its own unique function in governing the different functions of the human body. As the central nervous system (CNS) for the whole body, brain tissues mainly contain neuronal and glia cells which interact through electric and ionic signaling and neurotransmitters. The mechanical properties of neurons and glia cells play a key role in neuronal growth and development (Spedden and Staii, 2013). Studying local and global brain topography and mechanics noninvasively can lead to a better understanding of the development of various diseases such as neurodegenerative diseases and cancer. Previous rheological studies on brain tissues were mostly conducted non-destructively on a macroscopic scale of centimeter to millimeter. Investigation into micro- and nano-scale range regions of living brain samples may allow distinguishing between cell and ECM properties and their correlation.

The main goal of this mini review is to introduce readers to the working principle of AFM and its application in tissue imaging and the mapping of mechanical properties of tissues. Finally, we discuss the possibility of using AFM in brain tissue biomechanics.

# AFM – WORKING PRINCIPLE – IMAGING AND MECHANICAL MAPPING

Atomic force microscopy is conceptually a simple technique, employing the interaction between a tip whose shape can be tuned according to the application (sharp tips for high resolution imaging and pyramidal or spherical tips for mechanical mapping) attached to a soft cantilever spring and the sample. There are four main components (**Figure 1A**) in AFM: a cantilever, which acts as a spring with an integrated tip; a laser beam focused onto the very end of the cantilever where the tip is attached; a position-sensitive photo-detector to detect the reflected laser beam, which can measure the horizontal and vertical deflection of the cantilever; and finally, a xyz piezo scanner for moving the sample or the cantilever in all three directions. In our example schematics, the piezo scanner setup has been designed in such a way that the z piezo controls the cantilever movement in the z-direction and the xy piezo controls the sample movement in the xy-direction.

Different imaging modes such as contact (DC) and noncontact tapping (AC) modes are used in AFM to measure the sample topography. In contact mode, the AFM tip is brought into physical contact with the sample and the cantilever deflection is measured. In the constant height mode, the sample is kept at a constant height while the tip raster scans the sample. The topographic information is inferred from the deflection of the cantilever as the tip scans over areas of different heights. This particular mode is generally used for flat and rigid samples, since, due to the deflection of the cantilever the loading force will change. For soft biological samples, especially for cells, this mode will damage the cells as they will be exposed to large loading forces. In order to image soft samples, a feedback is introduced to adjust the z height such that the deflection, and therefore the loading force, is held constant. This mode is called a constant force or constant deflection mode. **Figure 1B** shows the height and error signal images of the extracellular matrix topography of the decellularized dermal matrix. In constant deflection mode, the output of the feedback corresponds to the height signal image which shows the overall sample topography. Since the feedback will react with a finite response time, the main time limiting factor will be the piezo transducers used in AFM, there are some residual changes in deflection, which are not perfectly compensated. In control theory this behavior is called the error (of the feedback loop); therefore, in AFM the phrase error signal image is also often used. To reduce lateral forces exerted to the

sample in contact mode, which can be substantial and destroy or detach samples, the tip is periodically retracted from the sample and the cantilever height is modulated at the cantilever's resonance frequency. This mode is called the tapping mode and is used largely in imaging biomolecules such as DNA, proteins, and lipids. Like in contact deflection mode, tapping mode produces two images: a height and an amplitude error image. A novel variant of the tapping mode, the peak force mode, where the data during one oscillation cycle are captured and analyzed online to control the maximum force, seems to be favorable for cell imaging (**Figure 1B**; Schillers et al., 2016).

In a force curve (**Figure 1C**) the interaction forces between the tip and sample are measured while the tip is approached and retracted from the sample. This can be performed over a region of interest of the sample, generating a force map or force volume (**Figure 1D**) in which each pixel in the map represents a force curve. Both the approach and retract curves reflect information on the mechanical, or more precise viscoelastic properties of the sample, as well as adhesion properties between the tip and sample, e.g., a cell or the ECM. The elastic properties of the sample can be inferred by fitting the data with an appropriate geometric model of the tip and sample to yield the Young's modulus. Different models are used from continuum mechanics depending on the shape of the AFM tip. In most cases, the AFM tip shapes are pyramidal, conical, and spherical. According to the tip geometry, the Hertz model (Hertz, 1882) for spherical indenters, Sneddon model (Sneddon, 1965) for conical indenters, and Sneddon extended model (Rico et al., 2005) for pyramidal indenters are used to describe the elastic behavior of the biological samples. The force can be calculated by Hooke's law from the deflection of the cantilever, if the spring constant is known. The sample indentation is calculated from the z movement of the z piezo and the cantilever deflection.

# TISSUE IMAGING AND MECHANICAL MAPPING IN AFM

The simplicity of the working principle of AFM allows users to obtain the fine microstructures of biological tissue with good resolution. Biological tissues are comprised of different cells and ECM, whose interplay facilitates tissue dynamics and maintains homeostasis. Investigating biomechanical properties and imaging of cells and ECM are studied individually and cells are mostly cultured in hydrogels, matrigels, or threedimensional (3D) matrices in order to evaluate the substrate

stiffness or composition-dependent cell elastic properties. Whereas decellularized ECM is evaluated for the ECM component arrangement and stiffness. Intracellular actin cytoskeleton arrangement and dynamics reveal that the cell stiffness and actin stress fibers interact with and transmit mechanical information to the ECM through the transmembrane protein focal adhesion complex. This adhesion complex consists of the transmembrane protein integrin, whose extracellular domain binds to the RGD (Arg-Gly-Asp) sequence of any of the ECM proteins and its intracellular domain binds to the adaptor proteins which further bind to the actin cytoskeleton. This combined complex transfers both the extracellular and intracellular force generated by respective ECM protein fibers and actin stress fibers in the cells, resulting in signaling in both directions: from the cell to the ECM environment and back from the environment into the cells (Discher et al., 2005). ECM imaging and biomechanical properties are so far performed in decellularized tissue samples. The surface topology of acellular ECM scaffolds provides information on the ECM protein fibers' orientation, spacing, diameter, and also records mechanical maps which enable their stiffness and surface roughness. Collagen fibers are mostly dominant and abundant in these decellularized tissue samples and sometimes the collagen fibers are also seen with other ECM proteins. AFM measurement can be combined with fluorescence microscopy in order to study different ECM proteins by tagging them with different fluorophores which make them easier to visualize in conducting decellularized ECM imaging and mechanics (Jorba et al., 2017).

Ensemble investigations of cells and ECM at the tissue or subtissue level provide information on the cell–ECM mechanical crosstalk and disease-related alterations in tissue morphology and mechanics. The AFM sample preparation for tissue investigation starts with the immobilization of tissue blocks which is quite challenging. Tissues are normally immobilized to a coverslip or any other suitable support in several ways. Tissue adhesives such as Histoacryl tissue glue (Stolz et al., 2009; Plodinec et al., 2010) or ethyl cyanoacrylate (Chan et al., 2010) are mostly used and care has been taken so that adhesives do not make contact with the investigated region. Nevertheless, this immobilization procedure will have some effect on the tissue as substances released from the adhesives may diffuse, either through the surrounding air or water, or directly through the neighboring tissue. Sectioned tissue specimens are often immobilized to microscope glass slides coated with poly-lysine (Stolz et al., 2004; Grant et al., 2012). This procedure is a good option but only for thin tissue sections. The main aim of immobilization of tissues is to avoid sample movement while recording images or mechanical maps. An alternative to adhesives is using Thermanox coverslips punctured in the center and used for holding down the sample in such a way that the tissue can be accessed by the AFM tip. The edge of the Thermanox coverslip can then be glued to the support (**Figure 2A**; Morgan et al., 2014). This setup avoids contact of tissue and glue or any other adhesive materials and serves as a better way to immobilize tissue samples. This approach paves way for investigating the topographical and mechanical changes of the mouse skin tissue. **Figure 2B** shows the presence of thick ECM fibers in mouse skin tissue before and after the addition of collagenase, which leads to the disappearance of fibers and correlatively decreases their elastic properties (Joshi et al., 2017). Biological tissues, cells and the ECM, are composed mainly of water (around 70% for the cytosol), even though for the biological function usually only macromolecules such as proteins or small organic and inorganic molecules are discussed. Therefore, it is very important to study tissues in a hydrated environment such as selecting the suitable medium which brings the utmost native environment to tissues that affect their topography and mechanical properties. It was reported (Zhu and Fang, 2012) that hydration and dehydration of cartilage affects the collagen fiber distribution and its roughness. Standardizing the tissue immobilization protocol and selecting the correct liquid medium are simple to set up and efficient in measuring the morphological and mechanical alteration in the tissues. Tissue samples are normally very soft. As a consequence, they are very difficult to cut into thin slices in their native state even with state of the art vibratomes; therefore, they are often frozen to prepare thin sections (cryosectioning). Normally thin tissue sections of 5–50 µm in thickness are generated for imaging and recording mechanical maps in AFM studies. However, the cryo-procedure not only decreases the cell viability dramatically, but also changes ECM and cell mechanics. Thus, the interplay of cell and ECM mechanics will be difficult or even impossible to investigate. A recent report (Xu et al., 2016) demonstrated the ability to show the difference in mechanical properties of vibratome and cryotome tissue sections. Vibratomed tissue sections show good cell viability and in mechanical maps, cell and ECM regions can be distinguished. Thus, in vibratomed sections nearly all the properties of the living tissue sample are preserved, e.g., for AFM measurements. In contrast, in mechanical maps recorded from cryotomed sections, cell and ECM regions could not be distinguished, because the freezing process increased the stiffness of the entire tissue, possibly because cells were not viable anymore.

For AFM imaging, cryotomed tissue sections or samples chemically fixed in paraformaldehyde are mostly used. Both preparations increase sample stiffness and decrease adhesion to the cantilever tip (Joshi et al., 2017), which makes them suitable only for imaging; mechanical data from chemically fixed samples will be strongly affected by this sample preparation and show no resemblance with the mechanical data of live cells or tissue (Braet et al., 1998). Tissue sample preparation, immobilization, and hydration procedures are important for imaging and for obtaining mechanical maps of native samples by AFM. In a pioneering work it has been demonstrated that AFM can be used for in vivo nanomechanical imaging in living mammals by capturing real-time changes in nanomechanical properties during vasodilation and vasoconstriction in blood vessels (Mao et al., 2009). Although it is not clear how this could be adopted to other applications, especially in humans, it clearly demonstrates the potential of AFM directly following dynamics in living mammalian tissues. AFM was also used for capturing ultrastructural high-resolution imaging of native biomolecules, specifically intracellular organelles and extracellular matrix structures in mammalian connective tissue cryo-sections (Graham et al., 2010). AFM nanoindentation experiments

are able to characterize the stiffness and elastic modulus of soft tissue scaffolds such as collagen–chitosan biopolymers (Zhu et al., 2011). 3D mechanical properties of the skin epidermis and dermis at the nanoscale resolution are measured using AFM nanoindentation (Kao et al., 2016). By selecting the right AFM probes and by standardizing tissue section preparation and immobilization, one can measure the morphological and biomechanical changes in the tissue of any animal.

# AFM IN NEUROBIOLOGY – PROSPECTS AND CHALLENGES

Here we discuss some of the reports where AFM was used in neurobiology, concentrating on neuronal and glial cells' biomechanics, including also brain tissue mechanics. For a detailed review on AFM usage in neuron biomechanics, readers are pointed to this review (Spedden and Staii, 2013). In neuronal growth and development, both neuronal cells and ECM function in a very coordinated manner. For example, the intracellular actin cytoskeleton helps neuronal growth cone movements. During development, neuronal growth cones are found at the apex of each axon and move in a direct way toward the target cell through the detection of extracellular signals. High resolution imaging and by studying the mechanical properties of these neuronal growth cones using AFM show varying protein motilities between different growth cone regions (Grzywa et al., 2006). AFM imaging of neurons grown on laminin micropatterns shows the laminin-guided neurite growth and the role of actin cytoskeleton in growth cone dynamics (Xing et al., 2010). AFM combined with confocal fluorescence microscopy have been used to analyze the morphology of neuronal growth cones of

rat dorsal root ganglia (Laishram et al., 2009). 3D imaging from AFM and confocal evaluates the 3D architecture of living chick dorsal root ganglia and sympathetic ganglia (McNally et al., 2005). 3D AFM imaging of neurons and glia cells provides a wealth of information on 3D cell structures and sub-cellular structures of organelles such as the mitochondria or the nucleus (Parpura et al., 1993). Viscoelastic properties of individual neuronal and glial cells in the CNS show a large mechanical heterogeneity because of the distribution of cell organelles. Generally, both cell types are very soft compared to other eukaryotic cells (Lu et al., 2006). This work also reveals that glial cells do not serve as a structural support or as a glue cell for neuronal cells. In all the reports mentioned above, chemically fixed or living cells are used for AFM imaging and mechanical measurements.

The discussion of AFM application in brain cells opens up the possibility of mechanical characterization of brain tissues using AFM. Before we further discuss this topic, the focus on tissue sample preparation for such application has to be elaborated on as this provides varied techniques along with their advantages and disadvantages. Tissue extraction, embedding, and slice preparations largely fall into deformations, due to the loss of the native environment and dehydration. This causes global shrinkage from the earlier primary and secondary deformations and greatly affects the tissue structures (Dauguet et al., 2007). This demands a unique embedding and slice preparation method to maintain the brain tissue integrity. Concerning 2D brain tissue preparation, conventional slice preparation methods for AFM investigation largely alter the tissue structure. Therefore, agar embedded tissue blocks were used for slide preparation which maintains the tissue structure and helps in carrying out mechanical measurements (Iwashita et al., 2014). Regarding 2D tissue section mechanical preservation, AFM mechanical characterization of living tissues showed cryotomed sections preventing the mechanical measurements of cells in the tissues. Comparatively, vibratomed sections are able to distinguish between the cell and ECM regions in the AFM-generated submicron resolution mechanical maps (Xu et al., 2016).  Earlier, our lab tried to prepare cancer tissue sections from vibratome and due to the softness of the cancer tissue, we could not succeed in obtaining reproducible and useful results. The preparation of brain tissues from techniques other than cryotome may be advantageous to measure living tissue's mechanical properties.

Sub-tissue level nanomechanical imaging of both cells and ECM could possibly demonstrate the elastic properties as well as fine details of biomolecular structures of different brain regions. AFM measurements of the hippocampal and cortex regions of a rat brain show mechanical heterogeneity in subregions and also age-dependent tissue stiffness correlation (Elkin et al., 2007; Elkin et al., 2010). AFM indentation, with a spherical indenter, shows significant mechanical differences between white and gray matter of the rat cerebellum (Christ et al., 2010). AFM spatiotemporal tissue mechanical profiling shows the gradual stiffness increase in ventricular and subventricular zones of a mouse brain during embryonic development (Iwashita et al., 2014). Spatial mechanical mapping of the living guinea pigs' retinae, using scanning force microscopy, determines the elastic modulus of retinal regions and finds the contribution of neuronal cell bodies to the mechanical properties of inner retina (Franze et al., 2011). At the microscale, AFM time lapse in vivo imaging was performed in a live Xenopus laevis embryo in order to follow their local stiffness, which changes during embryonic brain development. This change in brain local stiffness is largely due to brain cell proliferation (Thompson et al., 2019). This time-resolved stiffness measurement can be conducted in other tissue development and is able to capture the tissue dynamics in varying temporal resolutions. Other than following embryonic brain development, AFM mechanical maps show the mechanical heterogeneity in a mouse primary somatosensory cortex and their age-dependent increase in tissue stiffness. Furthermore, AFM topographical imaging of thin sections of the different layers of the sensory cortex at different ages (weaning to adulthood) shows the continuing smoothing of the cortex surfaces (Smolyakov et al., 2018). AFM is a useful tool that can elucidate stiffness maps which correlate to brain development and many neurodegenerative diseases. AFM mechanical measurement shows Alzheimer's disease associated reduced brain tissue stiffness in mice comparatively to their wild type in both normoxia and hypoxia conditions (Menal et al., 2018). The molecular mechanism of amyloid beta (Aβ) fibril formation and toxicity in Alzheimer's disease is well characterized by AFM imaging (Moores et al., 2011). Pathological conditions such as acidosis stiffens the cerebellar gray matter when brain tissue is exposed to CO<sup>2</sup> thus decreasing the pH (Holtzmann et al., 2016). Comparative to other mammalian tissues, CNS tissue softens after injury. The glial intermediate filaments and ECM composition such as laminin and collagen IV contribute to the rat brain neocortex tissue softening assessed by AFM microindentation experiments (Moeendarbary et al., 2017). The elastic modulus color maps from gray and white matter regions show a decrease in elastic modulus values in the 7 days post-injured area, as compared to the control (**Figure 2C**). AFM mechanical characterization of a rat brain subjected to thromboembolic focal vertebral ischemia shows a decreased Young's modulus compared to the wild type (Michalski et al., 2015). Lewy bodies from postmortem brain tissue samples of Parkinson's disease human patients were imaged by AFM. These images show the aggregated fibrillary nanostructures in Lewy bodies and also show disconnected neurons which are located in the substantia nigra (Tercjak et al., 2014). The mechanical fingerprint of human glioblastoma and meningothelial meningioma tissues was measured by AFM mechanical maps. These potential applications of AFM in brain development and diseased tissue characterization help to better understand the sub-tissue level mechanosensitivity and its implications in the cells and ECM mechanical interactions. The recent advancement in AFM makes it an ideal tool to understand the role of mechanical cues in brain tissues and to correlate these cues to histopathological features. Further advancement in brain tissue sample preparation methods, together with AFM bioimaging techniques, widely covers the neuroscientific

network by using AFM for diagnosing and analyzing neurodegenerative diseases.

# CONCLUSION

We reviewed the application of AFM bio-imaging and mechanical mapping of soft tissue samples like brain tissue and discussed the ability of AFM to work under near physiological conditions, which is essential for mechanical mapping. With the simple working principle, scientists from different disciplines can solve arising questions in tissue biology with the aid of AFM. The challenges in tissue sample immobilization, using the right liquid medium and tissue sectioning for AFM experiments were also discussed. As outlined and reviewed, medical engineers and scientists have to keep the challenges mentioned above in mind, and design experiments accordingly to study different tissues of varying animals. Concerning the brain tissue, there is great demand to use AFM in brain tissue imaging to

# REFERENCES

visualize the micro scale arrangements of cells together with ECM. Additionally, obtaining mechanical maps of different regions of the brain enables one to study varying stiffness within brain tissues.

# AUTHOR CONTRIBUTIONS

PV performed the AFM experiments, data analysis, and manuscript preparation. MR designed the content and was involved in data analysis and preparation of the manuscript.

# ACKNOWLEDGMENTS

We thank the Bruker Nano Incorporation for their support and helpful discussion. The AFM probes were a kind gift from the Bruker Nano Incorporation, Santa Barbara, CA, United States.

Clinical Applications : Proceedings of the International Conference on Physics of Cancer: Interdisciplinary Problems and Clinical Applications (PC IPCA'17) Tomskes, (Melville, NY: AIP Publishing LLC), doi: 10.1063/1. 5001607

**Conflict of Interest Statement:** The authors declare that the research was conducted in the absence of any commercial or financial relationships that could be construed as a potential conflict of interest.

Copyright © 2019 Viji Babu and Radmacher. This is an open-access article distributed under the terms of the Creative Commons Attribution License (CC BY). The use, distribution or reproduction in other forums is permitted, provided the original author(s) and the copyright owner(s) are credited and that the original publication in this journal is cited, in accordance with accepted academic practice. No use, distribution or reproduction is permitted which does not comply with these terms.

# Tissue Response to Neural Implants: The Use of Model Systems Toward New Design Solutions of Implantable Microelectrodes

Maurizio Gulino1,2,3, Donghoon Kim<sup>4</sup> , Salvador Pané<sup>4</sup> , Sofia Duque Santos1,2 and Ana Paula Pêgo1,2,3,5 \*

1 i3S – Instituto de Investigação e Inovação em Saúde, Universidade do Porto, Porto, Portugal, <sup>2</sup> INEB – Instituto de Engenharia Biomédica, Universidade do Porto, Porto, Portugal, <sup>3</sup> FEUP – Faculdade de Engenharia, Universidade do Porto, Porto, Portugal, <sup>4</sup> Multi-Scale Robotics Lab (MSRL), Institute of Robotics and Intelligent Systems (IRIS), ETH Zurich, Zurich, Switzerland, <sup>5</sup> ICBAS – Instituto de Ciências Biomédicas Abel Salazar, Universidade do Porto, Porto, Portugal

#### Edited by:

Jeffrey R. Capadona, Case Western Reserve University, United States

#### Reviewed by:

Abhishek Prasad, University of Miami, United States Takashi D. Y. Kozai, University of Pittsburgh, United States Janak Gaire, University of Florida, United States

\*Correspondence:

Ana Paula Pêgo<EMAIL_ADDRESS><EMAIL_ADDRESS>
#### Specialty section:

This article was submitted to Neural Technology, a section of the journal Frontiers in Neuroscience

Received: 15 February 2019 Accepted: 18 June 2019 Published: 05 July 2019

#### Citation:

Gulino M, Kim D, Pané S, Santos SD and Pêgo AP (2019) Tissue Response to Neural Implants: The Use of Model Systems Toward New Design Solutions of Implantable Microelectrodes. Front. Neurosci. 13:689. doi: 10.3389/fnins.2019.00689 The development of implantable neuroelectrodes is advancing rapidly as these tools are becoming increasingly ubiquitous in clinical practice, especially for the treatment of traumatic and neurodegenerative disorders. Electrodes have been exploited in a wide number of neural interface devices, such as deep brain stimulation, which is one of the most successful therapies with proven efficacy in the treatment of diseases like Parkinson or epilepsy. However, one of the main caveats related to the clinical application of electrodes is the nervous tissue response at the injury site, characterized by a cascade of inflammatory events, which culminate in chronic inflammation, and, in turn, result in the failure of the implant over extended periods of time. To overcome current limitations of the most widespread macroelectrode based systems, new design strategies and the development of innovative materials with superior biocompatibility characteristics are currently being investigated. This review describes the current state of the art of in vitro, ex vivo, and in vivo models available for the study of neural tissue response to implantable microelectrodes. We particularly highlight new models with increased complexity that closely mimic in vivo scenarios and that can serve as promising alternatives to animal studies for investigation of microelectrodes in neural tissues. Additionally, we also express our view on the impact of the progress in the field of neural tissue engineering on neural implant research.

Keywords: neural tissue response, microelectrodes, foreign body reaction, brain slice cultures, neural tissue engineering, deep brain stimulation

# IMPLANTABLE ELECTRODES IN NEUROLOGICAL AND NEUROPSYCHIATRIC DISORDERS

# Clinical Applications

Recent technological progress in the field of brain-machine interfaces boosted the development of innovative tools and electrodes for neurophysiological research and neurostimulation applications to treat neuro disability-related conditions.

Deep brain stimulation (DBS) is an invasive neurosurgical operation consisting in the delivery of electrical impulses to specific areas of the brain by means of implantable electrodes connected to a pulse generator. The concept of DBS is the generation of action potentials resulting in beneficial neurochemical effects, such as the recovery of disrupted neural circuits and physiological brain function. For specific applications, DBS is a U. S. "Food and Drug Administration" (FDA) approved technique which is already applied in the clinic for a vast number of neurological dysfunctions (Coenen et al., 2015; Revell, 2015). In patients affected by Parkinson's disease, DBS at the level of the globus pallidus or subthalamic nucleus is able to reduce bradykinesia, dystonia, as well as walking problems, allowing for substantial improvements in the quality of life (Ramirez-Zamora and Ostrem, 2018). In patients affected by dystonia, globus pallidus DBS has been shown to reduce tremors and involuntary motor contraction, with persistent effects after several years (Meoni et al., 2017). For essential tremors, the principal targets for DBS to reduce tremors with lower stimulation amplitudes and fewer side effects than previous treatments are the posterior subthalamic area and the zona incerta (Holslag et al., 2017; Barbe et al., 2018). In patients with medically refractory epilepsy, long-term DBS at the level of the anterior nucleus of the thalamus is an U. S. FDA-approved therapy that is able to reduce epileptic episodes and ensure perdurable improvements in the quality of life for years (Salanova, 2018). For the treatment of chronic and neuropathic pain, DBS is a supported therapy in Europe, although not currently approved by the U. S. FDA. Several clinical studies report a reduction of pain in patients with amputations, post-stroke pain, cranial and facial pain (Farrell et al., 2018). DBS was also found to be a valid therapeutic approach for the treatment of psychiatric disorders. Ventral capsule/ventral striatum DBS is another U. S. FDA-approved treatment under a humanitarian device exemption for patients affected by obsessive-compulsive disorder (Karas et al., 2019). A relevant efficacy of DBS has been also observed for the treatment of refractory depression. Clinical application of DBS in patients non-responsive to anti-depressant treatments reported a remission of depression after chronic stimulation of various brain targets, with a decrease in negativity and sadness, reduction of cerebral blood flow at the level of the limbic-cortical circuitry, and improvements in memory and motor function (Dandekar et al., 2018; Drobisz and Damborská, 2019).

Spinal cord stimulation (SCS) has shown to be an effective strategy for the treatment of different diseases. SCS is being successfully used to treat angina pectoris pain, low back, and leg pain and peripheral limb ischemia (Song et al., 2014). Clinical studies reported beneficial effects of SCS combined with activity-based training in the recovery of motor function and muscle activation patterns in patients that suffered spinal cord injury (Rejc et al., 2017).

Besides the applications described above, the use of implantable electrodes for neurostimulation therapy is continuing to expand toward many other medical conditions. DBS has been recently proposed for the treatment of pain, dystonia and motor symptoms in post-stroke, although additional investigations are necessary to identify specific brain districts to improve the effectiveness of the treatment (Elias et al., 2018). In a study involving patients affected by Tourette syndrome, DBS of the anterior and posterior globus pallidus, centromedian thalamus and anterior limb of internal capsule showed common positive results after 1 year of treatment in the reduction of motor tic symptoms (Martinez-Ramirez et al., 2018). Partial improvements have been described in clinical trials involving patients affected by Alzheimer's disease, with positive effects in cognition, reversal of memory and reduction of altered glucose metabolism (Mao et al., 2018).

As seen above, the progress in neurostimulation technology and the increased knowledge of the neurophysiology of the central nervous system (CNS) opened the way to new therapeutic approaches for the treatment of a vast number of neurological disorders and neuropsychiatric conditions. Although well established and approved for some diseases, additional trials and experimental work need to be conducted to better define the ideal brain targets, stimulation variables, and electrode design in order to ameliorate the clinical outcomes. Due to the impact of the inflammatory response and tissue encapsulation elicited by traditional DBS macroelectrodes, the field of neuroengineering is progressing toward the employment of implantable microelectrodes (Daniele and Bragato, 2014). Traditional DBS electrodes present several drawbacks such as the rigidity of the materials employed for fabrication as well as a high size which exacerbate the neuroinflammatory response and tissue damage. Although they have a higher activation radius compared to microelectrodes, macroelectrode implantation is often associated to a wrong positioning in the interested area causing a decrease in therapeutic efficacy of DBS (Kloc et al., 2017; Morishita et al., 2017). In the case of small target areas, the reduced dimensions of microelectrodes can provide a better targeting accuracy ensuring an increased therapeutic efficacy of DBS and tissue integration for chronic applications (Desai et al., 2012; Du et al., 2017; Ganji et al., 2017). These miniaturized devices offer additional advantages such as reduced tissue damage and impedance, increased signal-to-noise ratio and neuronal activation compared to traditional electrodes (Desai et al., 2014). A key component contributing to microelectrode design and, ultimately, to clinical performance is the selection of the materials for the device. Such materials should not only satisfy the mechanical and the structural requirements for the efficient electrochemical performance but also provide a durable and biocompatible interface with the brain tissue. In the following section, the current materials used in microelectrode fabrication are presented.

# Current Materials for Microelectrode Fabrication

For the long-lasting efficient microelectrodes, the material should not only function properly in vivo but also be biocompatible and durable for the protection of both the patient and the device (Szostak et al., 2017). As different materials show dissimilar behavior in the tissue environment the choice of the implant material is crucial. To mitigate foreign body

reactions and corrosion/degradation of the structures, electrical, chemical, and mechanical properties of materials, such as chemical composition, crystallinity, surface morphology, the electrode microstructure, and Young's modulus (a measurement of elasticity) need to be carefully considered (Cogan, 2008; Williams, 2008, 2009). With the advance of the micro-fabrication techniques, silicon and polymers have been widely employed as the substrate materials, while metals, carbon nanotubes, conductive polymers as electrode site materials (Samba et al., 2015; Yi et al., 2015; Antensteiner and Abidian, 2017).

Typical microelectrodes are designed to have either an array of microwires or micro-electromechanical system (MEMS) arrays. Microwires are generally composed of metals such as gold, tungsten, and stainless-steel, coated with insulators. Two different types of silicon substrate-based MEMS micro-machined electrodes, i.e., the Utah array (Normann and Fernandez, 2016; Wendelken et al., 2017) and the Michigan array (Kipke et al., 2008; Kiss et al., 2015), have been significantly exploited for decades. However, a huge mechanical mismatch between hard metals/silicon (E ∼ 10 to 100 GPa) and soft brain tissue (0.4–15 kPa) results in substantial strain at the tissue-electrode interface, causing local physical damage that result in inflammation and neural degeneration (Polikov et al., 2005; Harris et al., 2011; Merrill, 2014). The inflammatory process may hinder the stimulation of neuronal cells, as well as it may contribute to device failure as a result of electrode degradation (Kozai et al., 2015a; McCreery et al., 2016). However, it must be highlighted that not all the microelectrode types exhibit the same degradation profile or that there is a direct correlation between electrode failure and the acute inflammatory response (Gaire et al., 2018a). Several efforts have been reported to overcome the drawbacks from the mechanical mismatch between electrode and tissue by implementing materials with lower Young's modulus, e.g., flexible and biocompatible polymer substrates (Trevathan et al., 2019). Polyimide- (Lai et al., 2012) and parylene-based MEMS (Hess et al., 2011) electrodes have been heavily investigated for their improved mechanical properties, easy access to fabricate, and capability to introduce bioactive molecules at the interface to facilitate long-term interaction with the tissue. As chronic stimulation electrode site materials, metals including tungsten, platinum, iridium, tantalum pentoxide, and titanium nitride have been extensively used for their electrical charge-injection properties and biocompatibility (Cogan, 2008; Fattahi et al., 2014; Meijs et al., 2015). For its remarkably increased charge storage and injection capacity and high corrosion resistivity, iridium oxide has been also widely utilized as a coating material (Meyer et al., 2001; Hasenkamp et al., 2009) to enhance the performance and the durability of the electrode. Carbon nanotubes (Jiang et al., 2011; Schmidt et al., 2013) and conducting polymers such as poly(3,4-ethylene dioxythiophene) (PEDOT) and poly(styrene sulfonate) (PSS) (Cui and Martin, 2003; Pranti et al., 2018) are attracting considerable attention as alternatives to the metal electrodes and coatings for their biocompatibility and tunable electrical properties.

An important challenge in microelectrode fabrication for neural stimulation is the identification of smart materials that are able to provide enhanced biocompatibility. So far, all the current microelectrodes are recognized, in the long run, as foreign bodies by the nervous tissue. As the difference in Young's modulus between the electrode and tissue is the main factor that causes damage and inflammation, most research is focused on using materials with low Young's modulus for both substrates and electrode sites. However, it should be also taken into account that mechanical strain from the motion artifacts, such as bending of conducting material, can cause changes in resistance or capacitance of materials. As a result, this can affect the electrical signals of the electrode and can result in unintended performance (Michelson et al., 2019). Thus, the balance between Young's modulus and the electrical properties need to be carefully considered in designing microelectrodes for neural stimulation.

# FOREIGN BODY RESPONSE AS A CAUSE OF IMPLANT FAILURE

The nervous tissue response to implantable microelectrodes is a complex process characterized by a cascade of biochemical alterations and chemical reactions occurring at the level of the tissue-material interface. These biochemical and chemical alterations may culminate in an undesired foreign body response. Additionally, changes in the inherent properties of the electrode after long-term implantation, for example due to corrosion, may further impair its tissue compatibility and durability. Body fluids and tissues are highly corrosive environments characterized by an elevated presence of oxygen, saline electrolytes, macromolecules and dissolved ions that can cause the electrochemical detachment of microelectrode surface. Once surgically implanted, microelectrodes must remain intact for several years to ensure the efficacy of the therapy and device functionality. To provide successful integration, reliability, and durability once implanted in the brain tissue, microelectrodes must fulfill the following requirements:

Microelectrode implantation causes unavoidable damage to the tissue, triggering a series of neuroinflammatory reactions, which are part of the natural wound healing process that can seriously affect the stimulating site integrity and hamper the electrochemical performance in long-term implantations (Salatino et al., 2018). The complexity of such a process can be described by dividing it into two coupled factors: biotic factors, represented by the effects of cells and tissue reactions occurring

at the surface; and the abiotic factors, related to the characteristic of the material itself.

Biotic factors include the blood-brain barrier (BBB) rupture, protein absorption at the material surface, immune cells and fibroblast recruitment, increased production of radical ions, cell death and the formation of the insulating glial scar around the electrode surface, which hampers blood supply and ionic equilibrium at the injury site. The abiotic factors are represented by the physicochemical surface modifications such as the dissolution of passive films, material-related impedance, the failure of the stimulating site integrity, the formation of electrochemical cells at the level of the surface that can evolve in crevices or pits. Biotic and abiotic factors cannot be considered as two separated processes, as they are strictly dependent and occur simultaneously interacting in different manners during the lifetime of the electrode. The complexity of such a process is not totally understood and further research is needed to clarify whether the contributions of these interrelated factors occur and what are the most effective intervention strategies.

In this section, we will provide a description of the main cascade of biochemical and cellular events occurring upon brain microelectrode implantation with a focus on the biotic reactions.

The process that leads to glial scar formation due to implantation and can culminate in the encapsulation of an implant can be divided into two phases (**Figure 1**). An acute phase that starts immediately after device implantation and characterized by BBB dysfunction and glial cell activation, followed by a chronic phase characterized by an immune response and the development of a glial scar around the implant.

The first and one of the most critical events occurring during device implantation is the rupture of the BBB. The implantation causes a break at the level of the endothelial vessels, with a reduction of blood flow and oxygen supply, accumulation of plasma proteins and pro-inflammatory factors, and myeloid cell infiltration (Kozai et al., 2015b). Cell membrane damage by mechanical stress causes an increase in Ca2<sup>+</sup> concentration either by its release through the pores in the cell membrane and by disturbances in the electrochemical potential of Na<sup>+</sup> channels, which lead to membrane depolarization (Eles et al., 2018; Salatino et al., 2019). Membrane depolarization, in turn, leads to the increase in intracellular Ca2+, neurotransmitters release from presynaptic terminations (Eles et al., 2018), resulting in excessive production of reactive oxygen species (ROS) due to mitochondrial damage (Ereifej et al., 2018). The BBB breach has been shown to be crucial in the triggering of biochemical pathways responsible for neuronal degeneration and glial activation (Saxena et al., 2013). Some plasma proteins such as globulins, fibrinogen, thrombin, plasmin, and albumin can accumulate at the injury site through the BBB gap and can be adsorbed at the electrode surface.

# Microglia

Glial activation represents the main cellular event involved in the neuroinflammatory response. As the resident macrophage cells of the brain, microglia are ubiquitous in the CNS, and they become activated to carry out their neuroprotective functions immediately after electrode implantation. Once activated, they act as principal effectors of the neuroinflammatory response and can orchestrate the process through cross-talk with astrocytes and oligodendrocytes. It is already been accepted that microglia exist in "pro-inflammatory" and "anti-inflammatory" phenotypes. The former is the "classical activation" phenotype, in which cells secrets pro-inflammatory cytokines and contribute to neuronal injury; in the case of the latter phenotype cells secret anti-inflammatory cytokines and contribute to tissue remodeling and repair, phagocytosis of cell debris, as well as antagonize pro-inflammatory activity. In the early hours post-implantation, pro-inflammatory microglia secrete pro-inflammatory cytokines and chemokines such as interleukins IL-1α, IL-1β, IL-6, tumor necrosis factor (TNF-α), monocyte chemoattractant protein 1 (MCP-1) (Sawyer et al., 2014), ROS and reactive nitrogen species (RNS), determining massive immune cell recruitment and additional cytokine production (Hermann et al., 2018a). The lack of oxygen redox homeostasis acts directly on microglia, astroglia and endothelial cells causing activation of metalloproteinase, downregulation of tight junctions and adherens junction genes in the first hours after BBB injury (Bennett et al., 2018), facilitating the entrance of infiltrating macrophages, which will also have a crucial role in neurodegeneration (Ravikumar et al., 2014).

# Astrocytes

Astrocytes are another type of neuroglia that is very affected after implantation. Astrocytes perform many functions, including biochemical support of endothelial cells that form the BBB, supplying of nutrients to the nervous tissue, maintenance of extracellular ion balance, having a key role in the repair and scarring process of the brain. In analogy to microglia, astrocytes exist in a pro-inflammatory phenotype and an anti-inflammatory phenotype. Pro-inflammatory astrocytes are activated by pro-inflammatory microglia and secret neurotoxins creating a hostile environment for neuronal and oligodendrocytes regeneration. Pro-inflammatory astrocytes are activated by Il-1β, TNF-α and complement component 1q (C1q) from microglia, responding immediately to electrode implantation and changes in neuronal activity, accumulating in the vicinity of the microelectrode during the first week after implantation (Liddelow et al., 2017). At this level, astrocytes alter the neuronal viability causing neuronal loss, reduction of fiber density and overexpression of glial fibrillary protein (GFAP) and vimentin, which are critical for their change in morphology and extension of the protrusions at the injury site (Woolley et al., 2013; Moeendarbary et al., 2017).

# Neurons

The neuronal loss also occurs immediately after implantation. The mechanical stress caused by electrode entry into the tissue leads to axonal morphological changes, neuronal membrane disruption with the formation of axonal blebs as an indication of neuronal damage. Oxidative stress in neurons is caused by the increase in intracellular calcium through glutamate-N-methyl-D-aspartate receptor (NMDA) activation, as well as by ROS and RNS produced by microglia and astrocytes, causing mitochondrial dysfunction. Neuronal degeneration and neuroinflammation are exacerbated by the persistent secretion

of proinflammatory cytokines and glial fibrillary proteins deposition by microglia and astrocytes, finally forming the glial scar (**Figure 1**). As a consequence of this neuronal death at the tissue-electrode interface, the distance between electrode and synapses grows in time, hampering electrical stimulation performance.

# Oligodendrocytes

Oligodendrocyte cell death will also occur at the implantation site. Either due to cell membrane damage or as a result of neuronal cell death or axonal degeneration. Oligodendrocytes play the important function of ensuring axonal support and myelin production and maintenance. In the acute phase of foreign body reaction, oligodendrocytes become highly sensitive to oxidative stress by ROS and RNS as well as excitotoxic damage by glutamate oversignaling. Apoptosis of oligodendrocytes leads to demyelination and can also culminate in neuronal death depending on the extent of the event. In the adult brain, one can find not only myelinating oligodendrocytes but also cells in the form of neuron-glia antigen 2-expressing glial cells (NG2) precursors, which are present from development to the adult phase, denominated oligodendrocyte progenitor cells. In vivo studies showed that new NG2 precursors become activated by proinflammatory factors secreted by reactive microglia and can be seen migrating to the injury site 12 h post-implantation (Wellman and Kozai, 2018). But there, NG2

phase of inflammation, a glial fibrotic scar surrounds the microelectrode impeding material and stimulating site integrity that, ultimately, may result in implant failure.

precursors preferentially differentiate to astrocytes and further move toward the implant participating in the formation of the glial scar, not contributing to the turnover of new myelinating oligodendrocytes (Wellman et al., 2018a).

# The Glial Scar

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Over 2 weeks post-implantation, in the chronic phase of the process, it has been observed that astrocytes and microglia have their protrusions extended toward the material surface creating a non-permeable barrier between the implant and the tissue over 2 weeks post-implantation (Wellman and Kozai, 2017). At this stage, fibroblasts have reached the inflammation core from meninges and secret ECM proteins such as fibronectin, type IV collagen, laminin, and chondroitin sulfate proteoglycans, also contributing to the formation of the glial scar and the encapsulation of the microelectrode at the parenchymal level (Dias and Göritz, 2018). This insulating barrier constitutes a hostile environment that hampers electrophysiological performance due to the absence of contact between microelectrode and neurons, leading to the failure of the implant over extended periods of time (**Figure 1**).

Despite all the efforts that have been carried out to study the dynamics of glial scar in injury and disease, additional investigations are required to understand the specificities of the foreign body response in the context of electrode implantation, to uncover the most effective intervention strategies to promote microelectrode integration in the CNS (Salatino et al., 2018). Besides, it is important to take in consideration that other aspects can influence glial scar heterogeneity, such as the type of microelectrode material used, the cerebral anatomical district of implantation, as well as the pathological context in which it is applied. Hence, the use of modeling systems that can mimic specific in vivo pathological conditions is a great opportunity to move toward the establishment of innovative approaches on which to base future microelectrode design.

# EXPERIMENTAL MODELS TO STUDY FOREIGN BODY RESPONSE TO NEURAL IMPLANTS

Despite the great advances achieved in neural interface technology, some questions related to the molecular and cellular events involved in nervous tissue response to implantable microelectrodes remain unanswered. Several in vivo studies have been performed to identify critical aspects and design solutions to inhibit glial encapsulation in chronic applications. However, due to their cost, time consumption and complexity in vivo models are not ideal systems to investigate the detailed cues of tissue-electrode interactions. With this purpose, substantial research has been focused on the development of relevant in vitro biological platforms of increased complexity to test new materials and biosurfaces, which can offer a controlled and reproducible platform for high-throughput screenings. In the following paragraphs, we provide an overview of the current in vitro/ex vivo/in vivo models employed in microelectrode research, as well as a description of new promising 3D in vitro technologies with increased complexity that can be of added value to future investigations in this field (**Figure 2**).

# In vitro Models

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One of the main goals in the design and testing of new materials as well as coatings for microelectrodes is to reduce glial cell activation while allowing/inducing neuronal synaptic activity. In vitro 2D cultures represent the simplest model to investigate the impact of materials properties on the cellular response (**Table 1**).

#### Immortalized Cell Line Cultures

The use of relevant immortalized cell lines can provide significant insights regarding material biocompatibility and can contribute to the study of cell-microelectrode material interactions. The experimental conditions are controlled in terms of cell identity, adaptability, and reproducibility. Immortalized cell lines are simple to culture, can be grown for indefinite periods of time, maintaining genotypic stability and allowing the readily generation of large amounts of cells for analysis. Fibroblasts are one of the most well-characterized cell types to study the biocompatibility and the cell-adhesion properties of metals and coatings for biomedical devices. These cells play a critical role in the formation of the fibrotic scar in the late phases of inflammation. The modulation of their adhesion and interactions with the implants is crucial for ensuring a stable device performance. Fibroblast cell lines such as L929 and NIH/3T3 have been widely used to conduct standard material biocompatibility and cytotoxicity testing in accordance with the International Organization for Standardization (ISO) norm 10993-5. The latter defines a series of test methods employing cell monolayers in contact with the material or with material extracts to assess toxicity. In particular, the mouse embryonic NIH/3T3 fibroblasts are a well-characterized cell type used by FDA for biocompatibility testing of materials and coatings for neural devices. Namely, these were employed to assess surface properties of various preparations of polymeric conductive materials for neural devices applications (Mantione et al., 2016; Rejmontová et al., 2016; Hadler et al., 2017; Morin and He, 2017; Wang J. et al., 2018).

Given their central role in orchestrating nervous tissue response to microelectrodes, glial cell were also employed to investigate the effects of new surfaces and designs on cell adhesion, morphology and activation (Persheyev et al., 2011; Ereifej et al., 2013b; Lee et al., 2014). Bérces et al. (2018) employed the immortalized murine microglial cell line BV-2 to investigate the effects of nanotopography on silica and platinum surfaces and compared their behavior with neural stem cells. They showed that while BV-2 cells grew indifferently on nanostructured and non-coated samples, neural stem cells grown on nanostructured surfaces displayed a decrease in cell viability, adhesion and a tendency to adhere to each other instead of to the surface. C6 glioma cells were employed to investigate the biocompatibility of Pt-grown carbon nanofibers coatings for enzymatic glutamate biosensors and compared to Ni-grown nanofibers, showing that cells exhibit different cell adhesion and morphology at different dimensions of nanofibers (Isoaho et al., 2018). Several studies have been conducted to investigate the biocompatibility and the effect of surface properties, like roughness and topography, on neuronal cells. Rat pheochromocytoma PC12 neuronal cell lines are the most used to study the ability of new biomaterials to promote neuronal adhesion and neurite outgrowth (Klymov et al., 2015; Li et al., 2015). Wandiyanto et al. (2018) have recently shown that PC12 cells grown on anti-bactericidal titanium nanostructures displayed enhanced proliferation, differentiation and neurite outgrowth compared to non-nanostructured surfaces. Tasnim et al. (2018) recently investigated, using the SH-SY5Y cell line, the biocompatibility of graphene oxide coating for commercially available 316 stainless steel. They showed that graphene oxide coating enables cell adhesion, proliferation and viability, as well as reduces ROS production compared to bare 316 stainless steel. Nissan et al. (2017) also employed SH-SY5Y to investigate the nanotopographical effects of silver nanoline coatings, showing that cells positively respond by increasing neurite outgrowth and branching points compared to unmodified silica wafers.

While immortalized cell lines are a versatile and readily available tool in the early phases of material testing (**Table 1**), these do not have the same biological relevance and response of their primary counterparts or even. Immortalized cell lines display evident phenotypic and physiological differences from the cell type of origin. These differences can be due to the cell source (many times tumor samples), immortalization process, the propagation and differentiation protocols, as well as the culture conditions and medium composition (Kaur and Dufour, 2012; Lorsch et al., 2014). Consequently, despite being from a similar cell type, immortalized cells can display different viability, metabolic and adhesive properties, as well as different expression profiles and cytotoxic responses to materials. In the context of neurophysiological investigations, they can display different electrophysiological responses to stimulations/recordings, thus they are not the best candidates to serve as models to recapitulate the pathophysiology of diseases (Xicoy et al., 2017). Therefore, results obtained with these cells require validation and comparison with more relevant experimental models. Primary cell lines are a more reliable cell type as they do not have a tumor origin or were not manipulated, and, therefore, more closely recapitulate the characteristics of neural cells in vivo.

### Primary Cell Cultures

Primary cells represent the most used and reliable cell type for in vitro studies because they are similar to cells involved in the tissue response in vivo (**Table 1**). These cultures are not always the first choice due to experimental constraints, like ethics and economic issues. However, they are an excellent tool to study cell behavior prior to in vivo studies. CNS neural primary cells are obtained by dissociation of excised CNS tissue explants and subsequent isolation and plating. For in vitro studies, CNS primary cells are most commonly obtained from animal models like rat and mouse, as one has very limited access to human CNS biopsies. With the recent advances in neuronal cell derivation from (human) pluripotent stem cells, neuronal cultures derived from these cell sources are emerging as a powerful tool for in vitro modeling (Song et al., 2016; Chen W. et al., 2018).

The use of single isolated cell types, despite being useful for the investigation of specific biochemical and morphological responses once in contact with a surface, excludes the possibility to study cell interaction with the material in the presence of the glial and neuronal cells crosstalk. A common approach to improve the in vitro assays is the use of mixed glial cultures with neuronal cells, obtained in a single isolation procedure, as a strategy to mimic the environmental characteristics and cellular events involved in astrogliosis. The use of mixed glial cells allowed to increase the physiological relevance of in vitro testing and to monitor the neuroinflammatory process and microelectrode modifications under manageable and reproducible conditions. Mixed glial and neuronal cultures were used as an efficient in vitro glial scar model for the screening of new design coatings for microelectrodes (Achyuta et al., 2010; Sommakia et al., 2014) (see **Table 1** for examples). However, despite useful and cost-effective if compared to in vivo studies, primary neural cell cultures also present limitations. The isolation procedures are challenging and require appropriate expertise (Uysal et al., 2018). In addition, primary cell lines do not divide (as in the case on neurons) or do not divide indefinitely as immortalized cell lines, hence the number of cells obtained for each isolation is substantially reduced, limiting the number of experiments and the amount of sample for molecular studies (Gordon et al., 2014). Besides the difficulty in their manipulation, primary neural cells, are isolated in early stages of development, therefore they can result unapt for the study of processes that are only observed in the adult or lead to unaccurate results, as some cellular responses can only be observed in early stages of development. The neural differentiation protocols, including for embryonic stem cells (ESCs) or induced pluripotente stem cells (iPSCs), are also laborious and expensive and can lead to different maturation properties (Verpelli et al., 2013; Engel et al., 2016).

Common to conventional immortalized and primary cell cultures conducted in 2D substrates, is the loss of the ECM composition and structure, cell-ECM and cell-cell interactions (namely, the neuronal network), and cell mechanics of the tissue of origin, which, inevitably, results in a different cell behavior compared to in vivo (Tekin et al., 2018). Thus, conventional cell cultures cannot provide detailed information about the interactions of electrode materials with the neural tissue in the initial phases of acute injury or the process of foreign body response in pathological environments. Whence, the necessity of developing complementary models in order to properly study the host reaction to microelectrodes.

#### 3D in vitro Cell Cultures

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With progress in the field of tissue engineering, one is observing an increase in the number of reports of mixed culture systems conducted in 3D scaffolds (**Table 1**). The use of 3D matrices provides additional dynamics to the application of in vitro platforms for glial scar modeling, offering a valid and reproducible system to implement microelectrode research prior to in vivo studies. Several types of 3D scaffolds for neural cultures have been developed. These engineered scaffolds can be based on natural or synthetic materials (mostly polymers). These constitute a great improvement for in vitro studies in terms of increase in complexity and open the way to a vast window of applications in nervous system modeling (Ko and Frampton, 2016). Jeffery et al. (2014) developed a photocrosslinkable and tunable hyaluronic acid-based hydrogel scaffold for mixed glial cultures and high throughput screening of microelectrode materials. A commercially available synthetic polystyrene scaffold was shown to support neuronal cell growth and differentiation. It has been already tested for the development of a 3D model of neuroinflammation employing embryonic primary cortical neurons that are able to grow, interact and form networks possessing electrical activity in the presence of mixed glial cultures (Smith et al., 2015). We have proposed the use of a alginate-based simple and reproducible astrocyte 3D culture system that mimics many features of astrogliosis (Rocha et al., 2015). Using this platform, we established the ECM mechanical properties as a key modulator of astrogliosis. Spencer and coworkers developed a type-1 collagen gel with mixed primary embryonic neural cultures as an in vitro model of glial scar to investigate the effects of micromotion around neural implants (Spencer et al., 2017a). Koss et al. (2017) recently developed a hyaluronic acid-based 3D hydrogel model to study the process of glial scar formation in response to implantable microelectrodes. The biocompatibility of this system allows the encapsulation of primary oligodendrocytes, microglia and astrocytes and has shown to reproduce the typical features of the in vivo glial scar process.

The additional advantage of using 3D systems is the possibility to manipulate and tune scaffold composition through the incorporation of different matrix components and bioactive factors to promote cell survival, migration, and differentiation in a 3D context. The objective is to generate 3D structures with mechanical and structural properties as similar as possible to the ones of the CNS tissue (see **Table 1** for examples). Despite the great advances in this field, these systems still present some constraints. Some biomaterials used for scaffold production are characterized by a high modulus compared to the neural tissue and can lead to altered cell viability, proliferation, and differentiation. Conversely, soft biomaterials are more difficult to handle. The design of matrices with a nanosized microstructure and topographical cues that fully mimic the one found in the nervous tissue was still not attained. The procedures of cell extraction for molecular analysis after testing, scaffold processing, and imaging become more challenging in 3D, limiting high-throughput studies. Additional complications are related to the cell culture conditions: 3D scaffolds can constitute a physical barrier that limits oxygen perfusion, nutrient supply and accumulation of toxic compounds that can cause cellular alterations or apoptosis. Optimizations are still required to ensure a versatile cell encapsulation for different cell types and controlled culture conditions as close as possible to the in vivo environment. This is even more challenging in the context of disease modeling, where cells must be induced to display specific pathological profiles.

It is expectable in the future that the continuous progresses attained in the tissue engineering field, particularly related to improved scaffolds/matrices development, will allow the development of better in vitro systems that recapitulate neural tissue architecture, in an effort to minimize the gap between in vitro and in vivo experiments.

# Ex vivo Models

The development of 3D in vitro models of brain tissue, as stated above, represents an attractive tool for researchers working in the neurosciences field and their use in microelectrode research would have a great impact on the screening of new biomaterials for biomedical applications. However, additional optimizations are still required to allow a consistent application in neuroscientific research. A potential alternative that enables researchers to get closer to in vivo conditions is the use of ex vivo excised brain/spinal cord tissues.

### Organotypic Cultures

Tissue explants can be extracted from euthanized animals or obtained from human biopsies and cultured in vitro. The great advantage of organotypic cultures compared to artificial in vitro systems is the preservation of the native cytoarchitecture with the maintenance of intact neuronal networks. Although several types of ex vivo models have been described in literature (Mii et al., 2013; Nery et al., 2015; Jones et al., 2016; Neville et al., 2018), brain slice cultures from rodents are the most established and widely used as a system of election for neurophysiological investigations, neuropharmacology and as a model of disease. The procedure consists in the isolation of specific districts from the whole brain, their dissection in slices and incubation under controlled conditions. The possibility to obtain several slices from a single animal constitutes an additional advantage in terms of reduction of the number of animals for experiment and of related costs.

Brain slice cultures have been well established from different brain regions (**Table 2**). Two types of brain slices preparations exist: acute slices from the adult brain, with a short life and

TABLE 2 | Organotypic cultures as a model of neurological and neurodegenerative diseases.

APPsdl, amyloid precursor protein gene mutation; A53T, point mutation human α-synuclein protein; CA1, "Cornu Ammonis" 1 hippocampal region; MPP+, (1-Methyl-4 phenylpyridin-1-ium); 6-OHDA, 6-hydroxidopamine [5-(2-aminoethyl)benzene-1,2,4-triol]; P301S, microtubule associated protein Tau (OMIM 157140 genetic mutation); tau4R0N, human microtubule associated protein tau, transcript variant 3, mRNA; 3xTg-AD, transgenic Alzheimer disease mouse model.

mainly used for electrophysiological recordings, and organotypic slices from neonatal animals. The latter are the most diffused ex vivo platforms for the study of many physiological and pathological conditions, thanks to the possibility to reproduce, by external intervention, the hallmarks of diseases that occur in vivo (Magalhães et al., 2018). The great success of this system is due to the simplicity of the procedure and manipulation by mechanical or pharmacological treatment, as well as the possibility to perform electrophysiological recordings on bioelectric activity. Based on these features, organotypic cultures can constitute an ideal model for the long-term assessment of the complex host reaction to microelectrodes or for high-throughput biocompatibility studies of new materials and surfaces. Nevertheless, few works can be found in the open literature (Kristensen et al., 2001; Huuskonen et al., 2005; Ereifej et al., 2013a; Usmani et al., 2016; Leclerc et al., 2017). A possible explanation for this fact could be related to the preferential use of in vivo models as the gold standard for microelectrode testing. Although necessary for the translation of new materials to the clinic, in vivo experiments have ethical issues, they are expensive, time-consuming and unapt for screening studies due to their complexity. The great advances achieved by in vitro/ex vivo systems can be the successful strategy to accelerate microelectrode research prior to in vivo testing. The controllable and reproducible conditions make them suitable to identify strategies to mitigate neuroinflammation, to prevent the early biochemical and corrosion-related events at the interface electrode-neural tissue and to impair foreign body response. Nevertheless, although organotypic cultures maintain the 3D cytoarchitecture, slice preparation causes an unavoidable axotomy of the brain tissue and neuronal death. This physical damage is accompanied by loss of blood flow, and consequently jeopardize oxygen perfusion and nutrient supply. As occurs in primary neural cultures, organotypic cultures are derived from animals in early stages of development and require extensive periods of culture for their maturation

for use in post-developmental studies and assessment of pathophysiological processes. Another important caveat is the lack of the BBB and circulating immune cells. In the context of testing materials for neural devices, we previously showed that these factors play a crucial role in the process of foreign body response. BBB dysfunction and cell infiltration are also associated with several neuropathological processes. This can limit the physiological relevance of organotypic cultures for disease modeling (Humpel, 2016). Hence, researchers are developing innovative approaches combining microfluidic technologies with cellular vascular structures to mimic BBB microarchitecture and improve culture conditions for long term studies (Xu et al., 2016). These authors proposed a new and dynamic in vivo-like three-dimensional microfluidic system to replicate the BBB in vivo. Despite these limitations, organotypic cultures are still one of the most relevant models and can represent a fascinating tool to reduce the differences between in vitro and in vivo studies (some examples are present in **Tables 1**, **2**).

Finally, as a duty of each and every scientist, the use of in vitro/ex vivo models must be encouraged in order to improve the ethical acceptability of research in the fulfillment of the principles of Replacement, Reduction, and Refinement (3R's) (Lossi and Merighi, 2018).

# In vivo Models

Different types of in vivo studies have been carried out in microelectrode research to evaluate therapeutic efficacy, durability and safety of microelectrodes. In this type of studies rats and mice are the most common model of choice. The animal disease models that are used to assess the efficacy of neurostimulation therapies are several (**Table 1**). The main categories are represented by animal models of neurodegenerative disease such as Parkinson's disease (Badstuebner et al., 2017; Musacchio et al., 2017), Alzheimer's disease (Leplus et al., 2018), epilepsy (Desai et al., 2016), sensory-motor deficits due to spinal cord injury (Capogrosso et al., 2018), blindness (Tang et al., 2018), hearing loss conditions (Allitt et al., 2016) and ischemic models (Yang et al., 2017). Large animals such as cats, dogs, sheep, pigs, and non-human primates, are used for chronic studies on the efficacy and safety of neural stimulators. They concern the analysis of both biotic and abiotic reactions on the tissue-electrode interface in long-term experiments (Shepherd et al., 2018). The employment of large animals for these types of investigations is recommended because their anatomy perfectly mimics the environment in which microelectrodes will be applied, allowing the use of all the device components in their real size. Moreover, the full inflammatory component is present in vivo as opposed to in vitro/ex vivo models. The surgical procedure in in vivo experiments generally consists in the exposition of the skull in a deeply anesthetized animal, the production of one or more drills in the vicinity of the target region, the insertion of the neural implant in a specific site and fixation of the plugs with dental cement, followed by continuous monitoring to assess the recovery of the animal, integrity of device and/or efficacy of the therapy (Fluri et al., 2015). Besides the damage caused by electrode implantation, the majority of in vivo experiments are

carried out by tethering the animal to an external component through cables, causing severe discomfort. To overcome this, new wireless microstimulation technologies were developed to ensure better freedom of movement, allowing the reduction of distress (Fluri et al., 2017; Pinnell et al., 2018). Besides the use of animal models of disease, chimeric mice models have found great utility for answering important biological questions concerning the role of different cell types in the process of foreign body response. Chimeras are animal models with two or more different genotypes experimentally obtained by transplanting cells or organs from another organism. Bone marrow chimeric mice were employed to investigate the contributions of different cells in the mechanism of foreign body response (Ravikumar et al., 2014). Sawyer et al. (2014) generated chimera mice between wild type and MCP-1 knock out mice, assessed the key-role of MCP-1 in the enhancement of neuronal loss and showed that its inhibition can be an effective strategy to prolong the lifetime of implantable microelectrodes. Bedell and co-workers recently developed chimeric mice lacking cluster of differentiation 14 (CD-14) genes in myelinating cells and blood-derived macrophages. They demonstrated that targeting CD-14 in blood-derived macrophages improved microelectrode performance in long term experiments (Bedell et al., 2018a). Genetically engineered animal models are another successful tool in microelectrode research. Mice lacking specific genes involved in neuroinflammation and immunity were employed to investigate the biochemical pathways involved in the foreign body response, the identification of pharmacological targets (Kozai et al., 2014; Bedell et al., 2018b; Hermann et al., 2018a,b) and material testing (Lee et al., 2017). Mice carrying cell specific fluorescent tags have been shown to provide great advantages in the study of the contribution of different cells in a single animal as well as a valuable alternative to immunostaining steps (Sunshine et al., 2018; Eles et al., 2019). Gaire and colleagues recently developed a quadruple-labeled mouse with specific fluorescent tags for oligodendrocytes, microglia, neurons, and astrocytes and investigate the process of nervous tissue response to "Michigan" array silicon microelectrodes (Gaire et al., 2018b).

Substantial improvements have been implemented in the quality and quantity of in vivo investigations: new bio-imaging tools were applied in microelectrode research to make in vivo works more explanatory. Two-photon laser scanning microscopy allows imaging of living animals with elevated resolution. It has been used to investigate the nervous tissue response to new electrode coatings (Eles et al., 2017), glial cells characterization (Wellman and Kozai, 2018) or live calcium imaging (Kondo et al., 2017). Optical coherence tomography is a minimally invasive technique that has been proposed to be used in combination with two-photon laser scanning microscopy to provide high resolution angiography of damaged tissues around microelectrodes (Hammer et al., 2016). As an alternative to laborious histological and staining procedures on sectioned brains, x-ray micro CT scanning has been proposed as a high-resolution, time and cost saving procedure that allows a 3D x-ray scanning of the entire brain to quantify and characterize the lesions caused by electrode implantation (Masís et al., 2018).

The growing scientific interest in neural interfaces in the last decades is confirmed by the multitude of in vivo works focused on the testing of new and fully biocompatible coatings (Du et al., 2017; Spencer et al., 2017b; Shen et al., 2018; Vitale et al., 2018), less invasive implantation strategies (Tawakol et al., 2016; Shoffstall et al., 2018b) and new designs with improved electrical performance (Ferlauto et al., 2018; Xu et al., 2018) and with longer durability.

Despite being considered the ultimate model to test microelectrode prior clinical tests, in vivo models present some important drawbacks as previously mentioned. An important aspect to consider is the elevated costs and time required for animal experiments, as well as the resulting ethical constraints. In vivo experiments are complex and demand adequate facilities and technical expertise. As discussed above, transgenic mice offer great advantages, however, they have an extremely high cost due to their production and maintenance. More importantly, the effects of such modifications can lead to altered phenotypes that depart from the real scenario. In the context of device testing, in vivo experiments are quite laborious and require long periods of time to assess the longterm performance of microelectrodes or the biological and behavioral effects of specific neurostimulation therapies. The complexity and invasiveness of the experimental techniques do not always enable scientists to identify the early biochemical and material-related events at the interface electrode-neural tissue and the strategies to mitigate neuroinflammation.

Despite these weaknesses, in vivo studies constitute the gold-standard for the investigations on neural implants, where the results obtained by previous in vitro testing find their effective validation. They represent the final step in the long process of microelectrode design and testing before the application in the clinic. A rigorous progression along all the steps determines the success of new technologies. This is even more important in the case of new materials or designs, where a careful preclinical assessment is necessary to minimize the risk of failure once applied to human patients.

# FUTURE PERSPECTIVES

# Increasing Complexity

Preclinical studies are paramount in the development and testing of new materials for neural implants. As a consequence, the demand for more reliable in vitro/ex vivo models is growing to satisfy the need for assessing of the increasing number of new materials being proposed for this application, improve the quality of device testing and reduce the time between prototyping and commercialization of new products. As discussed in the previous section, several in vitro models and new platforms have been described in the literature (**Table 1**), but they still need to be explored, tested and eventually adapted for microelectrode research.

The use of ex vivo platforms from tissue explants can represent a valuable solution that fits perfectly these purposes. Brain organotypic cultures have been widely used in the last years as an excellent model for a great number of applications. They were employed to study physiological (Svensson and Chen, 2018), and pathological conditions (Tan et al., 2017) or for the screening of new therapeutics (Minami et al., 2017). Furthermore, several organotypic cultures have been established as models of neurological and neurodegenerative diseases (**Table 2**), for which neural implants have been proposed as therapeutic strategies. This places organotypic models that mimic a pathological environment in a privileged position to serve as platforms for microelectrode testing and a potential strategy to move closer to the in vivo scenario. Moreover, thanks to the preservation of intact neural circuitry, organotypic cultures are particularly suitable to perform electrophysiological studies, analyze microelectrode performance and assess astrogliosis.

Developments in the stem cell biology field have also contributed to the establishment of new in vitro models of human disease, aiming at an increase of complexity to reach the relevance of the in vivo environment, while maintaining the controllability and manageability of in vitro systems. Organoids, self-organized 3D tissue cultures derived from stem cells, are currently leading these technologies and have already been developed for the majority of human tissues, including the brain. More recently, the advances in hiPSC reprogramming techniques are also contributing to a better performance of organoid platforms in mimicking human disease and serve as testing platforms for personalized medicine (Perkhofer et al., 2018). Lancaster and Knoblich (2014) the preparation of cerebral organoids prepared from hiPSC. The described methodology allows cell aggregates cultured in Matrigel to mimic native brain tissue, originating different developing brain regions, namely, cerebral cortex, ventral telencephalon and retinal, among others, within 1 to 2 months. iPSC derived organoids represent an innovation in the field of in vitro disease modeling, offering a great opportunity to investigate pathophysiological mechanisms of neurological diseases with elevated reliability (Ho et al., 2018; LaMarca et al., 2018; Sun et al., 2018) and can also contribute to the field of implantable microelectrodes. In fact, Ormel et al. (2018) demonstrated that organoids can innately develop microglia and have a response to inflammatory stimuli that recapitulates neurons-glia interactions in vivo. This is an important aspect since glial cells, particularly microglia, are involved in a great variety of pathophysiological mechanisms.

Great advances in vitro modeling have been also achieved thanks to the application of microfluidics combined with 3D in vitro cultures (Rocha et al., 2016). Microfluidic platforms consist of polymer-based platforms for in vitro culture of cells that allow control and manipulation of microenvironment and fluids (see **Figure 3** for relevant examples). The use of these microdevices brings in vitro models to a whole new level, thanks to the possibility of modifying spatial organization by isolating specific districts and simulate 3D tissue architecture of the native tissue. These systems allow the continuous control of external conditions, conferring an added value to in vitro technology, and giving the possibility to reproduce new biological features that are not possible to achieve with conventional culture systems (as discussed in section "Organotypic Cultures" for the case of the BBB). Wang and colleagues developed an

perforating multi-electrode array (MEA) integrated in a PDMS device for long-term culture, live imaging, recording and stimulation of brain tissues and 3D cultures

organ-on-a-chip system for long-term culture of brain organoids under controlled conditions. Brain organoids were cultured on Matrigel scaffolds with a sided channel for the culture medium supply and a central perfusion channel, allowing a continuous culture medium flow and providing an improved proliferation and neural differentiation compared to static culture conditions (Wang Y. et al., 2018). Liu et al. (2018) combined multielectrode array technology with a microfluidic perfusion system for organotypic hippocampal slices as a platform for high throughput drug discovery. Microfluidic vascular models have been developed and applied to brain-on-a-chip platforms, enabling scientists to improve the quality of culture conditions and get even more close to in vivo dynamics (Osaki et al., 2018; Wang, 2018). New microfluidic devices that model the BBB were fabricated and tested on 2D and 3D cultures, showing that BBB integrity and permeability simulates in vivo characteristics (Chin and Goh, 2018). Adriani et al. (2017) developed a 3D neurovascular chip composed by a central hydrogel co-culture of rat primary neurons and astrocytes, and two lateral channels hosting human umbilical vein endothelial cells and human cerebral microvascular endothelial cells. Bang et al. (2017) developed a 3D microfluidic BBB platform with a vascular channel (VC) composed by a co-culture of human umbilical vein endothelial cells and primary human lung fibroblasts directly interfacing with a neural channel (NC) composed by a co-culture of primary rat neurons and astrocytes to simulate the neurovascular unit. They showed that this platform displayed permeability, cellular contacts and synaptic structures comparable to the in vivo BBB, suggesting its great potential for the drug screening for neurological diseases. Microfluidic technology can conduct in vitro culture system to a more complex and realistic level providing many advantages and details that cannot be extracted with conventional in vivo models, such as the easy manipulation, low cost, and the

(Killian et al., 2016).

possibility to investigate more intimately key mechanisms of diseases. These characteristics perfectly fit with ideal biological platforms for the testing of microelectrode materials developed in the last decades. 3D microfluidic systems can eliminate some of the limitations of 3D in vitro technology, creating new high-fidelity throughput systems that can improve the testing performance and reduce the cost and time for pre-clinical assessment. An additional advantage is the possibility to induce specific pathological features by external treatment for long-term experiments, offering the possibility to investigate in advance microelectrode performance and nervous tissue response under disease conditions.

# Design Solutions and New Materials for the Improvement of Microelectrode Durability and Biocompatibility

Despite remarkable developments in implantable microelectrodes for neuroprosthetics and DBS, additional investigations are still required to address the biocompatibility

and the long-term durability issues. A critical issue is reducing the physical stress, local inflammation and electrode degradation caused by the reaction between electrode and tissue interface while maintaining the electrical sensitivity of the electrode (Prodanov and Delbeke, 2016). To tackle these issues, multiple material-based strategies regarding this problem have been suggested, including (i) chemical modification of the electrode materials, (ii) new design of electrode structures, and (iii) non-invasive and wireless approach using functional nanoparticles.

Biological and non-biological electrode modifications, especially through the surface coating of substrates and electrode sites, are the most commonly used strategies to improve interfacial mechanical mismatch (Aregueta-Robles et al., 2014). Advances in fabrication approaches for integrating conductive polymers (Kim et al., 2018), shape-memory polymers (Shoffstall et al., 2018a), hydrogels (Crompton et al., 2007; Frampton et al., 2007) and carbon nanotubes (Baranauskas et al., 2011; Bareket-Keren and Hanein, 2012) onto complex electrode structures, provide not only a chronically stable neural interface, but also an improvement in the electrode performance. The reduced surface area combined with low impedance and sensitivity provided by such materials make them suitable for either stimulation and recording applications (Vitale et al., 2015; Du et al., 2017; Pancrazio et al., 2017; Wang et al., 2019).

The additional advantage is that the bioactive molecules can be attached to the coating surfaces to increase stimulating/recording sensitivity. Employing composite materials for electrodes and coatings has also emerged as a promising strategy for upgrading electrode functionalities and biocompatibility. Heo et al. (2016) have reported improved biocompatibility in polyimide-based microelectrodes by coating them with PEG hydrogels containing Poly lactic-glycol acid (PGLA) microspheres loaded with the anti-inflammatory drug. On the other hand, Zhou et al. (2013) proposed a carbon nanotube doped PEDOT composite coating material onto the Pt electrode. They showed that this coating makes the electrode more stable with enhanced charge transfer capacity and tissue-electrode interaction. While chemical modification of materials is still being suggested as an efficient way of protecting both electrode and brain tissue, the long-term stability issue caused by the degradation and delamination of coating materials still remains as the challenge that needs to be overcome (Green et al., 2008).

Another attempt to reduce the immune response while enhancing functionality is to introduce new microelectrode designs. The development of the fabrication techniques of soft materials has enabled the production of ultrasoft and ultrathin electrodes with complex designs that minimize the mechanical mismatch of the electrode-tissue interface (Weltman et al., 2016). Recently, Kim et al. have fabricated ultrathin polyimide-based polymer electrodes covered by bioresorbable silk film. They successfully demonstrated the integration of the ultrathin electrodes with a complex structure by allowing the silk to be dissolved and resorbed. This procedure encouraged the spontaneous wrapping process driven by the capillary effect at the material-tissue interface, generating greatly improved biocompatibility (Kim D.H. et al., 2010). Carbon nanotube-based soft fiber microelectrodes have also proved to have low impedance and effective therapeutic stimulation along with single-neuronal-unit signal detectable resolution, owing to their high surface area and electrical conductivity (Vitale et al., 2015). Compared to the similar dimension and surface environment, ultra-soft and ultra-thin electrodes have a great potential to significantly reduce inflammatory tissue response in the long-term scale (Du et al., 2017). However, as mentioned earlier, the balance between flexibility/softness and the electrical performance should be carefully considered when designing these type of electrodes (Wellman et al., 2018b). Implementing functional nanoparticles are attracting increasing attention as a non-invasive and remotely controllable method. Chen et al. (2015) succeeded in utilizing the magnetothermal effect of nanoparticles for DBS (**Figures 4A,B**). They injected Fe3O<sup>4</sup> magnetic nanoparticles in the ventral tegmental area of mice and exposed them to the external magnetic field. When magnetic nanoparticles are exposed to the AC magnetic field, stimulation of neurons at the targeted brain region was triggered by the dissipated heat from the magnetothermal effect. Wireless neural stimulation was successfully performed 1 month after injection. Moreover, lower glial activation, less macrophage accumulation and neuronal loss have been reported compared to a stainless steel implant. Chen S. et al. (2018) recently proposed optogenetic treatment by shining near infrared light to molecular tailored upconversion nanoparticles (**Figures 4C–E**). They injected nanoparticles into the ventral tegmental area of the brain to stimulate deep neurons and successfully demonstrated that light treatment on upconversion nanoparticles can induce dopamine release from dopaminergic neurons, activation of inhibitory neurons, inhibition of hippocampal excitatory cells, and memory recall. Magnetoelectric nanoparticles are another great candidate for neural stimulation. It has been already proven that the piezoelectric materials can generate electric signals under the acoustic wave and can induce neural cell differentiation (Chen et al., 2019). As ferromagnetism and ferroelectricity are coupled to each other, applied external magnetic field can induce variation in electric polarization of nanoparticles, causing a change of the electronic structure at the particle surface and therefore facilitating stimulation of the tissue deep inside the brain (Kargol et al., 2012; Guduru et al., 2015). As a non-invasive stimulating method, implementing magnetoelectric material is attracting significant attention. While these approaches using nanoparticles have great potential, biocompatibility and cellular uptake of these functional particles still remain as a problem to be solved (Adjei et al., 2014; Behzadi et al., 2017).

# CONCLUDING REMARKS

The field of the brain-machine interface is exponentially growing and comprises an important source of progress in many aspects of neurosciences. The application of bionic systems, neural prosthetics and neurostimulation for restoring/treating severe neuro-debilitating conditions and neurological diseases has attracted the interest of many researchers and clinicians. All these technologies require the use of implantable microelectrodes

to interface with the CNS. They represent an essential tool that serves as a link between the electronic components and the neuronal networks, in order to ensure a stable electrochemical communication over time. Great success has been achieved by the clinical application of neural prosthetics in the improvement of the quality of life of patients that suffer from sensory-motor deficits.

The recent progress in microfabrication techniques made possible the development of microelectrodes capable of simultaneous recording and stimulation with improved cell selectivity and spatial resolution. Despite the improvements in device fabrication, biocompatibility and electrochemical performance for long-term applications, unfavorable nervous tissue response and microelectrode failure are still significant limitations. The process of nervous tissue response to microelectrodes has been described by an acute and chronic phase. The acute phase represents the most critical step characterized by a series of pathological reactions triggered by BBB dysfunction and glial activation. The persistence of this neuroinflammation is responsible for the immune response and the formation of the glial scar in the chronic phase, which may lead to microelectrode failure. A better understanding of the signaling pathways involved in the acute and chronic responses is required in order to develop new design strategies to mitigate neuroinflammation and promote a successful integration. Several surface modified microelectrodes have been designed to provide minimal damage and establish a minimally reactive interaction with the brain tissue. However, additional studies are necessary to comprehend some of the key-cellular mechanisms implicated in the process of the glial scar formation around microelectrodes that remain to be elucidated. An important aspect in which research must be focused on is the binary role of the glial scar: several studies report a neuroprotective function of the glial scar in many pathological conditions, and its modulation has been suggested as a therapeutic approach to improve neuronal recovery and tissue regeneration. In the case of chronically implanted microelectrodes, the participation of the various glial subtypes to the nervous tissue response and how their activation states can be influenced to soften tissue damage and avoid rejection are still unclear aspects. Toward this end, the use of appropriate experimental models can provide significant advantages in the development and testing of biocompatible and durable neural devices.

In this review, we provide an overview of the current and potential experimental in vitro, ex vivo, and in vivo models to investigate the mechanisms of foreign body response to implantable microelectrodes. The progress in 3D tissue engineering and disease modeling opened the way toward the development of in vitro biological platforms with increased complexity and physiological relevance to be used for high-throughput studies before moving to in vivo animals. Organotypic culture systems are widely established ex vivo platforms which offer the possibility to simulate several pathological conditions and to isolate specific cerebral regions, ensuring the preservation of tissue architecture and synaptic organization for electrophysiological studies. While the use of organotypic culture systems as screening platforms for novel microelectrodes is still limited, their application is expected to grow in the near future, not only for the reasons mentioned earlier but also because these systems can contribute to significantly minimize the use of animal models. Additional implementations have been also achieved in vivo studies. Despite being considered the gold-standard for microelectrode safety and efficacy studies, the principal limitation of in vivo experiments is the difficulty to monitor tissue response in the initial phases of injury. New advanced neuroimaging techniques open a new window of opportunities to improve the relevance of in vivo assessment thanks to the possibility to study biochemical processes, cell behavior and structural modifications in real time with elevated resolution. The advent of iPSC technology has enabled to simulate more closely the pathophysiological cues that occur in human diseases, offering the relevant advantage to recapitulate molecular and biological characteristics of the human brain. Success is also being achieved by the use of microfluidic systems combined with 3D cell cultures and/or iPSC-derived organoids, which allowed for integrating mechanical and physiological dynamics to simulate organ-like functions and responses. They represent a cost-effective compromise between the versatility of in vitro models and physiological relevance of in vivo models, offering the possibility to model pathophysiological cues under simulated conditions.

In conclusion, researchers have now a great variety of relevant models that can be adopted to improve microelectrode research in all the phases of development and to address the scientific unknowns related to the nervous tissue response to microelectrodes. Ultimately, with further improvements of these in vitro models, one can expect the creation of optimal milieus, which can substantially replace animal experimentation for large scale studies.

# AUTHOR CONTRIBUTIONS

All authors listed have made a substantial, direct and intellectual contribution to the work, and approved it for publication.

# FUNDING

This work was financially supported by the Projects NORTE-01- 0145-FEDER-000008 and NORTE-01-0145-FEDER-000012, supported by the Norte Portugal Regional Operational Programme (NORTE 2020), under the PORTUGAL 2020 Partnership Agreement, through the European Regional Development Fund (ERDF) and FEDER - Fundo Europeu de Desenvolvimento Regional funds through the COMPETE 2020 – Operacional Programme for Competitiveness and Internationalisation (POCI), Portugal 2020, and by Portuguese funds through FCT/MCTES in the framework of the project "Institute for Research and Innovation in Health Sciences" (POCI-01-0145-FEDER-007274).

# ACKNOWLEDGMENTS

fnins-13-00689 July 4, 2019 Time: 16:11 # 18

SP acknowledges funding from a Consolidator Grant from the European Research Council (ERC) under the European Union's Horizon 2020 Research and Innovation Programme (Grant Agreement No. 771565). MG and DK acknowledge their Ph.D. grants in the framework of the project mCBEEs funded by the European Union's Horizon 2020 Research and

# REFERENCES

Innovation Programme under the Marie Skłodowska-Curie grant agreement No. 764977. SS would like to acknowledge national funds through the FCT – Fundação para a Ciência e a Tecnologia, I.P., provided by the contract program and according to numbers 4, 5, and 6 of art. 23 of Law No. 57/2016 of 29th August, as amended by Law No. 57/2017 of 19th July. AP and SP would like to acknowledge the contribution of the COST Action CA16122.

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**Conflict of Interest Statement:** The authors declare that the research was conducted in the absence of any commercial or financial relationships that could be construed as a potential conflict of interest.

Copyright © 2019 Gulino, Kim, Pané, Santos and Pêgo. This is an open-access article distributed under the terms of the Creative Commons Attribution License (CC BY). The use, distribution or reproduction in other forums is permitted, provided the original author(s) and the copyright owner(s) are credited and that the original publication in this journal is cited, in accordance with accepted academic practice. No use, distribution or reproduction is permitted which does not comply with these terms.

# Insights From Dynamic Neuro-Immune Imaging on Murine Immune Responses to CNS Damage

R. Dixon Dorand<sup>1</sup> , Bryan L. Benson<sup>2</sup> , Lauren F. Huang<sup>3</sup> , Agne Petrosiute2,3,4 and Alex Y. Huang2,3,4 \*

<sup>1</sup> Department of Medicine, Vanderbilt University Medical Center, Nashville, TN, United States, <sup>2</sup> Department of Pathology, Case Western Reserve University School of Medicine, Cleveland, OH, United States, <sup>3</sup> Department of Pediatrics, Case Western Reserve University School of Medicine, Cleveland, OH, United States, <sup>4</sup> Angie Fowler Adolescent & Young Adult (AYA) Cancer Institute/University Hospitals (UH) Rainbow Babies & Children's Hospital, Cleveland, OH, United States

#### Edited by:

Ulrich G. Hofmann, University Medical Center Freiburg, Germany

#### Reviewed by:

Naoto Kawakami, Ludwig Maximilian University of Munich, Germany Renaud Blaise Jolivet, Université de Genève, Switzerland

> \*Correspondence: Alex Y. Huang<EMAIL_ADDRESS>
#### Specialty section:

This article was submitted to Neural Technology, a section of the journal Frontiers in Neuroscience

Received: 15 February 2019 Accepted: 02 July 2019 Published: 17 July 2019

#### Citation:

Dorand RD, Benson BL, Huang LF, Petrosiute A and Huang AY (2019) Insights From Dynamic Neuro-Immune Imaging on Murine Immune Responses to CNS Damage. Front. Neurosci. 13:737. doi: 10.3389/fnins.2019.00737 Evolving technologies and increasing understanding of human physiology over the past century have afforded our ability to intervene on human diseases using implantable biomaterials. These bio-electronic devices present a unique challenge through the creation of an interface between the native tissue and implantable bio-materials: the generation of host immune response surrounding such devices. While recent developments in cancer immunology seek to stimulate the immune system against cancer, successful long-term application of implantable bio-material devices need to durably minimize reactive immune processes at involved anatomical sites. Peripheral immune system response has been studied extensively for implanted bio-materials at various body sites. Examples include tooth composites (Gitalis et al., 2019), inguinal hernia repair (Heymann et al., 2019), and cardiac stents and pacemaker leads (Slee et al., 2016). Studies have also been extended to less well-studied immune reactivity in response to CNS neural-electronic implant devices. Recent technological advances in 2-Photon Laser Scanning Microscopy (2P-LSM) have allowed novel insights into in vivo immune response in a variety of tissue microenvironments. While imaging of peripheral tissues has provided an abundance of data with regards to immune cell dynamics, central nervous system (CNS) imaging is comparatively complicated by tissue accessibility and manipulation. Despite these challenges, the results of dynamic intravital neuro-immune imaging thus far have provided foundational insights into basic CNS biology. Utilizing a combination of intravital and ex vivo 2P-LSM, we have observed novel pathways allowing immune cells, stromal cells, cancer cells and proteins to communicate between the CNS parenchyma and peripheral vasculature. Similar to what has been reported in the intestinal tract, we have visualized myeloid cells extend dendritic processes across the blood brain barrier (BBB) into pial blood vessels. Furthermore, transient vessel leaks seen during systemic inflammation provide opportunities for cellular protein to be exchanged between the periphery and CNS. These insights provide new, visual information regarding immune surveillance and antigen presentation within the CNS. Furthermore, when combining intravital 2P-LSM and microfluidic devices complexed with mathematical modeling, we are gaining new insights into the intravascular behavior

of circulating immune cells. This new knowledge into the basic mechanisms by which cells migrate to and interact with the CNS provide important considerations for the design of neuro-electronic biomaterials that have the potential to connect the peripheralneural microenvironments into a unique, artificial interface.

Keywords: two-photon microscopy, microglia, CNS, immune response, innate and adaptive immunity, tissue microenvironment, blood-brain-barrier, vascular crawling

# INTRODUCTION

The use of implantation devices dates back to ancient Egypt with archeological discovery of a toe prosthesis (Nerlich et al., 2000; Brier et al., 2015). More recently, technological advances have allowed clinical application of implantable electronic devices such as pacemakers and deep brain stimulators, once thought to belong to the realm of science fiction, to become routine in clinical practice. These devices have advantages due to their macroscopic scale and relatively large output voltages, which allow them to overpower the potential negative functional effects of tissue fibrosis and host immune reactions in response to the presence of foreign substances.

Meanwhile, challenges remain in the field of implantable neural devices, whether in applying fine sensor arrays for recording neural activity, or in reconstituting neural activities into motor commands to circumvent deficits resulting from spinal cord injury or stroke. Numerous factors contribute to this ongoing challenge, but two are classically identified as being of major importance: tissue stiffness mismatch and breakdown of blood-brain barrier (BBB). Critically, both of these issues interrupt normal neural physiology, implying that simple strategies to prevent tissue fibrosis, such as scavenging of cytokines like TGF-β, will not be sufficient. Such an approach may reduce impedance from tissue fibrosis, but would result in faithfully reporting deranged neural activity.

First and foremost, these devices do not match tissue stiffness of the brain, an organ with the lowest stiffness second only to blood. The importance of tissue stiffness is underscored by the broad success of bony prostheses dating back before recorded history, whereas for blood, the most pliable tissue, even state-of-the art intravascular devices such as extracorporeal membrane oxygenation remain plagued by tissue reaction. This issue is well established in neural technology and extensively reviewed elsewhere (Lotti et al., 2017; Salatino et al., 2017; Campbell and Wu, 2018).

Second, and unique to brain tissue, device implantation invariably disrupts the BBB. This disruption leads to dramatic alterations in the function of all intra-parenchymal cells, whose physiology is predicated on residing within an environment separated from the intravascular milieu. Microglia, astrocytes, neurons, pericytes, and brain endothelial cells are all known to trigger inflammatory programs in response to serum proteins present within the brain parenchyma. In turn, this leads to further inflammation as intra-parenchymal cytokine release promotes further increase in BBB permeability, resulting in accumulation of peripheral blood immune cells such as monocytes, a cell type that is typically excluded from the immune-privileged brain parenchyma. This inflammation occurs via the coordinated synergistic efforts of multiple mechanisms, most of which follow patterns well established in other tissues.

# FAILURE OF NEURO-ELECTRONIC DEVICES

In a quest to restore function to damaged neuronal circuits resulting from trauma, stroke, or neurodegenerative processes, implantable neuro-electronic devices have shown some efficacy. However, the success of these neuro-electronic devices have been limited by inflammatory responses, scarring, and associated neuronal cell death in response to immune cell activation, leading to a progressive neurodegenerative state and ultimate failure of device function (McConnell et al., 2009). In one study, of 78 silicon microelectrode arrays (MEA) implanted in 27 nonhuman primates, 56% of devices failed within 1 year with a mean recording duration of 387 days. Approximately 15% of devices failed due to the development of a dense, fibrous meningeal encapsulation with extrusion of the device from the cortex (Barrese et al., 2013). Additional studies demonstrated that CCL2- and TNFα-secreting CD68<sup>+</sup> cells were responsible for MEA device failure that was associated with the destruction of neuronal circuits in close proximity to the implanted biomaterials (Biran et al., 2005). More recently, peripheral bloodderived monocytes bearing surface CD14 antigen has been implicated as a dominant innate immune cell population that could potently mediate neurodegenerative pathologies associated with microelectrode implantation (Ravikumar et al., 2014b), and serve as a potential therapeutic target to alleviate this problem (Bedell et al., 2018; Hermann et al., 2018). Further review of the acute and chronic tissue response to electrode implantation is reviewed elsewhere (Fernandez et al., 2014; Kozai et al., 2015).

# IMMUNE CELLS IN THE CNS

It is critical to decipher which cell types are responsible for generating and potentiating immune responses in the CNS in order to potentially address the issue of immune-mediated device failure. In peripheral tissues, dendritic cells (DCs) and circulating monocytes serve as the major source of antigen presenting cells (APCs), which recognize damage-associated molecular patterns (DAMPs) or pathogen-associate molecular patterns (PAMPs), while presenting processed antigenic peptides in the draining lymph nodes (Itano and Jenkins, 2003).

However, in the CNS, APCs are made up of both CNS-resident microglia and blood-derived DCs (Ransohoff and Cardona, 2010). Microglia, characterized as CX3CR<sup>1</sup> <sup>+</sup>/Ly6Clo/CD45lo/Iba-1 <sup>+</sup>, serve as resident phagocytes and immune sentinels of the non-inflamed CNS parenchyma, as they utilize their dendritic extensions to survey the entire CNS multiple times per day (Nimmerjahn et al., 2005). Microglia are present in both white and gray matter in varying densities and may possess different activation states based on their anatomic location (Lawson et al., 1990; Mittelbronn et al., 2001). Microglia have been shown to play a functional role in enhancing the activity of neuronal synapses through physical contacts. Disruption of these synapses by inflammatory stimuli could result in dissynchronization of neural networks (Akiyoshi et al., 2018). Because of these important Microglial functions, understanding microglia physiology is essential as they are major contributors to the success or failure of implantable bio-materials. Microglia begin to show morphologic changes from as far as 130 microns from probe insertion sites in as soon as 6 h post implantation (Kozai et al., 2012b), implying that these immune cells' reactivity may extend beyond direct physical contact.

# STRATEGIES FOR MODIFYING MICROGLIA FUNCTION

Since microglia play an integral role in the CNS immune response, both biological interference and mechanical device design play a central role in manipulating the neuroelectronic interface. For example, anti-Parkinson's disease monoamine oxidase B inhibitors have been shown to dampen the inflammatory response of microglia by regulating voltage-gated sodium channels, reducing the neuro-inflammatory reaction to MPTP (Hossain et al., 2018). Another novel target includes the THIK-1 potassium channel, which was recently shown to play an integral role in microglia ramification, surveillance, and was shown to be necessary for IL-1b release. Injection of tetrapentylammonium reduced microglia surveillance by 60% (Madry et al., 2018). Additionally, coating electrodes with neuronal adhesive molecules can help dampen the acute effects of the innate immune response by 80% (Eles et al., 2017). Lastly, reducing electrode size and enhancing their mechanical compliance with brain tissue also reduces the immune response to implanted bio-materials, as evidenced by diminished microglia and astrocyte activation (Kozai et al., 2012a). Whether these methods of immune suppression will lead to overall improved device longevity and decreased scar formation remains to be seen. Nevertheless, these and other approaches certainly bear further investigation.

# PERIPHERAL-DERIVED MYELOID CELLS

Another important immune cell type to consider is the CX3CR<sup>1</sup> <sup>+</sup>/Ly6C+/CD45hi/CCR2<sup>+</sup> blood derived macrophages that can be found in the perivascular and parameningeal spaces and can sample cerebrospinal fluid (CSF) in the arachnoid and Virchow-Robin spaces (McMahon et al., 2006; D'Agostino et al., 2012). Peripheral monocytes can be further divided into an inflammatory subset (CX3CR1loCCR2hiGR1+Ly6Chi) that are short-lived and home to inflamed tissue, or a patrolling subset (CX3CR1hiCCR2loGR1-Ly6Clo) that are long-lived and are recruited to non-inflamed tissue (Geissmann et al., 2003). Common to microglia and both monocyte subsets is the surface expression of CD14, a glycosylphosphatidylinositol-anchored protein notable for its role as a co-adaptor protein for toll-like receptor 2 (TLR-2), toll-like receptor 4 (TLR-4), and damage-associated molecular patterns (DAMPS) including heat shock protein 70 (hsp70) and S100A9 (Haziot et al., 1988; Yang et al., 1999; Asea et al., 2000; Beschorner et al., 2002; He et al., 2016). Using bone marrow chimera approach to distinguish the contribution of brain-resident microglia versus bone marrow-derived infiltrating monocytes, we observed that blood-derived monocyte/macrophages, not CNS-resident microglia, dominated the infiltrating cell population following microelectrode implantation site at 2 and 16 weeks post implantation, and their densities correlated with neuron loss at the microelectrode-tissue interface (Ravikumar et al., 2014b). In particular, functional CD14 on blood-derived monocytes and the presence of endotoxin contamination on intracortical microelectrode devices contribute to the decline in overall device performance (Ravikumar et al., 2014a; Bedell et al., 2018), so much so that specific inhibition of CD14 signaling pathway were shown to improve long-term performance of the implanted microelectrode (Hermann et al., 2018).

# APPLICATION OF INTRAVITAL 2P-LSM TO STUDY THE CNS

As host immune response is a dynamic process involving the recruitment, retention, and functional differentiation of highly motile immune cells to damaged CNS tissue sites, traditional static histologic examinations often do not fully capture the evolution of the inflammatory process. To this end, 2P-LSM allows for the visualization of fluorescently tagged structures deep within undisturbed living tissues in a time-resolved, dynamic fashion (Denk et al., 1990). Two photon excitation also enables second harmonic generation to illuminate ordered structures such as collagen and myosin without exogenous fluorescence probes, thus allowing for morphologic observation in collagen-rich structures within tissues such as the lymph node and the brain (Campagnola et al., 2002; Williams et al., 2005; Kawakami and Flugel, 2010). While the field of immunology has taken advantage of two photon technology in the past two decades (Germain et al., 2012), neuroscientists first began using this technology in the late 1990s to monitor calcium flux in organotypic brain slice cultures (Yuste and Denk, 1995). Since the early 2000s, intravital CNS imaging in the brain has flourished and scientists have been able to capture a variety of biological processes including but not exclusively: astrocyte reactivity, dendritic spine

turnover, formation of plaques in Alzheimer disease, microglial dynamics, and immune cell trafficking. Using intravital 2 photon laser scanning microscopy (2P-LSM), our group has identified three additional mechanisms of immune surveillance in various physiological and pathological states that may have unique implications for brain tissue reactivity to intracortical implantation devices.

For the remainder of this review, we will focus on these novel insights provided by a combination of disease models and application of intravital 2P-LSM to reveal synergistic mechanisms that maintain BBB disruption and promote ongoing inflammation within the brain parenchyma. First, similar to what has been reported in the intestinal tract, we have visualized inflamed myeloid cells extending dendritic processes across the BBB into pial blood vessels. Second, we have demonstrated that cytokine-induced BBB permeability increase is temporal-spatially heterogenous and includes prominent but transient vessel leaks. Third, we have demonstrated that geometric changes to the pial vasculature in response to inflammation promote the arrest of peripheral blood immune cells, promoting their recruitment into the brain. We propose that these additional mechanisms must be addressed to achieve faithful, fail-safe and durable recording and control by neural implants.

# INTRAVASCULAR DENDRITIC PROJECTIONS BY MYELOID CELLS THROUGH INTACT BBB

Recently, we described a unique mechanism whereby microglia extend their dendritic processes through the basement membrane and endothelium into the surrounding pial vessels, thereby providing a new potential avenue for immune cell interactions (Barkauskas et al., 2013). While extending dendritic processes between distinct anatomic compartments is not unique to the CNS (Niess et al., 2005; Girard et al., 2012), this is the first-time microscopy has captured this type of interaction between the immune privileged CNS and peripheral tissues. Using electron microscopy, we were able to confirm astrocyte end feet on either side of the extension, indicating that the astrocytes are actively participating in this breach of the BBB. At baseline in the gray matter of the cortex, we found a density of 175 CX3CR<sup>1</sup> <sup>+</sup> cell projections per mm<sup>2</sup> of vessel wall surface compared with only 75 projections in the dorsal column white matter in the spinal cord. Using three distinct pathologic models, we compared the frequency of projections in different inflammatory states. During the induction phase of experimental autoimmune encephalomyelitis (EAE), we found that the number of extensions doubled in the cortex by day 9 of induction and doubled in the spine by day 12, even though clinical symptoms usually only appear by day 12. In the cortex, projections could be seen in vessels of both small and large caliber, while in the spinal cord, they were only seen in association with small caliber vessels. In contrast to this EAE model, studies of traumatic spinal cord injury showed a 40% reduction in projections that slowly recovered to 85% of baseline by day 8 after injury. Projections in spinal cord injury were associated with both large and small caliber vessels. Of interest, in a third model of primary CNS malignancy in the gray matter of the cortex, we observed dendritic extensions into the tumor neo-vasculature with a frequency of less than half of baseline (73/mm<sup>2</sup> vs. 175/mm<sup>2</sup> ). These intriguing observations required further investigations in order to understand the molecular and cellular signaling between microglia, surrounding CNS parenchyma, and endothelial tissues that regulate this behavior under physiological and different disease states.

Differential contribution of resident microglia and bloodderived monocytes/macrophages to neuron loss during CNS trauma was also demonstrated by direct intravital observation by 2P-LSM. Using a traumatic spinal crush injury model, we investigated the role of inflammatory monocytes in the process of post-traumatic axonal dieback (Evans et al., 2014). Immediately after crush injury, CX3CR1+ cells extravasated from blood vessels into the lesion and increased their abundance by sixfold in 5 days. Parenchymal microglia, on the other hand, did not accumulate at the injury site. While imaging animals on days 2, 5, and 8 post-injury, we observed axon-macrophage contacts that were associated with morphologic changes of the axon, ultimately resulting in either thinning of the axon or destruction within an hour.

# TRANSIENT CNS VESSEL LEAK AND DC ACCUMULATION

Peripheral monocytes can also differentiate into DCs, characterized as CD11c+, and have the potential to stimulate CD8+ T-cells (Geissmann et al., 2003). Using an EAE model, we utilized intravital 2P-LSM to elucidate the initiating events in EAE and found that in fact multiple cell types are required to induce durable immune responses. First, we found transient focal blood vessel leaks from post-capillary venule junctions that occur within hours to days immediately following EAE induction before any clinically evidence symptoms can be appreciated. These leaks lasted as long as 30 min and where rapidly cleared by CSF flow near the pial surfaces but persisted when vessel contents leaked into the parenchyma. We visualized CX3CR<sup>1</sup> microglia phagocytose blood vessel contents, which lead to their activation in the first 3 days following induction of EAE. The phagocytic CX3CR1+ microglia were then associated with an increase in CD11c+ DCs by day 6 post induction, which was then followed by infiltration and accumulation of antigen-specific T cells by day 9. Both antigen-specific T-cells and DCs congregated around areas of previous vessel leak, forming immune clusters within the CNS parenchyma. Following these observations of immune clusters, clinical EAE symptoms then started around day 10–12. Of interest, treatment with hydroxyzine ultimately decreased microglia activation and phagocytic capacity by day 3 and resulted in decreased accumulation of peripheral immune cells by day 6 (Barkauskas et al., 2015).

BBB Permeability has been implicated in the progression of multiple pathologic processes including Alzheimer's disease (AD), psychosis, and systemic lupus erythematosus (Bowman et al., 2007; Hirohata et al., 2018; Pollak et al., 2018). Indeed, an inverse correlation has also been demonstrated between CNS electronic implant performance and the severity and frequency of BBB breach (Saxena et al., 2013; Fernandez et al., 2014). Histamine, bradykinin, fibrin, and clotting pathways have all been implicated in modulating BBB permeability (Ryu and McLarnon, 2009; Passos et al., 2013; Chen et al., 2017). In fact, the accumulation of CNS infiltrating DCs was further demonstrated recently in a murine model of Alzheimer's disease, a process that was directly related to plasma Factor XII. Reduction in Factor XII correlated with reduced inflammation, neuronal damage, and improved cognitive function in the early stages of AD (Chen et al., 2017). Utilizing intravital 2P-LSM in these disease models would be helpful in establishing whether similar vessel leaks as observed in our EAE studies (Barkauskas et al., 2015) could also serve as a primary mechanism of cell accumulation in these other models of neurodegenerative diseases. **Figure 1** summarizes the above-mentioned findings regarding the functional roles of CD14<sup>+</sup> monocytes, DC, macrophages and microglia in traumatic CNS injury and associated BBB breach such as intracortical electrode implantation.

# INTRAVASCULAR CELLULAR DYNAMICS

While our understanding of the role of different immune subsets involved in implantable bio-material device failure will continue to evolve, one constant parameter is the absolute requirement for circulating cells to successfully navigate the vast vasculature as they traverse and home to sites of insult; either in the periphery or in the CNS. When combining 2P-LSM with resonance scanning, we were able to visualize cellular flow within post-capillary venules in the CNS in real time (Benson et al., 2018). Our initial intravital observations revealed that cellular arrest was most prominent in areas of sudden vasculature volume expansion. As such, we designed biomimetic microfluidic devices that allowed efficient modeling of the CNS vasculature with different characteristic widths. Once coupled with a physiologic hematocrit, purified leukocytes exhibited slow rolling and adhesion that was dependent on

Intracellular adhesion molecule-1 (ICAM-1). The modeling of CNS vasculature with microfluidic devices and microscopy allowed us to demonstrate the critical importance of dynamic vessel geometry changes in promoting recruitment of peripheral blood leukocytes. This is likely to critically influence the early response to device implantation, since vessel dilatation is seen as an early response to implantation (Wellman and Kozai, 2018; Eles et al., 2019). Parallel work in suppressing vascular remodeling in lung inflammation suggests angiopoietin 2 (Ang2) as a potential additional target for this process (Le et al., 2015), in addition to TNF which is already known to contribute to gliosis and neuronal death after implantation.

# ADDITIONAL CONSIDERATIONS