Dr Lionel Dahan

Elena Ludovic

Ahuva

James

Angelica Lawrence

Eugene Curtis

Tony Dinis

Eric Leclerc

Maxime, Mathieu J, Cécile Angelo Hélène

Gilles

Arthur, Séraphin Romain Paul

Emilie Christophe, for all the in vivo aspects. Lionel, you've been an incredible guide in the streams of electrophysiology, I'll preciously keep all those perfect drawings of mouse brains and evoked potentials... I've enjoyed very much confronting myself to the neuroscience community, and I've learned a great deal from it. Thanks also, to my little friends with whiskers, Jack the 1rst, Aladdin, Baloo, Crochet, Dumbo and Eurêka, I've loved bringing you chocolate cereal from time to time... I would also like to thank Valentina Castagnola, the PhD student before me, you did an incredible work starting this project from scratch, it really inspired me. Many thanks to Amélie Degache, former intern and my friend, for going out of your way to help me with the preparation of implantable devices with such a crazy deadline, I loved working with you side by side and wish you the best for your own PhD ! This work could not have been possible without all the technical help I received in both the biochemical lab and the Clean Room. First of all, Charline, for cell cultures and everyday experiments, for giving a great musical mood in the lab, for all our talks. Thanks to Sandrine for her help with electrochemistry, Nicolas Mauran for your amazing work with the mechanical bench automation, Xavier Dollat for your energy and countless tool manufacturing and Xavier Bouquet for your help with packaging and welding. Also, many many thanks to all TEAM members for their help in the Clean Room. More specically, Aurélie and David B for providing so much time and energy in the development of Parylene C plasma etching, Adrian and Véronique for all Parylene-related issues, Benjamin for his help with SEM and other characterization methods, David C and Samuel for the packaging and integration issues and Fabien for inkjet attempts and recent OLED development. Having such a strong and helpful team for lab and Clean Room processes is, in my opinion, one of the greatest assets of working at LAAS.

National conferences

Journées Nationales du Réseau Doctoral en Micro-nanoélectronique (JNRDM), may 2015 A. Lecomte, V.Castagnola, L. Dahan, E. Descamps and C. Bergaud, Rigidication par la soie de microélectrodes implantables souples pour la mesure de l'activité cérébrale, Oral presentation.

Plasma Etch and Strip in Microtechnology (PESM), mai 2016

A. Lecestre, A. Lecomte

Introduction

Neural interfaces aim at bridging the brain to the outside world, through the recording and transducing of neural activity. As an example, Brain Computer Interfaces (BCI) aim at restoring or bypassing functional skills for patients with paralysis, motor dysfunction or limb loss. A variety of devices can be used for the recording of the Central Nervous System, where their invasiveness is directly linked to the application of the BCI. Macro-electrodes on the scalp can record slow (~100 Hz) variations of brain waves, and might be used for simple computer transduction, such as spelling devices. On the other hand, implantable electrodes are able to detect high quality Local Field Potentials and single neuron activity with great temporal resolution (0.2-7 kHz), that could be used to control more complex external, such as a robotic hand. Moreover the use for implantable devices are emerging in the treatment of numerous diseases, from Parkinson's Deep Brain Stimulation to dystonia and depression.

If implantable electrodes have been widely used for neurophysiological aims, the turn towards chronic implantation has brought many issues yet to be addressed. Unstable electrical recordings over time stem from body acceptance and implant material lifetime. The foreign body response brings device migration and encapsulation that eventually results in impedance increase of even loss of electrical signal. Metal corrosion and degradation of passivation layers also play a great role in device failure over long periods of time. To increase device acceptability, biocompatibility and material life time have become areas of interest, as well as insertion conditions and design optimization. One of the emergent approach brings out the use of exible, polymer-based neural shanks that would partially relieve the strain between the soft tissues and the implanted device. To that end, Parylene C, polyimide and SU-8 are currently being evaluated as substrate candidates for the next generation of chronically implanted neural devices.

The implantation of neural devices for long periods of time, as in the case of chronic neural interfaces, sets the background to this PhD project.

The rst Chapter aims at giving an overview of the related works to this project, from the understanding of neural network to the implantable devices that are currently used both commercially and for research purposes. A brief summary of chronic neural interfaces will be described, as well as the issues brought up by long-implanted devices in the brain, such as the body reaction against the neuroprosthesis. This enables us to understand the choice of a soft substrate and the actual challenges facing chronic neuroprosthetic devices nowadays.

In a second Chapter, the fabrication of our polymer-based neuroprosthesis is presented, focusing on the technological locks that remain for widespread use. The device was developped in the clean room at LAAS-CNRS, and special attention was brought to the biocompatibility of both chosen materials and resulting sensor.

Then, the third Chapter focuses on the mechanical properties of such neuroprosthesis, and more particularly on the implantation diculty that arises from probe exibility. After a thorough state-ofthe-art presentation on the subject, dierent insertion problematics are discussed, such as the insertion of single shank and multiple shanks in the cortex. The latter was carried out at USC Los Angeles, in the lab of Dr Ellis Meng, as part of a short visiting scholarship during the course of this thesis.

Furthermore, the in vitro and in vivo studies in mice are presented in Chapter 4. These preliminary tests are a stepping stone to assess the functionality and reliability over time of the implants developped in this thesis. All in vivo tests were carried out at the CRCA Lab (Research Center for Animal Cognition), with the help of Dr Lionel Dahan.

In the nal Chapter, the main results achieved during the course of this PhD will be summarized and short term perspectives will be introduced. Preliminary tests towards possible optogenetic studies are also given as a long-term perspective of this work.

Chapter 1

General introduction

1.1 The brain: cradle of body command

1.1.

The brain and the Central Nervous System

The brain is the most important organ in all vertebrate animals. With the spinal cord, it forms the Central Nervous System (CNS), that regulates the other organs of the body and constitutes the chair of cognitive functions. Indeed, all behaviours except for reexes, require the brain to incorporate the information transmitted through sensorial systems and to act accordingly. This information is transmitted via large and complex networks of neurons, that represent the most basic and core component of the brain. The peripheral nervous system, on the other hand, is responsible for the tranmission of information to the organs, via long networks of nerves.

As seen on Figure 1.1, dierent regions dene the brain. The cerebrum, majority component of the brain, is divided in two hemispheres, right and left, each of them split in four lobes : temporal, frontal, parietal and occipital lobes. The cerebellum, brain stem and the limbic system complete the brain structure. Each of these regions is associated with distinct brain functions. As an example, the occipital lobe is associated with visual processing, the cerebellum with coordination of movement, posture and balance and the hippocampus, as part of the limbic system, is involved in memory and spatial navigation. These regions of the brain communicate with each other via electrical signals propagating within the most fundamental cerebral cell: neurons. Indeed, the brain is a complex structure that can contain billions of neurons connected to one another. The neurons are cells that communicate via long protoplasmic bers called axons. The axon of a neuron transmits nervous inux called action potentials, to the target cells located in distant regions of the brain or the organism. The glial cells, even more numerous than neurons, are another cell type of the brain and insure the support of neurons and their function. Examples of the rich variety of neurons morphologies found in the human nervous system [3] A neuron is composed of a cellular body (soma), multiple dendrites and an axon, both dendrites and axon being referenced as neurites (Figure 1.3).

-The soma contains mainly the nucleus and cytoplasm, and its diameter varies from 5 to 120 µm.

-The dendrites are numerous, short and ramied from their origin. Their role is to direct the incoming nervous inux.

-The single axon's length varies from 1mm to 1m, and is responsible for the generation and propagation of the nervous inux.

Neurons communicate with each other by sending chemicals, called neurotransmitters, across a tiny space, the synaptic connections. A typical synapse is a contact between the axon of one neuron and a dendrite or soma of another. Synaptic signals may be excitatory or inhibitory. If the excitation from a rst neuron is powerful and long enough, the target neuron will generate a brief pulse called an action 1.1. THE BRAIN: CRADLE OF BODY COMMAND potential, which starts at its soma and propagates rapidly along the axon, activating synapses onto other neurons.

Depending on their relation to the outside world, three classes of neurons are established:

-sensory neurons, or aerent neurons, carry information from the periphery toward the brain or spinal cord.

-motor neurons or eerent neurons, carry information away from the brain or spinal cord.

-interneurons, with short axons, participate in the local circuitry of the brain.

1.1.2.2 The neuroglial cells Neuroglial cells (or glial cells) represent the other cell types present in the CNS. They might look similar to neurons, but do not participate directly in the general electrical network that denes brain communication, although their supportive functions help dene synaptic contacts and maintain the signaling abilities of neurons [3]. There are three main categories of glial cells (Figure 1.4):

-Astrocytes have for main function to maintain an appropriate chemical environnement for neurons.

-Oligodendrocytes are involved in the myelin formation of neural axons.

-Microglial cells (or simply microglia) are the equivalent of macrophages in the CNS: their role is to defend the organism against exterior bodies.

The role of glial cells in the body defense mechanisms will be explained further along in this Chapter.

The generation of electrical signal

Neurons produce electrical signals that carry information from one point to another. This electrical generation is based on the ow of ions across their plasma membranes. Action potentials are then propagated along the axons and are the fundamental signal that carries information from one place to another in the nervous system. The four steps of action potential propagation can be described as follows (Figure 1.5):

Resting state The resting membrane potential depends on diering concentrations of potassium (K+)

and sodium (Na+) ions inside (cytoplasm) and outside the neuron (extracellular uid). In the rest state, ion channels are closed and neurons generate a slightly negative potential (-60mV), called the resting membrane potential, that can be measured by recording the voltage between the inside and outside of nerve cells.

Depolarization The action potential corresponds to the change of this membrane potential to a positive value, with the mechanism of ion exchange along the membrane. Indeed, when the stimulus is higher than the threshold, specic voltage-gated ion channels will open and let passage of sodium inside the membrane: it is the depolarization. An action potential is created when the neuron is depolarized by about 20 mV to its threshold potential. The potential then rapidly becomes positive, reaching a value of around 40mV.

Repolarization The sodium channels are inactivated and the potassium channels open. In this phase, the potassium will ow outside the cell membrane and reinstore the balance of the membrane potential.

Hyperpolarization Often, the potential becomes more negative than the resting state, because calcium channels are close and potassium channels open. It is a temporary phase, before the membrane potential goes back to its original resting value. Retrieved from [4] It is important to notice that the intensity of the neural signal transmitted in a neuron is not coded by the magnitude of the membrane potential, as all action potentials are of the same intensity, but by the number of action potentials generated per unit time, called the ring rate.

Recording the brain activity 1.2.1 Brain recording methods

The information encoded in the transmission of the action potentials can be looked at from dierent levels (Figure 1.6). From a microscopic level, a single unit can be detected and recorded when the recording device is very close from the targeted neuron, that is to say from an implanted device in the brain structure. From a macroscopic level, the activity of a whole population of neurons can be detected as brain oscillations by an electrode positioned on the scalp, therefore without the need to implant the recording device.

The dierent types of brain signals can be recorded by several devices, classied depending on their level of invasiveness.

Non-invasive recording: Electroencephalography (EEG)

Electroencephalography (EEG) is a cerebral exploratory method that measures the electric activity of the brain via macro-electrodes placed on the scalp (Figure 1.6a and 1.7a). The electrical signal obtained is the result of large numbers of synchronous post-synaptic action potentials. The potentials occur primarily in cortical pyramidal cells, and thus must be conducted through several layers of the cerebral Figure 1.6: Representation of the placement of implanted and non-implanted recording devices at dierent depth, over, onto or inside the brain and the related recorded signals (adapted from [5,6]) cortex, cerebrospinal uid (CSF), pia mater, and arachnoid mater and nally the skull [START_REF] Schalk | Brain-computer interfaces using electrocorticographic signals[END_REF]. Therefore, one drawback in this technique is the level of signal attenuation.

1.2.1.2 Semi-invasive recording: Electrocorticography (ECoG) Electrocorticography (ECoG), or intracranial EEG, is an electrophysiological monitoring that uses electrodes placed between the scalp and the brain in order to record electrical activity from the cerebral cortex (Figure 1.6b and 1.7b). Even though the brain tissue is not penetrated, a craniotomy (hole in the skull) is required to place the electrode grid, which places the procedure as a semi-invasive approach.

Like EEG, ECoG signals are composed of synchronized postsynaptic potentials recorded from the exposed surface of the cortex. However, the skull is here absent, thus the signal is not attenuated by the low conductivity of bone [START_REF] Libenson | Practical approach to electroencephalography[END_REF]. The spatial resolution of an ECoG is therefore much higher, which can be benecial when a critical imaging is required (eg: epilepsy presurgical planning). ECoG oers a temporal resolution of approximately 5 ms and a spatial resolution of 1 cm [START_REF] Asano | Origin and Propagation of Epileptic Spasms Delineated on Electrocorticography[END_REF].

Invasive technique: Implanted microelectrodes

The ultimate method to record the brain activity is to use microelectrodes implanted directly in the brain, called cerebral intraparenchymal or intracortical electrodes, close to the signal source (Figure 1.6c). These electrodes will detect electrical signals from smaller populations of neuron, with greater resolution. The use of implanted microprobes range from pure neuroscience and elecrophysiology research to epilepsy diagnostic and Brain Computer Interfaces, as we'll developp further along.

Besides resolution, another advantage for implanted electrodes resides in the ability to record or stimulate cortical layers. For instance, Deep Brain Stimulation of burried hippothalamus layers has proven ecient for otherwise treatment-resistant movement and aective disorders such as Parkinson's disease, essential tremor and dystonia, among others [START_REF] Kringelbach | Translational principles of deep brain stimulation[END_REF]. The dierent existing cerebral microelectrodes and their use in neural interfaces will be dealt further in the next section.

Neural signal types

The sensor location determines the type of recorded neural signal that can be recorded, which is directly linked to the amount of information that can be retrieved from the recording. The two main categories of neural signals are eld potentials (FPs), that are rhythmic oscillations produced by large numbers of neurons, and single units, coming from neurons discharging (spikes) (Figure 1.8).

Figure 1.8: Categories of neural signals that can be recorded in the brain: spikes and eld potentials represent the two main categories. Field potentials are dierenciated into rhythmic signals such as slow cortical potentials (SCPs) or sensorimotor rhythms (SMRs, like delta, theta, alpha, beta and gamma waves), or into event-related potentials (ERPs) such as P300 or N100 (adapted from [START_REF] Donoghue | Bridging the brain to the world: a perspective on neural interface systems[END_REF][START_REF] Wolpaw | Braincomputer interfaces for communication and control[END_REF])

Field Potentials (FPs)

A eld fotential (FP) is an electrophysiological signal coming from the electric current owing from multiple nearby neurons. It varies as a result of the simultaneous activity of numerous postsynaptic potentials.

To measure eld potentials, the electrode needs to be placed suciently far from individual local neurons to prevent any particular cell from dominating the electrophysiological signal. This signal is then low-pass ltered, cut o at ~300 Hz. An FP can be measured using the EEG, ECoG or implanted technique, but the spatial resolution of the FP is all the better as the electrode is close to the source.

Implanted electrodes are able, for instance, to record local eld potentials (LFP) resulting from smaller populations of neurons, thus with better spatial resolution and higher amplitude (Figure 1.9). However, the temporal resolution only depends on the targeted type eld potential, not the position of the sensor. Figure 1.9: Simultaneously recorded LFP traces from the supercial (`surface') and deep (`depth') layers of the motor cortex in an anaesthetized cat and an intracellular trace from a layer 5 pyramidal neuron [START_REF] Buzsáki | The origin of extracellular elds and currents: EEG, ECoG, LFP and spikes[END_REF]-Note that the LFP presents with theta oscillations (frequency 4-10 Hz) and gamma oscillations nested in the theta oscillations (frequency 30-100 Hz).

Slow cortical potentials (SCPs)

Slow cortical potentials are gradual changes in the membrane potentials of cortical dendrites that last from 300 ms to several seconds. These slow FPs can be recorded with simple EEG, where a helmet with macroelectrodes is put on the scalp. They can be consciously auto-regulated via a closed loop referred to as biofeedback, where the user visualizes his own recordings and learns to modulate them. Because the method is non-invasive and relatively easy, it is starting to expand in the non-medical area, as a relaxation method or as part of a video game [START_REF] Buzsáki | The origin of extracellular elds and currents: EEG, ECoG, LFP and spikes[END_REF][START_REF] Wolpaw | Braincomputer interfaces for communication and control[END_REF].

Event-related potentials (ERPs)

ERPs are also eld potentials, but in this case they result directly from a specic sensory, cognitive, or motor event. The P300, for example, is a positive ERP that occurs 300ms after neuron ring. ERPs provide excellent temporal resolution, as they are triggered by a time-locked event. However the spatial resolution remains low, and it is dicult to obtain spatial information on these signal sources [START_REF] Luck | An introduction to the Event-Related Potential technique[END_REF].

Sensorimotor rhythms

They represent the most commonly recorded Field Potentials. They form large oscillations between 1 and 100 Hz popularly referred to as brain waves. Certain frequency range in the brain waves codes for various conscious and subconscious states. For example, large, slow oscillations in the range of 10-13 Hz (alpha waves) is typical of a relaxed, resting state of mind, whereas the theta waves in the range of 4-10 Hz, are essentially emitted when the subject is in active or exploratory behaviour (Figure 1.8). EEG is most often used to diagnose epilepsy, which causes abnormalities in EEG readings. It is also used to diagnose sleep disorders, coma, encephalopathies, and brain death. However, the poor spatial resolution of the method might involve the use of an additional diagnosis procedure [START_REF] Libenson | Practical approach to electroencephalography[END_REF].

Single units (spikes)

Recording action potentials from single neurons ring require great signal quality only reached when the probe is implanted in the brain. Indeed, theoretical models suggest that the distance required to maintain a recording between an electrode site and a neuron cell body is of the order of cell dimensions [START_REF] Polikov | Response of brain tissue to chronically implanted neural electrodes[END_REF]. Measuring ring neurons confers the most information compared to Field Potentials, as the spatiotemporal resolution is highly increased, thus are of use when precise electrophysiology recordings are required. The activity of a single neuron can be recorded either intracellularly or extracellularly.

Intracellular recording

Intracellular recording is the measurement of voltage or current across the membrane of a cell. It typically involves an electrode inserted in the cell and a reference electrode outside the cell in a saline solution.

The renowned patch-clamp method (Figure 1.10), developed by Neher and Sakmann in the 1970s allowed for the rst time to record currents of single ion channels [START_REF] Neher | Single-channel currents recorded from membrane of denervated frog muscle bres[END_REF]. Intracellular recording with glass pipette electrodes remains widely used for the study of the electrophysiological properties of single neural cells. Figure 1.10: llustration and microscopic image of the patch clamp technique (adapted from [START_REF] Molleman | Patch clamping: an introductory guide to patch clamp electrophysiology[END_REF])

Extracellular recording

Extracellular recordings from implanted microelectrodes allow for the recording of both individual cell and neural networks. Depending on the position of the probe, a single action potential or group of neuron ring will be recorded with dierent proles. For this reason, MultiElectrode Arrays (MEAs) and polytrodes are used to obtain multiple recordings of the same region: a single neuron ring can then be very precisely positioned geographically, which increases all the more the inherent spatiotemporal resolution of extracellular recordings.

Focus on implantable devices

Existing implantable electrode sensors include single recording methods (glass pipette, single wire) and multichannel recordings (microwires, planar silicon probes, and platforms with microelectrode arrays) [START_REF] Donoghue | Bridging the brain to the world: a perspective on neural interface systems[END_REF] (Figure 1.11). A tiny glass cone contains embedded neurotrophic factors and two wires. When the cone electrode is implanted into the brain, the neurotrophic factors favours neurite outgrowth inside the cone, so that when these neurons re, the electrical pulses are recorded by the two wires [START_REF] Kennedy | Restoration of neural output from a paralyzed patient by a direct brain connection[END_REF][START_REF] Brumberg | Brain-Computer Interfaces for speech communication[END_REF] (Figure 1.12). This type of single unit system has been implanted into the Broca's (speech) area of several advanced stage ALS (Amyotrophic Lateral Sclerosis) patients to enhance speech by identifying the specic neural ring patterns that correspond to dierent speech producing sounds [START_REF] Dilorenzo | Neuroengineering[END_REF]. Figure 1.12: Illustration a microscopic image of a Kennedy Neutrophic Electrode, also known as glass cone. The cone is lled with neurotrophic factor to promote neurite outgrowth inside the cone, where the two wires will record their impulses (adapted from [START_REF] Dilorenzo | Neuroengineering[END_REF])

Microwires and tetrodes

The rst implantable recording devices were made of metal wires such as tungsten, insulated except for the small tip, that are able to penetrate the brain and reach buried structures (Figure 1.13). These insulated microwires, developed in the 1950s, allowed neurophysiologists to study the activity of individual neurons in awake, behaving animals. This simple method is still used nowadays for both acute and chronic extracellular recordings. A tetrode of microwires (four microwires attached to each other) is also considered a useful tool in electrophysiology for in vitro and in vivo studies, because it enables to locate more precisely a neuron of interest with triangulation. One major drawback of these systems is that the overall geometry and setup is hardly reproducible. Besides, the biocompatibility of the materials used and the delamination of the passivation layers make this technique unreliable for chronic implantation [START_REF] Prasad | Comprehensive characterization and failure modes of tungsten microwire arrays in chronic neural implants[END_REF].

Utah array

Multi-neuronal recording allows for simultaneous recording of both neuronal populations and individual neurons, thus more information can be extracted from the brain. The Utah array is a multielectrode array (MEA) employed routinely by electrophysiologists [START_REF] Suner | Reliability of signals from a chronically implanted, silicon-based electrode array in non-human primate primary motor cortex[END_REF][START_REF] Donoghue | Bridging the brain to the world: a perspective on neural interface systems[END_REF]. It consists of a 10x10 conductive sharpened silicon needles, each one electrically isolated from its neighbours in which the tip is coated with platinum and the shanks are insulated with polyimide (Figure 1.14). Some of these devices have positionable electrodes, while others have modied single electrodes (with larger site sizes and/or reduced impedances) which are capable of recording neural activity without precise positioning [START_REF] Rousche | Chronic recording capability of the utah intracortical electrode array in cat sensory cortex[END_REF].

Michigan array

In the 1960s, the growing development of photolithographic techniques and silicon etching technology provided a new tool set for fabricating recording electrodes, permitting the microfabrication of structures with high precision in which the spatial resolution between the electrode sites can be precisely designed.

The technology was developed over the next two decades as part of the more general development of integrated sensors and microelectromechanical systems (MEMS) [START_REF] Wise | Wireless implantable microsystems: High-density electronic interfaces to the nervous system[END_REF]. One of the rst silicon-based multielectrode arrays were made by Wise and Angell [START_REF] Wise | A low capacitance multielectrode probe for use in extracellular neurophysiology[END_REF]. These structures have then evolved in what we know nowadays as the Michigan array. In this array several microelectrode sites are patterned on each shank of the structure, providing higher density of sensors while avoiding material addition. Besides, the patterns can be designed to match a specic brain structure (Figure 1.14).

These silicon-based intracortical interfaces have been studied for chronic implantation, but the recording quality and number of functioning electrodes have often shown to reduce over time [START_REF] Hochberg | Neuronal ensemble control of prosthetic devices by a human with tetraplegia[END_REF][START_REF] Suner | Reliability of signals from a chronically implanted, silicon-based electrode array in non-human primate primary motor cortex[END_REF][START_REF] Rousche | Chronic recording capability of the utah intracortical electrode array in cat sensory cortex[END_REF].

They are coated with a biocompatible polymer, such as polyimide [START_REF] Wise | Wireless implantable microsystems: High-density electronic interfaces to the nervous system[END_REF][START_REF] Nicolelis | Chronic, multisite, multielectrode recordings in macaque monkeys[END_REF], in order to provide electrical insulation but also to increase device biocompatibility, thus reducing the immune response post-insertion [START_REF] Polikov | Response of brain tissue to chronically implanted neural electrodes[END_REF]. Then, Rousche and coworkers eliminated silicon completely and created a exible electrode out of polyimide and gold, which would provide better neural contact and less mechanical tears into the brain: this was the beginning of a new generation of exible neural probes [START_REF] Rousche | Flexible polyimide-based intracortical electrode arrays with bioactive capability[END_REF].

Polymer-based probes

There has been a growing interest toward polymer-based implant materials that could be exible enough to mimic living tissue and reduce the mechanical damage without evoking any adverse tissue reaction. In the last 20 years, polymer-based implants using polyimide [START_REF] Mercanzini | Demonstration of cortical recording using novel exible polymer neural probes[END_REF][START_REF] Xiang | Ultra-thin exible polyimide neural probe embedded in a dissolvable maltose-coated microneedle[END_REF][START_REF] Sankar | A highly compliant serpentine shaped polyimide interconnect for front-end strain relief in chronic neural implants[END_REF], SU8 [START_REF] Altuna | SU-8 based microprobes for simultaneous neural depth recording and drug delivery in the brain[END_REF][START_REF] Cho | Biocompatible SU-8-based microprobes for recording neural spike signals from regenerated peripheral nerve bers[END_REF], BCB (benzocyclobutene) [START_REF] Singh | Improving mechanical stiness of coated benzocyclobutene (BCB) based neural implant[END_REF] and Parylene C [START_REF] Castagnola | Parylene-based exible neural probes with PEDOT coated surface for brain stimulation and recording[END_REF][START_REF] Kuo | Novel exible Parylene neural probe with 3D sheath structure for enhancing tissue integration[END_REF][START_REF] Takeuchi | Parylene exible neural probes integrated with microuidic channels[END_REF][START_REF] Kim | 3D Parylene sheath neural probe for chronic recordings[END_REF] as both the structural and insulation material have been micromachined for both acute and chronic recording (Figure 1.15). The fabrication of these devices typically involves metal sites sandwiched by thin lm of polymer using standard planar photolithographic techniques. Figure 1.15: Photographs of a polyimide [START_REF] Mercanzini | Demonstration of cortical recording using novel exible polymer neural probes[END_REF], a SU-8 [START_REF] Altuna | SU-8 based microprobes for simultaneous neural depth recording and drug delivery in the brain[END_REF] and a Parylene C [START_REF] Kim | 3D Parylene sheath neural probe for chronic recordings[END_REF] exible intracortical probes for chronic implantation

Chronic neural interfaces

All these methods and devices used for the recording of brain activity respond to a variety of needs, both in the neurology research and for the treatment of neurological disorders. Indeed, they are used in theoretical neuroscience to study physiological processes at the cellular level, to understand basic brain functions, and to analyse causes of their dysfunctions. The recent developments made at the intersection between neuroscience, computer science, engineering, and medicine have allowed the creation of new technologies to evaluate and treat nervous system disorders and to restore lost neural functions [START_REF] Donoghue | Bridging the brain to the world: a perspective on neural interface systems[END_REF].

Numerous diseases that impair neural pathways aect millions of lives worldwide. Amyotrophic lateral sclerosis (ALS), locked-in syndrome (LIS) brainstem stroke, brain or spinal cord injury, cerebral palsy and multiple sclerosis, to name a few, attack the sensorymotor functions and deprive the patient from motor or communication abilities. Therapeutic options are limited for the restoration of some losses. Dierent approaches are used in order to bypass the issue. For example, people largely paralyzed by massive brainstem lesions can use eye movements to answer questions, give simple commands, or even operate a word processing program [START_REF] Birbaumer | A spelling device for the paralysed[END_REF]. An alternative option is to provide the brain with a non-muscular communication channel for conveying messages and commands to the external world.

The term Neural Interfaces (NI) was proposed by Pr Donoghue in 2008 as a generic term to dene a system that senses neural signals and transmits to the outside world via a variety of means. In this sense, NIs include more specic terms such as Brain-Computer Interfaces (BCI), Brain-Machine Interfaces (BMI) or even Neuromotor Prosthesis (NMP). For example, NI research might focus on providing a kind of replacement part, or prosthesis, for the motor system. Also, a system that senses brain signals may have other roles in evaluating disease states, for example to predict seizure onset in epilepsy, or to guide therapy [START_REF] Donoghue | Bridging the brain to the world: a perspective on neural interface systems[END_REF].

Treating neurological disorders: Deep Brain Stimulation

One of the goals of NIs is to provide a treatment for certain neurological disorders. As a highly promising example, Deep Brain Stimulation (DBS) is an invasive method that consists in implanting long and thin intracortical probes in the deep layers of the brain, such as the subthalamic nucleus. These electrodes are connected to an electrical generator implanted under the epidermis, and that delivers a pulsed (130 Hz), low-amplitude electrical current (Figure 1.16). It is currently designed for the treatment of Parkinson's disease, essential tremor, dystonia and more recently, Obsessive-Compulsive Disorder (OCD) and major depression [START_REF] Kringelbach | Translational principles of deep brain stimulation[END_REF].

The work of Dr Benabid, with the help of Dr Pollak, in the late 1980s has largely contributed to the development of the technique [START_REF] Benabid | Deep brain stimulation for Parkinson disease[END_REF]. The procedure is reversible, adjustable and have a great spatial resolution, which explains its overwhelming success over ablation procedures [START_REF] Groothuis | Physiological challenges for intracortical electrodes[END_REF]. In 2002 DBS (Activa® Parkinson's Control System) was approved by the FDA as medical therapy for essential tremor in Parkinson's disease, and current research is investigating its ecacy in treating epilepsy, depression, and OCD [START_REF] Tarsy | Deep Brain Stimulation in Neurological and Psychiatric Disorders[END_REF]. Approximately 90,000 patients are nowadays implanted with a DBS setup. However, despite the long history of DBS, its underlying principles and mechanisms remain unclear. Another aim for these brain implants is the restoration of sensory functions, such as hearing. For instance, the cochlear implant is a widespread electronic implant for severely hearing-decient patients. The idea is to bypass the normal hearing pathway. The sound is captured via a microphone outside the skin, generally behind the ear, which will transmit a signal to an array of electrodes placed in the cochlea (Figure 1.17). These electrodes, when stimulating the cochlear nerve, will give the patient a sense of sound [START_REF] Roche | On the Horizon: Cochlear Implant Technology[END_REF]. More specically, the electrode array has platinum contacts and is inserted into the scala tympani of the cochlea, stimulating the auditory nerve through the bone [START_REF] Nicolelis | Actions from thoughts[END_REF]. In 2013, approximately 324,000 people over the world were implanted with cochlear implants, both adults and children. Although results vary from case to case, even slight improvements in auditory performance have helped people to communicate better and to become more aware of their surrounding environment. A Brain Computer Interface (BCI) is a system that allows a person to control an external device without using nerve or muscles: the command comes directly from the brain (Figure 1.18). First, the signal is recorded via an invasive or non-invasive technique as decribed earlier. Then, this signal is digitalized and translated into machine command. Finally, the user is able to get feedback from his thoughts.

BCIs are primarily investigated to assist individuals whose motor functions are severely impaired because of stroke, spinal cord injury, cerebral palsy and other severe motor-impairments. In this sense, therapeutic applications of BCIs oer users access to means of communication, independent locomotion and increased control over their environment. Their use, therefore, has the potential to improve signicantly users quality of life, independence and self esteem [START_REF] Hildt | Brain-Computer Interaction and Medical Access to the Brain: Individual, Social and Ethical Implications[END_REF]. Amyotrophic Lateral Sclerosis (ALS) is a progressive motor disease resulting in a complete destruction of the peripheral and central motor system but only aecting sensory or cognitive functions to a minor degree. The disease progresses until the patient has no control whatsoever on his muscles, resulting in a Locked-in State (LIS). In this case, a BCI that provides means to communicate with the outside world, such as cursor control or spelling device, is a priority [START_REF] Birbaumer | Breaking the silence: brain-computer interfaces (BCI) for communication and motor control[END_REF][START_REF] Brumberg | Brain-Computer Interfaces for speech communication[END_REF]. On the other hand, a partially motor-impaired patient will be aiming for a more complex external device, such as a prosthetic limb, to regain full motor ability.

Spelling letters and moving a robotic arm in 3D require dierent levels of complexity. Spelling letters calls for one-dimension input with less constraint in temporal resolution, that can be achieved through user modulation of slow waves. A robotic arm necessitate high spatial and temporal resolution for precise movements in a three-dimensional space, so that the command can only be reached through high resolution implanted probes (Figure 1.

19).

The key issue is selecting the optimal neural signal to provide the targeted control. The chosen neural control signal would need to achieve the quality of the communication link between the brain and the able body. Greater information content, speed of transfer (information rate), reliability, and signal accessibility are parameters that will guide the optimal neural signal selection [START_REF] Donoghue | Bridging the brain to the world: a perspective on neural interface systems[END_REF]. positive or negative SCPs upon the command of an auditory cue. They were then able, to some extent, to select letters appearing from a screen [START_REF] Birbaumer | A spelling device for the paralysed[END_REF]. Sensorimotor rhythms [START_REF] Wolpaw | Braincomputer interfaces for communication and control[END_REF][START_REF] Vaughan | EEG-Based Communication: Prospects and Problems[END_REF][START_REF] Lauer | EEG-based control of a hand grasp neuroprosthesis[END_REF] and event-related potentials [START_REF] Birbaumer | Breaking the silence: brain-computer interfaces (BCI) for communication and motor control[END_REF][START_REF] Farwell | Talking o the top of your head: toward a mental prosthesis utilizing event-related brain potentials[END_REF] ECoG-based BCI present with higher spatial resolution, signal delity, resistance to noise, and substantial robustness over long recording periods than EEG-based BCI. In this way ECoG recordings might represent a balance between delity and clinical practicality [START_REF] Schalk | Brain-computer interfaces using electrocorticographic signals[END_REF]. They achieve higher speed for similar spelling devices than EEG (17 characters per minute [START_REF] Brunner | Rapid Communication with a P300 Matrix Speller Using Electrocorticographic Signals[END_REF]) and can be involved in slightly more complex BCIs such as two-dimensional cursor command [START_REF] Schalk | Decoding two-dimensional movement trajectories using electrocorticographic signals in humans[END_REF]. Recently, a trial involving ECoG-based BCI was successful in decoding the hand movements of a post-stroke patient and used the resulting commands to control a prosthetic hand in real time [START_REF] Yanagisawa | Real-time control of a prosthetic hand using human electrocorticography signals[END_REF].

Implantable electrodes, on the other hand, have shown great promise in Local Field Potentialbased BCI. Kennedy's Neurotrophic Elecrode was successfully implanted in the premotor cortex of LIS patients, with the aim of moving a cursor mouse. Signal stability over time and accuracy were very encouraging [START_REF] Kennedy | Restoration of neural output from a paralyzed patient by a direct brain connection[END_REF][START_REF] Kennedy | Direct control of a computer from the human central nervous system[END_REF][START_REF] Kennedy | Computer control using human intracortical local eld potentials[END_REF].

Moreover, spikes and multi-unit activity are of special interest because they represent the richest source of movement information [START_REF] Donoghue | Bridging the brain to the world: a perspective on neural interface systems[END_REF]. Single-unit recordings have shown better accuracy for linear predictions of hand position, velocity and grasping force in non-human primate [START_REF] Carmena | Learning to control a brain-machine interface for reaching and grasping by primates[END_REF], therefore their recording is of special interest in the design of high-resolution BCIs. For instance, results of a study using electrode arrays implanted in monkey dorsal premotor cortex, indicate that a fast and accurate key selection system is possible (up to 15 words per minute) with the implantation of a 96 electrode array [START_REF] Santhanam | A high-performance braincomputer interface[END_REF]. [START_REF] Kim | Multistate decoding of point-and-click control signals from motor cortical activity in a human with tetraplegia[END_REF] (c) Self-feeding task for BrainGate patient Cathy Hutchinson [START_REF] Hochberg | Reach and grasp by people with tetraplegia using a neurally controlled robotic arm[END_REF] Impressive work has been done for the development of high resolution Brain Computer Interfaces, involving trials in both non-human primates [START_REF] Velliste | Cortical control of a prosthetic arm for self-feeding[END_REF][START_REF] Carmena | Learning to control a brain-machine interface for reaching and grasping by primates[END_REF][START_REF] Wessberg | Real-time prediction of hand trajectory by ensembles of cortical neurons in primates[END_REF][START_REF] Suner | Reliability of signals from a chronically implanted, silicon-based electrode array in non-human primate primary motor cortex[END_REF] and human patients [START_REF] Hochberg | Neuronal ensemble control of prosthetic devices by a human with tetraplegia[END_REF][START_REF] Collinger | High-performance neuroprosthetic control by an individual with tetraplegia[END_REF][START_REF] Hochberg | Reach and grasp by people with tetraplegia using a neurally controlled robotic arm[END_REF][START_REF] Kim | Multistate decoding of point-and-click control signals from motor cortical activity in a human with tetraplegia[END_REF].

As a stepping stone, in 2003 the BrainGate Neural Interface System presented a trial that signicantly contributed to the expansion of implanted BCI. In this trial, Matthew Nagle, a C3 tetraplegic, paralyzed from the neck down after being stabbed, was implanted with a 96-electrodes Utah array in the area of his motor cortex responsible for his hand movements. The implanted device was connected to a computer, and Matthew was able, after little training, to move a cursor mouse, check his emails and reproduce drawings with considered precision [START_REF] Hochberg | Neuronal ensemble control of prosthetic devices by a human with tetraplegia[END_REF] (Figure 1.21). Following FDA regulations and the study protocol, the implanted device was removed after one year. In 2012, another patient, Cathy Hutchinson, beneted from similar trial with BrainGate. She was able to move a robotic arm with the control of her thoughts, in order to self-feed a drink from a bottle, the rst time she was able to drink unaided in 15 years [START_REF] Hochberg | Reach and grasp by people with tetraplegia using a neurally controlled robotic arm[END_REF]. Around the same time, another tetraplegic patient, Jan Sheuermann, was implanted with 2 Utah arrays in her motor cortex in a trial with Motorlan, from the University of Pittsburg. Her training was completed in 13 weeks with the goal of controlling an anthropomorphic prosthetic limb with seven degrees of freedom (three-dimensional translation, three-dimensional orientation, one-dimensional grasping) [START_REF] Collinger | High-performance neuroprosthetic control by an individual with tetraplegia[END_REF]. Within weeks she could reach out, and change the position of the hand to pick up objects on a table, including cones, blocks and small balls, and put them down at another location.

In the future, BCIs will have the potential to impact not only individual users but also society as a whole. The research and development of future BCIs applications inevitably raise ethical and societal issues and a public debate on rights and restrictions must be encouraged. As the neurotechnologies continue to advance, it will be increasingly important to assess the neuroethical ramication of their deployment in society [START_REF] Montgomery | Novel neurotechnologies: intervening in the brain[END_REF]. 1.4 The Foreign Body Response to implanted neural probes Implanting an external device inside the brain inevitably brings out the question of body defense mechanisms. Even when the probe does not elicit a violent rejection from the brain, certain body response is to be expected, therefore a key aspect towards chronic implantation resides in the neural probe's ability to be well-accepted by neurological tissues. The body response in form of glial scar formation at the electrode/biological tissues interface is considered a primary source of electrode failure, for it isolates the recording sites from the neurons of interest [START_REF] Polikov | Response of brain tissue to chronically implanted neural electrodes[END_REF]. If the mechanisms for such impact remain unclear, studies have shown that a reduction of inammatory response involves greater recording quality and longevity [START_REF] Rennaker | Minocycline increases quality and longevity of chronic neural recordings[END_REF]. The foreign body response has shown to present two time scales: the acute reaction the rst few days following the implantation, and the chronic reaction, between two and six weeks post-insertion (Figure 1.22).

The acute reaction

The early reactive response is directly linked to the mechanical trauma of insertion, for the electrode severs capillaries, extracellular matrix and neuron cells as it makes its path into the cortex. Starting from implantation and during the rst week that follows, activated microglia proliferate at several hundreds 1.4. THE FOREIGN BODY RESPONSE TO IMPLANTED NEURAL PROBES of micrometers around the implant [START_REF] Mcconnell | Implanted neural electrodes cause chronic, local inammation that is correlated with local neurodegeneration[END_REF]. To lower this acute inammatory response, tip geometry, shaft diameter and mechanical compliance are studied [START_REF] Edell | Factors inuencing the biocompatibility of insertable silicon microshafts in cerebral cortex[END_REF]. Fluid displacement, vessel rupture, vessel severing and dragging of blood vessels by the device are to be avoided, for they lead to greater foreign body response [START_REF] Fernandez | Acute human brain responses to intracortical microelectrode arrays: challenges and future prospects[END_REF][START_REF] Bjornsson | Eects of insertion conditions on tissue strain and vascular damage during neuroprosthetic device insertion[END_REF] (Figure 1.22a). Slow insertion allows for neuronal tissue to adjust to the implant, and so will minimize the damage caused by the electrode [START_REF] Nicolelis | Chronic, multisite, multielectrode recordings in macaque monkeys[END_REF] but on the other side, fast insertion appear to cause the least amount of uid displacement along transcerebral arteries [START_REF] Bjornsson | Eects of insertion conditions on tissue strain and vascular damage during neuroprosthetic device insertion[END_REF]. Blunt tips might drag microvascular networks whereas sharp tips tend to sever microvessels [START_REF] Edell | Factors inuencing the biocompatibility of insertable silicon microshafts in cerebral cortex[END_REF]. Essentially, high speed insertion and a rened tip present with the advantage of decreasing pial tissue deformation (Figure 1.23).

The chronic reaction

The sustained reactive response is characterized by glial brillary acidic protein (GFAP)-secreting astrocytes also known as glial scar [START_REF] Turner | Cerebral Astrocyte Response to Micromachined Silicon Implants[END_REF]. After two-week implantation and until reaching stabilization at around 6 weeks, these astrocytes form a dense sheath that encapsulates and isolates mechanically and electrically the prosthesis from the surrounding neurons (Figure 1.22b). This glial scar appears to play a substantial role in the implant's performance: it creates both insulating sheath for the electrode and increased distance between the electrode and its targeted neurons, thus decreasing or even losing signal quality [START_REF] Liu | Stability of the interface between neural tissue and chronically implanted intracortical microelectrodes[END_REF][START_REF] Williams | Complex impedance spectroscopy for monitoring tissue responses to inserted neural implants[END_REF]. Astrogliosis may also be cause to rejection displacements of the device [START_REF] Groothuis | Physiological challenges for intracortical electrodes[END_REF], resulting in similar signal loss. Neurite extension as well have shown to be inhibited around the sheath, which leads neural regeneration to develop further from the implantation site [START_REF] Bovolenta | Nervous system proteoglycans as modulators of neurite outgrowth[END_REF] (Figure 1 .24).

Many studies reported unstable recordings or a decrease in action potential amplitude over time, consistent with either device migration or encapsulation [START_REF] Chestek | Long-term stability of neural prosthetic control signals from silicon cortical arrays in rhesus macaque motor cortex[END_REF][START_REF] Nicolelis | Chronic, multisite, multielectrode recordings in macaque monkeys[END_REF]. Either way, the chronic foreign body response largely contributes to electrode failure. Interventions from both mechanical and biological aspects have elaborated to promote electrode viability.

Biotic interventions

Biotic strategies have emerged to reduce this device encapsulation, often in the form of a coating that locally releases bioactive drugs. Dexamethasone (DEX), a corticosteroid anti-inammatory drug, has antibody, become activated and migrate toward the electrodes. Note the electrode track lled with blood cells (asterisk) and a nearby blood vessel. Calibration bar = 50µm (adapted from [START_REF] Fernandez | Acute human brain responses to intracortical microelectrode arrays: challenges and future prospects[END_REF]) been reported to signicantly attenuate astroglial response and reduce neuronal loss after silicon-based nitrocellulose-DEX probe implantation in rats [START_REF] Zhong | Dexamethasone-coated neural probes elicit attenuated inammatory response and neuronal loss compared to uncoated neural probes[END_REF]. Laminin-1-coated silicon electrodes have also shown improvements in chronic neural recordings through dispersion of the astroglial scar [START_REF] He | Nanoscale laminin coating modulates cortical scarring response around implanted silicon microelectrode arrays[END_REF]. Finally, neural regeneration at the biotic-abiotic interface can be promoted using chondroitinase ABC (ChABC) [START_REF] Filous | Immature astrocytes promote CNS axonal regeneration when combined with chondroitinase ABC[END_REF] that disrupts the neural growth inhibition by proteoglycans (e.g. CSPGs). Delivering small amounts of regeneration-enhancing drugs, such as Nerve Growth Factor (NGF), brain-derived neurotrophic factor (BDNF) or neurotrophin-3 (NT-3) [START_REF] Oudega | Neurotrophins promote regeneration of sensory axons in the adult rat spinal cord[END_REF], also promotes axion regeneration towards the implant site.

Preventing protein adhesion through non-fouling surface coating, as well as using bio-inspired surface coatings, appears to contribute to body acceptance [START_REF] Sommakia | Resistive and reactive changes to the impedance of intracortical microelectrodes can be mitigated with polyethylene glycol under acute in vitro and in vivo settings[END_REF][START_REF] Faveri | Bioinspired hybrid microelectrodes: a hybrid solution to improve long-term performance of chronic intracortical implants[END_REF].

Abiotic interventions

Concerning the size, a non-quantitative immunohistochemical study by Szarowski and coworkers in 2003 pointed out that dierent sizes, shapes, surface textures and tip geometry of neural probes produced very similar foreign body response and glial scar. This would suggest that strategies to improve device acceptance in the long run through purely structural changes may be ineective [START_REF] Szarowski | Brain responses to micro-machined silicon devices[END_REF]. However, two studies later pointed out contradictory ndings: Stice et al showed that a low cross section (diameter below 50 µm) in electrode wire reduced GFAP expression at four weeks implantation in rat somato-sensory cortex [START_REF] Stice | Thin microelectrodes reduce GFAP expression in the implant site in rodent somatosensory cortex[END_REF]. Thelin and coworkers then proved that after twelve weeks post-implantation, 50 µm diameter wires presented with reduced GFAP staining and higher neuron density around the shaft compared to 200 µm diameter wires. Specialized reviews on the subjet of biotic and abiotic interventions on the foreign body response are available for further understanding [START_REF] Polikov | Response of brain tissue to chronically implanted neural electrodes[END_REF][START_REF] Kozai | Biomaterials mechanical failure modes of chronically implanted planar silicon-based neural probes for laminar recording[END_REF][START_REF] Sommakia | Materials approaches for modulating neural tissue responses to implanted microelectrodes through mechanical and biochemical means[END_REF][START_REF] Groothuis | Physiological challenges for intracortical electrodes[END_REF].

In the light of these ndings, attention has recently turned towards material science, and more particularly towards biocompatible polymers as an insulate layer for standard rigid electrodes or as the device substrate in itself. 1.5 The turn towards exible implants 1.5.1 Why exible implants ?

Historically, silicon has been extensively exploited in neuroprosthetics. If an insulating layer oers an improvement in electrode viability, other parameters are yet to be taken into account. The mechanical mismatch between a standard silicon probe (modulus 150-170 GPa) and the surrounding tissues (modulus 3-100 kPa) still prompts a great deal of strain (Figure 1.25). Indeed, the brain micromotions respiration, vascular pulsatility, head movements-induce displacements in the order of tens of microns around the silicon or metal shank, which leads to an enhanced mechanical stress around the shank [START_REF] Muthuswamy | Tissue micromotion induced stress around brain implants[END_REF]. Stress amplitude has recently shown to be reduced by 30-50% for compliant probe (12 MPa) compared to silicon probes [START_REF] Sridharan | Compliant intracortical implants reduce strains and strain rates in brain tissue in vivo[END_REF]. Moreover a recent study based on immunohistochemistry suggested that compliant probes induces similar acute response during the rst few days after insertion but reduced inammatory response when compared to similar sti probes from two to sixteen weeks post-implantation [START_REF] Nguyen | Mechanically-compliant intracortical implants reduce the neuroinammatory response[END_REF].

Figure 1.25: Schematic scaling representing the mechanical properties (Young's modulus) of both brain tissues and neural probe substrate-Adapted from [START_REF] Ware | Smart polymers for neural interfaces[END_REF] Besides, this stress can be exacerbated as a result of mechanical tethering of the device to the skull [START_REF] Sommakia | Materials approaches for modulating neural tissue responses to implanted microelectrodes through mechanical and biochemical means[END_REF]. Studies have shown that compared with untethered, freely oating implants, skull-anchored microelectrodes elicited signicantly greater macrophage/microglial activation and increased astrogliosis [START_REF] Kim | Chronic response of adult rat brain tissue to implants anchored to the skull[END_REF]. Silicon-based microelectrode arrays with exible cable were able to relief part of the strain induced on the device, by decreasing the relative motion between the brain and the skull [START_REF] Ledochowitsch | Nanoex for neural nanopobes[END_REF][START_REF] Huang | Integrated parylene-cabled silicon probes for neural prosthetics[END_REF][START_REF] Sankar | A highly compliant serpentine shaped polyimide interconnect for front-end strain relief in chronic neural implants[END_REF] (Figure 1.26).

Figure 1.26: An illustration of tethered vs un-tethered (or anchored vs un-anchored) implantation techniques. Un-tethered implant is completely sealed beneath articial dura, leaving no connection to the articial bone or skull. In contrast, tethered implant is glued to the sealant materials. [START_REF] Thelin | Implant size and xation mode strongly inuence tissue reactions in the CNS[END_REF] In this context, exible polymers as all-substrate for implantable neuroprosthetic devices have three main advantages:

1) biocompatibility: chemically speaking, the device's biocompatibility is enhanced compared to traditional silicon neuroprosthetics [START_REF] Polikov | Response of brain tissue to chronically implanted neural electrodes[END_REF] 2) biostability: it allows for a certain degree of strain relief in the tissues thus a diminution in glial scarring [START_REF] Nguyen | Mechanically-compliant intracortical implants reduce the neuroinammatory response[END_REF][START_REF] Subbaroyan | A nite-element model of the mechanical eects of implantable microelectrodes in the cerebral cortex[END_REF] 3) eciency: the compliant contact favors adjacent neuron proximity and improves signal quality [START_REF] Kim | Dissolvable lms of silk broin for ultrathin conformal bio-integrated electronics[END_REF][START_REF] Harris | Mechanically adaptive intracortical implants improve the proximity of neuronal cell bodies[END_REF] Several polymers have been tried out as a soft substrate for intracortical recording electrodes, Parylene C [START_REF] Takeuchi | Parylene exible neural probes integrated with microuidic channels[END_REF], polyimide [START_REF] Mercanzini | Demonstration of cortical recording using novel exible polymer neural probes[END_REF], benzocyclobutene (BCB) [START_REF] Singh | Improving mechanical stiness of coated benzocyclobutene (BCB) based neural implant[END_REF] and SU-8 [START_REF] Huang | In vitro and in vivo characterization of SU-8 exible neuroprobe: From mechanical properties to electrophysiological recording[END_REF] being the most popular (Figure 1.25). Parylene C has proven to be a prominent choice due to its lower mechanical properties (Young's modulus 2.76 GPa), outstanding biocompatibility (ISO 10993 USP, Class VI biomaterial) and its chemical inertness. Both polyimide and Parylene have reported successful integration of buried microchannels in neural probes for local drug delivery system (DDS) [START_REF] Ziegler | Fabrication of exible neural probes with built-in microuidic channels by thermal bonding of parylene[END_REF][START_REF] Takeuchi | Parylene exible neural probes integrated with microuidic channels[END_REF][START_REF] Metz | Flexible polyimide probes with microelectrodes and embedded microuidic channels for simultaneous drug delivery and multi-channel monitoring of bioelectric activity[END_REF] which makes them highly competitive candidates for chronic neural recording. Thermo and chemo-responsive polymers are also primarly considered for such application, for their elastic modulus drop in physiological uids. [6].

Flexible probes are typically 100 to 500 µm wide and 10 to 30 µm thick, with a length varying from 1 to 6 mm in average [START_REF] Xiang | Ultra-thin exible polyimide neural probe embedded in a dissolvable maltose-coated microneedle[END_REF][START_REF] Chen | A three-dimensional exible microprobe array for neural recording assembled through electrostatic actuation[END_REF][START_REF] Mercanzini | Demonstration of cortical recording using novel exible polymer neural probes[END_REF][START_REF] Castagnola | Parylene-based exible neural probes with PEDOT coated surface for brain stimulation and recording[END_REF]. The increased width compared to traditional siliconbased or wire-based multielectrode arrays allows for the support of numerous recording sites on the device, patterned with thin layers of gold [START_REF] Ware | Thiol-ene/acrylate substrates for softening intracortical electrodes[END_REF][START_REF] Wester | Development and characterization of in vivo exible electrodes compatible with large tissue displacements[END_REF][START_REF] Castagnola | Parylene-based exible neural probes with PEDOT coated surface for brain stimulation and recording[END_REF], platinum [START_REF] Metz | Flexible polyimide probes with microelectrodes and embedded microuidic channels for simultaneous drug delivery and multi-channel monitoring of bioelectric activity[END_REF][START_REF] Kuo | Novel exible Parylene neural probe with 3D sheath structure for enhancing tissue integration[END_REF] and other biocompatible metals.

Flexibility versus compliancy

Flexibility scaling remains an unresolved issue. The designation of exible (or soft) has not yet been universally dened, and the polymers used as soft substrate for neural application involve a wide range of Young's moduli, as seen in gure 1.25. This lack of universal terminology might reveal misleading. For instance, a recent study showed that both microglial cells and astrocytes respond to mechanical cues in vitro as well as in vivo, where the sti probes used in the study had an elastic modulus of 10-30 kPa and exible probes a modulus of 100 Pa [START_REF] Moshayedi | The relationship between glial cell mechanosensitivity and foreign body reactions in the central nervous system[END_REF]. On the other hand, tens and maybe hundreds of studies imply that polyimide and Parylene C are soft material for neural applications, when their modulus reaches 2-5 GPa, that is to say 100,000 times stier. Other teams prefer the term compliancy, for the modulus is not the only parameter that needs to be taken into account: material density, surface chemistry and size are also predominant parameters [START_REF] Nguyen | Mechanically-compliant intracortical implants reduce the neuroinammatory response[END_REF][START_REF] Skousen | A strategy to passively reduce neuroinammation surrounding devices implanted chronically in brain tissue by manipulating device surface permeability[END_REF]. Reducing substrate thickness undoubtfully leads to greater compliance, as the buckling force will be signicantly reduced, so that the resulting device will t more closely to the tissues. Besides, in a study by Lind et al, low-density probes (1.35 ± 0.12 g/cm 3 ) presented with very little glial scarring, compared to high-density probes (21.45 g/cm 3 ) with same shape, size and isolating material [START_REF] Lind | The density dierence between tissue and neural probes is a key factor for glial scarring[END_REF]. According to the authors, as brain tissue density is estimated at 0.99 g/cm 3 , the gap in density leads to inertial forces when the implanted animal change speed or direction.

Therefore, the reduction of glial scarring through material means is a more global approach than the reduction in Young's modulus, and the biomimetism needs to be taken to the next level to increase implant acceptability.

Stretchable devices have recently shown great promise for strain-relief in subdural implantation, as the device is dynamically adaptive. A comparative study by Minev and coworkers has recently shown that a polyimide implant induced a pronounced attening of the simulated spinal cord, whereas the soft, elastomeric implant showed great compliance when the model was bent, avoiding local compressions along the hydrogel cord [START_REF] Minev | Electronic dura mater for long-term multimodal neural interfaces[END_REF].

Polymeric implants, however, have repeatedly shown to fail within months as a result of polymer swelling and layer delamination. Water uptake leads to lm cracks, swelling and hole formation that eventually aects the device mechanical and electrical properties. Strategies are emerging to overcome this moisture-barrier issue, such as atomic-layer-deposited aluminium oxide or silicon oxide on top of the encapsulation layer [START_REF] Minnikanti | Lifetime assessment of atomic-layer-deposited Al2O3-Parylene C bilayer coating for neural interfaces using accelerated age testing and electrochemical characterization[END_REF][START_REF] Xie | Long-term bilayer encapsulation performance of atomic layer deposited Al2O3 and Parylene C for biomedical implantable devices[END_REF][START_REF] Hogg | Protective multilayer packaging for long-term implantable medical devices[END_REF], or the annealing of the resulting device to increase the adhesion between the layers [START_REF] Von Metzen | The eects of annealing on mechanical, chemical, and physical properties and structural stability of Parylene C[END_REF][START_REF] Kim | Enhancement of the barrier performance in organic/inorganic multilayer thin-lm structures by annealing of the parylene layer[END_REF]. Jor and coworkers have acurately reviewed the issues linked to the presence of water over long periods of time [6].

The buckling issue

As explained in this section, the use of soft materials as substrate for neural probes aims at achieving better compliance with the surrounding neurons while maintaining minimal rejection. However, an important issue with exible neural probes is that they tend to fail penetrating the pia mater and reach their location goal. Indeed, a device that is too soft tends to bend when pressed against a surface, such as the brain. The buckling of such implants has been largely mentioned in the literature, which lead to a special focus this last decade over the best way to obtain an implant sti enough during surgical insertion, while maintaining its exibility once positioned in the cortex.

When soft polymers rst started to be used as probe substrate, the insertion could not be performed as such. In their experiment in 2001, Rousche and coworkers implanted their polyimide-based multi-Figure 1.27: Strategies to induce more device compliancy without acting on material exibility (Young's modulus) (a) Optical image of an electronic dura mater tailored for the spinal cord, made of stretchable elastomer [START_REF] Minev | Electronic dura mater for long-term multimodal neural interfaces[END_REF] (b) Illustration of the conformity induced by the reduction of thickness in the case of a polyimide ECoG implant [START_REF] Kim | Dissolvable lms of silk broin for ultrathin conformal bio-integrated electronics[END_REF] (c) A schematic and picture of the insertion in vivo of a Z-shape electrode array. The design is meant to match that of brain structures to enable some stress relief on the electrodes [START_REF] Agorelius | An array of highly exible electrodes with a tailored conguration locked by gelatin during implantation : initial evaluation in cortex cerebri of awake rats[END_REF].

Figure 1.28: The buckling of a exible polyimide-based neural implant on the surface of sh brain [START_REF] Wu | A exible sh-bone-shaped neural probe strengthened by biodegradable silk coating for enhanced biocompatibility[END_REF] channel intracortical MEMS for their in vitro and in vivo experiments on rat brains by performing a pial incision with a scalpel or a 100 μm-tungsten wire [START_REF] Rousche | Flexible polyimide-based intracortical electrode arrays with bioactive capability[END_REF], but then the risk of mechanical trauma increases considerably. Strategies have multiplied over the last decade over the best way to implant a exible probe in the brain. Some teams focus on the use of a retractable support [START_REF] Felix | Removable silicon insertion stieners for neural probes using polyethylene glycol as a biodissolvable adhesive[END_REF][START_REF] Barz | Novel technology for the in-plane transfer of multiple interconnection lines in 3D neural probes[END_REF], while others promote the integration of a permanent sti layer in the probe fabrication [START_REF] Kim | A hybrid silicon-parylene neural probe with locally exible regions[END_REF]. Bioresorbable polymers might also be added to the polymeric implant, providing a rigid structure that dissolves into brain tissues once the probe is implanted [START_REF] Takeuchi | Parylene exible neural probes integrated with microuidic channels[END_REF][START_REF] Xiang | Ultra-thin exible polyimide neural probe embedded in a dissolvable maltose-coated microneedle[END_REF].

The buckling issue, and strategies to overcome it, will be explored in more details in Chapter 3 of this thesis.

1.6 Challenges and perspectives for chronically implanted neural devices 1.6.1 Overcoming probe failure over time

In spite of the success and growing development of neural implants, many studies have shown chronic cortical recordings to be unreliable over time. As early as 1974, a progressive decline in unit recordings was shown in cat cerebral cortex after implantation, with only 8% of the electrodes functioning after ve months [START_REF] Burns | Recording for several days from single cortical neurons in completely unrestrained cats[END_REF]. Forty years later, recording instability is still a commonly documented problem.

For instance, Liu and coworkers reported that implanted electrodes are unstable during the acute phase of foreign body response, which leads to a continual decrease in recording ability over the following months [START_REF] Liu | Stability of the interface between neural tissue and chronically implanted intracortical microelectrodes[END_REF]. Recently two recent studies have both described uctuations in recording stability that agree well with these previous ndings [START_REF] Freire | Comprehensive analysis of tissue preservation and recording quality from chronic multielectrode implants[END_REF][START_REF] Ludwig | Chronic neural recordings using silicon microelectrode arrays electrochemically deposited with a poly(3,4-ethylenedioxythiophene) (PEDOT) lm[END_REF].

Scaling electrode number

Attention has focused over the last decades on Multielectrode Arrays (MEAs) with various transducers techniques, for either or multiple of these purposes [START_REF] Obien | Revealing neuronal function through microelectrode array recordings[END_REF]:

-to increase the number of neurons observed -to increase the reliability of spike sorting -to perform source localization Thanks to the recent advances in lithography techniques and material science, these MEA benet nowadays from a gradual increase in performance and reliability.

However, increasing the number of electrodes for implanted devices means having to drastically downsize single electrode dimensions, if the overall device is to match anatomic constraints. Besides, the integration of functional connectivity able to multiplex high-density arrays of electrode also remains a challenge.

Electrode/neuron interface optimization

The devices introduced in this chapter all present with numerous recording sites, principally made of conductive material such as gold, platinum, iridium oxide and glassy carbon [START_REF] Green | Conducting polymers for neural interfaces: Challenges in developing an eective long-term implant[END_REF]. For implanted neural recording devices, electrode downsize is a critical requirement for recordings with good spacial resolution, that is to say restricted to small populations of neurons. Besides, as explained in a previous section, high-density electrode arrays are of interest for resoluted Brain Computer Interfaces. Unfortunately the reduction in electrode size greatly increases dramatically electrode impedance, limiting recording sensitivity [START_REF] Groothuis | Physiological challenges for intracortical electrodes[END_REF]6]. A high electrode impedance results in a low signal-to-noise ratio (SNR), when the targeted SNR if 5:1 or higher if single action potentials are to be recorded and sorted [START_REF] Obien | Revealing neuronal function through microelectrode array recordings[END_REF].

Conductive polymers and nanostructuration To obtain small electrodes of a micrometer scale with low impedance, increasing the eective surface area of electrodes is a worthy solution (Figure 1.29).

It can be achieved by modication of a metal electrode with porous conductive materials such as black Platinum [START_REF] Rodger | Flexible parylene-based multielectrode array technology for high-density neural stimulation and recording[END_REF], conductive polymers such as polypyrrole [START_REF] Ates | A review study of (bio)sensor systems based on conducting polymers[END_REF][START_REF] George | Fabrication and biocompatibility of polypyrrole implants suitable for neural prosthetics[END_REF], PEDOT [START_REF] Castagnola | Parylene-based exible neural probes with PEDOT coated surface for brain stimulation and recording[END_REF][START_REF] Ludwig | Chronic neural recordings using silicon microelectrode arrays electrochemically deposited with a poly(3,4-ethylenedioxythiophene) (PEDOT) lm[END_REF] and nanomaterials (carbon nanotubes [START_REF] David-Pur | All-carbonnanotube exible multi-electrode array for neuronal recording and stimulation[END_REF][START_REF] Zhang | Tissue-compliant neural implants from microfabricated carbon nanotube multilayer composite[END_REF], nanowires [START_REF] Suyatin | Nanowire-Based Electrode for Acute In Vivo Neural Recordings in the Brain[END_REF], graphene [START_REF] Liu | Implantable graphene-based neural electrode interfaces for electrophysiology and neurochemistry in in vivo hyperacute stroke model[END_REF]). Indeed, they are able to detect biological activity by monitoring their electrical properties, and can be used in numerous electrochemical biosensors [START_REF] Nambiar | Conductive polymer-based sensors for biomedical applications[END_REF][START_REF] Green | Conducting polymers for neural interfaces: Challenges in developing an eective long-term implant[END_REF][START_REF] Ates | A review study of (bio)sensor systems based on conducting polymers[END_REF][START_REF] Liu | Conducting polymers with immobilised brillar collagen for enhanced neural interfacing[END_REF]6,[START_REF] Fattahi | A review of organic and inorganic biomaterials for neural interfaces[END_REF]. Contrary to smooth metal surfaces, they provide a high surface area appropriate for an intimate contact with the surrounding cells.

CMOS-based technology

The goal to integrate thousands of parallel single-unit recordings can be achieved thanks to the advances in novel active multi-electrode arrays [START_REF] Berdondini | Chapter 7: Active Pixel Sensor Multielectrode Array for High Spatiotemporal Resolution[END_REF]. Decreasing electrode size to the nanometer scale with reasonable impedance is however hard to achieve with simple passive metal electrodes. Field-eect transistors (FETs), on the other side, are able to detect the potential of the solution independently on the device/electrolyte interface impedance [START_REF] Basham | Design optimization of transistors used for neural recording[END_REF]. In this case, the signal transduction come from the transmembrane current from a neuron in the electrolyte that will polarize the gate dielectric and eventually lead to a change in the conductance of the underlying silicon channel. However, because of the poor biocompatibility of the oxide layer, FET-based devices are often limited to in vitro and acute in vivo sensing [START_REF] Ruther | Recent progress in neural probes using silicon MEMS technology[END_REF][START_REF] Duan | Chapter 2: Nanoscale Field-Eect Transistors for Minimally Invasive, High Spatial Resolution, and Three-Dimensional Action Potential Recording[END_REF].

Another solution resides in the use of organic electrochemical transistors (OECTs). Contrary to FETs, where the channel is separated from the electrolyte, the channel of an OECT is in direct contact with the latter, so that the channel/electrolyte interface constitutes an integral part of the operation mechanism of OECTs (Figure 1.30a). Organic semiconductors such as poly (3,4-ethylenedioxythiophene) doped with polystyrene sulfonate (PEDOT:PSS) has shown great promise for high-SNR recordings in vivo [START_REF] Khodagholy | In vivo recordings of brain activity using organic transistors[END_REF][START_REF] Khodagholy | NeuroGrid: recording action potentials from the surface of the brain[END_REF]. In general, transistors are starting to be used for a variety of electrochemical sensing in the neural eld, for instance to detect the eect of a stimulation treatment on the production chemical substance (e.g. dopamine) [START_REF] Van Dersarl | Integration of 2D and 3D thin lm glassy carbon electrode arrays for electrochemical dopamine sensing in exible neuroelectronic implants[END_REF][START_REF] Tawk | Deep brain stimulation results in local glutamate and adenosine release: Investigation into the role of astrocytes[END_REF][START_REF] Johnson | Implantable microelectrode arrays for simultaneous electrophysiological and neurochemical recordings[END_REF], to characterize brain glucose mechanisms [START_REF] Wei | Simultaneous recording of brain extracellular glucose, spike and local eld potential in real time using an implantable microelectrode array with nano-materials[END_REF] or to follow ion dynamics during seizures [START_REF] Raimondo | Ion dynamics during seizures[END_REF].

Reliable Connectic and packaging

Connection with a long-term implanted neural recording device is not trivial, as the packaging must meet certain criteria :

low noise: the overall connection must provide the smallest resistance possible to keep up with recording quality.

inertness in biological uids: the implant xation must be so that the connectic is not altered by the presence of biological tissues (corrosion, biofouling).

mechanical reliability: the connection must withstand the stress induced by subject movements over time.

electrical reliability: (especially for stimulating electrode) the constant current charge passing through the dierent channels may lead to cracks, overheating and passivation defaults to be prevented.

architecture customization: addressing large numbers of electrodes requires an appropriate multiplexing architecture.

While passive electrodes connecting relies on individual metal wiring of each recording site, the emerging generation of CMOS technologies are able to address high-density matrices that can put up to thousands of electrodes [START_REF] Berdondini | Chapter 7: Active Pixel Sensor Multielectrode Array for High Spatiotemporal Resolution[END_REF][START_REF] Ruther | New approaches for CMOS-based devices for large-scale neural recording[END_REF]. These techniques have just started to be combined with the use of a exible device. Viventi and coworkers have for instance recently proposed an subdural implant with CMOS silicon circuits assembled on a polyimide substrate 1.30. Obien and coworkers [START_REF] Obien | Revealing neuronal function through microelectrode array recordings[END_REF] recently compared the architectures used in MEA devices and the advantages/drawbacks associated with each of them (Figure 1.31). For small numbers of electrodes, the xed wiring with direct adressing is most commonly used, but when higher densities are considered, special care must be taken on the architecture that will allow low noise multiplexing at a speed that is suitable compared to the sampling frequency (generally between 1 and 10 kHz).

Scope of the thesis

As presented in the introduction, the implantation of neural devices for long periods of time, as in the case of chronic neural interfaces, sets the background to this PhD project. Since the body reaction to the implant is a preponderant factor in neuroprosthesis reliability over time, our work focuses on the fabrication of a bio-friendly implant that matches the brain tissues. The choice of soft materials and the fabrication of this probe is described in Chapter 2 of this thesis. Special attention is brought on biocompatibility, process repeatability and modularity of such implant.

Another technological lock that arises from this general introduction, is the implantation of such exible probe in the brain, as the exible shank will deform and buckle over the brain surface. This issue will be tackled in Chapter 3, where we describe solutions that enable the insertion of two dierent probe designs in the cortex.

Finally, the assessment of device reliability is crucial for future use in humans. For this reason, Chapter 4 is dedicated to the in vitro and in vivo tests that we carried out to target electrode failure modes and potential issues for chronic purposes. Little studies are available on chronic failures of exible probes, and with our 6-months study, we intend to bring a rst piece of the puzzle.

Chapter 2

Fabrication of a exible polymer-based neural implant

In this Chapter, the fabrication of a exible, Parylene C-based neural probe will be presented.

As explained previously, one of the main goal of this thesis is to design and fabricate an implantable electrode on soft substrate that would benet from improved eciency and lifetime over traditional neuroprosthetic devices. The choice of the polymer was developed as part of a previous thesis, as well as the very rst microfabrication protocols [START_REF] Castagnola | Implantable microelectrodes on soft substrate with nanostructured surface for stimulation and recording of brain activities[END_REF]. Parylene C is a transparent, conformably-deposited polymer that has gained attention over the years in the biomedical eld. Indeed, it presents with outstanding biocompatibility, excellent chemical inertness and low Young's modulus. The conception of a Parylene-based neural implant should come with good fabrication reproducibility, modularity and biocompatibility.

After an introduction of Parylene C and its remarkable yet challenging features, the dierent microfabrication steps are presented, with increased focus on the technological locks that such fabrication infers. Depending on the targeted in vivo procedure, dierent designs and corresponding packaging were developed during this thesis. Finally, primary biocompatibility assessment of the resulting devices is evaluated through in vitro cytotoxicity assays. -It is physically and chemically neutral and inert, with a conformal, pinhole-free thickness over the substrate [START_REF] Chun | Evaluation of sub-micrometer parylene C lms as an insulation layer using electrochemical impedance spectroscopy[END_REF][START_REF] Hogg | Protective multilayer packaging for long-term implantable medical devices[END_REF].

-It withstands the highest class of biocompatibility for plastics in the US: ISO 10993, USP (United States Pharmacopeia) class VI , the highest criteria for implantable materials [START_REF] Schmidt | Long-term implants of Parylene-C coated microelectrodes[END_REF].

-It has a low Young's modulus, which substantiates its obvious choice as an implant material and substrate [START_REF] Loeb | Parylene as a chronically stable, reproducible microelectrode insulator[END_REF].

Parylene C deposition

The deposition process, known as the Gorham process, begins with a granular dimer precursor, dip-xylylene [START_REF] Gorham | A new, general synthetic method for the preparation of linear poly-p-xylylene[END_REF]. This precursor is vaporized via sublimation and then pyrolyzed at a temperature above 550°C to cleave the dimer into its reactive radical monomer. Within the deposition chamber (at room temperature), the reactive monomer adsorbs to all exposed surfaces and begins to spontaneously polymerize to form conformal, poly(p-xylylene) lms (Figure 2.1). This process enables the control of (1) The precursor dimer is evaporated. (2) In the furnace, the dimer is pyrolyzed to cleave the dimer into monomer units. (3) The reactive monomers adsorb to the substrate within the deposition chamber and coats the samples. (4) A cold trap is also used within the setup to trap unreacted monomer from entering the vacuum pump. deposition parameters (vaporizer, pyrolysis temperature and chamber pressure) and also allows for the full conversion of the dimer into the polymer lm without any byproducts. An advantage of this process is that the substrate is coated at room temperature, allowing compatibility with thermally sensitive materials.

Parylene C for MEMS applications

Parylene has ideal properties for barrier applications as it is chemically inert, and the conformal, uniform deposition process allows for optimal coverage for encapsulation. Parylene C rst found use as a coating for implantable electronics, after a set of early experiments conrmed the polymer's biocompatibility [START_REF] Loeb | Parylene as a chronically stable, reproducible microelectrode insulator[END_REF][START_REF] Schmidt | Long-term implants of Parylene-C coated microelectrodes[END_REF]. This material rapidly gained popularity as a MEMS material due to the advantages of its deposition process and compatibility with standard micromachining and photolithographic processes (Figure 2.2).

Its pinhole-free polymerization at room temperature and eective gap ll made the coating process compatible with a variety of MEMS materials and structures for applications as a coating [START_REF] Hogg | Protective multilayer packaging for long-term implantable medical devices[END_REF][START_REF] Lo | Characterization of parylene as a water barrier via buried-in pentacene moisture sensors for soaking tests[END_REF] and a structural material [START_REF] He | Ion liquid chromatography on-a-chip with beads-packed parylene column[END_REF]. The ability to form smooth lms that incur little optical scattering and high transmittance in the visible spectrum, makes Parylene C suitable for applications requiring optical transparency [START_REF] Ledochowitsch | A transparent ECoG array for simultaneous recording and optogenetic stimulation[END_REF][START_REF] Yamamura | Flexible organic leds with parylene thin lms for biological implants[END_REF]. Currently, free lm Parylene C devices are increasing in popularity compared to other polymer (e.g. polyimide (PI) and poly (di-methyl siloxane (PDMS)) devices due to its simpler deposition process, and combination of high exibility and mechanical strength. However, as explained further in this section, one of the key-challenges in the use of Parylene resides in the diculty to obtain highly dened patterns of thick Parylene structures. [START_REF] Jean | In situ vapor-deposited parylene substrates for ultra-thin, lightweight organic solar cells[END_REF] (b) Force sensor array [START_REF] Kim | Parylene-based electrochemical-MEMS force sensor array for assessing neural probe insertion mechanics[END_REF] (c) Intracortical probe [START_REF] Kim | 3D Parylene sheath neural probe for chronic recordings[END_REF] and (d) Retinal implant [START_REF] Rodger | High-density exible parylene-based multielectrode arrays for retinal and spinal cord stimulation[END_REF] 2.1.4 Roughness and surface properties Roughness and surface properties are an important aspect to take into account in microelectronics, because they play a critical role in layer-by-layer adhesion. For instance, it is crucial that metal patterns on the Parylene lm present with good adhesion on the substrate, especially when the device is meant to withstand prolonged contact in physiological uids. and Atomic Force Microscope (AFM). The principal prole roughness parameters are displayed (R a , R q and R max ). Parylene presents with moderate roughness, though at some points on the substrate, lumps might be found, linked to the conformal deposition of Parylene on occasional dust on the substrate. SEM images with dierent magnications are also displayed.

Surface aspect

The surface aspect of Parylene depends on the substrate on which it is deposited. For a traditional, silicon wafer, the surface aspect of a 20µm-thick Parylene layer is shown in Figure 2.3. The corresponding surface roughness was measured both via interferometry and Atomic Force Microscopy (AFM), where:

-R a is the mean roughness, that is to say the arithmetic average of the absolute values of the roughness prole ordinates. R a is the usual parameter adopted in engineering, because it gives a general description of the height variations in the surface.

-R q , also called Root Mean Square (RMS), is the root mean square average of the roughness prole ordinates. It gives information on the spatial distribution of the roughness. When R a and R q are very dierent values, it means that the roughness is not well distributed, with signicant dierences in prole ordinates.

-R max is the largest single roughness depth within the evaluation length. It gives information on eventual big lumps or holes present within the surface sample.

The AFM measurement is more focused, with a surface sample of 5x5mm, when the interferometer scans on greater surfaces (~100x120µm). The results displayed on Figure 2.3 show that Parylene presents with moderate, well-distributed roughness (low and similar R a and R q ) with occasional lumps when greater surfaces are considered (R max = 150µm). This is due to the deposition process, because the Parylene coater used is not in the Clean Room. Eventually, dust and small particles remain in the cham-ber, and are conformally coated with Parylene, leading to these lumps on the surface. The phenomenon is therefore all the more prevailing as the coating is thick.

Surface modication

The wettability of a surface can be estimated by measuring the contact angle of a water drop on said surface: its wettability will induce a low contact angle, and conversely. After being deposited, Parylene is usually hydrophobic, with a droplet contact angle of θ ∼ 90°. However, it is possible to obtain a hydrophilic structure, through either slight plasma or UV ozone treatment of the surface.

Oxygen plasma A plasma treatment can be recommended to form low contact angle. Indeed, oxygenated plasma (such as O 2 plasma) induces functionalized oxygenated groups (e.g. hydroxyl, peroxide etc.). Such surface functionalization leads to a raise in its wettability, because the surface energy will be all the higher that it is clean [START_REF] Trantidou | Oxygen plasma induced hydrophilicity of Parylene-C thin lms[END_REF]. An oxygen plasma also interferes mechanically on the surface (slight etching due to energetic rays), which will lead to greater surface roughness.

In our case, a light plasma (200 W, 2 minutes, at 20 sccm) makes the contact angle drop to θ ∼ 20°. in wet surrounding is one of the most crucial properties for implant survival [START_REF] Seymour | The insulation performance of reactive parylene lms in implantable electronic devices[END_REF]. Whenever Parylene is used as a passivation layer, the coating must be awless, void free, and stable with good adhesion that is impermeable for ions [START_REF] Von Metzen | The eects of annealing on mechanical, chemical, and physical properties and structural stability of Parylene C[END_REF].

A common strategy relies in the use of an adhesion promotor, the silane A-174, known to improve to some extent the adhesion of Parylene C on silicon, on silicon compounds and passivated metal surfaces [START_REF] Hassler | Characterization of parylene C as an encapsulation material for implanted neural prostheses[END_REF]. To apply the adhesion promoter, the substrate is usually exposed to silane vapor, although it can also soak in the compound. The silane then chemically bonds to the surface, thereby providing something to which Parylene can mechanically adhere. However, this method doesn't prove as ecient for other polymers noble metals such as gold, commonly used in implanted devices [START_REF] Hassler | Characterization of parylene C as an encapsulation material for implanted neural prostheses[END_REF]. Moreover, the eect of silane on the biocompatibility of the resulting device has yet to be demonstrated [START_REF] Hsu | Eect of Thermal and Deposition Processes on the Surface Morphology, Crystallinity, and Adhesion of Parylene-C[END_REF][START_REF] Kim | Enhancement of the barrier performance in organic/inorganic multilayer thin-lm structures by annealing of the parylene layer[END_REF].

As explained previously, oxygen plasma pre-treatment of the substrate might also be used, where surface functionalization and slight roughness increase has shown to improve layer-to-layer adhesion of Parylene C [START_REF] Nguyen | Characterization of Biocompatible Parylene-C Coatings for BioMEMS Applications[END_REF]. But other studies suggested that the structure of Parylene C would adhere more to hydrophobic surfaces [START_REF] Chang | Adhesion-enhancing surface treatments for parylene deposition, 16th International Solid-State Sensors[END_REF][START_REF] Hassler | Characterization of parylene C as an encapsulation material for implanted neural prostheses[END_REF]. In this way, hexane, toluene, propylene carbonate and tetrauoromethane (CF4) plasma elicited a higher peeling force and better performance in soaking tests [START_REF] Chang | Adhesion-enhancing surface treatments for parylene deposition, 16th International Solid-State Sensors[END_REF].

The most ecient method yet achieved involves a thermal treatment of the Parylene layer to increase the bonding between the layers [START_REF] Kim | Enhancement of the barrier performance in organic/inorganic multilayer thin-lm structures by annealing of the parylene layer[END_REF], as explained in the next section. Post-fabrication thermal annealing of Parylene leads to an increase in crystallinity, which induces a slight change in stiness [START_REF] Hsu | Eect of Thermal and Deposition Processes on the Surface Morphology, Crystallinity, and Adhesion of Parylene-C[END_REF][START_REF] Von Metzen | The eects of annealing on mechanical, chemical, and physical properties and structural stability of Parylene C[END_REF] and reduces water vapor permeation [START_REF] Menon | Characterization of water vapor permeation through thin lm parylene C[END_REF] in the annealed lm. The annealing of Parylene consists in exposing the polymer to a temperature greater than the glass transition point (Tg=60-90°C), but below the melting temperature (Tm=290°C) during a certain amount of time.

Annealing of Parylene

In this way, the proportion of crystallinity versus amorphology in the Parylene structure is increased, and bonds between the layers are formed (Figure 2.6). The addition of a bonding pressure increases the phenomemon, so that solid, water-resistant Parylene-Parylene structures can be formed [START_REF] Ziegler | Fabrication of exible neural probes with built-in microuidic channels by thermal bonding of parylene[END_REF][START_REF] Yen | A pneumatic neural probe structure fabricated by parylene thermal bonding technique[END_REF][START_REF] Noh | Wafer bonding using microwave heating of parylene intermediate layers[END_REF].

Finally, the annealing must occur under vacuum or nitrogen ow, because oxygenated Parylene becomes brittle and less transparent [START_REF] Kuo | Novel exible Parylene neural probe with 3D sheath structure for enhancing tissue integration[END_REF][START_REF] Kim | Formation of three-dimensional Parylene C structures via thermoforming[END_REF][START_REF] Hassler | Characterization of parylene C as an encapsulation material for implanted neural prostheses[END_REF].

Overall, the annealing of Parylene C layers improves Parylene-to-Parylene adhesion, which should present as better device biostability in aqueous environment. The development of microfabrication techniques and technologies began in 1947 within the semiconductor industry with the demonstration of the rst transistor at Bell Laboratories. To fabricate a microdevice, the starting point is commonly a silicon wafer, although other types of substrate are expanding (glass, silica, plastic etc.). On this substrate, many processes are carried out, one after the other, and often need to be repeated. These processes usually involve depositing a lm, patterning and removing (or etching) portions of it. In this sense, the three main process categories are: photolithography, deposition and etching (Figure 2.7).

Photolithography Photolithography is the process by which a photo-sensitive material, also known as photoresist (PR), is deposited and patterned on a substrate. The photoresist typically serves as a mask for etching/deposition steps or to create sacricial or structural elements of the device. It is rst spun onto a wafer, usually using a spin-coating method, and patterned with UV light. Depending on the type of the photoresist, positive or negative tone, the UV light can either make exposed regions more soluble (positive) or insoluble (negative) within a developer solution. After UV exposure, the PR-coated wafer is developed within this solution to realize patterns for subsequent fabrication steps.

Deposition processes Deposition (or additive) processes involve steps to deposit a layer of material on the wafer. The layer formation is achieved through a variety of means, including evaporation (e.g. for metals) and chemical vapor deposition (CVD, e.g. for Parylene). The choice of deposition process varies on the compatibility of the process within the material to be coated (e.g. thermal compatibility) and the intended thickness. Adhesion of the deposited material on the underlying layer is also important to consider. Etching Subtractive processes involve steps to remove material from a wafer using either physical of chemical etching. In wet etching, the substrate is immersed in a solution that will attack specic layers, while in dry etching, physical attack (e.g. plasma) will dig onto the substrate. Some plasmas can link the physical and chemical reaction, such as RIE (Reactive Ion Etching), where the ions will have both chemical and physical interference with the sample. The choice of etching method depends on the desired material to be etched, its thickness and the desired sidewalls (isotropic or anisotropic).

In Reactive Ion Etching plasma (RIE), a gas is introduced into the vacuum chamber. The gas is ionized in reactive plasma using a radio frequency generator that excites a coil placed around the chamber. The plasma therefore produced will engage a physico-chemical attack of the wafer. In the mean time, the volatile reactions products are extracted by a pumping system (Figure 2.8). The overall fabrication steps are represented in Figure 2.10 and will be explained step by step in this section. A detailed version of these fabrication steps and the related parameters are also available in Table 2.1.

Parylene C micromachining

Parylene substrate deposition

The rst step is to deposit the Parylene layer that will act as a substrate for the implantable device.

As explained before, Parylene deposition involves the use of poly(p-xylylene) dimer as a precursor which is sublimated at 220°C in vaccum. It then diuses into the pyrolysis chamber, where at 670°C it is dissociated into a monomer. The latter nally goes into the process chamber where it condensates and deposit itself at 30°C conformly all over the chamber. The maximum pressure in the chamber is xed at 60.10 -3 mbar so that the coating is formed slowly and uniformly. Whenever this pressure was too high, the resulting Parylene layer resulted milky and brittle, because the monomer didn't have enough time to difuse properly in the chamber and formed amorphous aggregates.

The CVD machine used during this work was designed by the Swiss company COMELEC and purchased by the CIRIMAT lab at Labège, therefore the wafers were coated outside of the Clean Room

a = 5000rpm/sec v = 3600rpm t = 30sec V = 3mL -soft bake 45sec at 90 • C 2nd coating: a = 5000rpm/sec v = 850rpm t = 30sec V = 3mL -soft bake 90sec at 90 • C PR: MicroChemicals Coater: Gyrset SUSS Microtec UV exposure λ = 405nm (CH2) 15sec at 20mW/cm²-PEB 90sec at110 • C SUSS Microtec MA150 Development Microposit MF CD-26
t 30sec -Pre-etching bake The Parylene-covered wafers are then annealed at 150°C during 4 hours under nitrogen ow, and slowly cooled to atmosphere temperature, so that the Parylene achieves a higher degree of crystallinity. First, a slight plasma will act as an adhesion promotor for both the metal and the photoresist. Then, a negative photoresist (AZ-nLOF 2035, Mi-croChemicals) is patterned via photolithography, and the metal is evaporated and patterned via lift-o (Figure 2.13a.) A 200 nm-thick gold (Au) layer was evaporated on the substrate and a thin titanium (Ti) layer (50 nm) was used to improve adhesion between Parylene and gold.

V = 6mL a = 1500rpm/sec v = 200rpm t = 20sec a 2 = 1500rpm/sec v 2 = 1200rpm t 2 = 30sec

Microelectrode patterning

Passivation and contact opening

A thin Parylene layer (~1.3 µm) is deposited on top of the wafer to act as a passivation layer for the gold traces. This passivation then needs to be partially etched, so that the 50µm-circular recording areas and the contact pads be opened (Figure 2.13b.). The electrode surfaces and the contacts are opened by dry plasma etching of Parylene using a photoresist mask (AZ-ECI 3027, MicroChemicals) with a thickness of about 2.7 μm. The etching rate of Parylene with these parameters was found to be ~220 nm/min, when the etching rate of AZ-ECI photoresist is about ~340 nm/min. The remaining photoresist is then stripped in acetone. Note that this method enables us to etch thin layers of Parylene C with good prole cut and minimum residues (Figure 2.13c.).

Inuence of mask

As the overall Parylene device is thick (~25µm) and the material highly inert, the etching rate is considerable slow (between 0.1 and 1µm/min [START_REF] Meng | Parylene etching techniques for microuidics and bioMEMS[END_REF][START_REF] Selvarasah | A reusable high aspect ratio parylene-C shadow mask technology for diverse micropatterning applications[END_REF]). As a result, the photoresist usually etches faster than the targeted material, and so the selectivity (i.e. the ratio between Parylene and photoresist etching rate) can be lower than 1. In this case the chosen mask material must meet the following criteria:

-It must withstand long etching rates in the plasma conditions (high temperature etc.)

-It must provide conformal coverage during the time of plasma.

-It must be coated in the thickness required by the selectivity (i.e. if the selectivity is 0.5 and the Parylene layer to etch is 10 µm thick, then we must be able to form a mask coating over 20µm thick).

-The remaining residues in the openings must be lowest, so that we can easily retrieve the unitary designs from the silicon.

-The cut prole must be as vertical as possible.

Three dierent kinds of masks were tried out: metallic mask, positive and negative photoresists.

Metallic mask The metal mask used was a 50nm-thick Nickel mask, patterned via lift-o over the Parylene substrate. The main advantage of a metal mask is that it etches very slowly with an oxygen plasma, so that the layer to be deposited does not need to be thick. However, in this case the plasma is strong and induces sputtering where the metal was supposed to mask (Figure 2.14a.). This is a critical drawback, because the Parylene layer underneath is partially attacked, which leads to passivation defaults. Moreover, the Parylene residues found in the openings are troubling, often the unitary probes are hard to detach from the wafer, resulting in occasional tearing of the devices. Positive photoresist In the previous section, positive photoresist AZ ECI 3025 was used for the etching of thin layers of Parylene. It showed great results in the masking of the Parylene passivation layer, but etches twice as fast as Parylene, so that the etching of 23µm of the material would require at least 46µm of the PR. Yet, it is a very liquid resist that can only be coated up to 1.5µm at a time.

Therefore, photoresist AZ40XT was chosen as a positive PR candidate, because it can be patterned in thick layers. Unfortunately, this PR quickly overheated in the plasma chamber, and within minutes formed cracks and bubbles on its surface (Figure 2.14b.).

Negative photoresist Finally, a negative photoresist was tried out as a plasma mask. Intervia Bump Plating Negative (BPN) is a thick, high aspect ratio negative photoresist, that has shown at LAAS to withstand a 17:1 aspect ratio with convential photolithography processes [START_REF] Bourrier | BPN a new thick negative photoresist with high aspect ratio for MEMS applications[END_REF]. In our case, a 50 to 60 µm-thick BPN layer showed great promise as a mask for our ICP-RIE process, as shown on Figure 2.14c.

Its conformal coverage is able to withstand long and powerful plasma etching, with minimal residues and good prole cut. Besides, as we'll develop in the next section, the power, gas ux and temperature increases have shown little eect on the selectivity.

Inuence of Plasma parameters

As part of this study, two equipments were tried out for the etching of thick Parylene structures: the Aviza Technology Omega 201 (hereafter referred to as ICP3) and the STS Multiplex (hereafter referred to as P4). The characteristics for each test carried out with these pieces of equipment are listed in Table 2.2.

Gas type, gas ow, pressure, ICP power, bias power and temperature were assessed as to their inuence on the resulting Parylene etching rate and selectivity with comparison to BPN photoresist mask. (c) Bias and ICP power The increase in bias power, ranging from 0 to 80W, produced little eect on Parylene etching rate. As a second step, the P4 module was used, because it produced higher gas ow but more importantly higher ICP power: from 500W for the ICP3 module, to 2800 W for the P4 module (Table 2.2). Increasing the power means increasing the physical parameter in the RIE, therefore should have more impact on Parylene etching rate. Indeed, in these conditions, Parylene was etched twice as fast (at a rate of ~0.8µm/min) while maintaining a selectivity of ~0.5 over BPN photoresist. Indeed, a uoropolymer layer has shown to protect feature sidewalls from lateral etching [START_REF] Meng | Parylene etching techniques for microuidics and bioMEMS[END_REF]. By doing so, it also prevents some residues to deposit back in the openings during the plasma (Figure 2.15).

This Parylene etching trial allowed us to obtain an optimum method for the etching of over 20µm-thick Parylene layer with an etching rate of almost 1µm/min, which is at the state-ofthe-art [START_REF] Meng | Parylene etching techniques for microuidics and bioMEMS[END_REF][START_REF] Selvarasah | A high aspect ratio, exible, transparent and low-cost parylene-C shadow mask technology for micropatterning applications[END_REF][START_REF] Pang | Parylene Technology for Neural Probes Applications[END_REF]. The Star recipe described in Table 2.2, along with the use of a conformal, high-aspect ratio photoresist, the Intervia BPN, was further used in this thesis to cut our Parylenebased neural probes (Figure 2.16). The main drawback in the technique developped is the diculty to remove the hardened BPN remains after plasma. Indeed, it needs to be soaked in BP Stripper for 40 to 60 minutes at 80°C, with continuous stirring, which is time-consuming. The resulting unitary devices are then thoroughly rinsed in DI water (10 rinses at least) to make sure that no stripper residues are to be found on the devices.

Design and packaging

The dierent steps towards the fabrication of single Parylene-based neural probes was described in the previous section. These process steps remain the same whatever the chosen design (tetrode, multielectrode array etc.). During this thesis, dierent probe designs were fabricated, depending on the application targeted : rst tests, process optimization, in vivo recordings etc. In this section, we will describe the two latest designs (Generation 1 and 2, hereafter referred to as GEN1 and GEN2) that were fabricated along with the purpose of implantation. As electrical recordings were intended for these applications, packaging was also part of the process. The connectics were specically designed for each purpose, with each its own requirements. The implantations goals and procedure setup will be further developed in Chapter 4, but for now the specications of such trial need to be stated in order to understand the design and packaging developed for each particular use. The rst procedure in vivo in mouse brain took place in september 2014. The goal was to implant a single mouse and to measure dierent signals during a unique recording session, with no intent to record over longer periods of time. The main specications for such trial were :

-a design with 6 recording sites at 3 dierent levels of implantation, with specic distances between the electrodes to match the structure of a mouse hippocampus.

-a thin and long shank able to reach deep hippocampus layers in mouse brain.

-a wired connectic and packaging compatible with the implantation setup in the CRCA electrophysiology lab.

-an overall design and packaging space consumption that are of appropriate scale for a mouse brain, and that allows the implantation on the side of a stimulation electrode.

GEN1 design

The design of GEN1 probes was established in agreement with Dr Lionel Dahan. The number of recording sites, the distance between them, the length and width of the shank were discussed to match the implantation procedure. For instance, the targeted region of the brain, the hippocampus, relies deep in a mouse brain, therefore the shank must be long enough to reach such layers. The lateral structure allows for easy manipulation and provides convenient space for the placement of a stimulation electrode on the right side of the probe, as we'll see in Chapter 4.

The design of the rst generation is displayed on Figure 2.17. ). This way, it is robust enough to be plugged many times into a ZIF connector, without damaging the connection. This solution was appraised convenient for the setup generally used by the electrophysiologists for this in vivo procedure.

Wired connectic integration

Generation 2: chronic in vivo wireless recordings 2.3.2.1 Second trial specications

The second procedure in vivo in mouse brain took place in september 2015. This time, the probes were meant to remain implanted in the mice for longer periods of time, so that regular wireless recordings could be performed over time. Therefore, the specications were :

-a design with 4 recording electrodes at 2 dierent levels of implantation, with the same specic distances between the electrodes matching mouse hippocampus structure.

-a wireless connectic and packaging compatible with the implantation setup in the CRCA electrophysiology lab.

-an overall design as small as possible (length <1cm) so that the connectic part that remains on top of the mouse skull does not bother the animal.

GEN2 design

The design of GEN2 electrode organization is similar to that of GEN1 probes: the distance between the probes and their diameter remain the same, since the implantation procedure aims at the same brain region (mouse hippocampus). The number of recording sites (4) was chosen because the CRCA wireless acquisition system was set for a maximum of 4 inputs. The structure was designed so that the overall length of the probe was as small as possible. Besides, an identication number was added on the side, so that the probe be manipulated using this protuberance, and be identied easily from other probes (Figure 2.18). The wireless connectic is a lot harder to implement on the probe than the previous wired connectic, because :

Wireless connectic and packaging

-the limitations in space are hard to meet: the length of the packaged probe must be as small as possible because it aims at staying implanted in mouse brain for months. Therefore, the connecting part that will stay attached to the mouse skull should be critically small.

-the wireless aquisition system from MultiChannel Systems (Headstage W4-HS) uses 6 pins in 3D as an input, whereas our probe is thin and at.

-the pitch of such pin is 1.27mm, which is not a standard pitch for commercially available pins and connectors.

-because it is dedicated to a chronic use, the packaging system must be solid and reliable.

The development of such connection was performed by Amélie Degache during her 3 months-internship in the summer of 2015.

The GEN2 probe is bonded on a customized exible polyimide cable with golden traces (or ex, from Shenzhen Sienta Industry Co) via Anisotropic Conductive Film (ACF). An ACF is an insulating adhesive tape containing conductive beads. When pressed and heated, the beads atten and form an anistropic link between the upper and lower substrate. Wherever these substrates are metallic, it ensures a strong, reliable electrical connection without any possibility of cross-over between the contact pads (Figure 2.21).

Once the probe is mounted onto the ex, the latter can easily be plugged into a ZIF connector. In this particular in vivo procedure, before ACF bonding, the ex is pierced and welded onto the 1.27mm-pitch pin connector (DigiKey) that ts the connecting part of the wireless acquisition system (Figure 2.20).

The overall connectic system is then passivated with insulating biocompatible glue (Polytec EP 601).

Electrode nanostructuration using conductive PEDOT polymer

As explained in Chapter 1, the miniaturization of electrode size is a critical requirement for electrical recordings involving a narrow population of neurons. Our gold electrodes have a small diameter of 40 µm, and therefore have a high impedance, leading to poor recording sensitivity. One strategy relies the use of conductive polymers to reduce electrode impedance without increasing electrode dimension. Among these conductive polymers, poly(3,4-ethylenedioxythiophene) (PEDOT) is of high interest in the eld, thanks to its high-conductivity, biocompatibility and excellent stability [START_REF] Aregueta-Robles | Organic electrode coatings for next-generation neural interfaces[END_REF]. If PEDOT lms can be obtained through a variety of means, electrochemical polymerization is considered more adequate for neuroprosthetic applications because it enables control over the lm thickness, surface properties, and electrical conductivity [START_REF] Cui | Surface modication of neural recording electrodes with conducting polymer / biomolecule blends[END_REF]. This work was the focus of a previous PhD program developped between 2011 and 2014 [START_REF] Castagnola | Implantable microelectrodes on soft substrate with nanostructured surface for stimulation and recording of brain activities[END_REF]. During this time, Valentina Castagnola has carried out a study on the electropolymerization of PE-DOT:PSS for neural probes applications. More particularly, her work focused on the inuence of electrochemical parameters on the morphology, conductivity and stability of PEDOT coatings on gold [START_REF] Castagnola | Morphology and conductivity of PEDOT layers produced by dierent electrochemical routes[END_REF]. Some results of this study are displayed on Figure 2.22a.

An optimized protocol was elaborated in this line of work to obtain repeatable, stable polymerized PEDOT coating over the gold electrode with a diameter of 40 µm. The so-formed electrodes see their electrical properties improve signicantly. For example, at 1kHz (frequency usually retained for neural recordings) the impedance drops from 700kΩ to 10kΩ (Figure 2.22b). In this thesis, this very same protocol was used to improve the electrical properties of the Parylene probes, and further comparison was assessed for in vivo studies in Chapter 4. However, it is important to clarify that no incremental studies were performed in this line of work during this PhD program, and that the PEDOT coatings were merely obtained following the recipe developped in the earlier PhD program [START_REF] Castagnola | Morphology and conductivity of PEDOT layers produced by dierent electrochemical routes[END_REF][START_REF] Castagnola | Parylene-based exible neural probes with PEDOT coated surface for brain stimulation and recording[END_REF].

Cytotoxicity and viability assessments 2.5.1 Biocompatibility and cytotoxicity

Biocompatibility is a relativist notion to describe the ability of a material to perform with an appropriate host response in a specic application [START_REF] Black | Biological Performance of Materials: Fundamentals of Biocompatibility[END_REF]. In the case of implanted materials, it relates to the limitation of the host response to an exterior material, and the limitation of material response to a biological environnement. In both cases, the assessment of device biocompatibility requires long-lasting studies that matches current regulations. The U.S. Food and Drug Administration (FDA) and the International Organization for Standardization (ISO), for instance, guide the manufacturer towards a series of pre-clinical and clinical trials specic to the targeted biological application.

In our case, the materials in the probe confection already benet to some extent from clearance for long-term purposes: Parylene is an FDA-approved, USP class VI biomaterial, which means that it has meet the highest criteria in biocompatibility and can be used for chronic implantation. Gold is also used as part of commercially available implants, such as eyelid surgical weight and coronary stent [START_REF] Edelman | Gold-coated NIR stents in porcine coronary arteries[END_REF][START_REF] Demann | Gold as an implant in medicine and dentistry[END_REF].

However, the fabrication process in itself might endanger this intrinsic material biocompatibility. For instance, traces of solvents and photoresists in biopolymers have shown to advocate immune reactions [START_REF] Vernekar | SU-8 2000 rendered cytocompatible for neuronal bioMEMS applications[END_REF]. Therefore, it is critical to assess the overall manufacturing process by performing cytotoxicity assays on the resulting devices. The preliminary objective is to establish the non-cytotoxicity of the produced devices, after the clean room process. This can be achieved through a cell viability assay : the MTT assay.

It is a colorimetric assay that enables simple quantication of viability percentage in a medium. In concrete termes, an enzyme called succinate deshudrogenase is capable of reducing the tetrazolium salt MTT 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide, present in living cells, to its insoluble formazan, which has a purple color. Then, a DMSO (dimethyl sulfoxide) solubilization solution is added to dissolve the formazan product into a homogeneous colored solution. The absorbance of this colored solution can be quantied by measuring at 570 nm by a spectrophotometer. In other terms, the level of this purple color is directly related to the proportion of living cells in the medium [START_REF] Gerlier | Use of MTT colorimetric assay to measure cell activation[END_REF][START_REF] Ciapetti | In vitro evaluation of cell/biomaterial interaction by MTT assay[END_REF] (Figure

2.23).

Generally, the MTT assay is performed on cells in culture in empty wells (control) and wells with the device. The absorbance is compared between the control and the test, to see if there's a dierence in cell viability, at dierent times.

The cell line used here is SH-SY5Y, a thrice-cloned subline of the neuroblastoma cell line SK-N-SH, which was established in 1970 from a metastatic bone tumor. They are often used as in vitro models of neuronal function and dierentiation and adhere easily without the need for a protein adhesion promotor [START_REF] Biedler | Morphology and growth, tumorigenicity and cytogenetics of human neuroblastoma cells in continuous culture[END_REF].

Procedure

Sample preparation The probes are from Generation 2, small enough to t into a 24-wells culture plate. Three culture plates are used, one for each release time (24, 48 and 72 hours), and each of them containing four control wells and four wells with the devices. These Parylene devices tend to oat in culture medium, and have to be attached to the bottom of the well. We tried to ensure an hydrophilic surface through UV ozone on the probes but they eventually came back oating on top. The best option was to glue the devices using a drop of biocompatible glue (Polytec EP 601) and the glue was let dried in oven at 70°C overnight. First culture cells on wells containing a drop of biocompatible glue and control wells showed that if the glue is properly mixed (base and curing agent) and let dried enough time, it indeed doesn't interfere with cell viability (data not shown).

All the wells (control and samples) are sterilized in 70% ethanol and rinsed three times. The absorbance for sample and control wells is plotted on Figure 2.24. The graph shows that the population of living cells increases over time in a similar fashion for both control and sample wells, meaning that the SH-SY5Y adhered and proliferated in all wells. The viability percentage, as dened by the ratio between the absorbance measured from a well containing the device and the mean absorbance of control wells, is also calculated and plotted on the graph. At 24, 48 and 72h, the devices have shown a positive test, since the viability percentage highly exceeds 75% (between 93 and 99%).

Results

The variability in measurement is similar for both control and sample wells, which means it is more consistent with the manual proceedings of the MTT assay (cell seeding, solution homogeneity, volume distribution) than with sample fabrication.

Overall, in the light of these outcomes, the samples are considered non-cytotoxic for a biological environment.

Fluorescent viability assay

Principle of the uorescent viability assay

The uorescent viability assay, carried out using the LIVE/DEAD® Viability/Cytotoxicity Assay Kit (Invitrogen L3224) completes the quantitative MTT assay with a qualitative touch.

The kit is composed of a two-color uorescence cell viability assay, using dierent dyes for live and dead cells. The calcein AM dye penetrates intact, live cells to produce a uorescent compound retained in the cells cytoplasm. When excited at 485nm, this compound uoresces at 530nm, giving a green color to all living cells in the well. On the other hand, the Ethidium homodimer-1 dye (EthD-1) only enters cells with damaged membranes and undergoes an enhancement of uorescence upon binding to nucleic acids, hence the nuclei of dead cells produce a red (635nm) uorescent compoung when excited at 495nm. Therefore, the viability kit assay enables us to distinguish live and dead cells through uorescent imaging, where the live cells will appear in green and the nuclei of dead cells in red.

Procedure

The SH-SY5Y cells are seeded following the same procedure as for the MTT assay. This time, Parylene probes are not glued to the bottom of their petri dish, because the glue uoresces in red with the Ethidium dye. Besides, the medium used (DMEM) is without phenol red, as it interferes with the uorescent dyes.

After 24, 48 and 72 hours, the petri dish is retrieved from the oven and the DMEM medium replaced by a solution at 0.2mM Calcein/0.2mM Ethidium. Twenty minutes later, the uorescent images are taken using a BX51TF uorescence microscope (Figure 2.25).

Results

The uorescent images are displayed on Figure 2.26. Between 24 and 72 hours, the SH-SY5Y live cells proliferated signicantly around the probe, with spreading shape, which attests to the cells well being around the Parylene implants. In this case, no dead cell was present in the probe surroundings, but a few were indeed dyed by the ethidium in the well, conrming that the assay worked ne. This uorescent viability enables us to validate once more that our implants have bio-friendly, non-cytotoxic features.

Conclusion

In this chapter, the microfabrication process of Parylene C-based implants has been presented. This choice of substrate intends on increasing body acceptance, through a better biocompatibility and greater brain tissue compliance compared to traditionally-used silicon implants.

The Clean-room fabrication process was described, starting from a silicon wafer to the release of unitary implants. If standard microfabrication techniques benet from decades of knowledge improvements for usual substrates such as silicon, the turn towards polymer-based sensors still lacks of general expertise. Microfabrication processes on Parylene C have only been studied for a little over a decade, and technological locks are remaining nowadays to achieve the same level of perspective and hindsight as standard substrates. More particularly, Parylene C micromachining techniques are still considered a challenge, and special attention was brought in this thesis to obtain optimized, reproducible microfabrication protocols for the achievement of a Parylene C functional device.

Besides, the modularity of these fabrication steps was illustrated in a second phase. Indeed, two implant designs and corresponding packaging and connectics were detailed in this chapter. Depending on the targeted application, this modular approach could as easily be transferred to a variety of aims.

Finally, the non-cytotoxocity features of the resulting devices were shown by both quantitative and qualitative culture cell viability assays. In both tests, the Parylene C probes manufactured at LAAS showed no sign of biological toxicity.

Chapter 3

Mechanical insertion of exible implants in the brain

In this Chapter, the mechanical insertion of exible neural probes in the brain will be discussed.

As stated earlier in this manuscript, if substrate exibility presents with the advantage to adapt to the inherent soft tissues, this very same property involve some drawbacks in practical handling. Indeed, a soft implant will tend to deform and buckle on the brain surface, which makes surgical implantation quite challenging.

First, the requirements for proper insertion in the cortex are explained, along with a presentation of the strategies that have developed over the years to allow insertion of a soft implant in brain tissues. This section is part of a more detailed review on the subject, whose aim is to provide tools for potential researchers in the eld to adapt a specic strategy for insertion of their probes [START_REF] Lecomte | A review on soft polymer-based neural probes and insertion techniques for chronic implantation[END_REF].

Then, in the following sections, two insertion-based issues are investigated:

the insertion of a single Parylene shank into a mouse hippocampus

The insertion of the Parylene probes designed and fabricated in Chapter 2 is discussed. We focused on the integration of a sti yet bioresorbable polymer coating made of silk broin, that supports the probe during implantation before resorbing in the brain.

-the insertion of multi Parylene shanks into a rat hippocampus One of the goals of research in chronically implanted neuroprosthetic devices is to design probes with numerous recording sites, which involves an increase in the number of shanks. In this section, the insertion of longer Parylene multi-shanks for implantation into a rat's hippocampus will be studied.

More particularly, the eect of the number of shanks on insertion mechanisms is assessed. This study was carried out as part of an international exchange program with the University of Southern California, Los Angeles.

Insertion strategies for exible neural probes in the litterature

Many strategies have been developed over the past few years to obtain exible probes that are able to withstand implantation. All the ideas revolve around the same criteria to ensure a viable solution for chronically implanted neural electrodes: full biocompatibility, reduced mechanical injuries, optimal structure compatibility. If promising results have been established, no miracle solution has yet been achieved and further investigations need to be done before clinical trials can be considered.

Surgical considerations

For surgical implantation in the cortex, the skin needs to be incised and a craniotomy is performed to allow access to the meninges (Figure 3.1a). The dura mater is then punctured and the pia mater is revealed.

The pia mater consists of a brous impermeable membrane that retains cerebrospinal uid. With an elastic modulus reaching 70-100 MPa, it is highly stier than the white and grey matter underneath [START_REF] Ozawa | Mechanical properties and function of the spinal pia mater[END_REF][START_REF] Runza | Lumbar dura mater biomechanics: experimental characterization and scanning electron microscopy observations[END_REF][START_REF] Aimedieu | Tensile strength of cranial pia mater: preliminary results[END_REF], and can be nicked for easier implantation. Collagenase has been proposed in the literature as a structure breakdown of the collagen present in the pia mater, therefore allowing for easier implantation with minimum dimpling [START_REF] Kralik | Techniques for long-term multisite neuronal ensemble recordings in behaving animals[END_REF][START_REF] Paralikar | Collagenase-aided insertion of intracortical microelectrode arrays: evaluation of insertion force and chronic recording performance[END_REF].

Brain phantom

A common method to assess insertion ability is to perform insertions in brain phantom. Agar gels of dierent concentrations are commonly used as tissue phantoms, and more particularly gel at 0.6% in DI water (Figure 3.1b). At this concentration, the agarose gel has an elastic modulus of 10 kPa, value similar to white and grey matter of the brain [START_REF] Hassler | Characterization of parylene C as an encapsulation material for implanted neural prostheses[END_REF][START_REF] Hsu | Encapsulation of an integrated neural interface device with parylene C[END_REF]. Despite the lack of anisotropy modeling, this brain phantom has been widely used as a model of implantation for grey and white matter [START_REF] Mccracken | Mechanical transient-based magnetic resonance elastography[END_REF]. The pia mater is rarely considered for such model, probably because exible probes stand little chance to penetrate it [START_REF] Hosseini | Comparative study on the insertion behavior of cerebral microprobes[END_REF], and so puncturing might be applied.

Insertion model and theory

From an ideal mechanical point of view, a system is able to be implanted if the maximum compression force it can withstand before bending (called buckling force) is higher than the minimum force required for insertion in the soft tissues (called insertion force) (gure 3.2). The buckling force is a function of probe dimensions and Young's modulus, as seen in Euler's formula below (Equation C.1) :

F buckling = π 2 I x E (KL) 2 (3.1)
where F buckling is the buckling force, E is the beam Young's modulus, L its length (dimension along the z axis) and K the beam eective length factor, whose value depends on the beam xation. The moment of inertia I x along the x axis depends on the shape and dimension of the beam cross-section. In the case of a rectangular beam, it is equal to I x = ab 3 12 , where a is the beam thickness (dimension along the y axis) and b its width (dimension along the x axis, where the buckling occurs).

At this point, it is important to stress out that probe length plays a major role in device mechanical failure. As seen in equation C.1, the buckling forces is proportional to 1 L 2 , and a slight increase in length results in a substantial drop in probe's stiness. For this reason, the targeted brain area is a predominant factor in the choice of a stiening strategy. Diagram not to scale.

The emergence of insertion techniques

The insertion force for white and grey matter is commonly estimated around 0.5-2 mN [START_REF] Tian | Monitoring insertion force and electrode impedance during implantation of microwire electrodes[END_REF][START_REF] Jensen | Assessment of subdural insertion force of single-tine microelectrodes in rat cerebral cortex[END_REF][START_REF] Molloy | Experimental determination of the force required for insertion of a thermoseed into deep brain tissues[END_REF][START_REF] Howard | Measurement of the force required to move a neurosurgical probe through in vivo human brain tissue[END_REF]. However, from a pratical point of view, insertion forces in agarose gel can reach up to 10-50 mN before penetration [START_REF] Hosseini | Comparative study on the insertion behavior of cerebral microprobes[END_REF][START_REF] Harris | In vivo deployment of mechanically adaptive nanocomposites for intracortical microelectrodes[END_REF][START_REF] Lecomte | Silk and PEG as means to stien a parylene probe for insertion in the brain: toward a double time-scale tool for local drug delivery[END_REF]. Depending on tip shape, implant size, material used and insertion speed, certain dimpling is to be expected [START_REF] Bjornsson | Eects of insertion conditions on tissue strain and vascular damage during neuroprosthetic device insertion[END_REF][START_REF] Welkenhuysen | Eect of insertion speed on tissue response and insertion mechanics of a chronically implanted silicon-based neural probe[END_REF]. Low insertion force and dimpling are not the only parameters to take into account : the maximum shear force the brain will endure along the insertion track also represents a major concern (gure 3.2), and highly depends on the depth of insertion [START_REF] Andrei | A response surface model predicting the in vivo insertion behavior of micromachined neural implants[END_REF].

A study by Ruther's group showed signiciant changes in insertion and retraction forces for silicon, glass, polyimide and tungsten probes with varied shapes [START_REF] Hosseini | Comparative study on the insertion behavior of cerebral microprobes[END_REF]. In this study, the 460 µm wide, 10 µm thick probe was able to sustain insertion in agarose gel, but could not undergo pial insertion in vivo. In general, typical exible neural probes have shown to barely withstand implantation [START_REF] Wester | Development and characterization of in vivo exible electrodes compatible with large tissue displacements[END_REF].

First studies recount the use of extra tools such as a scalpel or a tungsten wire to create incisions meant to match the shaft spacing pattern before actual probe insertion [START_REF] Rousche | Flexible polyimide-based intracortical electrode arrays with bioactive capability[END_REF], but this enduces greater insertion trauma.

Strategies to enable insertion of exible intracortical probes

For the last ten years, strategies have emerged to overcome this exibility/stiness paradox. These strategies are represented in gure 3.3 and divided in four groups:

1) the integration of a permanent sti structure in the exible probe 2) the addition of a removable prop 3) the turn toward mechanically-adaptative polymers 4) the use of biodegradable polymer coating on a exible substrate [START_REF] Kim | 3D Parylene sheath neural probe for chronic recordings[END_REF]). (d) Schematic frames of insertion of a silicon stiener attached to the probe using Polyethylene Glycol (PEG) as glue, and subsequently extracted after implantation via addition of PBS (reproduced from [START_REF] Felix | Removable silicon insertion stieners for neural probes using polyethylene glycol as a biodissolvable adhesive[END_REF]). -Softening polymer: The exible substrate might also be made of a mechanically-adaptative polymer that will soften upon exposure to physiological tissues.

(e) Microscope view of a laser-micromachined cortical probe made of poly(vinyl acetate) (PVAc) matrix (reproduced from [START_REF] Hess | Development of a stimuli-responsive polymer nanocomposite toward biologically optimized, MEMS-based neural probes[END_REF]). (f ) Photograph of a 16-channel intracortical electrode array based on Thiol-ene acrylate substrate (reproduced from [START_REF] Ware | Thiol-ene/acrylate substrates for softening intracortical electrodes[END_REF]). -Bioresorbable coating: Finally, the integration of a sti biodegradable layer to the exible device might reveal promising: the supporting layer (Maltose, PEG, Silk broin, Collagen/Gelatin, chitosan, PLGA, to name a few) resorbs once the probe is positioned in the brain. (g) Frames of insertion in vivo of a maltose coated ultra-thin polyimide-based exible neural probe (reproduced from [START_REF] Xiang | Ultra-thin exible polyimide neural probe embedded in a dissolvable maltose-coated microneedle[END_REF]). (h) Image of a dyed silk-coated sh-bone shaped polyimide probe with tip shown (reproduced from [START_REF] Wu | A exible sh-bone-shaped neural probe strengthened by biodegradable silk coating for enhanced biocompatibility[END_REF]). (i) Parylene C-based neural probe back-coated with gutter-shaped dyed silk broin (reproduced from [START_REF] Lecomte | Silk and PEG as means to stien a parylene probe for insertion in the brain: toward a double time-scale tool for local drug delivery[END_REF]). (j) Parylene C uidic channel lled with PEG is inserted into a gelatin tissue model (left) and recovers exibility after insertion (right) (reproduced from [START_REF] Takeuchi | Parylene exible neural probes integrated with microuidic channels[END_REF]). increases its buckling force to 1.3mN, but the pia mater needs to be nicked for insertion in lamb cadaver brain [START_REF] Egert | Parylene microprobes with engineered stiness and shape for improved insertion[END_REF]. Lee & coworkers suggested a 5-10 µm silicon backbone enclosed to a polyimide-based probe [START_REF] Lee | Polyimide-based intracortical neural implant with improved structural stiness[END_REF]. Recently, a solid silicon/hollow Parylene C tube hybrid design provided adjustable buckling force up to 500 mN (Figure 3.3a) [START_REF] Kim | A hybrid silicon-parylene neural probe with locally exible regions[END_REF].

The addition of a permanent rigidier might however be considered counterproductive, for the strain relief it brings to cerebral tissues is questionable. The denition of tissue compliance is yet to be dened as function of the improvements it brings on foreign body response, before such hybrid systems can be considered as candidates for chronic implantation.

Removable aid

A rigid backbone or shuttle can also be temporarily attached to the compliant device and removed afterwards. For instance, the shuttle can be separated from the implant through the customization of the probe/shuttle interface. Using an hydrophilic self-assembled monolayer, Kozai and Kipke showed rst results in deep (8.5mm) insertion of a 12.5 µm thick, 200 µm wide polyimide-based probe [START_REF] Kozai | Insertion shuttle with carboxyl terminated self-assembled monolayer coatings for implanting exible polymer neural probes in the brain[END_REF].

However, the two surface have a high chance of undergoing sliding when surgically manipulated. The use of degradable PolyEthylene Glycol (PEG) as a temporary glue guarantees a more reliable adhesion [START_REF] Felix | Removable silicon insertion stieners for neural probes using polyethylene glycol as a biodissolvable adhesive[END_REF][START_REF] Sohal | The sinusoidal probe: a new approach to improve electrode longevity[END_REF] (Figure 3.3d). PEG is a highly biocompatible, fast-degrading polymer, that dissolves quickly with the addition of water, as discussed further in this chapter. The shuttle can also take the form of a cylindrical needle where the device to implant remains inside [START_REF] Wang | Micro-wing and pore design in an implantable FPC-based neural stimulation probe for minimally invasive surgery[END_REF], and vice versa: the needle can be placed inside a tube-shaped probe [START_REF] Tamaki | Development and evaluation of tupe-shaped neural probe with working channel[END_REF] or a hollow tip (Figure 3.3c) [START_REF] Kim | 3D Parylene sheath neural probe for chronic recordings[END_REF].

Nevertheless, no matter the shuttle geometry or the adhesion type, a removable sti substrate undoubtedly would cause displacement and insertion trauma issues. Some shuttles increase over ve times the probe dimension [START_REF] Barz | Novel technology for the in-plane transfer of multiple interconnection lines in 3D neural probes[END_REF][START_REF] Sohal | The sinusoidal probe: a new approach to improve electrode longevity[END_REF], extending considerably the print during implantation. Moreover, the alignment step between the probe and its support makes the system doubtfully reliable for mass reproducibility.

Mechanically-adaptative substrate

Smart polymers such as shape memory polymers developed essentially in the 21rst century as mechanicallytunable polymers upon dierent classes of stimulus, such as light, temperature, humidity, etc. [START_REF] Ward | Thermoresponsive polymers for biomedical applications[END_REF][START_REF] Ware | Smart polymers for neural interfaces[END_REF].

Chemoresponsive polymer A research group from Case

Western developed a new class of chemoresponsive polymers, with a structure that mimicks that of sea cucumber dermis (Figure 3.4). It is composed of sti cellulose brin nanobers encased in a soft, polymeric poly(vinyl acetate) (PVAc) matrix [START_REF] Capadona | Stimuliresponsive polymer nanocomposites inspired by the sea cucumber dermis[END_REF], that acts as a dually-triggered mechanical system. Upon exposure to physiological uids, the water at 37°C both dismantles the ber network and reduces the PVAc glass transition to 20°C, therefore will doubly soften the structure (E~12 MPa) [START_REF] Shanmuganathan | Bio-inspired mechanicallyadaptive nanocomposites derived from cotton cellulose whiskers[END_REF][START_REF] Harris | In vivo deployment of mechanically adaptive nanocomposites for intracortical microelectrodes[END_REF][START_REF] Hess | Development of a stimuli-responsive polymer nanocomposite toward biologically optimized, MEMS-based neural probes[END_REF]. Fittingly, this PVAc nanocomposite has demonstrated great features in terms of mechanical tissue compliance, enhanced implantation and body acceptance.

However, the neural probe itself should be evaluated, and the obvious next step in line would involve the study of neural recordings quality on the long run, as well the mechanics involved on probe stability.

The swelling of the device is such (60-70%, [START_REF] Harris | Mechanically adaptive intracortical implants improve the proximity of neuronal cell bodies[END_REF]) that it raises issues about material degradation, metal delamination and structural changes in vivo on the long term.

Thermoresponsive polymer A thermally reactive copolymer made of monomers methyl acrylate (MA) and isobornyl acrylate (IBoA) crosslinked with poly(ethylene glycol) diacrylate was presented as a highly customizable neural probe substrate, where the ratio of each compound determines the glass transition. The device thus obtained was able to soften from a Young modulus of 700 MPa down to 300 kPa, 24 hours post-insertion [START_REF] Ware | Three-dimensional exible electronics enabled by Shape Memory polymer substrates for responsive neural interfaces[END_REF]. In this structure, the rapid water uptake was a predominant factor in probe softening, which might prove inconvenient for the stability of electrical features. The authors then suggested a thiol-ene/acrylate substrate that showed a drop in modulus from 1 GPa to 18 MPa upon exposure to physiological conditions, and succeeded in recording neural activity from the primary auditory cortex of rats during four weeks [START_REF] Ware | Thiol-ene/acrylate substrates for softening intracortical electrodes[END_REF] (Figure 3.3f ). These encouraging outcomes need to be intensively investigated in order to have more perspective on their suitability for chronic neural implantation.

Bioresorbable polymers

The integration of a sti biodegradable polymer remains the most widespread strategy for exible probe implantation. The polymer acts as sti prop during the implantation procedure, providing enough stiness to overcome the buckling force. When the probe is implanted, the polymer is in contact with physiological uids and starts its resorption. A variety of bioresorbable polymers has been tried as candidates for the enhancement of the probe mechanical strength, as seen in Table 3.1. Many of them are commercially available, such as poly(lactic-co-glycolic acid) (PLGA), Polyethylene Glycol (PEG), collagen, chitosan and maltose/sucrose. They provide easy access for research groups, and benet from certain hindsight considering their use in biomedical applications. The others, like tyrosinederived polycarbonate and silk broin, need to be synthesized, and so some skills in chemistry and materials might be required for appropriate use.

Criteria for polymer-based strategies

In this section, criteria for the selection of a strategy for temporary enhancement through polymeric means will be discussed. We deliberately leave aside the integration of a permanent rigid structure in the probe, as intended exibility is negated. Note that a more detail discussion is available in our review [START_REF] Lecomte | A review on soft polymer-based neural probes and insertion techniques for chronic implantation[END_REF].

Stiness and morphology

The most obvious criterion to compare these strategies is the stiness of the system. If the Young's modulus is high, the device overall dimensions may be decreased, which leads to smaller insertion print and dimpling. It is also a predominant parameter when probe length is increased for deeper implantation.

However, inherent Young's modulus is not the only parameter to take into account. Stiness is classically dened as a ratio between the elastic modulus, the sample dimensions and the material structure.

Therefore, crystalline structure plays a major role, as the overall structure can be weakened, leading to minor buckling force. For instance, PEG and sugar coatings can have a weak crystalline structure which leads to fractures during mechanical insertion [START_REF] Mullarney | The powder ow and compact mechanical properties of sucrose and three high-intensity sweeteners used in chewable tablets[END_REF]. Highly porous chitosan and PLGA can be achieved through freeze-gelation [START_REF] Ho | Preparation of porous scaolds by using freeze-extraction and freeze-gelation methods[END_REF], but then decreased stiness is inevitable.

Sti-to-exible transition and degradation

The fate of bioresorbable polymers involves three main steps. First, the liquid uptake and temperature rising in vivo will induce certain polymer swelling as well as the loss of its mechanical properties.

Then, the degradation occurs, with a loss of global molecular weight due to bond breaking through [START_REF] Kato | Preliminary study of multichannel exible neural probes coated with hybrid biodegradable polymer[END_REF][START_REF] Kim | Sustained release of dexamethasone from hydrophilic matrices using PLGA nanoparticles for neural drug delivery[END_REF] Tyrosine-derived polycarbonate [START_REF] Lo | Coating exible probes with an ultra fast degrading polymer to aid in tissue insertion[END_REF][START_REF] Lewitus | The fate of ultrafast degrading polymeric implants in the brain[END_REF][START_REF] Lewitus | Ultrafast resorbing polymers for use as carriers for cortical neural probes[END_REF] 393 MPa [START_REF] Lewitus | Ultrafast resorbing polymers for use as carriers for cortical neural probes[END_REF] + [START_REF] Bourke | Polymers derived from the amino acid L-tyrosine: Polycarbonates, polyarylates and copolymers with poly(ethylene glycol)[END_REF] At research state -clinical study state for coronary stent application- [START_REF] Lewitus | Ultrafast resorbing polymers for use as carriers for cortical neural probes[END_REF] Hours [START_REF] Lo | Coating exible probes with an ultra fast degrading polymer to aid in tissue insertion[END_REF] Integrable DEX [START_REF] Lewitus | Ultrafast resorbing polymers for use as carriers for cortical neural probes[END_REF] Collagen/Gelatin [START_REF] Lind | Gelatin-embedded electrodes -a novel biocompatible vehicle allowing implantation of highly exible microelectrodes[END_REF][START_REF] Shen | Extracellular matrix-based intracortical microelectrodes: Toward a microfabricated neural interface based on natural materials[END_REF][START_REF] Agorelius | An array of highly exible electrodes with a tailored conguration locked by gelatin during implantation : initial evaluation in cortex cerebri of awake rats[END_REF] 2-200 MPa [START_REF] Grant | Tuning the elastic modulus of hydrated collagen brils[END_REF] ++ [START_REF] Gorgieva | Collagen vs gelatine based biomaterials and their biocompatibility: Review and perspectives[END_REF] Articial skin, heart valves, hemostatic agent, blood vessels prosthesis, nerve repair [START_REF] Gorgieva | Collagen vs gelatine based biomaterials and their biocompatibility: Review and perspectives[END_REF] Minutes to hours (tunable) [START_REF] Paige | Real-time enzymatic biodegradation of collagen brils monitored by atomic force microscopy[END_REF] Haemostatic and antibacterial [START_REF] Lind | Gelatin-embedded electrodes -a novel biocompatible vehicle allowing implantation of highly exible microelectrodes[END_REF] Chitosan [START_REF] Ceyssens | Neural implants containing a resorbable chitosan matrix[END_REF][START_REF] Ceyssens | Resorbable scaold based chronic neural electrode arrays[END_REF][START_REF] Ceyssens | Long term LFP measurements with ultrane neural electrodes embedded in porous resorbable carrier[END_REF] 1.2 GPa [START_REF] Amin | Reinforced materials based on chitosan, TiO2 and Ag composites[END_REF] +++ [START_REF] Kim | Chitosan and its derivatives for tissue engineering applications[END_REF] Skin & bone tissue engineering, nerve & blood vessels regeneration [START_REF] Kim | Chitosan and its derivatives for tissue engineering applications[END_REF] Weeks (tunable) [START_REF] Khor | Implantable applications of chitin and chitosan[END_REF] Haemostatic and antibacterial -integrable BNDF [START_REF] Kim | Chitosan and its derivatives for tissue engineering applications[END_REF][START_REF] Shi | BDNF blended chitosan scaolds for human umbilical cord MSC transplants in traumatic brain injury therapy[END_REF] Maltose/sucrose [START_REF] Singh | Improving mechanical stiness of coated benzocyclobutene (BCB) based neural implant[END_REF][START_REF] Xiang | Ultra-thin exible polyimide neural probe embedded in a dissolvable maltose-coated microneedle[END_REF][START_REF] Hassler | Chronic intracortical implantation of saccharose-coated exible shaft electrodes into the cortex of rats[END_REF][START_REF] Jeon | Partially exible MEMS neural probe composed of polyimide and sucrose gel for reducing brain damage during and after implantation[END_REF] 3-5 GPa [START_REF] Mullarney | The powder ow and compact mechanical properties of sucrose and three high-intensity sweeteners used in chewable tablets[END_REF][START_REF] Jeon | Partially exible MEMS neural probe composed of polyimide and sucrose gel for reducing brain damage during and after implantation[END_REF] + [START_REF] Mullarney | The powder ow and compact mechanical properties of sucrose and three high-intensity sweeteners used in chewable tablets[END_REF] Food sweetener, pharmaceutical coating [START_REF] Mullarney | The powder ow and compact mechanical properties of sucrose and three high-intensity sweeteners used in chewable tablets[END_REF] Seconds to minutes [START_REF] Xiang | Ultra-thin exible polyimide neural probe embedded in a dissolvable maltose-coated microneedle[END_REF] -PEG [START_REF] Barz | Novel technology for the in-plane transfer of multiple interconnection lines in 3D neural probes[END_REF][START_REF] Felix | Removable silicon insertion stieners for neural probes using polyethylene glycol as a biodissolvable adhesive[END_REF][START_REF] Cheng | Three-dimensional exible polyimide based probe array with stiness improvements by using biodegradable polymer[END_REF][START_REF] Takeuchi | Parylene exible neural probes integrated with microuidic channels[END_REF] 0. Minutes [START_REF] Shanmuganathan | Bio-inspired mechanicallyadaptive nanocomposites derived from cotton cellulose whiskers[END_REF] Antioxydant curcumin release [START_REF] Potter | Curcuminreleasing mechanically adaptive intracortical implants improve the proximal neuronal density and blood-brain barrier stability[END_REF] Table 3.1: Comparison table of biodegradable and mechanically-adaptive polymers features reported in the literature as means to stien a exible neural probe enzymatic reactions [START_REF] Horan | In vitro degradation of silk broin[END_REF] or hydrolysis [START_REF] Rao | Polyethylene glycol-containing polyurethane hydrogel coatings for improving the biocompatibility of neural electrodes[END_REF]. Finally, the decomposition products are slowly evacuated by the metabolism [START_REF] Gomes | Biodegradable polymers and composites in biomedical applocations: from catgut to tissue engineering[END_REF]. Therefore, the overall resorption kinetics highly depend on water uptake capability, dimensions, crystallinity, molecular weight and hydrophilic anity.

The loss of mechanical properties is an important drawback in the use of these bioresorbable or smart polymers. The task of surgical implantation is harder, because there is a small time window of probe placement in the cortex. For instance, all fast-resorbing polymers, such as PEG [START_REF] Barz | Novel technology for the in-plane transfer of multiple interconnection lines in 3D neural probes[END_REF], gelatin [START_REF] Paige | Real-time enzymatic biodegradation of collagen brils monitored by atomic force microscopy[END_REF] and sugar-derived polymers [START_REF] Hassler | Chronic intracortical implantation of saccharose-coated exible shaft electrodes into the cortex of rats[END_REF], as well as mechanically-adaptative PVAc nanocomposite [START_REF] Shanmuganathan | Bio-inspired mechanicallyadaptive nanocomposites derived from cotton cellulose whiskers[END_REF], necessarily need to be implanted within minutes, therefore are not suitable for precise, fastidious implantation.

The degradation rate, that is to say the time before all the polymer has gone through enzymatic or non-enzymatic hydrolisis, is often evaluated through in vitro experiments [START_REF] Jin | Waterstable silk lms with reduced β-sheet content[END_REF][START_REF] Lewitus | The fate of ultrafast degrading polymeric implants in the brain[END_REF]. However, in vitro studies do not take into account brain size, uid availability and renewal, and may underestimate the degradation rate. Kozai and coworkers recently observed that the bioresorbable carboxymethylcellulose coating implanted in rat brains degraded in days, when they expected complete dissolution within 20

minutes [START_REF] Kozai | Chronic tissue response to carboxymethyl cellulose based dissolvable insertion needle for ultra-small neural probes[END_REF]. The authors suggest uorescent labeling of their polymer to track its fate.

With regards to the bioresorption itself, a recurring issue aects cytotoxicity of the decomposition products and long-term side eects. The remaining low-molecular weight products of hydrolisis can interact with the surrounding cells, leading to eventual malfunction. It has been noticed that bioderived polymers such as collagen and chitosan tend to induce less inammatory response linked to degradation compounds than synthesized polymers [START_REF] Gomes | Biodegradable polymers and composites in biomedical applocations: from catgut to tissue engineering[END_REF]. However, the fate and underlying mechanisms of the resulting products remain unclear.

Biocompatibility and sterilization

The polymer candidates all have obviously demonstrated non-cytotoxic features. However, special care must be drawn to polymer synthesis and integration, for their biocompatibility might be altered by the dierent process steps. More particularly, solvent residues has shown to create considerable toxic environment [START_REF] Vernekar | SU-8 2000 rendered cytocompatible for neuronal bioMEMS applications[END_REF]. Fortunately, the majority of the considered polymers, like PEG, silk broin, collagen and sugar-derived materials, are obtained through simple evaporation of water solvent [START_REF] Takeuchi | Parylene exible neural probes integrated with microuidic channels[END_REF][START_REF] Rockwood | Materials fabrication from Bombyx mori silk broin[END_REF][START_REF] Gorgieva | Collagen vs gelatine based biomaterials and their biocompatibility: Review and perspectives[END_REF][START_REF] Jeon | Partially exible MEMS neural probe composed of polyimide and sucrose gel for reducing brain damage during and after implantation[END_REF].

As mentioned earlier, commercially available polymers have generally more hindsight as to their biocompatibily, for they have been used for decades in other biomedical applications. PLGA, PEG, chitosan and collagen recount many possible applications in the medical eld, including tissue engineering, drug delivery and grafts [START_REF] Nair | Biodegradable polymers as biomaterials[END_REF][START_REF] Gorgieva | Collagen vs gelatine based biomaterials and their biocompatibility: Review and perspectives[END_REF]. Besides, chitosan and collagen both present with intrinsic haemostatic and antibacterial properties, that helps soothing foreign body response [START_REF] Kim | Chitosan and its derivatives for tissue engineering applications[END_REF][START_REF] Gorgieva | Collagen vs gelatine based biomaterials and their biocompatibility: Review and perspectives[END_REF].

Sterilization also plays a part in device biocompatibility. Sterilization must be performed after fabrication to reduce the risk of infections and possible complications. It can be achieved either through chemical means (ethanol 70%), dry heating (160-190 °C), autoclaving (120-135°C), ethylene oxide gas or UV radiation, in agreement with biomedical device regulation [START_REF] Booth | Sterilization of medical devices[END_REF].

Bioactive function and integration

One of the greatest advantage of bioresorbable polymers as stiening materials is that they are able to soothe foreign body reaction, either through their chemical features or via the integration of bioactive agents. Drugs can be blended with bioresorbable polymers through non-covalent binding, such as adsorption, physical entrapment or electrostatic interactions, and covalent binding [START_REF] Liu | Current sustained delivery strategies for the design of local neurotrophic factors in treatment of neurological disorders[END_REF]. Unfortunately, release kinetics are usually hard to quantify, because a number of parameters needs to be taken into account. The time during which the drug is released and functional from the host polymer depends on binding strength, host degradation rate and diusion mechanisms, among others. For this reason, only a percentage of the total amount of drug implanted is actually bioactive. For instance, silk broin co-dissolved with nerve growth factor (NGF) lowered the availability of the protein to a concentration of 5070% of the actual amount [START_REF] Uebersax | Silk broin matrices for the controlled release of nerve growth factor (NGF)[END_REF].

Conclusion

The insertion of a exible polymer-based probe in the brain is not a trivial issue. Numerous studies have looked upon this issue over the past decade, coming up with inventive ideas over the best way to overcome this problem.

It cannot be stressed enough that the perfect method does not exist, and that an adequate solution might be provided only in accordance with the targeted application. For this reason, implant type and implantation procedure must be well-dened to nd a matching technique that will enable the procedure.

In the following sections, two insertion-based issues are investigated:

the insertion of a single Parylene shank into a mouse hippocampus In this issue, the insertion of the Parylene probes developped in Chapter 2 is discussed, through the study of a bioresorbable polymer integrated to the shanks..

the insertion of multi Parylene shanks into a rat hippocampus Over a second phase, I will discuss the insertion of longer Parylene multi-shanks for implantation into a rat's hippocampus, and more particularly, the eect of the number of shanks over the trauma of insertion. This study was carried out as part of an international mobility program with the University of Southern California, Los Angeles.

Silk as a bioresorbable stiener for a single Parylene shank 3.2.1 Introduction

The two generations of Parylene probes fabricated in Chapter 2 consists of thin, single shank probes that are to be implanted in mice hippocampi. First in vivo procedures showed that these probes were hard to insert without buckling, and that a stiening strategy was required for a proper and reliable insertion.

Originally, an optic ber integrated to the backside of the probe via biocompatible cyanoacrylate glue elicited implantation and electrical recordings from a Parylene-based probe for a pre-trial that occurred at the very beginning of my thesis [START_REF] Castagnola | Parylene-based exible neural probes with PEDOT coated surface for brain stimulation and recording[END_REF] (Figure 3.5). It is obvious that this precarious solution was not meant to last, because then we would loose the advantage of exibility. In this section, we will focus on the insertion of the rst generation of Parylene probes. Indeed, they are longer than the GEN 2 probes (6mm compared to 2mm) and the implantation procedure called for much more accuracy, thus the requirements in terms of mechanical reliability were far harder to meet than for the second generation probes, as we will detail into further detail in the next Chapter.

Therefore, we intend to establish in this section a method for the insertion of a GEN1 probe, a fortiori that works on smaller, simpler GEN2 probes. This insertion method focused on the use of bioresorbable polymers, because they present with many attractive features, as explained in this section. Among these polymers, silk broin (SF) appears as a prominent candidate. SF is among the most biocompatible polymer used for this application (Table 3.1), with extensive use and studies on human body reaction, and happen to recount more studies for cerebral implantation [START_REF] Hopkins | Silk Hydrogels as Soft Substrates for Neural Tissue Engineering[END_REF][START_REF] Kim | Dissolvable lms of silk broin for ultrathin conformal bio-integrated electronics[END_REF][START_REF] Metallo | Silk coating as a novel delivery system and reversible adhesive for stiening and shaping exible probes[END_REF][START_REF] Tien | Silk as a multifunctional biomaterial substrate for reduced glial scarring around brain-penetrating electrodes[END_REF][START_REF] Wu | A exible sh-bone-shaped neural probe strengthened by biodegradable silk coating for enhanced biocompatibility[END_REF]. Bombyx Mori silkworm silk broin has been used for centuries as suture material for its strength and renowned biocompatibility [START_REF] Altman | Silk-based biomaterials[END_REF]. Besides, its tunable degradation rate [START_REF] Jin | Waterstable silk lms with reduced β-sheet content[END_REF] might prove very useful for a wide range of implantation procedure, and for the potential integration of bioactive agents in the SF structure [START_REF] Tien | Silk as a multifunctional biomaterial substrate for reduced glial scarring around brain-penetrating electrodes[END_REF]. Sericiculture, or silk farming, dates back to -2500 BC in China and remains nowadays a major industrial sector in this country. In the medical eld, silk bers have been used for centuries as a resorbable suture material.

Introduction to silk broin (SF)

A brief history of silk

Silk structure

Raw silk is composed of 65% broin and 25% sericin, the rest being water and fat compounds (Figure 3.6c).

-Fibroin: the core of raw silk consists of folded β-sheet polypeptide, piled up through Van der Walls interactions. More precisely they are alternation between amino acids such as glycine (38%), alanine (25%), serine (13%) and tyrosine (12%). The result is a self-assembled hydrophobic protein insuring a strong material [START_REF] Rockwood | Materials fabrication from Bombyx mori silk broin[END_REF].

-Sericin: It is a glycoprotein that surrounds broin bers. It has shown to be responsible for eventual allergic reactions and reduces silk stiness, thus should be extracted from the cocoons when silk is used for biomedical applications [START_REF] Rockwood | Materials fabrication from Bombyx mori silk broin[END_REF][START_REF] Vepari | Silk as a biomaterial[END_REF].

Silk properties

Silk broin has recently gained attention in biomedical elds [START_REF] Altman | Silk-based biomaterials[END_REF] because of its several desirable properties. Among them, we can cite:

-Biocompatibility: it has shown great compatibility for many in vivo applications, including the neural eld.

-Strength: with an intrinsic tensile strength of 740 MPa, B. Mori bers is tougher than commonly used polymeric degradable biomaterials such as collagen and poly(l-lactic acid) (PLA) (respectively 0.9-7.4MPa and 28-50Mpa) [START_REF] Rockwood | Materials fabrication from Bombyx mori silk broin[END_REF]. -Tunable resorption: when submitted to a shear force (annealing, methanol treatment), the β-sheet content of silk broin is reduced and the overall structure becomes less soluble in water. Therefore, depending on the treatment, the resorption of silk broin ranges from minutes to months [START_REF] Jin | Waterstable silk lms with reduced β-sheet content[END_REF].

-Availability and cost: given the large scale of world sericiculture for textile industry, cocoons are easy to obtain and at low cost. However, the broin extraction procedure raises a little the overall expense for biomedical applications [START_REF] Vepari | Silk as a biomaterial[END_REF].

-Materials forms: after broin extraction, the solution can easily be processed into a variety of forms and structures (Figure 3.7), such as gels, lms sponges etc.

[261] -Sterilization compatibility: silk broin is compatible with a number of sterilization processes, which is convenient for in vivo uses [START_REF] Yucel | Silk-based biomaterials for sustained drug delivery[END_REF].

Silk broin extraction

The extraction of silk broin is the rst step to obtain an aqueous broin solution that will form the base of our silk structure. The recipe is almost entirely based on the protocol recommended by Rockwood and coworkers [START_REF] Rockwood | Materials fabrication from Bombyx mori silk broin[END_REF], and is depicted in Figure 3.8. Note that the silk protein is very sensitive to contamination, and during the whole process ultrapure water with controlled resistivity should be used, as well as extra clean glass recipients.

Cocoon preparation First, the cocoons are cut in 6 -8 pieces and the dried worms are disposed.

Cocoons must be intact, with no hole or yellow stains.

Sericin removal The cocoons are then boiled in a solution of 0.02M of N a 2 CO 3 during 30 minutes. The boiling enables ber dismantling from the cocoon, and the sodium carbonate allows for the dissolution of the sericin present. Boiling time has an inuence over chain length and mechanical properties, therefore extra attention should be paid to boil for exactly 30 minutes. After boiling, the degummed bers are rinsed in ultrapure water for 20 minutes, three times, before letting out to dry.

Dissolution The dissolution of degummed bers involve a 9.3M solution of lithium bromide (LiBr).

The volume of LiBr depends on the mass of dried bers, therefore the latter should be particularly dry when weighed. Be careful when manipulating, as lithium bromide and water form a exothermic reaction.

The solution is poured on the silk bers and left to work for 4 hours at 60°C. The volume of LiBr solution is small compared to that of the bers, and occasional stirring might be required.

Purication The resulting silk solution is critically dense and sticky. To perform purication, this solution needs to be dialyzed against ultrapure water for three days. Here, for budgetary reasons, we use a dialysis membrane of 3.5kDa xed with clean string instead of a dialysis cassette, left in ultrapure water during 3 days. The water is changed at 1h, 2h, 3h, 6h then on day 2, in the morning and in the afternoon, and nally on day 3, in the morning then 6 hours later.

Clean and storage

The resulting silk solution is then retrieved from the membrane and centrifuged twice at 4300 rpm to remove small dust and particles. The solution, containing about 7-8% w/v of silk, can nally be stored at 4°C for about one month.

Concentration After all steps of the fabrication, the silk can be concentrated to form dense, concentrated silk broin. To do that, a portion of the solution might be transferred to a humid dialysis membrane and let hang in ambiant air for a certain time, depending on the targeted concentration.

In our case, we try to have a very concentrated silk solution of about 20% w/v (about the maximum percentage before the dense solution forms aggregates and solidies within minutes), and the membrane is let hang for about 24 to 48 hours. Weighing the sample allows the user to track the concentration process.

Silk coating integration

The most common way to obtain a polymer coating has been reported as the dip-coating, where the probe is immersed in a diluted polymer [START_REF] Cheng | Three-dimensional exible polyimide based probe array with stiness improvements by using biodegradable polymer[END_REF][START_REF] Xiang | Ultra-thin exible polyimide neural probe embedded in a dissolvable maltose-coated microneedle[END_REF][START_REF] Altman | Silk-based biomaterials[END_REF]. This method enables the control of the coating thickness by dening the up and down speed, the solution concentration, and the number of coatings. However, it requires the recording areas to be immersed as well, which considerably delays probe adjustment in vivo.

An insulator layer on the electrodes compromises the signal recording, and thus the device positioning.

The use of a polymeric mold is useful, because it allows to dene the patterns of the polymer coating.

Mold fabrication Here, the rigid silk coating is obtained using a Polydimethylsiloxane (PDMS) mold. The mold is fabricated by pouring un-reticulated PDMS onto a rigid micropatterned mold with negative prints and let dried at 60°C overnight. PDMS is often used for the transfer of accurate patterns onto a polymeric substrate (Figure 3.9). First tests allowed to assess that the coating should be in the order of a few hundreds of micrometers thick, so that the negative print mold is fabricated using a thick photoresist, the SU-8 photoresist, patterned via photolithography on a silicon substrate. Multiple PDMS molds can be obtained from the same silicon-SU-8, which is cost-ecient. Besides, the patterns dimensions can simply be adjusted using dierent SU-8 photolithography masks, and SU-8 thickness. Silk coating integration As explained in the previous chapter, Parylene surface tends to be hydrophobic after device fabrication. Therefore, to emprove the adhesion with our aqueous silk broin, the Parylene devices are previously submitted to 15 min UV-ozone treatment.

We want the silk coating to be integrated on the backside of the probe, so that when the probe is implanted, the silk insulating layer does not interfere with electrical recordings. In order to do that, the probes are set face down against the mold and remain so with electrostatic interactions while the polymer is syringe-jetted on the back (Figure 3.10a,b). The positioning and jetting steps are carried out under binocular. At this point, contrary to the Parylene surface, the PDMS mold is highly hydrophobic, and so the silk solution will preferably stay inside the PDMS channels without advocating spills. After a few hours at room temperature, the silk solution dries and the probes can easily be removed from the mold (Figure 3.10c).

Note that aqueous silk broin solution is intrinsically transparent, but for the sake of visualization, a drop of blue food dye was used in the concentrated SF solution. This allowed us to dene more precise coatings, and to keep track of eventual residues throughout the characterization process. Thickness measurements and microscope characterization (Figure 3.12 and 3.11) show that this coating protocol is rather repeatable, as the dierence in thickness between the coating structures are small. For instance, using a 300µm-thick mold with a 20%w/v silk solution, the dried lm is 240 ± 15 µm thick (N coatings, N = 10). . Besides, the thickness can be tuned if needed, adjusting the polymer dilution percentage and the mold dimensions in agreement. As an example, we measured the dimensions of a 13% w/v and 20% w/v silk broin solution cast in a 300 µm and 150 µm thick PDMS mold. Figure 3.11 shows that the mold thickness is a predominant factor in the control of our thickness compared to the silk dilution percentage. Depending on the brain area of interest, the animal implanted, and the precision of surgical positioning, requirements might dier for dimensions of the probe (thickness and length).

For the following mechanical characterization of the tests, we chose to use thick coatings (~250 µm) because the in vivo procedure for GEN1 probes was delicate, and thus great strength over time was needed. However, smaller dimensions can be considered for other uses.

In vitro insertion tests

From a qualitative point of vue, an important data to collect when working with softening, resorbable polymers resides in their ability to be manipulated during the insertion process. Indeed, depending on the surgical procedure, the probe might need to be adjusted for a critical time of several minutes or more, during which it must remain sti. This critical time is linked both to the material degradation rate and to the loss of its mechanical properties post-insertion. For this reason, cycled insertions of SF-coated probes in a brain phantom were performed using a motorized test bench coupled with a digital force gauge in order to measure the force in relation to the position. The instruments were controlled using an internally-developed NI LabVIEW program. Brain phantom was made of agarose gel 0.6%, which best mimic the brain's mechanical features, and each cycle lasts 1 minute. These tests give us an overview of the reliability of our devices for practical surgical handling. Blunt Parylene probes are barely or not at all able to penetrate the gel as the mean compression force in vitro approaches 1.32 mN, when the insertion force is estimated at 1.53 mN [START_REF] Tian | Monitoring insertion force and electrode impedance during implantation of microwire electrodes[END_REF] (Figure 3.13).

On the contrary, silk-coated devices succeed in penetrating the brain phantom without buckling, with a rather low insertion force. With a staple-shaped beam, silk broin benets from an optimized design for an enhanced insertion, therefore decreasing the chance of vessel rupture during probe implantation.

Besides, in terms of mechanical holding,we observed SF-coated probes still penetrate the gel after tens of insertion (12.3 in average, N = 6). This property could be of use for extended surgical procedures requiring fastidious positioning of the probe.

Standard compression tests

Standard compression tests against a hard substrate are performed to assess the mechanical properties of our SF-coated probes. Indeed, as explained in the previous section, axial compression of beams eventually results in its buckling, and the measurement of this buckling force is a good parameter to assess a sample stiness. For that experiment, the same motorized bench with digital force gauge was used, with a hard tile instead of the agarose gel. All compression tests were performed with a slow speed of 20mm/min, because exible electrodes tend to buckle all the more as the insertion is fast [START_REF] Paralikar | Collagenase-aided insertion of intracortical microelectrode arrays: evaluation of insertion force and chronic recording performance[END_REF].

The silk coating improves more than 100-times the probe strength, with a resulting buckling strength of about 270 mN. The strength of a probe is here shown to be linked to its length, the Young's modulus, and the form factor. Standard compression models usually involve many errors when compared to experience. The beam fabrication protocol and their manual clamping to the apparatus lead to a variety of error sources that aect the measured buckling force. However, the compression forces depicted in Figure 3.14 shows rather repeatable experiments in which the standard deviation does not exceed 10% for SF samples; therefore, the tendancies appear reliable.

Young's modulus determination

One way to assess if the standard compression tests described above are meaningful is to determine silk's Young's modulus and compare to the literature. As explained in the rst section of this chapter, Euler's formula links the buckling force of a beam with its dimension and material Young's modulus. where F buckling is the buckling force, E is the beam Young's modulus, L its length (dimension along the z axis) and K the beam eective length factor, whose value depends on the beam xation.

F buckling = π 2 I x E (KL) 2 (3.2)
Moment of inertia The moment of inertia I x along the x axis depends on the shape and dimension of the beam cross-section. In the case of a rectangular beam, it is equal to I x = ab 3 12 , where a is the beam thickness (dimension along the y axis) and b its width (dimension along the x axis, where the buckling occurs).

In the case of a U-beam, the moment of inertia can be determined using parallel axis theorem (Equation 3.3). A U-beam is represented as three rectangles shapes with areas A i where i = 1, 2, 3 (Figure 3.15c).. The moment of inertia is determined using each rectangle shape's moment of inertia I C along an x axis parallel to the centroidal x axis. A i represents the area of each rectangular shape, y i is the mean distance between the rectangles i and the centroid y 0 , and d i is the perpendicular distance between the two axes :

I x = [I Ci + A i .d i ²] (3.3) with d i = ȳi -y 0 and y 0 = A i . ȳi A i
Error modelization Therefore framing of the polymers of Young's moduli values can be deduced from the physical modelization of the beams (plane section and length) and their measured buckling force.

The plane section was considered using a U-shape model (SF), as depicted in Figure 3.15. In this case, the uncertainty in the plane section modeling leads to little variation in the second moment of inertia Ix, thus it was not considered prevailing. Origin 8 software was used to plot the Young's moduli of PEG and SF coating as cast on a PDMS wafer with associated errors. It appears that the predominant parameters come from the error in the measured buckling force (its standard deviation) as well as in the shank length.

Our beam samples are manually clamped to the test bench, which leads to uncertainty considering the beam length L of 6.25 ± 0.25mm. Results Our calculated silk coating's Young's modulus reaches 2.2 ± 0.2GP a (Figure 3.16), a value in direct agreement with the literature; cast lms of silk broin have already been characterized as having a modulus of 1.77 GPa [START_REF] Tien | Silk as a multifunctional biomaterial substrate for reduced glial scarring around brain-penetrating electrodes[END_REF] and 2.8 GPa [START_REF] Jiang | Mechanical properties of robust ultrathin silk broin lms[END_REF]. These mechanical try-outs thus were shown to be highly consistent with literature data, conrming our model's coherence.

This experiment also shows the importance of the probe form factor. With a U-beam, the moment of inertia is increased compared to a rectangular beam, which leads to greater buckling strength. If our silk coating was rectangular, then it would need a plane section surface of 6.3 × 10 -8 m 2 to reach the same buckling force as a U-beam, where the surface is 5.8 × 10 -8 m 2 . This means that at equivalent strength, a U-shaped probe is of smaller dimensions and would induce less insertion print than a rectangular shape.

Dissolution rate measurement

As explained in the previous section, the dissolution rate of a polymer is an important data to collect, so that it is of appropriate time-scale with the application considered.

Regenerated silk broin casted and dried is soluble but can be processed to become water-stable.

Exposure to methanol or to physical shear reduces the β-sheet content in the lms that will allow for a water-insoluble structure [START_REF] Jin | Waterstable silk lms with reduced β-sheet content[END_REF][START_REF] Horan | In vitro degradation of silk broin[END_REF]. This structure can then be degraded with slow enzymatic reactions in vivo.

The proportion of the structure that becomes water-insoluble depends on the strength and length of the treatment. For instance, the time of a water-annealing treatment has already shown to induce tunable degradation rate [START_REF] Jin | Waterstable silk lms with reduced β-sheet content[END_REF]. Indeed, the proportion of the silk that becomes insoluble increases as the length of treatment increases.

To show control of annealing time to obtain silk of dierent dissolvability, enzymatic degradation of the protein was assessed.

Experimental protocol

The dissolution rate of our samples was measured via loss of weight over time, in both a solution containing proteases and a PBS solution. The protease solution represents a model of enzymatic uid in which the brain is commonly surrounded [START_REF] Jin | Waterstable silk lms with reduced β-sheet content[END_REF].

Samples were obtained by casting and drying 200 µl of a 7% w/v silk broin solution in 5 mm Ö 5 mm Ö 2 mm PDMS molds. Initial mass was 15.9 ± 0.9 mg (N = 18). Nine samples were water-annealed for, respectively, 1, 3, and 6 h, and incubated at 37 °C in a 1 ml solution of 1 mg ml=1 protease XIV bent with milli-Q water at pH 7. Nine other samples, water-annealed for 6 h, were immersed in PBS as control. Culture wells were protected from evaporation using paran lms and a water-lled beaker to make the atmosphere saturated in water steam. At days 1, 2, 3, 4, 7, 10, and 15, samples were rinsed and allowed to dry for 1 h at 50 °C. The weighing was performed with a precision balance, before renewing the aqueous media for the next day.

Results

Figure 3.17: Enzymatic degradation in protease XIV (PXIV) solution of 1 h, 3 h, and 6 h water-annealed (wa) silk lms. PBS is used as control. Mass at t = 0 is 15.9 ± 0.9 mg (N = 9/group) In a proteolytic medium, silk broin samples degrade at dierent rates depending on the waterannealing time treatment: 1 h and 3 h treated lms degrade within a few days and up to two weeks for 6 h annealed lms (Figure 3.17). This control of degradation rate is consistent with the literature. Jin and colleagues showed proteolytic degradation in vitro of water-stable silk lms after approximately two weeks [START_REF] Jin | Waterstable silk lms with reduced β-sheet content[END_REF].

Depending on the targeted application, the tunability of silk broin degradation rate might reveal a crucial tool. For instance, when precise surgical positioning is required, it is best if the silk lm conserves its stiness for longer period of time, therefore a longer annealing treatment will be applied.

Besides, silk broin has shown to be compatible with the blending of bioactive agents such as Nerve Growth Factor (NGF) or chondroitinase ABC (ChABC), that helps soothe the immune response to the implanted object [START_REF] Tien | Silk as a multifunctional biomaterial substrate for reduced glial scarring around brain-penetrating electrodes[END_REF][START_REF] Uebersax | Silk broin matrices for the controlled release of nerve growth factor (NGF)[END_REF]. In this case, the bioresorbable polymer acts as a a time-dependent drug delivery system, triggered by its own degradation. Therefore, depending on the drug dispensing method (chronic or spontaneous) and its mode of action, the tunable degradation rate of silk broin might be useful.

Comparison with PolyEthylene Glycol

PolyEthylene Glycol (PEG) is a resorbable polyether that has proven its remarkable features in the medical eld. Recently, this polymer was involved in several implantable devices for neural engineering [START_REF] Chen | A three-dimensional exible microprobe array for neural recording assembled through electrostatic actuation[END_REF], either as a temporary fast-degrading stiener [START_REF] Felix | Removable silicon insertion stieners for neural probes using polyethylene glycol as a biodissolvable adhesive[END_REF][START_REF] Takeuchi | Parylene exible neural probes integrated with microuidic channels[END_REF] or as a biodegradable glue for a removable prop [START_REF] Hosseini | Comparative study on the insertion behavior of cerebral microprobes[END_REF][START_REF] Felix | Removable silicon insertion stieners for neural probes using polyethylene glycol as a biodissolvable adhesive[END_REF][START_REF] Seymour | Fabrication of polymer neural probes with sub-cellular features for reduced tissue encapsulation[END_REF]. So far, it was used in the majority of strategies developped for the temporary stiening of exible neural probes. We thought that a comparison with the silk method would be relevant.

In order to do that, the same experiments as described in this section were carried out with PolyEthylene Glycol 100, 000g.mol -1 , diluted at 30 in DI water and mixed with a red dye. Other molecular weights were tried (3000, 6000, 8000, and 10, 000g.mol-1) but PEG 100, 000 was chosen because it was less brittle The dierent trials show clear superiority of silk in terms of mechanical strength. PEG has a very crystalline structure, which makes it more likely to break under pressure. Therefore, the eective Young's modulus and buckling force are considerably inferior to that of silk broin structure.

Besides, PEG tends to dissolve very quickly in aqueous solution, hence the mechanical holding is also very low: the PEG structure allows for implantation of the probe in the agarose gel, but starts dissolving within seconds, and the probe cannot be withstand multiple insertions in the gel. However, it is a very cheap, commercially-available material, and can be useful when unprecise insertion of short probes are demanded.

Conclusion

Silk broin has proven to present with very attractive features for the temporary stiening of our single shank Parylene probe. When deposed on the backside of the probe, it confers an enhanced structure upon the Parylene device to facilitate insertion with optimized cross-section, which should minimize the acute response linked to the trauma of insertion. Its high buckling force combined with its tensile strength in vitro and in vivo makes silk broin a reliable stiener that shows no alteration of the electrode electrical properties. Given the exceptional mechanical strength of the silk coating, a diminution in the lm thickness could reduce the implantation print while maintaining sucient stiness for the probe to withstand cortex penetration. Finally, its degradation rate can be tuned to either match implantation specications or to promote neural outgrowth after implantation. Indeed, silk broin can be blended with a variety of drugs that showed to alleviate the foreign body response [START_REF] Tien | Silk as a multifunctional biomaterial substrate for reduced glial scarring around brain-penetrating electrodes[END_REF][START_REF] Uebersax | Silk broin matrices for the controlled release of nerve growth factor (NGF)[END_REF]. probes were designed to be 5.5 mm in height, to allow for free movement and laxity of the probe between its attachment to brain tissue and to the skull.

Each group of electrodes on a single probe was spaced at a depth appropriate to reach CA1, DG, and/or CA3.

Fabrication

The fabrication of these devices follows a protocol that has been used by Dr Meng's lab for a few years, and has been detailed in numerous articles [START_REF] Kim | 3D Parylene sheath neural probe for chronic recordings[END_REF][START_REF] Lee | Matrigel coatings for Parylene sheath neural probes[END_REF][START_REF] Kuo | Novel exible Parylene neural probe with 3D sheath structure for enhancing tissue integration[END_REF]. Briey, a 10 μm layer of vapor deposited Parylene will form the base of each shank. Platinum electrode recording sites, traces, and contact pads will be lithographically patterned on top of the base layer using e-beam or sputter deposition at a thickness of 200nm, followed by lift-o. The metal layer is then encapsulated by a 10 μm thick CVD deposited Parylene insulation layer. Electrodes and contact pads will be subsequently exposed by DRIE and the probes will be cut out from the substrate. Devices are released from the silicon carrier wafer with an acetone swab.

During my time at USC, I was able to follow some fabrication steps and get valuable feedback for our own process. Indeed, the fabrication for their devices and for ours is very similar and sharing our respective experience has enabled both parties to widen our knowledge and perspectives on these issues (see Appendix A).

Insertion strategy developed for USC hippocampal prosthesis

As explained before, the buckling force of a system is given by Euler formula :

F buckling = π 2 I x E (KL) 2 (3.4)
In the case of the USC implants, if we represent each shank by a 100 µm wide, 20 µm thick, 5.5 mm long Parylene shank (Young's modulus 3.6 GPa), the buckling force would be around 0.37 mN, which is not enough to overcome the insertion force at ~1.5mN.

The probes need to be stier to be inserted. The dimensions at play are far more challenging than the probe fabricated at LAAS. Once again, the targeted application guide the requirements:

-As the number of shanks and electrodes increases, the targeted in vivo region does not need to be as precise as for the single shank with four electrode developped at LAAS. Therefore, the mechanical holding does not need to be as strong or long-lasting.

-The thin width of the shanks prevents them from being properly backed with eight thin individual layers of bioresorbable coating, without bringing displacement and trauma-inducing issues.

For these reasons, the Meng's lab has developed another strategy based on the reduction of shanks eective length.

Indeed, a reduction in the length L of these probes would drastically increase the buckling force, as seen in Equation 3.4. To that end, a bioresorbable brace reduces the probe length by half: in this case, the buckling force would theoretically be four times higher, that is approximately 1.5 mN, which would suce for insertion in the brain. After insertion of the rst half, the polymer dissolves with the addition of saline. Polyethylene glycol (PEG) 3,350 g/mol (melting temperature 54-58°C) was chosen as the bioresorbable polymer, as it is biocompatible and dissolves within a few minutes [START_REF] Felix | Removable silicon insertion stieners for neural probes using polyethylene glycol as a biodissolvable adhesive[END_REF][START_REF] Nie | Local modication of elastic properties of polystyrene-polyethyleneoxide blend surface[END_REF][START_REF] Takeuchi | Parylene exible neural probes integrated with microuidic channels[END_REF].

The PEG brace is integrated to the device using a PDMS mold, as seen on Figure 3

Insertion mechanisms of arrays of exible neural probes 3.3.3.1 Motivation

The insertion mechanisms of traditional, silicon-based single neural probe or tungsten wires have been the object of many studies [START_REF] Tian | Monitoring insertion force and electrode impedance during implantation of microwire electrodes[END_REF][START_REF] Jensen | Assessment of subdural insertion force of single-tine microelectrodes in rat cerebral cortex[END_REF][START_REF] Molloy | Experimental determination of the force required for insertion of a thermoseed into deep brain tissues[END_REF][START_REF] Howard | Measurement of the force required to move a neurosurgical probe through in vivo human brain tissue[END_REF].

It is generally admitted that a ne tip decreases tissue dimpling and vessel dragging [START_REF] Bjornsson | Eects of insertion conditions on tissue strain and vascular damage during neuroprosthetic device insertion[END_REF][START_REF] Edell | Factors inuencing the biocompatibility of insertable silicon microshafts in cerebral cortex[END_REF], and that probe dimension plays a major role in the reduction of shear force and inammatory response [START_REF] Stice | Thin microelectrodes reduce GFAP expression in the implant site in rodent somatosensory cortex[END_REF][START_REF] Skousen | A strategy to passively reduce neuroinammation surrounding devices implanted chronically in brain tissue by manipulating device surface permeability[END_REF][START_REF] Thelin | Implant size and xation mode strongly inuence tissue reactions in the CNS[END_REF].

However, insertion speed does not benet from such consensus, as seen on Table 3.3. On the one hand, a fast insertion is meant to compensate for poor tip design, and will transect membranes and vessels instead of dragging them, while on the other hand a slow insertion will allow for tissue relaxation and accomodation of compression forces, with minimal shock [START_REF] Welkenhuysen | Eect of insertion speed on tissue response and insertion mechanics of a chronically implanted silicon-based neural probe[END_REF][START_REF] Edell | Factors inuencing the biocompatibility of insertable silicon microshafts in cerebral cortex[END_REF]. Hosseini [START_REF] Hosseini | Comparative study on the insertion behavior of cerebral microprobes[END_REF] 85-160 1670 targeted) or setup (manual or pneumatic insertion tool) might not leave room for much of a choice. For instance, thin, compliant, or exible probes cannot be inserted at high speed, for they are fragile and will buckle or break when the insertion is too fast [START_REF] Paralikar | Collagenase-aided insertion of intracortical microelectrode arrays: evaluation of insertion force and chronic recording performance[END_REF]. Most of these studies focused on the insertion of a silicon shank or tungsten wire, and a lack of knowledge is to be regretted concerning the mechanical insertion of exible neural probes. This new generation of compliant, polymer-based recording devices has expanded for the last decade, for they represent a worthy alternative to traditional hard devices.

Reference

Hosseini and coworkers [START_REF] Hosseini | Comparative study on the insertion behavior of cerebral microprobes[END_REF] published in 2007 a small study on the dierent insertion, retraction forces and dimpling of silicon, glass, tungsten wire and polyimide probe. Their study showed that the measured insertion force varied depending on the probe used, but remained stable with the dierent speed used.

They also presented rst results on the inuence of the number of shanks in the case of 2D arrays (also known as combs). However, in their study, the probes diered signicantlly in terms of width, thickness, length and tip geometry, and so the force measurements are dicult to compare and interpret.

To my knowledge, no study has yet focused on the increase of the number of shanks. The insertion speed was set as 100µm/sec, in the medium speed range (Table 3.3). The eect of speed was not taken into account in this study. Insertion depths were set to 2.5mm. The 2D planar sham probes are taped to an acrylic holder so that the eective length corresponds to that of probes with a PEG brace, that is around 2.8 mm. Initially at 8, the number of shanks is reduced by cutting them two by two and proceeding the same insertion experiment in agarose gel (Figure 3.24).

Sham probes preparation

The 3D matrix array of probes consist of 2D planar arrays with 8 shanks stacked together to form a 3D structure. The spacing between the 2D probes is made of a plastic polyether ether ketone (PEEK) layer sandwiched between two pieces of double sided tape. The overall spacing items are 250µm-thick, the same pitch as between the probes of a same 2D structure (Figure 3.25).

Data treatment and interpretation

Once the mechanical test is over, the data are treated using Origin, and certain parameters can be deduced from the standard compression/force graph:

insertion force: it corresponds to the maximum force experienced by the sample before it starts penetrating the gel.

dimpling: it is the maximum compression the surface of the gel undergoes before the sample penetrates inside of it shear force: it corresponds to the force that the gel experiences as the probe makes its path, and can be read as the maximum force experienced before relaxation (when the compression is over)

Insertion of 2D, planar arrays of exible probes

Dimpling

The dimpling is included between 180 and 240 µm for the ve samples (Figure 3.26b). It increases slightly with the number of probes, which means that the brain would undergo a little more compression before actual probe insertion, but not in a signicant manner.

Insertion force

The insertion force increases linearly and considerably more as a function of the number of probes (Figure 3.26c). As the probes touch the surface of the gel, each of them will apply theoretically the same local pressure on it. But if some of the shanks are not perfectly straight, or if the gel surface is not smooth, this force might be altered.

Shear force

As seen on Figure 3.26d, the shear force, on the other hand, expectedly appears to increase proportionnally to the number of probes. Indeed, the shear force comes from the addition of matter in the tissue:

as they make their way into the tissues, the shanks will bring more and more compression forces. This means that doubling the number of shanks will double the surface in contact with the gel that brings these compression forces, thus the resulting force will be twice as high.

We may now wonder if these insertion mechanisms linked to an increased the number of probes are dierent when the system is in three dimensions.

Insertion of 3D, matrix arrays of exible probes

In order to give an overview, Figure 3.27a shows a standard compression/force graph obtained with three 3D samples. The overall force is once again as expected higher as the number of shank stackings increases.

The parameters (dimpling, insertion force and shear force) for ve trials carried out on these 3D samples, were extracted and integrated in the results of the 2D experiments, as depicted on Figure 3.27b-d.

Dimpling

Dimpling might prove an issue when many probes are inserted. Indeed, when the pitch between arrays of probes is too small, the gel experiences a general distortion that results in high compression and dimpling:

all happens as if the accumulation of probes close to each other act as a unique, bigger device. In our case, the dimpling as seen on Figure 3.27b still increases lightly with the number of shanks : the dimpling value doubles as the number of shanks is eight times higher (0.201 mm for 4 shanks, 0.388 mm for 32 shanks). No apparent dierentiation is noticed between the combs-like and the brush-like sham probes. This all suggests that the pitch chosen here (250 µm) is high enough so that the gel has freedom to relax in between the probes.

Insertion force

The insertion force still increases linearly, as previously noticed for planar probes (Figure 3.27c). Its value is four times higher when the number of probes is multiplied by seven (0.916 mN for 4 shanks, compared to 3.897 for 28 shanks). The curve suggests no maximum insertion force being reached as the number of shanks increases, but other experiments involving more shanks should be performed to verify this. Once again, no signicant dierence is noticed between the insertion of plana or 3D probes.

Shear force

As seen on Figure 3.27d, the shear force still clearly increases proportionally to the number of probes, wether the devices are planar or matrix-like. This again is to be expected, as the shear force should only depend on the volume of material compressing the tissues along the insertion track.

Conclusion

Sham probes made of Parylene and resembling combs and brush-like traditional designs were successfully manufactured and tested to assess their insertion mechanisms. In light of the rst outcomes, the increase in the number of shanks contributes to a lesser extent to the compression that the tissue surface undergoes before probe penetration (also called dimpling). This result may be promising for the design of future matrix implants that do not induce an extended insertion trauma. The insertion force, on the other, increases linearly with the number of probes, but at a rate smaller than the shear force. The latter was demonstrated to be proportional to the number of shanks, as the addition of matter intrinsically increases the compression of the tissues along the probes track. Overall, the design of the multi-shank devices (planar probes and 3D matrix arrays) did not alter signicantly the insertion mechanisms.

However, attention is required over the limits of this experiment. First of all, shank design, pitch and insertion speed were not evaluated in this trial, and extra care should be taken when extrapolating these results to another setup conguration. For instance, the shear force is said to be reduced when the insertion speed is lower, because the tissues can accomodate more easily along the insertion track [START_REF] Andrei | A response surface model predicting the in vivo insertion behavior of micromachined neural implants[END_REF].

Besides, shank spacing (or pitch) should be properly chosen to allow for such relaxation.

Finally, these experiments were carried out with an in vitro model, where tissue homogeneity, age and specy dierences do not bring extra experiment errors and deviation.

Conclusion

In this chapter, the insertion of exible neural implants was discussed.

From a theoretical point of view, the surgical and mechanical requirements for inserting a probe into the brain were shown. Flexible probes are typically 100 to 500 µm wide and 10 to 30 µm thick, with a length varying from 1 to 6 mm in average [START_REF] Xiang | Ultra-thin exible polyimide neural probe embedded in a dissolvable maltose-coated microneedle[END_REF][START_REF] Chen | A three-dimensional exible microprobe array for neural recording assembled through electrostatic actuation[END_REF][START_REF] Mercanzini | Demonstration of cortical recording using novel exible polymer neural probes[END_REF][START_REF] Castagnola | Parylene-based exible neural probes with PEDOT coated surface for brain stimulation and recording[END_REF]. Combined with a low Young's modulus, these dimensions cause for a buckling force in the order of magnitude of the theoretical brain insertion force, therefore certain dimpling or breaking is to be expected. To overcome this issue, many strategies have developed, such as the integration of biodegradable or smart polymers that are sti enough for insertion, but will resorb/soften in the tissues. Bioresorbable polymers have become most popular in this area, because they are easy to use and can be blended with potential neuron-promoting factors to help even more implant acceptance.

In the case of a single shank Parylene probe such as the one developed in chapter 2 of this thesis, the integration of a silk broin coating was demonstrated as a mechanical-enhancing polymer. The elaborated coating method is repeatable, conned and stress-reducing, and was successfully demonstrated in rst in vitro and in vivo trials.

The insertion mechanisms of multi-shank arrays was also appraised. When the number of shanks is increased, the compression of tissue surface is not signicantly altered, whereas the shear force along the shank track increases linearlly with the number of shanks. This preliminary in vitro study is a rst step towards stress-reducing insertion of 2D and 3D exible arrays.

Chapter 4

In vitro & in vivo results towards chronic implantation

In this Chapter, a rst discussion on Parylene devices lifetime is given. In other terms, we intend to show preliminary evidence that the Parylene neuroprosthesis developed in this thesis is functional for actual in vivo purposes, and give primary results of its biostability in both articial and actual cerebral environment.

As an introduction, a denition of biostability and lifetime is discussed. Particular emphasis is put on the understanding of known failure modes involving material degradation for intracortical probes, especially in the case of polymeric implants such as Parylene implants, where swelling and delamination has often shown to advocate device failure over time.

First, the results of an acute recording trial in a mouse brain is given. This rst trial allows us to present and validate the functionality of the Parylene probe, the use of silk coating and the electrical improvements made by the PEDOT nanostructuration on the gold electrodes.

The question of long-term functionality is then approached via both in vitro and in vivo experiments. During six months, some Parylene implants are immersed in physiological-like environment, while others are implanted in the brains of mice. The monitoring of their electrical and mechanical properties gives us a rst hindsight on the performance of our devices in the long run.

All in vivo experiments are carried out in the CRCA (Research Center of Animal Cognition) lab, with the collaboration of Dr Lionel Dahan's group from the REMEMBeR team (REvealing MEmory Mechanisms of the BRain).

Lifetime issues and biostability

Biostability and failure modes of neural interfaces

The body can be a challenging environment for many types of materials and devices in contact with tissues or body uids in the long term. This environment produces modications in the chemical, physical, mechanical and structural properties of biomaterials. The biostability of a device refers to its ability to fulll its intended function in a specic physiological environment as a function of time. The mechanisms leading to changes in the properties of biomedical polymers are considered as part of a biodegradation aecting its biostability. In the case of chronically-implanted neuroprosthesis, the implant must be able to record and/or stimulate neural tissues over long periods of time without adversely aecting body uids, tissues and organs [START_REF] Anderson | Biostability of biomedical polymers[END_REF].

Implant unstability stems either from the host response, that is to say the local and system response of the living systems to the material, or from the material response to said living systems [START_REF] Black | Biological Performance of Materials: Fundamentals of Biocompatibility[END_REF]. In the case of a neuroprosthetic device, as stated by Jor et al [6], a number of failure modes likely inuence chronic recording stability and quality (Figure 4.1).

Host response

Regrettably, the neuro-inammatory response has shown to be predominant in the failure of chronic neural recordings [START_REF] Polikov | Response of brain tissue to chronically implanted neural electrodes[END_REF]. This section has been previously examined in Chapter 1. Briey, the biological environment that is the brain reacts to both implant mechanical and biochemical properties. The formation of an extended inammatory response, called glial scar, is said to cause device encapsulation and migration, which results in poor implant biostability and unstable recording quality [START_REF] Sommakia | Materials approaches for modulating neural tissue responses to implanted microelectrodes through mechanical and biochemical means[END_REF]. Therefore a reduction in inammatory response has shown to induce greater recording quality and longevity [START_REF] Rennaker | A comparison of chronic multi-channel cortical implantation techniques: manual versus mechanical insertion[END_REF]. The design of a exible, polymer-based neural probe with a high degree of biocompatibility developped in this thesis is meant to alleviate said host response to ensure greater implant biostability.

Material response

Material response to a biological environment presents with many failure modes. For intracortical devices, these failure modes include : [6]

Direct mechanical damage of insertion

Implantation mechanisms can severe the device regardless of its nature. A study by Ward and coworkers showed that seven of nineteen commercially-available intracortical devices suered from mechanical failure upon insertion [START_REF] Ward | Toward a comparison of microelectrodes for acute and chronic recordings[END_REF]. Microshanks made of brittle silicon can break upon insertion in the cortex, glass pipette tips are also quite fragile, and exible microprobes can experience compression forces that might generate connectic breakage or faulty contact. More generally, surgical handling of small, fragile devices with unreliable packaging leads to a certain percentage of electrode failure after surgical implantation [START_REF] Prasad | Comprehensive characterization and failure modes of tungsten microwire arrays in chronic neural implants[END_REF][START_REF] Prasad | Abiotic-biotic characterization of Pt/Ir microelectrode arrays in chronic implants[END_REF].

Corrosion of electrode sites

While the phenomenon is predominant for stimulating electrodes compared to recording electrodes, metal corrosion is a commonly attributed issue in recording stability. Corrosion of a metal stems from the presence of oxidative species and acidic pH in the physiological environment. The release of ions can aect tissue response some materials likely experience faradaic charge transfer and corrosion over time [START_REF] Merrill | Electrical stimulation of excitable tissue: Design of ecacious and safe protocols[END_REF]. Corrosion rate, however, is hard to quantify, because the phenomenon tends to be uniform on the electrode surface, therefore it might be noticed only when signicant amounts of metal are lost [START_REF] Black | Biological Performance of Materials: Fundamentals of Biocompatibility[END_REF].

Besides, corrosion products might be cytotoxic for the environment, and should be carefully examined. If in absolute, no material is utterly stainless, platinum and gold are privileged material due to their noble properties. Biofouling is a term used to dene the adhesion of proteins and other biological matter on a surface in contact with biological uids [START_REF] Bixler | Biofouling: lessons from nature[END_REF] (Figure 4.2a). More particularly, surface biofouling due to the adhesion of microbial or thrombotic agents is a consequence of the foreign body response [START_REF] Harding | Combating medical device fouling[END_REF]. The aggregation of biological components on the recording or stimulating site alters its sensitivity. Surface morphology, wettability and chemical properties are ways to prevent unwanted cell adhesion on the sensor.

Membrane biofouling

For instance, increasing surface hydrophobicity is an ecient way to obtain a smooth, clean surface even after long periods of implantation [START_REF] Wisniewski | Characterization of implantable biosensor membrane biofouling[END_REF][START_REF] Wisniewski | Methods for reducing biosensor membrane biofouling[END_REF][START_REF] Harding | Combating medical device fouling[END_REF].

Degradation of passivation layers and insulation coatings

The passivation layers and insulation coatings commonly incorporated in microelectrodes may degrade over time. This phenomenon is exacerbated as passivation layers are often composed of polymeric chains, that tend to swell and lift in physiological conditions [START_REF] Prasad | Comprehensive characterization and failure modes of tungsten microwire arrays in chronic neural implants[END_REF]. Whenever a polymer is used as a passivation layer, the coating must be awless, void free, and stable with good adhesion that is impermeable for ions [START_REF] Von Metzen | The eects of annealing on mechanical, chemical, and physical properties and structural stability of Parylene C[END_REF].

The polymer-metal and polymer-polymer adhesion strength is essential to limit layer-by-layer delamination. Besides, the choice of the polymer is crucial to limit water uptake and swelling, that increases delamination eects (Figure 4.2b). Parylene C is said to be chemically inert, with a good water barrier resistance, therefore is a good candidate as an insulating layer [START_REF] Hassler | Characterization of parylene C as an encapsulation material for implanted neural prostheses[END_REF][START_REF] Hsu | Characterization of Parylene-C lm as an encapsulation material for neural interface devices[END_REF].

Lifetime of a Parylene neuroprosthesis

Many in vitro studies focused on the performance of Parylene C as an encapsulation layer for neuroprosthetic devices. These studies usually include soaking tests, accelerated ageing and/or peeling force measurement to assess adhesion-promoting strategies of Parylene C to a substrate, or its water-barrier capability [START_REF] Hassler | Characterization of parylene C as an encapsulation material for implanted neural prostheses[END_REF][START_REF] Minnikanti | Lifetime assessment of atomic-layer-deposited Al2O3-Parylene C bilayer coating for neural interfaces using accelerated age testing and electrochemical characterization[END_REF][START_REF] Seymour | The insulation performance of reactive parylene lms in implantable electronic devices[END_REF][START_REF] Lo | Characterization of parylene as a water barrier via buried-in pentacene moisture sensors for soaking tests[END_REF]. Though valuable to improve Parylene hermetic properties towards a reliable neural interface, very little studies included the evaluation of Parylene as implant substrate, and not just insulating layer, for chronic in vivo purposes, or the appraisal of Parylene-based devices lifetime in the brain. In fact, few studies relate the implantation and monitoring of brain activity over long period of time. Some majority involve short-term implantation, between 3 weeks [START_REF] Agorelius | An array of highly exible electrodes with a tailored conguration locked by gelatin during implantation : initial evaluation in cortex cerebri of awake rats[END_REF] and 5 weeks [START_REF] Wu | A multi-shank silk-backed parylene neural probe for reliable chronic recording[END_REF][START_REF] Kim | 3D Parylene sheath neural probe for chronic recordings[END_REF].

Recently, a study by Ellis Meng's team at USC showed proof of electrical recordings in three rats for a period of 3 months, and primary investigation over the signal degradation via Signal-to-Noise Ratio (SNR) measurement and Electrochemical Impedance Spectroscopy (EIS) [START_REF] Lee | Matrigel coatings for Parylene sheath neural probes[END_REF].

Overall, if Parylene C is largely being investigated as a substrate for chronically-implanted neural probes, the stability of Parylene-based devices lacks crucial perspective. In order to achieve such hindsight, in vitro appropriate models and in vivo accurate experiments need to be intensely pursued.

In the following sections, we intend to demonstrate preliminary evidence of the biostability and lifetime of Parylene-based neural probe.

Acute recording in mouse hippocampus

First trial goal

The rst step is to show that the probe developped in the previous chapter is indeed functional for neural recordings in vivo. More particularly, the following characteristics must be appraised :

-the silk coating's ability to support the probe during penetration of the brain -the ability of our Parylene probe to record both Local Field Potentials (LFP) and single unit activity (spikes) with adequate SNR -the comparison between PEDOT-coated and pristine electrodes in terms of signal quality For these purposes, a preliminary trial allowed us to perform an acute in vivo study in the hippocampus of a single mouse. The hippocampus is part of the limbic system, located in the medial temporal lobe (Figure 4.3).

Acute implantation in the hippocampus

The hippocampal formation is composed of the hippocampus per se, also called Cornu Ammonis (CA, literally the egyptian god Ammon's horn) and the Dentate Gyrus (DG). It plays major role in long-term memory, as well as spatial representation and navigation.

The areas CA1 and CA3 are mainly composed of pyramidal neurons, with the somas aligned to form a dense, compact layer of cell bodies, and with the dendrites tangent to this curved layer (Figure 4.3). Schaer collaterals are axon collaterals given o by CA3 pyramidal cells and projecting to area CA1, where they make synapses on the dendrites of CA1 pyramidal neurons. Therefore, stimulating articially these axons allows to trigger simultaneously a large set of synapses. Schaer collaterals are essential in activity-dependent synaptic plasticity and information processing in the hippocampus. In other terms, they play a substantial role in the coding of long-term memory, and are highly investigated in research involving learning mecanisms and potential deciency, such as Alzheimer's disease and dementia [START_REF] O'keefe | The hippocampus as a cognitive map[END_REF][START_REF] Sokolowski | Wired for behaviors: from development to function of innate limbic system circuitry[END_REF]. A stimulation electrode is inserted in the CA3 region and delivers pulses that will trigger the schaer collateral pathway. The dendrite of stimulated neurons project to the CA1 pyramidal neurons where they make synapses. The implantation of our GEN1 recording electrode in the CA1 region will then be able to detect evoked post-synaptic potentials issues from the transmission of signal between the CA3 and CA1 region (adapted from [START_REF] O'keefe | The hippocampus as a cognitive map[END_REF])

Experimental setup

The procedure is detailed in Lecomte et al 2015 [START_REF] Lecomte | Silk and PEG as means to stien a parylene probe for insertion in the brain: toward a double time-scale tool for local drug delivery[END_REF]. A 14 weeks-old male C57BL/6 J mouse is anaesthetized by an intra-peritoneal injection of ketamine/xylasine and set to a stereotaxic apparatus.

Lidocaine is used as a local anesthetic. After skin incision, two craniotomies are performed, one for the stimulation electrode and one for the recording probe. The dura mater of the meninges is incised to allow access to the pia mater. The implantation procedure is delicate and time-consuming, because both the stimulating electrode and the recording probe need to be adjusted in the brain to allow proper propagation of the evoked signal. This means that the Parylene implant should be sti enough to allow reliable manipulation throughout the procedure. To achieve such stiness, the probe is back-coated with a thick (250µm) silk structure, and the device water-annealed during 6 hours in a dessicator, so that the silk does not dissolve during implantation procedure.

Stimulation electrode and reference electrode

Results

As explained, the Parylene probe is composed of pairs of pristine gold and PEDOT-modied microelectrodes placed at three dierent depths across area CA1. This allows us to obtain a direct comparison between the modied and non-modied microelectrode, because the recordings are made at the same depth and therefore are from the same population of neurons.

First of all, the silk-backed Parylene probe was able to withstand insertion in the cortex without buckling, and remained sti enough until the proper positioning of both the stimulation and the recording probe, which lasted a couple of hours.

In terms of electrical capability, the spontaneous activity of non-stimulated single neurons (spikes) was recorded as well as evoked local eld potentials (LFP) (Figure 4.7). Without the PEDOT nanostructuration, both spontaneous and evoked signal are mainly composed of 50 Hz electromagnetic noise, which is most likely a consequence of the small diameter of the electrode active area (40 µm), resulting in its high impedance (630 kΩ at 1 kHz, see Figure 4.5). The useful signal is caught up in high noise and cannot be extracted. However, the PEDOT nanostructuration reduces the impedance by a factor of almost 100 for the same diameter, with an impedance of 7 kΩ at 1 kHz (Figure 4.5). As a result, this electrode shows spontaneous single unit activity exhibiting bursts of action potentials, depicted in Figure 4.7. Moreover, the stimulated LFPs in CA1 region after a stimulation of 0.1, 0.05, and even 0.025 mA in the CA3 are of high quality. The signal shows a clear upward deection distinct from the stimulation artifact. This deection is an evoked post-synaptic eld potential, reecting the glutamatergic transmission between CA3 and CA1. Consequently, this means that the electrical improvements from the PEDOT nanostructuration compared to gold electrodes were not altered by the presence of silk broin, demonstrating that the silk layer is indeed conned on the probe backside. In this trial, we demonstrate the feasability of neural recording from the Parylene probe conceived and fabricated during this PhD. The enhancement of both electrical properties via PEDOT nanostructuration and mechanical properties via the silk coating is once again shown. The question of biostability is yet to be addressed, that can only be achieved through long-term experiments in vitro and in vivo.

In vitro ageing of Parylene C implants

The next step in line to evaluate if our Parylene probes are suitable candidates for chronic implantation is to see if they withstand prolonged contact with physiological uids. In this section, we will detail the procedure and results of a 6-months soaking test. Usually, the liquid consists of PBS or NaCl at 0.9% w/v, that mimics general physiological conditions. Accelerated ageing tests, where the medium is brought to 60°C and/or with ultrasonic stress, is a way to perform these tests [START_REF] Minnikanti | Lifetime assessment of atomic-layer-deposited Al2O3-Parylene C bilayer coating for neural interfaces using accelerated age testing and electrochemical characterization[END_REF][START_REF] Hsu | Characterization of Parylene-C lm as an encapsulation material for neural interface devices[END_REF].

Soaking test procedure

Here, we chose a more elaborate medium, the Articial Cerebral Spinal Fluid (ACSF), that resembles ionic concentrations and pH in brain environment (Table 4.1), and we considered a more biomimicking approach for the soaking conditions, so that all samples were soaked at 37°C in an oven during the length of the trial.

Four GEN2 Parylene implants with each four PEDOTmodied gold electrodes were soaked in ACSF for a period of over six months. Each device is bonded with anisotropic conductive lm to a long exible cable, and the connectic part is glued with biocompatible glue EP630 for isolation. The devices are soaked into the same volume of ACSF (25mL) in plastic cups, and the lid is sealed to prevent water from evaporating (Figure 4.8). Every week or few weeks, the samples are retrieved from the oven and let cool at room temperature until they reach 20°C. Electrochemical Impedance Spectroscopy (EIS, 10Hz-10kHz) is then performed at room temperature on all those electrodes. Since ACSF is basically a standard NaCl 0.9% w/v solution (among with other ions), the EIS is carried out directly in the soaking medium, with a platinum wire as counter electrode. Preliminary results showed no signicant dierence in EIS measurement in NaCl or in ACSF solutions. Throughout the 29-weeks study, the impedance of the 16 PEDOT-modied electrodes remained considerably stable. As a matter of fact, the impedance of the electrodes on 3 out of 4 soaked implants were either at stable value, that is to say within 10% of their initial value, or even decreased. A representative example of this slight decrease is represented on Figure 4.9b. This minor impedance decline might come from a slight change in the composition of the ACSF medium, but a renewal of the medium at around week 19 did not seem to alter impedance measurement. It can also be interpreted as PEDOT relaxation and stress relief as it is soaked for long periods of time, which could present with minor surface ratio rise (swelling) and decreased impedance. Finally, it could come as a consequence of the beginning of delamination of Parylene around the electrode, resulting in a subtle increase of electrode diameter thus lower impedance. A closer look at the morphology of the device allows us to better understand this phenomenon.

Morphology alteration

Chronic recording in mouse hippocampi

In vitro ageing test only brings out so much information, as the environment model does not take into account tissue density, liquid renewal and biological components. Preliminary trials in vivo in small animals are an essential step to validate device biostability over time for possible chronic implantation in humans. In this section, ve GEN2 Parylene probes were implanted in mouse hippocampi for a duration of 11 to 29 weeks. Neural signals were recorded regularly during the trial length, and impedance variation before and after implantation was measured. As explained earlier in this Chapter, the hippocampus is a region of brain involved in both memory consolidation and spatial navigation.

The sense of the environment and the ability to navigate in it are considered one of the most fundamental brain functions. American experimental psychologist Edward Tolman studied how animals learn to navigate and rst suggested the concept of a map-like representation of place in the brain [START_REF] Tolman | Cognitive maps in rats and men[END_REF]( Figure 4.11a). It is only in 1971 that O'Keefe and Dostrovsky related the concept of cognitive map to the existence of specic neurons of the hippocampus now known as place cells [START_REF] O'keefe | The hippocampus as a spatial map: preliminary evidence from unit activity in the freely-moving rat[END_REF]. Simply put, a place cell is a pyramidal neuron of the hippocampus that becomes active, ie that res action potentials at elevated frequency, whenever the subject enters a specic place of the spatial environment (Figure 4.11b) [START_REF] Girardeau | Selective suppression of hippocampal ripples impairs spatial memory[END_REF]. The activation of such place cells in the context of spatial navigation leads to a very characteristic brain signal composed of Theta oscillations at around 7-12 Hz, where smaller and faster Gamma oscillations are nested [START_REF] Lisman | The Theta-Gamma Neural Code[END_REF][START_REF] Zheng | Spatial Sequence Coding Diers during Slow and Fast Gamma Rhythms in the Hippocampus[END_REF][START_REF] Tallon-Baudry | Oscillatory gamma activity in humans and its role in object representation[END_REF] (Figure 4.11c).

The study of this Theta-Gamma neural coding is primordial to understand patterns of object representation and memory consolidation during sleep state [START_REF] Tallon-Baudry | Oscillatory gamma activity in humans and its role in object representation[END_REF][START_REF] Zheng | Spatial Sequence Coding Diers during Slow and Fast Gamma Rhythms in the Hippocampus[END_REF]. It can easily be recorded from electrodes placed in the hippocampus without the need of too precise electrode positioning. In this study, we decided to assess the biostability of our probes by recording the hippocampal brain waves produced by freely-moving mice discovering an unfamiliar environment.

Experimental setup 4.4.2.1 Recording Parylene probes

The Parylene probes used for this study are the GEN2 probes, specically designed for this application.

They are integrated with a wireless packaging (Figure 4.12) so that the signal can be recorded in freelymoving mice. Depending on the probes, two or three gold electrodes are PEDOT-nanostructured, in order to have a comparison between gold and PEDOT-modied electrodes in the long run.

In this trial, no neural recording is performed during the implantation procedure, because the wireless packaging was not adapted to the acquisition system in the implantation room. All probes are simply inserted in the hippocampus using only stereotaxic coordinates and an automatic micropositioner (LN Junior22, Luigs&Neumann, Ratingen, Germany). Therefore, the requirements in terms of mechanical reliability of the probes are far less challenging that from the rst in vivo trial, where acurate positioning during implantation procedure was demanded. For this reason, the silk coating used is only 150µm-thick and no water-annealing is performed, so that the silk dissolves fast into the mouse brain.

Implantation protocol

The ve Parylene devices were implanted following a protocol similar to the one described earlier in this Chapter, and represented on Figure 4.13. C57BL/6 mice (Charles River) are anaesthetized with isouorane (Veturane) and set to a stereotaxic apparatus. After a skin incision, a craniotomy is performed, and the dura mater of the meninges is incised to allow access to the pia mater. The Parylene shank is then inserted in the brain, and a rst level of dental ciment is mounted on the shank base to x its position.

Then, the reference electrode (microscrew with silver wire) is screwed to the occipial bone. Finally, the whole apparatus is xed in a second layer of dental ciment, so that only six pins are accessible on the top of the implant. Finally, the skin is sutured around the dental ciment. Silk or no silk ? With a short length of ~2mm, the non-coated Parylene shanks are theoretically supposed to withstand implantation in the brain. Indeed, the theoretical buckling force is approximately 6 mN for such Parylene beam, therefore higher than the theoretical insertion force at 1.5mN. Besides, insertion tests in agarose gel were successful, so that in this case, the silk coating should not be needed. With that in mind, both silk-coated and pristine devices were tried out for insertion in mice brains. Surprisingly, it was observed that non-coated Parylene shanks tended to buckle on the surface (Figure 4.14), probably because for a small mouse brain, it is uneasy to check if the pia mater is removed or not. Also, coagulation is very fast in mice and uids tend to dry quickly on the brain surface, producing a sligh crust that is hard to break for a Parylene shank. When the brain surface was slightly punctured using a needle gauge, the Parylene shank was able to be inserted quickly after, without the need of a silk coating. On the other hand, silk-coated shanks were all inserted without any sign of buckling. This highlights once more that the denition of a stiening strategy is highly dependent on the targeted application and experimental considerations. To let the mice recover, the wireless recording sessions start after one week post-implantation. Figure 4.15 depicts the principle of a recording session. During these sessions, a mouse is connected to a headstage (MultiChannel System W4-HS), mounted with a wireless battery on top (MultiChannel System W2100-B-30mAh-BB). The animal is then transferred to the recording cage, composed of a carton box with litter, where it stays for ve minutes. During this time, the animal discovers its new environment, and its brain activity is recorded via the MultiChannel System data acquisition system. After the ve minutes, the headstage is removed and the animal placed back in its cage. The experiment is then carried out on the following mouse.

Wireless recordings

Signal extraction and data treatment

If the electrode is working properly, the recorded signal should present with a pattern as seen on Figure 4.16a, whenever the animal is actively exploring its environment (and not defecating or scratching itself, for instance). This representation is therefore highly dependant on animal behaviour at the moment considered. To adequately quantify signal amplitude, a power spectrum is usually applied, so that the signal is represented as a function of the signal amplitude given at a specic frequency (Figure 4.16b).

Here, we chose to form the power spectrum during each session of the signal taken between minute 3 and minute 5 of the recording session, in order to avoid preliminary random animal behaviour caused by the stress of connecting the headstage, and to avoid animal habituation to its environment, which leads to a passive behaviour. Given that, in the hippocampus, most of the signal amplitude is comprised in the Theta range, comparing signals is then transduced in comparing their Theta power amplitude.

Experiment plan

The experiment plan is represented on Table 4.2. Three out of twenty implanted electrodes were not functioning since the rst recording session, with a signal suggesting a short-circuit in the circuitry, therefore they are removed from data treatment in the following section.

Besides, mouse C and mouse D presented with medical issues and are knowingly excluded from the trial. Their constant ghting in the cage resulted in improper wound healing around the prosthesis, and they had to be sacriced after only 7 weeks of implantation.

Evolution of neural recordings 4.4.3.1 Evolution of number of functioning electrodes over time

A functioning electrode is dened as an electrode that is able to acurately record hippocampal oscillations typical of a mouse being actively discovering its cage. It is relatively easy and uncontroversial to check if an electrode is in working conditions by analyzing the spectral representation of a recording session, as depicted on Figure 4.17. Table 4.2 shows the number of functioning electrode ('1'), and not functioning ('0') electrodes. Some electrodes (mouse E electrode C or mouse B electrode A for instance) are able to record local eld potentials but at some points appear to be unfunctional. Given the rate at which they are either stable or unstable and the signal amplitude at times where they are not working, it appears more likely to be linked to a faulty contact than a momentary increase in electrode impedance for biological reasons. Overall, the number of recording electrodes in vivo in this trial varies depending on the timing, but at any given moment, between 70 and 100% of all recorded electrodes are indeed functional. After approximately 14 weeks in vivo, the unstable electrode has lost the ability to acurately record brain activity: the signal is random, which leads to a variety of power amplitudes in the spectral representation.

On the other hand, the stable signal still displays a noticeable Theta oscillation with rather stable amplitude.

Evolution of SNR over time

We now focus on the electrodes that are indeed functional at a given time, and try to assess the evolution of signal quality of these functional electrodes over time. The Signal-to-Noise Ratio (SNR) represents a ratio between the meaningful signal and the measurement noise, and can be calculated using the denition below :

SN R = A 2 × SD noise
where A is the amplitude of the signal, that is the peak-to-peak voltage of the mean waveform, and SD noise is the standard deviation of the noise [START_REF] Suner | Reliability of signals from a chronically implanted, silicon-based electrode array in non-human primate primary motor cortex[END_REF]. To calculate the SNR, our recorded signal is smoothed using the Smooth Data Processing tool in Spike2 software and the noise is considered as the dierence between the raw signal and the smoothed signal, as seen on Figure 4.18.

It is important to stress out that the previous denition of the SNR is usually applied in electrophysiology to dene the quality of spike sorting [START_REF] Ward | Toward a comparison of microelectrodes for acute and chronic recordings[END_REF][START_REF] Ludwig | Chronic neural recordings using silicon microelectrode arrays electrochemically deposited with a poly(3,4-ethylenedioxythiophene) (PEDOT) lm[END_REF], where the amplitude of the signal is independant from experiment condition, since a spike amplitude is constant. In our case, the SNR is hard to dene, because depending on experiment conditions (animal behaviour, implant position, etc.) the signal amplitude will vary, independantly from electrode quality. Therefore, evaluating the evolution of device electrical capacity over time is troublesome. We tried to form here an adapted denition of the SNR, that stands out from experiment variable. For one animal and one electrode, the SNR is calculated First of all, the results show that the SNR is high enough (between 13 and 28) for proper LFP hippocampal recordings. Besides, it seems that for a functioning electrode, the SNR does not drop over the course of the trial, and remains at a stable value around 20. This means that the working devices seem to withstand well prolonged contact in vivo in brain mice. Besides, one could assume that the body reaction to our implant (even though we didn't assess it) did not interfere with recording capability.

Morphology and impedance alteration

The alteration of electrode impedance before and after implantation can be observed through EIS measurement and morphology assessment via SEM images, before and after implantation. -A number of electrodes (~8%) were not operating from the rst day of recording, consistent with probable faulty contacts or direct mechanical damage during insertion. As explained in the rst section of this Chapter, this issue is indeed reported as one of the material failure modes described in the literature [6]. Extra attention should be paid on connectic packaging repeatability to avoid failures in the future.

-During the length of this pre-trial, at least 70% of the electrodes were able to record LFP with an rather stable SNR. However, in order to properly measure the SNR, future in vivo trials should feature the recording of single action potentials, so that the signal amplitude is known and not experiment-dependant.

-No signicant dierence between the performance of gold and PEDOT-modied electrodes was observed. Future in vivo works should involve more animals and actual recordings during the implantation procedure, in order to optimize LFP amplitudes and allow eventual spikes recordings to better assess signal quality over time.

-Both impedance and morphology assessment before and after implantation showed no evidence of material degradation during the length of the trial, which is consistent with the in vitro trial.

Conclusion

Chronically-implanted neuroprosthetic devices have the ability to help patients suering from a wide range of neurological disorders, from Parkinson's disease to Alzheimer's. The very rst requirements of such neuroprosthesis rely in the stability of the electrode/tissue interface. Prolonged contact between an implant and the physiological environment around it may lead to mechanical, chemical and electrical alterations of the device, provoking its eventual failure in the long term. For this reason, the biostability of implanted devices should be investigated at dierent levels before chronic implantation in humans can be considered. In this Chapter, we intended to give a preliminary assessment of the stability of our Parylene-based neural probes for possible chronic purposes.

First, we showed the feasability of acurate, high quality neural recordings in the hippocampus of a mouse. During this acute recording, the enhancements of both electrical properties via PEDOT nanostructuration and mechanical properties via the silk coating were once more demonstrated.

Then, in the rst step towards biostability assessment, Parylene devices were soaked in articial brain physiological uid for a period of 6 months, during which both morphology and electrical properties were monitored. No signicant sign of alteration of neither the Parylene substrate nor the PEDOT nanostructuration were observed, correlated to stable impedance measurement over time.

Finally, ve mice were implanted with each a Parylene implant, and wireless recordings from freelymoving animals were gathered for up to 6 months. Many electrodes (~8%) presented with faulty contacts or other damaging issues, resulting in recoring incapacity. However, at least 70% of the electrodes were able to record LFP over the course of the trial. It seems that for a functioning probe, signal quality remained stable over time, with rather high SNR. Morphology and impedance comparison before and after implantation were encouraging, with once again no sign of Parylene degradation in time. Besides, at this time we are not able to conclude on PEDOT versus gold electrode biostability, but the Parylene C devices seem to withstand well prolonged contact with cerebral tissue during several months.

In other terms, in this Chapter we showed preliminary evidence that the devices in themselves seem stable in articial and actual physiological conditions during a period of 6 months. However, biostability is a function of both material reaction to its environment, and biological reactions to the implant. The next step in line towards biostability assessment should therefore involve the evaluation of tissue reaction to our devices in the long run via immunohistochemical analysis.

Chapter 5

Conclusion and perspectives

Main thesis outcomes

The implantation of neural devices for long periods of time, as in the case of chronic neural interfaces, sets the background to this PhD project. Indeed, the turn from acute electrophysiology recordings to chronic implantation is a key point for future clinical purposes, as implanted devices tend to deteriorate over time.

To achieve widespead use, the neuroprosthesis must be both innocuous and reliable. These two aspects are highly correlated : the inammation following device implantation not only constitutes an injury to the patient, but also encapsulate the recording microelectrodes in the form of a glial scar, which eventually results in signal loss over time. Besides, prolonged contact in physiological uids tends to aect the device in the form of corrosion of electrical sites, membrane biofouling, swelling and delamination of the passivation layer. In the case of chronically-implanted neuroprosthesis, implants must be able to record and/or stimulate neural tissues over long periods of time without adversely aecting body uids, tissues and organs.

Therefore, this work is dedicated to designing a long-term, reliable neuroprosthetic device for the recording of deep brain neural layers, that is as minimally invasive, as well-tolerated and as ecient as possible. Modularity of design and purposes was also considered a guideline in this work, so that the implant prototype could be seen as generic and adaptable.

How do we conceive a neuroprosthesis that promotes body acceptance ?

We guided our work towards the fabrication of a exible, polymer-based implant that benets from both enhanced material biocompatibility, compared to traditional silicon implants, and from a reduced stress around the implant. Indeed, the conformal implant molds to the surrounding tissues, thus reducing the shear forces between the two and alleviating the foreign body response.

Parylene C was chosen for the implant substrate, as it is a USP-classied material with the highest standard of biocompatibility, and gold was selected for the microelectrode materials for its renown biocompatibility. These implants were fabricated via standard microelectronic processes, starting from a silicon wafer to the release of unitary implants, as well as packaging and connectics. If standard microfabrication techniques benet from decades of knowledge improvements for usual substrates such as silicon, the turn towards polymer-based sensors still lacks of general expertise. Microfabrication processes on Parylene C have only been studied for a little over a decade, and technological locks are remaining nowadays to achieve the same level of perspective and hindsight as standard substrates. More particularly, Parylene C micromachining techniques are still considered a challenge, and special attention was brought in this thesis to obtain optimized, reproducible microfabrication protocols for the achievement of a Parylene C functional device. Moreover, the fabrication procedure is inherently modular, as implant design is only dependant on a three-masks photolithography-based steps. Depending on the targeted application, this modular approach could as easily be transferred to a variety of aims, including dense MicroElectrode Arrays or retinal implants. Besides, non-cytotoxocity features of the resulting devices were shown by both quantitative and qualitative culture cell viability assays.

How do we ensure a stress-reducing yet sucessful implantation of this neuroprosthesis ?

If substrate exibility presents with the advantage to adapt to the inherent soft tissues, this very same property involves some drawbacks in practical handling. Indeed, a soft implant will tend to deform and buckle on the brain surface, which makes surgical implantation quite challenging.

After having sorted and compared multiple techniques developed in the literature to perform such insertion, our work focused on the integration of a bioresorbable coating on the probe shank, that will assist insertion before resorbing innocuously in the tissues. Silk broin was appraised adequate for such purpose, rst because it is a very sti material that benets from certain hindsight over biomedical applications. Besides, its tunable degradation rate allows for the system to be adapted a variety of purposes. In this work, we developped a coating method that is repeatable, cost-ecient and designadaptable, and could easily be automated for widespread use. The Parylene probe with integrated silk layer was successfully used in in vitro and in vivo trials, and showed great superiority over most commonly-used PolyEthylene Glycol.

Besides, because silk broin in solution is not commercially available, its use until now has been reserved to a little number of laboratories around the world, and especially in France. We hope that the development of this technology at LAAS-CNRS could benet other research topics in biomedical engineering.

How do we assess the lifetime of this neuroprosthesis ?

The body can be an challenging environment for many types of materials and devices in contact with tissues or body uids in the long term. For this reason, the biostability of implanted devices should be investigated incrementally before chronic implantation in humans can be considered. This line of work is all the more revealing that very few studies have yet shown the eect of longer-term implantation (max 3 months) over polymer-based neuroprosthetic devices. This work involved acute in vivo recordings, chronic in vitro soaking tests and chronic in vivo recordings. First, the functionality of the Parylene probe and its silk coating was demonstrated over a single implantation in mouse brain. In this experiment, both implantation and highly quality recordings were successfully carried out in a mouse hippocampus. This signal quality stems from an enhancement of the gold electrode capability via electrochemical deposition of conductive polymer PE-DOT, which was the subject of a previous PhD project. Then, in the in vitro trial, several implants were soaked in physiologically-resembling uids during a period of six months, where both morphology and electrical properties were monitored. Results indicate that no signicant sign of alteration of neither the Parylene substrate nor the PEDOT nanostructuration were observed, correlated to stable impedance measurement over time. In parallel, wireless recordings from freely-moving mice were gathered during these six months. The morphology and impedance comparison before and after implantation were encouraging, with once again no sign of Parylene degradation in time. At this time we are not able to conclude on PEDOT versus gold electrode biostability, but the Parylene C devices seemed to withstand well prolonged contact with cerebral tissue during several months.

In other terms, we have gathered encouraging results of Parylene-based neuroprosthesis adequacy for chronic implantation.

Short-term perspectives

The work developped in this thesis is a stepping stone towards a series of improvements and tests towards chronic implantation in humans.

The integration of bioactive agents in the silk coating An aspect that has not been studied here is the possibility to mix the silk broin coating with bioactive agents. Indeed, implant acceptance by the body can be enhanced via the delivery of small amounts of drug that ensure neural viability and dierenciation around the shaft [START_REF] Sommakia | Materials approaches for modulating neural tissue responses to implanted microelectrodes through mechanical and biochemical means[END_REF]. Porous silk broinbased nerve conduits infused with GDNF (glial cell line-derived neurotrophic factor) were shown as potentially viable conduit for Schwann cell migration and proliferation in the regeneration of peripheral nerves [START_REF] Lin | Spatially controlled delivery of neurotrophic factors in silk broin-based nerve conduits for peripheral nerve repair[END_REF].

Besides, an in vitro model presented the feasability for silk lm to encapsulate and release the enzyme chondroitinase ABC, known to promote neural regeneration [START_REF] Tien | Silk as a multifunctional biomaterial substrate for reduced glial scarring around brain-penetrating electrodes[END_REF]. However, it is hard to quantitatively assess the inuence of such drug delivery system, because the bioresorption kinetics and molecular bonding between the silk and the drug to be released are hard to represent [START_REF] Yucel | Silk-based biomaterials for sustained drug delivery[END_REF].

In the lab, the mixture of Nerve Growth Factor with silk broin at 500 ng/mL was achieved with good homogeneity, but the tests stopped for a lack of good in vitro model for the dierenciation of nerve cells. In my opinion, the improvement on glial scar of nerve-enhancing drugs in the silk coating should be put in evidence via in vivo tests, because the diusion and resorption of silk broin is too complicated to be modeled in vitro.

Second validation step in vivo

In this thesis, we have carried out a rst step towards the validation of our Parylene probe for possible chronic implantation in humans. The last study involved the chronic wireless recording of hippocampal activity from ve mice during six months. It allowed us to pin-point several experimental issues that should be addressed for the next in vivo trial.

-First, the next study should involve an increased number of animals. Five mice is already a little number for valid statistics, but when random issues occur (error in implantation location, animal weakness, fellow ghting etc.), it leads to retrieve some mice from the study.

-Implant packaging is already somewhat optimized regarding the considered application, but attention should be especially brought to the welding steps. Possibly, it would benet from certain automatization of the process, so that the nal packaging is as reliable as possible.

-The neural signal recorded should involve single units activity (spikes), so that proper assessment of the Signal-to-Noise Ratio evolution over time would be performed. This involves greater precision in surgical positioning, which is troublesome and time-consuming.

-Finally, for obvious reasons the most important feature would be to perform in parallel an immunohistochemical analysis of the glial scar around the Parylene probes. Correlated to electrode functionality and signal quality over time, this would greatly improve the understanding of biotic failure from chronic implantation. Optogenetics refers to the integration of optics and genetics to manipulate (i.e. activate or inhibate) certain neural cells with extreme temporal and spatial resolution. This remarkable technology has started to develop a little over a decade ago, when in 2005 a report described that upon the introduction of a microbial opsin gene, some specic neurons started to be responsive to light [START_REF] Boyden | Millisecond-timescale, genetically targeted optical control of neural activity[END_REF].

Towards optogenetic studies using a exible probe

The principle is based on the use of light-gated ion channels known as rhodopsins, often found in prokaryotes and light-sensitive algae. The gene coding for the fabrication of such protein is then linked to a promoter and implanted in a virus. When the animal (for instance a mouse) is injected with the virus, it will target specic neural populations (gene therapy), that will express the algae gene and exhibit such opsins on their surface. Therefore, when a certain light is emitted (often with an optic ber coupled with a laser) these light-gater ion channels will either trigger a pulse or prevent one from propagating [START_REF] Deisseroth | Optogenetics[END_REF][START_REF] Pashaie | Optogenetic brain interfaces[END_REF] (Figure 5.2). These opsins are wavelength-dependent, and can either induce stimulation or inhibition of neural pathways. For instance, channel-rhodopsin 2 (ChR2) stimulates the host-neuron when illuminated by a blue light, whereas halorhodopsin inhibates it when coupled with a yellow light [START_REF] Pashaie | Optogenetic brain interfaces[END_REF][START_REF] Tye | Dopamine neurons modulate neural encoding and expression of depression-related behaviour[END_REF].

Overall, the optical control allows for millisecond-scale temporal resolution coupled with cell-type specic spatial precision, making optogenetic approaches a window for better understanding of brain disorders or for possible new solutions [START_REF] Yizhar | Optogenetics in neural systems[END_REF]. In the long run, the BCI community could benet from such technology, where a highly-resoluted sensory feedback could be obtained via light stimulation [START_REF] Pashaie | Optogenetic brain interfaces[END_REF].

ChR2

) is transplanted in a virus injected in the animal, that will express the gene on a certain neural population. When irradiated with a specic wavelength (e.g. blue light at 470 nm), the ChR2 channels will induce depolarization of the neuron, thus a spike will be formed (https://www.addgene.org/optogenetics/) ). (f ) An optrode consisting of a silica ber and a tungsten electrode [START_REF] Gradinaru | Targeting and readout strategies for fast optical neural control in vitro and in vivo[END_REF] (g) Metal microwire array equipped with a silica ber [START_REF] Kravitz | Optogenetic identication of striatal projection neuron subtypes during in vivo recordings[END_REF] (h) Optetrode combining four tetrode bundles and an optical ber [START_REF] Anikeeva | Optetrode: a multichannel readout for optogenetic control in freely moving mice[END_REF]. The inset shows the cross-section schematic. (i) Utah array incorporating a tapered silica ber [START_REF] Zhang | Integrated device for optical stimulation and spatiotemporal electrical recording of neural activity in light-sensitized brain tissue[END_REF](j) Silicon multitrode probes integrated with tapered silica bers [START_REF] Royer | Multi-array silicon probes with integrated optical bers: Light-assisted perturbation and recording of local neural circuits in the behaving animal[END_REF]. Reproduced from [START_REF] Pashaie | Optogenetic brain interfaces[END_REF].

Engineering neural probes for optogenetic studies

Since optogenetics have spread in the neuroscience community, designing a neural probe that could be used for both neural recording and light stimulation has become highly desirable. The turn towards such devices is represented in Figure 5.3. Most of the time, these tools are just an implementation of traditional silicon devices with an optic ber. Coupled with a LED or LASER, these systems allow for a high-speed trigger that matches the pace of biological phenomenons (spikes, in the order of the millisecond) [START_REF] Kim | 3D silicon neural probe with integrated optical bers for optogenetic modulation[END_REF] (Figure 5.1). Unfortunately, an optic ber can be a huge drawback when the experiment involves a freely-moving animal, as it is very fragile and cannot be bent multiple times. In this sense, an integrated light source in the probe such as µLEDs, improves spatial targetting and is compatible with unhindered animal movements [START_REF] Cao | An integrated microLED optrode for optogenetic stimulation and electrical recording[END_REF][START_REF] Mcalinden | Thermal and optical characterization of micro-LED probes for in vivo optogenetic neural stimulation[END_REF][START_REF] Kim | 3D silicon neural probe with integrated optical bers for optogenetic modulation[END_REF] (Figure 5.4). However, tissue heating during the operation of microLEDs remains a major challenge [START_REF] Fan | A polycrystalline diamond-based, hybrid neural interfacing probe for optogenetics[END_REF]. A Parylene-based neural probe with integrated µLED(s) would benet from device biocompatibility, exibility but also transparency for optical stimulation, and should reduce tissue heating thanks to its low thermal conductivity. It could be used for both surface and depth stimulation, as seen on Figure 5.4c. μOLED are thinner than ILEDs and optic ber (OLED active layers thickness is about 200 nm compared with 7 μm for ILEDS), they allow spatial precision, low cost and preserve the environment (low CO 2 footprint). To our knowledge, there has only been a preliminary study featuring an OLED fabricated on top of a Parylene substrate [START_REF] Yamamura | Flexible organic leds with parylene thin lms for biological implants[END_REF] (Figure 5.4c).

The beginning of a new collaboration ?

The CRCA group, with the recruitment of Dr Laure Verret, is currently investigating cognitive decline mecanisms in Alzheimer's disease [START_REF] Bezzina | Early onset of hypersynchronous network activity and expression of a marker of chronic seizures in the Tg2576 mouse model of Alzheimer's disease[END_REF]. It has been demonstrated that transgenic mouse models of Alzheimer's disease display aberrant brain network activities [START_REF] Verret | Inhibitory interneuron decit links altered network activity and cognitive dysfunction in Alzheimer model[END_REF]. This malfunctioning oscillatory activity depends on the synaptic activity of parvalbumin-expressing fast-spiking interneurons (PV cells).

Recently, it has been shown that enhancing gamma activity by optogenetic stimulation of PV cells ameliorates performance of circuit processing. Besides, these PV cells can be genetically modied via the use of ChR2 to respond to blue light at around 470nm.

On the other side, at LAAS-CNRS, in the MICA group, Dr Isabelle Seguy and coworkers are currently investigating the enhancement of blue OLED properties, that shine at around 470 nm [START_REF] Tsopela | Development of a lab-on-chip electrochemical biosensor for water quality analysis based on microalgal photosynthesis[END_REF]. -The displayed wavelength must be about ~470 nm, with little specications regarding the spectrum.

-The out power must be in the range of 10-15 mW/mm² with reduced input voltage, so that autonomous, wireless recordings in genetically-modied mice could be performed.

-It should withstand a pulsed activity at around 100 Hz.

-The thermal heating resulting from the pulsed activity should be minimal, so that it doesn't endanger the surrounding tissues.

During this PhD, a series of tests were performed to assess the feasability of the project. The principle of OLED fabrication on a Parylene layer is described in Figure 5.6. A typical OLED is composed of thin layers of organic materials situated between two electrodes, the anode and cathode, deposited on a substrate. Here, the anode is made of Indium Tin Oxide (ITO, transparent) and the cathode is an aluminium layer. The organic materials are what denes the OLED color. In this case, the blue light is obtained with the use of NPB (20 nm) as a HTL (Hole Transport Layer), PCAN (40 nm) as an EML (Emission Layer) and Alq 3 (20 nm) as an ETL (Electron Transport Layer). A PEDOT-PSS layer can also be used before organics evaporation, to smooth the substrate and to improve charge transfer between the layers.

First steps of Parylene-based OLED

The ITO layer on top of the Parylene substrate was performed by Solems Sa and characterized at LAAS to see if Parylene could be used as an OLED substrate. First, the transparency of Parylene and Parylene/ITO was measured using a spectrophotometer, because the blue light will be emitted from the backside. The roughness was also evaluated, because the ITO layer should be smooth and conformal in order to obtain good OLED performance. Results show that very similar roughness between the Parylene layer and the Parylene/ITO (see Part II of this thesis), with a root mean square average of the roughness prole of about 10 µm measured via AFM. Besides, the sheet resistance was measured using a four-point probe (R = 80 ± 10Ω/ ) and the ITO thickness was appraised via mecanic prolometry (e IT O 70nm). These parameters are summarized on Table ??, and compared to the optimum scenario, where ITO is spread on a glass substrate, as used traditionnally for OLED fabrication. The gure of merit of a Parylene/ITO layer is a lot lower than for the glass substrate, but isn't so bad for a plastic compound, especially considering that these parameters where only measured and not improved in any way whatsoever.

We were then able to fabricate a few OLEDs on a Parylene substrate using the process described on Figure 5.6. As a result of lack of time, we did not look into OLED encapsulation, though we are aware that this issue will be particularly troublesome in the future. The rst batches of OLEDs were Roughness R q (AFM)

10

Transmission T @ 470nm (%) 85 78 Sheet resistance R (Ω/ ) 8.9 80 There are many experiments to be expected before we could fabricate a Parylene-based neural probe with OLED integration. Among these steps, the reduction in OLED size, the reliability of the encapsulation and the control of both power and temperature during device operation should be thoroughly studied. However, in light of these rst outcomes, it seems that such device fabrication should be viable for future optogenetic applications. In the civilian world, transhumanism advocate the (ab)use of emerging technologies to enhance physical or mental functions of the human body. This extreme school of thought predicts that humans are meant to overcome their weaknesses by the means of science, towards a post-human condition. With their ideas of non-ageing, transfer of identity and intellectual enhancement, they are gaining more and more attention and support. For instance, Google has recently become one of the major sponsor of the transhumanist wave, among others through the massive nancial support of biotechnology and cognitive science labs. Disregarding the controversiality of a post-human race, the question of availability remains: who will benet from these technologies? Are we heading towards a world where the breach between the haves and the have nots will be exacerbated, where there will be two co-habitating races on our planet?

Legislation and standards

Bioethics aim at dening the limits of medical intervention on the human body while guaranteeing the respect of person dignity. At the world level, UNESCO includes an International Bioethics Committee, that takes into account medical progress and ethical and judiciary stakes in the life science eld. In France, the national consultative committee of ethics (CCNE), created in 1983, often gathers to publish bioethics guidelines. Besides, the french Bioethical law is revised every ve years to adapt to newest discoveries or technological advances. The law includes indications and prohibitions on subjects such as end-of-life decision, genetic manipulation and medically-assisted conception. However, anticipating probable misuse or abuse of these neurotechnologies is troublesome. The emergence of technically advanced neurotechnologies raises questions that are not necessarily explicitely covered by current legislation. Councils and focus groups both nationally and worldwide gather to discuss new trends in neurotechnologies and possible drifts, but their reection has unfortunately no judiciary value.

If the warranty of human integrity is theoretically in the hands of institutions and public forces, it has become obvious that responsability shall be shared with everyone that works in the eld of neurotechnologies. No matter the intended use for a process or a device, one shall always consider potential drifts.

C.2.4 Tests de cytotoxicité

La biocompatibilité est une notion relativiste pour décrire l'habilité d'un matériau à fonctionner avec une réaction immunitaire appropriée pour une application spécique [START_REF] Black | Biological Performance of Materials: Fundamentals of Biocompatibility[END_REF]. Dans le cas des implants, la biocompatibilité implique une limitation de la réaction de rejet à un implant, et la limitation de la réponse de l'implant dans cet environnement biologique.

Dans notre cas, les matériaux qui interviennent dans la confection de l'implant bénécient d'autorisations gouvernementales pour des applications chroniques. Le Parylène est approuvé par la FDA (Food and

F f lambement = π 2 I x E (KL) 2 (C.1)
où F f lambement est la force de ambement, E le module d'Young, L sa longueur (dimension selon l'axe z ) et K un facteur dépendant des conditions aux limites (poutre xée, semi-libre etc.). Le moment d'inertie I x selon l'axe x (l'axe de ambement) dépend de la forme et des dimensions de la section de la poutre. Dans le cas d'une section rectangulaire, il est égal à I x = ab 3 12 , où a est l'épaisseur de la poutre (dimension selon l'axey) et b sa largeur.

La force d'insertion, elle, est la force minimale requise pour l'insertion dans le cortex. Le cerveau, ou plus précisément la pia mater, a des propriétés mécaniques semblables à un gel agarose à 0.6% [START_REF] Hsu | Encapsulation of an integrated neural interface device with parylene C[END_REF](Figure C.12a). La force d'insertion dière selon les dimensions de l'implant, mais a été évaluée à 0.5-2 mN [START_REF] Tian | Monitoring insertion force and electrode impedance during implantation of microwire electrodes[END_REF][START_REF] Jensen | Assessment of subdural insertion force of single-tine microelectrodes in rat cerebral cortex[END_REF].

Figure 1 . 1 :

 11 Figure 1.1: The human nervous system and the principal regions of the Central Nervous System

1. 1 . 2

 12 Neurons and neural circuitry 1.1.2.1 The neuron structure A neuron is an excitable cell that constitutes the most fundamental functional unity of the nervous system. Its discovery dates back to the beginning of the nineteeth century, where italian pathologist Camillo Golgi invented a method of imaging nervous tissue that could stain a limited number of cells at random. Later, the work of Santiago Ramón y Cajal was leading in what came to be known as neuron doctrine. Cajal eventually demonstrated that the nervous system was made up of individual cells connected to each other by small contact zones (Figure 1.2).

Figure 1 . 2 :Figure 1 . 3 :

 1213 Figure 1.2: (a) Ramón y Cajal's drawing of the cells of the chick cerebellum, from Estructura de los centros nerviosos de las aves, Madrid, 1905 [1] (b) Brain neurons from the dental gyrus (hippocampus)of a mouse stained with the brainbow technique[2]

Figure 1 . 4 :

 14 Figure 1.4: Varieties of neuroglial cells in the Central Nervous System (from [3])

Figure 1 . 5 :

 15 Figure 1.5: Local changes in ion channels and membrane potential during the propagation of an action potential. Retrieved from[4]

Figure 1 . 7 :

 17 Figure 1.7: (a) EEG and (b) ECoG electrodes on human to record brain activity

Figure 1 . 11 :

 111 Figure 1.11: Schematic representation of the generation design of leading microelectrode array technologies, including (a) Cone electrode (Neurotrophic electrode) (b) Utah electrode arrays (EUA) (c) microwires and (d) Michigan probe (adapted from [6])

Figure 1 .

 1 Figure 1.13: (a) Photograph of a single tungsten microwire insulated with gelatin (b) Illustration of a tetrode structure. The closer a wire is from the ring neuron, the greater the amplitude of the recording. (c) Post-explant SEM of individual microwire to indicate deterioration in electrode insulation for Parylene C-coated Pt/Ir microwires. The deterioration occurs in the form of delamination and cracks (adapted from[START_REF] Prasad | Comprehensive characterization and failure modes of tungsten microwire arrays in chronic neural implants[END_REF][START_REF] Prasad | Abiotic-biotic characterization of Pt/Ir microelectrode arrays in chronic implants[END_REF][START_REF] Faveri | Bioinspired hybrid microelectrodes: a hybrid solution to improve long-term performance of chronic intracortical implants[END_REF]).

Figure 1 . 14 :

 114 Figure 1.14: The two principle silicon-based Microelectrode arrays. (a) The Utah Array: SEM micrograph of the Utah array and the insulated electrodes with exposed platinum tips [26] -(b) The Michigan Array: picture of the Michigan array with 3D-spread shanks, each shank contains multiple recording sites.A typical probe layout is displayed, with a thin section of the cerebral cortex to illustrate the basic spatial arrangement and scale[START_REF] Wise | Wireless implantable microsystems: High-density electronic interfaces to the nervous system[END_REF]

Figure 1 . 16 :

 116 Figure 1.16: Illustration of a Deep Brain Stimulation setup: the electrode is implanted in the subthalamic nucleus and the neurostimulator is placed under the skin (adapted from [46])

Figure 1 . 17 :

 117 Figure 1.17: Cochlear implant illustation and photograph

Figure 1 . 18 :

 118 Figure 1.18: Schematic representation of a Brain Computer Interface system (from [50] )

Figure 1 . 19 :

 119 Figure 1.19: Categorization of neural signals depending on the spatiotemporal resolution and sensors location (adapted from [20])

  have also been used in trials for communication-based BCI (Figure 1.20). If EEG-based BCI require long training for the users (over 300 sessions for an EEG in average, with 20 minutes per session) to nally aquire slow control of a spelling device (one or two characters spelled per minute in average) [51], they still represent an easy-access communication pathway for severely impaired patients.

Figure 1 .

 1 Figure 1.20: (a) BCI using slow cortical potentials (SCP depicted at the top). Patient selects one letter from the letter string on screen (right below) with positive SCPs, the spelled letters appear on top of the screen. (b) P300-BCI. Rows and columns of letter strings are lighted in rapid succession. Whenever the desired letter (P) is among the lighted string, a P300 appears in the EEG (adapted from [51])

Figure 1 .

 1 Figure 1.21: (a) The BrainGate pilot trial involving tetraplegic patient Matthew Naggle [29] (b) The mean paths of a human neural cursor from the center to 8 radial targets (session 3, 57 recorded single units)[START_REF] Kim | Multistate decoding of point-and-click control signals from motor cortical activity in a human with tetraplegia[END_REF] (c) Self-feeding task for BrainGate patient Cathy Hutchinson[START_REF] Hochberg | Reach and grasp by people with tetraplegia using a neurally controlled robotic arm[END_REF]

Figure 1 . 22 :

 122 Figure 1.22: Schematic representation of the trauma of insertion and chronic reaction involved in the implantation of a chronic probe in the brain (a) During surgical implantation, the probe induces dimpling of the brain surface before actual insertion, and severs capillaries and neurons as it makes its path, which eventually leads to greater acute reaction. (b) Several weeks post-implantation, activated microglia and astrocytes surround the implant. Eventually the glial scar is formed, and neurons develop further from implantation site.

Figure 1 .

 1 Figure 1.23: (a) Frames of insertions demonstrating severe tissue deformation when implanting a silicon device ex vitro in rat brain slices. Interest points are marked in red or yellow, and their motion is tracked by green lines in subsequent images (adapted from [71]) (b) Human temporal lobe acute implantations with a Utah Electrode array. Once the array has been removed there are some evident microhemorrhages (adapted from [70]).

Figure 1 . 24 :

 124 Figure 1.24: Photographs showing the silicon-based Utah Electrode Array (UEA) and representative results of its implantation into human cortex. (a) Astrocytes, labeled here with anti-glial brillary acidic protein antibody (GFAP) increase the thickness of their main processes, especially around electrode tracks (asterisk).(b) Resident microglial cells and blood-borne macrophages, labeled here with anti-CD45

Figure 1 .

 1 Figure 1.29: (a) Electrochemically polymerized polypyrrole blended with bronectin fragments (SLPF) deposited on gold electrodes[START_REF] Cui | Surface modication of neural recording electrodes with conducting polymer / biomolecule blends[END_REF] (b) Epitaxially grown gallium phosphide nanowires over a surface of 12µm diameter for acute neural recordings[START_REF] Suyatin | Nanowire-Based Electrode for Acute In Vivo Neural Recordings in the Brain[END_REF]

Figure 1 . 31 :

 131 Figure 1.31: Three examples of MEA connection architectures. (A) Electrodes are directly connected to signal pads with no active circuitry. (B) Electrodes are directly connected to on-chip active circuitry for signal conditioning. (C) Signals are multiplexed to the signal pads via column, row addressing in static mode (adapted from [121])

2. 1

 1 Introduction to Parylene C 2.1.1 The poly(chloro-p-xylylene), or Parylene C Parylene C (or simply Parylene) is a member of the poly(para-xylylene) group, with a single chlorine group per unit (Figure 2.1). It is deposited via CVD (Chemical Vapor Deposition) with the chamber at room temperature, and results in an ultrathin, transparent coating that presents with many attractive features:

Figure 2 . 1 :

 21 Figure 2.1: A schematic overview of the Gorham deposition of Parylene C: (1) The precursor dimer is evaporated.(2) In the furnace, the dimer is pyrolyzed to cleave the dimer into monomer units.(3) The reactive monomers adsorb to the substrate within the deposition chamber and coats the samples. (4) A

Figure 2 . 2 :

 22 Figure 2.2: Examples of Parylene-based MEMS -(a) Thin solar cell[START_REF] Jean | In situ vapor-deposited parylene substrates for ultra-thin, lightweight organic solar cells[END_REF] (b) Force sensor array[START_REF] Kim | Parylene-based electrochemical-MEMS force sensor array for assessing neural probe insertion mechanics[END_REF] (c) Intracortical probe[START_REF] Kim | 3D Parylene sheath neural probe for chronic recordings[END_REF] and (d) Retinal implant[START_REF] Rodger | High-density exible parylene-based multielectrode arrays for retinal and spinal cord stimulation[END_REF]

Figure 2 . 3 :

 23 Figure 2.3: Roughness of a 20µm-thick Parylene layer on a silica substrate, measured by interferometry

Figure 2 . 4 :

 24 Figure 2.4: Eect of UV ozone treatment time on a water droplet contact angle on a Parylene layer -Parylene is hydrophobic after deposition (θ ∼ 90°) but after 15 minutes of UV ozone treatment, the contact angle is at his lowest (θ ∼ 25°) (N=3)

Figure 2 .

 2 Figure 2.4 shows the eect of UV ozone treatment time on the hydrophilicity of a Parylene substrate.Within 15 minutes of treatment, the substrate reaches the lowest contact angle (θ ∼ 25°).The tunability of surface wettability plays a fundamental role for the improvement of layer adhesion during the microfabrication process of a Parylene neural probe. Besides, the presence of oxgenated groups on the Parylene surface might reveal promising for possible functionalization or integration of water-based compounds on the surface.

2. 1 . 5

 15 Layer adhesion and thermal treatment 2.1.5.1 Wet adhesion issues

Figure 2 . 5 :

 25 Figure 2.5: Parylene cracks occurs on the edge of the metal lines, which eventually causes delamination (from [160])

Figure 2 . 6 :

 26 Figure 2.6: (a) Illustration of semicrystalline, thermoplastic Parylene polymer with highlight on the crystalline and amorphous regions before and after annealing. (b) Example of thermoformed Parylene macro-scale coils (from [165])

2. 1 . 6

 16 Patterning and micromachining 2.1.6.1 Standard microfabrication techniques

Figure 2 . 7 :

 27 Figure 2.7: Cartoon representation of (a) photolithography involving the patterning of photoresist (positive or negative) on Parylene. (b) metal deposition on Parylene and (c) plasma etching (adapted from [169])

Figure 2 . 8 :

 28 Figure 2.8: Schematic representation of an ICP-RIE plasma setup

Figure 2 Figure 2 . 10 :

 2210 Figure 2.9: 20µm-thick Parylene conformably deposited over a silicon structure (scale bar 25 µm)

(Figure 2 .a

 2 Photolithography for lift-o

(

  Photolithography for release etching

Figure 2 .

 2 Figure 2.11: (a) Picture of the C30S Comelec used for the deposition of Parylene during this thesis (b) Relation between the mass of Parylene dimer PXC put in the sublimation chamber and the thickness obtained in the deposition chamber

Figure 2 . 12 :

 212 Figure 2.12: Cartoon representation of a lift-o process -(a) the photoresist is patterned via photolithography (b) the material is deposited on the surface (c) the wafer is put in a solvent until all the photoresis dissolves and (d) the material patterns remain on the Parylene substrate

Figure 2 . 13 :

 213 Figure 2.13: Representation and SEM images of a Generation 2 probe after (a) the lift-o process, (b) plasma etching of the second Parylene layer and (c) photoresist stripping -The resulting probe has metal traces with opened recording sites and contact pads. (black scale bar is 150µm, white scale bar is 10µm)

Figure 2 . 14 :

 214 Figure 2.14: SEM images of plasma-etched 23 µm-thick Parylene structures using dierent mask types (magnitude x400). (a) The 50 nm Nickel mask provides a thin and elegant mask but is subject to sputtering and important residues, eventually leading to passivation defaults and device tearing. (b) A 40µm-thick positive PR (AZ40XT from MicroChemicals) was also tried out, but quickly overheated with the plasma, and within minutes formed cracks and bubbles on the surface. (c) A thick 50-60µm negative PR (Intervia BPN) provides conformal coverage over the plasma length, with minimal residues, and was selected as a solution for plasma etching of thick Parylene C structures. Scale bar is 50µm.

  (a) Oxygen ow and pressure The ICP3 equipment was rst tried out and enabled us to perform a rst series of tests. The increase in oxygen ow (from 20 to 40 sccm) and pressure (from 20 to 60 mTorr) produced very little change in the etching rate of Parylene C (approximately 0.4µm/min). Parylene is indeed inert, and its eective etching is most likely rather due to the physical bombardment of ions than to the chemical interaction between the ions and the Parylene surface. For this reason, increasing the chemical parameter of such plasma have little inuence on Parylene etching rate. (b) Temperature The temperature, on the other side, had a signicant impact on the etching rate of BPN. Ranging from 10°C to 30°C, Parylene etching rate made a jump from ~0.4 µm/min to almost Table 2.2: Parameters and characteristics of the plasma tests used in this study. Gas type, gas ow, pressure, ICP power, bias power and temperature are modied to assess their inuence on the resulting Parylene (PaC) etching rate and selectivity with comparison to BPN photoresist mask the double (0.78µm/min). However, the photoresist overheated and cracks appeared quickly on its surface. Other failed experiments involving improper thermalization and clamping of the wafer on the chuck produced similar results, therefore the temperature must be carefully adjusted to avoid photoresist overheating.

Figure 2 .

 2 Figure 2.15: SEM images of plasma-etched 26 µm-thick Parylene C structure during 26 min of oxygen plasma, with and without the addition of a single minute of uorinated plasma in the middle of the process. The uorinated plasma protects the sidewalls prole and prevents the residues to deposit back in the channel, resulting in a clean cut prole.

Figure 2 .

 2 Figure 2.16: SEM images of a Generation 2 neural probe cut out of the wafer using the Star recipe developed in this etching trial (Table2.2) before and after the stripping of the BPN photoresist

Figure 2 .

 2 Figure 2.17: GEN1 probe designed for single acute, wired recordings with stimulation electrode in vivo in mouse hippocampus

Figure 2 . 19 :

 219 Figure 2.19: Electrical connection and packaging for a GEN1 neural probe

Figure 2 . 20 :

 220 Figure 2.20: Picture and schematic illustration of the connectic developed by Amélie Degache for the in vivo wireless trial

Figure 2

 2 Figure 2.21: Operating principle of Anisotropic Conductive Film

Figure 2 . 22 :

 222 Figure 2.22: Illustration of the PEDOT nanostructuration developped at LAAS-(a) SEM images of dierent morphologies of PEDOT:PSS layers related to dierent electrochemical deposition parameters: a) 10 µm diameter electrode not cleaned (above) and cleaned (below) in H2SO4, b) 0.1% of EDOT (above) and 0.025% of EDOT (below), c) upper maximum potential at 1.1 V (above) and 1 V (below) -(b) Electrochemical mean impedance spectroscopy over a range of 10-104 Hz and related error bar, for a 40µm-diameter gold electrode with and without PEDOT nanostructuration. The impedance is reduced by roughly a factor of 100 when the electrode is polymerized with PEDOT (from [38])

  2.5.2 The MTT cytotoxicity assay 2.5.2.1 Principle of the MTT assay

Figure 2 .

 2 Figure 2.23: MTT assay protocol -First the probes are glued to the bottom of a culture plate and cells are seeded in the well. Then, after 24, 48 and 72 hours, once the cells have adhered to the bottom of the well, the medium is replaced by the MTT solution. After 4 hours, the MTT solution is replaced by the solubilization solution of DMSO and after 20 minutes, the nal purple solution goes into a spectrophotometer to measure its absorbance at 570nm.

  Cell seeding SH-SY5Y cells are cultured in Dulbecco's Modied Eagle's Medium (DMEM) with 10% fetal bovine serum, 1% L-glutamine and 1% penicillin-streptomycin added. Cells are seeded in the wells containing the Parylene probes and control wells, with approximately 20,000 cells/well. All three culture plates are kept at 37°C with 5% CO 2 until their release time. MTT assay After 24, 48 and 72h, a culture plate is retrieved, the medium is replaced by a solution of MTT (Sigma, M2128 Thiazolyl Blue Tetrazolium Bromide) at 1mg/ml in DMEM with phenol red or serum and the plate placed back at 37°C during 4 hours. Then the MTT solution is removed, replaced by DMSO and the absorbance measured using a ThermoScientic Varioskan Flash 3001 spectrophotometer (Figure 2.23).

Figure 2 . 24 :

 224 Figure 2.24: Results of the MTT cell viability assay: absorbance and percentage of cell viability between the wells containing Parylene implants and the control wells, after 24, 48 and 72 hours. The absorbance is proportional to the number of living cells in each well, and the viability is determined by the ratio between the absorbance measured from a well containing the device and the mean absorbance of control wells.

Figure 2 . 25 :

 225 Figure 2.25: Fluorescent viability assay protocol -First the probes are put at the bottom of a culture plate and cells are seeded in the well. Then, after 24, 48 and 72 hours, once the cells have adhered to the bottom of the well, the medium is replaced by the Calcein-Ethidium solution, and each dye will color live and dead cells. After 20 minutes, the images are taken with a uorescence microscope.

Figure 2 .

 2 Figure 2.26: Fluorescent viability assay of a SH-SY5Y cell culture over GEN2 Parylene implants. Images were taken at 24 and 72 hours. Live cells appear in green (calcein dye) and dead cells appear in red (ethidium dye). No dead cell was present around the probe. Scale bar is 200µm.

Figure 3 . 1 :

 31 Figure 3.1: (a) The dierent layers in human CNS meninges (b) 0.6% agarose is commonly used as a brain phantom

Figure 3 . 2 :

 32 Figure 3.2: Representative force versus compression curve for a single neural probe compressed (a) in agarose gel 0.6% mimicking brain tissues and (b) against a hard substrate. (a) The compression force will increase as soon as the probe touches the gel surface (1) , and experience a drop (2) when the probe is inserted. The corresponding compression value corresponds to probe dimpling. After being inserted, the probe will cause increasing shear forces as it makes its path into the gel (3) -(b) The compression force increases (1) until the probe starts buckling (2). Ultimately, the compression leads to probe breaking (3).

Figure 3 . 3 :

 33 Figure 3.3: Strategies for inserting exible probes into brain tissues inventoried from the literature. Integrated sti layer: A device with locally sti regions may confer more permanent stiness to the probe. (a) Schematic of the hybrid silicon-Parylene C probe with a locally exible region (reproduced from [117]). (b) SEM image of a Parylene probe with integrated vertical stieners (reproduced from [202] -Removable aid: A removable insertion tool temporarily attached to the probe might also allow penetration. (c) In vitro insertion in agarose gel of a Parylene C probe using a tool nested in the hollow tip (reproduced from [41]). (d) Schematic frames of insertion of a silicon stiener attached

3. 1 .

 1 INSERTION STRATEGIES FOR FLEXIBLE NEURAL PROBES IN THE LITTERATURE 71 3.1.2.1 Permanent electrode support Tailoring probe design might to a certain extent enhance the overall insertion force. Structuring a Parylene C-based probe with dierent thicknesses of Parylene C on the backside (Figure 3.3b) slightly

Figure 3 . 4 :

 34 Figure 3.4: Top: pictures of a sea cucumber, in the threatened (sti ) and relaxed (soft) and state. Bottom: simplied schematic representation of the switching mechanism found in the sea cucumber dermis and used in physiologically responsive mechanically-adaptive nanocomposites (from [6]).

First

  trials recount the successful insertion of 100 µm thick, 200 µm wide, 3 mm long nanocomposite probes into rat cortex, with a maximum force of 5.6 mN [198]. Recently, a more thorough immunohistochemistry trial on these implants showed signicant improvements of PVAc nanocomposite on the long run (Figure 3.3e) with a reduced neuroinammatory response and less neurodegeneration 16 weeks post-implantation [88].

Figure 3 . 5 :

 35 Figure 3.5: (a) Parylene probe buckling on the surface of a mouse brain during pre-trial in vivo procedure (b) Parylene probe integrated with an optic ber permanently glued in order to perform insertion in a mouse brain

Figure 3 . 6 :

 36 Figure 3.6: (a) Bombyx Mori silkworm on a mulberry tree (b) Bombyx Mori cocoons (c) Structure of raw silk

Figure 3 . 7 :

 37 Figure 3.7: Schematic of material forms fabricated from silk broin using both organic solvent and aqueous-based processing approaches (from [261])

Figure 3 . 8 :

 38 Figure 3.8: Schematic of the silk broin extraction procedure, from the raw material (cocoons) to the nal aqueous-based solution (from [261])

Figure 3 . 9 :

 39 Figure 3.9: Transferring patterns from a silicon/SU-8 substrate to a PDMS mold: (a) the unreticulated PDMS is rst poured onto the substrate and (b) let dried at 60°C overnight before (c) nally be gently removed from the substrate.

Figure 3 . 10 :

 310 Figure 3.10: Schematic illustration and zoom of the protocol to obtain a silk broin coating on the back of a Parylene C probe using a PDMS mold: (a) Electrode positioning (b) Polymer syringe-jetting (c) 12 h drying step (not to scale). The resulting probe has a staple-shaped silk structure on its backside.

Figure 3 . 11 :

 311 Figure 3.11: Silk lm thickness depending on its dilution percentage and the PDMS mold dimensions (N = 10/group).

Figure 3 . 12 :

 312 Figure 3.12: Microscope views of (a) a GEN1 Parylene probe with a silk coating on the backside (b) a silk U-beam structure obtained with the same process and (c) side view of the U-beam structure for thickness measurement by the microscope software

Figure 3 . 13 :

 313 Figure 3.13: Frames of insertion and removal in agarose gel for sole Parylene and SF-coated probes. Peak forces are displayed (N=5/group)

Figure 3 . 14 :

 314 Figure 3.14: Compression forces on a hard substrate for sole Parylene and silk-coated probes (thickness ~250 µm, length 6.5mm). Buckling forces are displayed (N = 5/group).

Figure 3 .

 3 [START_REF] Neher | Single-channel currents recorded from membrane of denervated frog muscle bres[END_REF] shows the determination of SF Young's moduli from the buckling force equation when these two parameters are taken into account.

Figure 3 .Figure 3 . 16 :

 3316 Figure 3.15: (a) Microscopic view and (b,c) model of the plane section of our silk U-beam structure used for the determination of silk's Young's modulus

2 :

 2 and thus less likely to break under insertion forces. The same coating protocol, characterization, standard compression tests and insertion tests in agarose were performed with PEG instead of silk broin. Comparison of silk broin and PEG as candidates for the stiening material of a Parylene C exible probe

3. 3

 3 Insertion mechanisms of 2D/3D arrays of Parylene probes 3.3.1 The collaboration with Dr Meng's lab at USC This work is part of a collaboration between the LAAS-CNRS lab in Toulouse, France, and Dr Ellis Meng's laboratory at the University of Southern California (USC) in Los Angeles, USA. Dr Meng is a professor of biomedical engineering and the Dwight C. and Hildagarde E. Baum Chair of the Department of Biomedical Engineeering at USC California. She directs the Biomedical Microsystems Laboratory founded in 2004 which conducts research on biocompatible polymer technology and micromachining, sensors and actuators, microuidics, and implantable and biomedical microelectromechanical systems (MEMS). Recently, a collaboration with USC colleague Dr Theodore Berger also from the Biomedical Engineering, department has oriented part of Dr Meng's research towards chronically implanted exible neural probes for measurements in the hippocampus.Hippocampal prosthesis are meant to address malfunctions of the hippocampus, an area of the brain involved in short-term and long-term memory [267]. Dr Berger's group is pioneer in the eld of hippocampal prosthesis, and their collaboration might lead to great advances in the development of long-term stable hippocampal neuroprosthetics. Between September and December 2015, and thanks to a grant obtained by the School of Doctors of Toulouse, I have been attending the USC lab and have been working with her research team on the insertion mechanisms of 2D and 3D Parylene-based neural probes for deep insertion in the hippocampus. As explained in Chapter 1, numerous chronic neuroprosthetic applications, such as hippocampal prosthesis, require high-density arrays of electrodes on the same substrate. This allows for simultaneous recording of great number of neurons, which leads to better neural interface resolution, as it increases the reliability of spike sorting and the performance of source localization. This goal in mind, Dr Meng's lab conceived comb-shaped Parylene C-based probes with 64 recording channels, and developed a method to allow deep implantation of long imsy neural probes. My project at USC centered on the mechanical evaluation of forces encountered during mock surgical implantation and to decipher how testing dierent probe geometries might yield clues on the potential to scale this technology platform and surgical method to allow high density recordings from large numbers of neurons.3.3.2 Fabrication and insertion of a hippocampal prosthesis 3.3.2.1 USC arrays of probes

Figure 3 . 18 :

 318 Figure 3.18: Rat hippocampal prosthesis prototype developed at USC-(a) Photography and (b) microscopic view of a single neural hippocampal device (c) Probe dimensions for a sample array between CA1 and DG, in μm. (i) Array of 8 probes. Probe height 5.5 mm, distance between probes 250 μm. (ii) Probe width varies from 70 μm to 110 μm, to a nal width of 150 μm. (iii) Patterned electrodes are 50 μm in diameter, and have a 30 μm exposure diameter. Probe thickness is 20 μm.

Figure 3 . 19 :

 319 Figure 3.19: In vivo placement of the hippocampal prosthesis developed at USC -(a) Aerial view of brain (gray) and hippocampus (yellow) seen in top left. Hippocampal cross sections at various anterior-posterior dimensions (colored slices through hippocampus). (b) Alternate view. Eight neural probes inserted into the hippocampus at a depth that spans CA1 and DG.

  .20. The tips of the probes are sandwiched between two layers of PDMS, and molten PEG is poured at 60°C in the remaining cavity and dried at room temperature. The remaining device was successfully implanted in a rat hippocampus, as seen on Figure 3.21.

Figure 3 .

 3 Figure 3.20: (a) Microscope image of a sham recording device (without platinum electrode) sandwiched into two layers of PDMS. The cavity is then lled with PEG, which will hold the bottom half of the probes together. (b) Schematic representation of the dierent steps to obtain a PEG layer. The detached probe (1) is xed to an acrylic holder (2), before being molded with PEG (3).

Figure 3 . 21 :

 321 Figure 3.21: Photographs showing (a) before insertion and (b) after insertion of a sham device (no electrodes) into a rat's hippocampus. First the tip is implanted up to 2.5mm into the tissues, then the PEG is dissolved with the addition of saline, and the rest of the probes are descended until reaching appropriate depth at 5mm.

Figure 3 . 22 :

 322 Figure 3.22: 1D, 2D and 3D neural recording devices (a) 1D Vector Array from Neuronexus (b) 2D Planar Array from Neuronexus (c) 3D Matrix Array known as Utah Array

Figure 3 . 23 :

 323 Figure 3.23: Bose setup for force/compression monitoring during probe insertion in brain phantom. The setup is placed on a pressurized table to avoid external vibrations.

  Both 2D planar arrays and 3D matrix arrays of sham Parylene C probes were evaluated in this study. Sham probes were obtained by plasma-cutting 20µm-thick Parylene C layer on a silicon substrate into the shape of the hippocampal prosthesis shown earlier.

Figure 3 . 24 :

 324 Figure 3.24: Pictures of the 2D planar arrays of Parylene sham probes used for the experiment -The shanks are cut to obtain 8, 6, 4, 2 and 1-shank combs.

Figure 3 . 25 :

 325 Figure 3.25: Example of a 3x8 3D array made with three Parylene devices stacked together with a spacer of 250 µm.As the stacking is manual, it is important to make sure that the shanks are aligned at least in the (z) directions, and that they are at on the (x) direction. The alignment in the (y) direction is dicult to achieve, but has less inuence in the force testing.

Figure 3 .

 3 Figure 3.26a shows a standard compression/force graph obtained with one of our samples. The setup allows for many sources of errors, electrode manufacturing, positioning and gel concentration being prevalent factors. Besides, the recorded forces are small compared to the signal, which increases all the more the uncertainty. However, the dimpling, insertion force and shear force were able to be extracted from these graphs, and are displayed in Figure3.26b-d.

Figure 3 . 26 :

 326 Figure 3.26: Eect of shank number in the case of 2D planar Parylene sham probes -(a) Standard compression test results for a 2.8mm sham probe (Probe 3) with respectively 8, 6, 4, 2 and 1 shank(s). (b) Dimpling (c) Insertion force and (d) Shear force associated with the compression of such samples in agarose gel. While the dimpling increases lightly, both the insertion and shear forces increase linearly with the number of shanks (N=5 for each experiment)

Figure 3 . 27 :

 327 Figure 3.27: Eect of shank number in the case of 2D and 3D matrix arrays of Parylene sham probes -(a) Standard compression test results for three 2.8mm sham probes with respectively 1x8, 2x8 and 3x8 shanks (b) Dimpling (c) Insertion force and (d) Shear force associated with the compression of 2D and 3D samples in agarose gel. While the dimpling increases lightly, both the insertion and shear forces increase linearly with the number of shanks (N=5 for each experiment, except for the 32 shanks, where N=2)

Figure 4 . 1 :

 41 Figure 4.1: Cellular level failure modes for implantable neural devices: (A) Cellular level failure modes act (1) disrupting the neuronal signal source, (2) impeding charge transport through the extracellular space or (3) disrupting charge transfer at the electrode recording site. (B) Higher, tissue level failure modes that have not been as widely described or considered in the literature. Asterisks denote failure modes with an underlying inammatory component. As nearly all failure modes have an underlying inammatory component, strategies that reduce neuro-inammation will be critical to improving the biocompatibility and function of intracortical microelectrodes (from [6])

Figure 4 . 2 :

 42 Figure 4.2: (a) Surface biofouling of an implanted medical device due to the adhesion of bacterial colonies and/or thrombolytic agents (from [276]) (b) SEM image representing the poor adhesion of Parylene on a silicon substrate, leading to great delamination (from EPFL Center of MicroNanoTechnology CMi website)

Figure 4 . 3 :

 43 Figure 4.3: 3D representation (in pink) and cross section of a mouse hippocampal formation (adapted from[START_REF] Sokolowski | Wired for behaviors: from development to function of innate limbic system circuitry[END_REF] and[START_REF] Defelipe | The evolution of the brain, the human nature of cortical circuits, and intellectual creativity[END_REF])-CA = Cornu Ammonis, DG= Dentate Gyrus.

Figure 4 . 4 :

 44 Figure 4.4: Representation of the experimental setup for the rst in vivo trial in mouse hippocampus.

  All signals are ltered and amplied (0.3 Hz3 kHz) and fed to a 50 Hz noise eliminator before being digitalized at 10 kHz and analyzed using Spike2 software (CED Systems).

Figure 4 . 5 :

 45 Figure 4.5: SEM and microscope images of the implanted probe: three electrodes were PEDOTnanostructured (black disc) and three remained gold (white disc). Electrochemical impedance spectroscopy of the implant is displayed: Z gold = 630kΩ and Z P EDOT = 7kΩ at 1kHz.

  The stimulation electrode is constituted by two NiCr 50 µm diameter wires, passivated by a Teon® layer and disrobed for 250 µm high. It is implanted in the ventral hippocampal commissure, in a tract of bers made of axons of CA3 neurons innervating the contralateral CA1 region. A stainless steel screw is inserted across the occipital bone as a reference electrode (Figure 4.6).

Figure 4 . 6 :

 46 Figure 4.6: (a) Photography of the in vivo rst trial, with positions on the scalp of the GEN1 Parylene recording electrode, the reference electrode and the stimulation electrode. (b) Zoom of the 250 µm-thick silk coated Parylene probe inserted into the mouse hippocampus with no sign of buckling.

Figure 4 . 7 :

 47 Figure 4.7: Electrophysiological signal for gold and PEDOT-modied electrodes implanted in the hippocampus of a mouse. Both gold spontaneous single unit activity (left) and stimulated LFPs (right) are displayed.

Figure 4

 4 Figure 4.8: Soaking test experimental setup -WE/CE/RE: Working/Counter/Reference Electrode

Solution N aCl N aHCO 3

 3 KCl M gSO 4 N aH 2 P O 4 CaCl 2 Glucose Concentration (mM)

Figure 4 . 9 :

 49 Figure 4.9: (a) Box plot showing the impedance evolution measured at 1kHz at dierent soaking times in ACSF for the 16 PEDOT-modied electrodes (b) Example of EIS evolution of one PEDOT-modied electrode soaked in ACSF during 29 weeks.

Figure 4 .

 4 Figure 4.9a displays the statistics of impedance evolution of all the electrodes soaked during this study. To allow comparison, impedance is taken at 1 kHz, frequency often used in electrophysiology as a reference because it corresponds to the frequency of neuron spikes. The graph shows that the impedance median of all 16 electrodes lies between 10 and 25 kΩ at 1 kHz, and 90% of all impedance values are below 100 kΩ at all times. Most of the outliers, represented at outside crosses on Figure 4.9a, are above one order of magnitude compared to the impedance median, and are probably linked to awed electrodes after Clean Room fabrication (problems in passivation etching, for instance) or faulty connectic or packaging.

  Scanning microscope images of the four implants involved in the trial, before and after soaking in ACSF are shown in Figure4.10. Over the four devices, none showed any signicant sign of Parylene delamination or swelling. The passivation layer remains well-retained on the surface, and the device sides appear unaltered. The presence of few aggregates on the Parylene surface is most likely due to precipitations of ACSF ions on the devices. Besides, from a macroscopic point of view, the PEDOT coatings seem indierent to this wet ageing experiment. Indeed, it appears that some residues rst found on the electrode sites might have dissolved in the soaking medium, which could contribute to the slight reduction in impedance value observed with the EIS measurement.

Figure 4 .

 4 Figure 4.10: SEM images and zoom of a Parylene C probe with four PEDOT electrodes, taken before and after 29 weeks of soaking treatment in Articial Cerebral Spinal Fluid (ACSF) (black scale bar is 50µm, white scale bar is 10μm)

4. 4 . 1

 41 The hippocampus as a cognitive map

Figure 4 . 11 :

 411 Figure 4.11: Spatial navigation and hippocampal representation (a) Conceptual model of the hippocampal representation according to the cognitive mapping hypothesis (from [284]) and (b) Schematic representation of spatial navigation coding by place cells. When the rat is exploring along the dotted track, the hippocampus oscillates (blue trace) and place cells are successively activated as the rat enters their respective place elds (red, yellow and green circles) (courtesy of M. Zugaro, [285]) (c) The hippocampal signal recorded from the CA1 region when the animal is actively exploring its environment is mainly composed of a Theta oscillation (7-12 Hz) in which are nested Gamma oscillations (40-80 Hz).

Figure 4 . 12 :

 412 Figure 4.12: Schematic representation of a GEN2 Parylene C-based neural probe with three PEDOT electrodes and one gold electrode as control. The scale bar for the microscopic image is 200µm. The wireless packaging developped in Chapter 2 encloses a reference and ground pin connected to a microscrew.

Figure 4 . 13 :

 413 Figure 4.13: Frame of implantation of a GEN2 Parylene probe in the hippocampus of a mouse. 1. The mouse is anaesthetized and set to a stereotaxic apparatus. 2. Coordinates are taken for implantation in the hippocampus, and the craniotomy is performed. 3. The GEN2 probe is mounted to the stereotaxic apparatus and the shank is inserted in the brain. 4. The rst level of dental ciment is mounted (white, Superbond, Sun Medical) to maintain the probe in position, and the reference electrode is screwed to the occipital bone. 5. The overall probe and reference electrode are covered in a second dental ciment (red, Duralay, Edisco) to maintain everything in position, with the out pins remaining at air. The skin is then sutured around the ciment. A picture of the mouse with an implanted probe is represented on the far right hand side.

Figure 4 .

 4 Figure 4.14: Non-coated Parylene shank buckling on the brain surface

Figure 4 . 15 :

 415 Figure 4.15: Representation of a wireless recording session. The mouse is connected to an acquisition headstage and battery (blue) upon its prosthesis output (red), and the brain activity is recorded wirelessly while the mouse discovers a new environment (called acquisition cage). The data is collected with an adequate software and brain signals can be saved and treated.

Figure 4 .

 4 Figure 4.16: (a) Temporal and (b) spectral representation of a signal recorded from a mouse hippocampus while the animal is in active exploratory behaviour. Note that most of the signal power is comprised in the Theta range (7-12 Hz), therefore is considered as the meaningful signal amplitude for the rest of data treatment in this chronic recording trial.

Table 4 . 2 :

 42 Experiment plan of this ve mice chronic recording trial. Gold and PEDOT-modied electrodes are marked. We distinguish the electrodes (A, B, C, D) that are functioning at the considered timing, not functioning or that have never functioned from the the rst recording session.

Figure 4 . 17 :

 417 Figure 4.17: Examples of a stable and an unstable electrode implanted in the same mouse.

  at 6 dierent intervals of 10 seconds in each recording sessions, and the results are plotted on Figure 4.19.Therefore, the quantitative values of SNR presented here are objective, but they allow us to observe qualitatively the evolution of device electrical stability over 6 months.

Figure 4 . 18 :

 418 Figure 4.18: Demonstration of the signal processing for SNR measurement using Spike2 data processing tool. The smoothed signal is used to dene signal amplitude, and the noise is the dierence between the raw and the smoothed signal.

Figure 4 . 19 :

 419 Figure 4.19: Example of evolution of electrode signal quality in vivo over time.

  Figure 4.20 shows the example of electrodes from Mouse B. After 30 weeks implantation in a mouse brain, the impedance of both gold and PEDOT-modied electrodes increased by roughly a factor of two at 1 kHz. This raise in impedance is not signicant in the context of electrophysiology, and could be linked to a variety of parameters, among which a slight aggregation of biological tissues on electrode sites, or a mechanical trauma on both electrode and connectic parts when the probe is removed from the mouse brain. Besides, SEM images of the implanted devices show no apparent sign of either Parylene delamination or PEDOT alteration, consistent with the in vitro results and the SNR measurements.

Figure 4 .

 4 Figure 4.20: EIS evolution before and after 30 weeks implantation in mouse brain for a gold and three PEDOT electrodes, along with a representative SEM image of a PEDOT-modied electrode and implant before and after implantation (white scale bar is 10 µm, black scale bar is 100 µm)

5. 3 . 1

 31 Introduction to optogenetics

Figure 5 . 1 :

 51 Figure 5.1: Optogenetics was highlighted in many high-IF scientic journals as a revolutionizing technology

Figure 5 . 2 :

 52 Figure 5.2: Operating principle of optogenetics. A gene coding for a light-sensitive protein (e.g.

Figure 5 . 3 :

 53 Figure 5.3: Examples of commonly used single unit and LFP recording devices and their optogenetics-compatible modications. (a) Silica-encapsulated tungsten electrodes (Thomas RecordingGmbH). (b) Tungsten microwire array (Tucker-Davis Technologies Inc.). (c) Tetrode microwire bundle (Stanford University). (d) Silicon Utah multielectrode array (Blackrock Inc.). (e) Silicon multitrode probes (Neuronexus Inc.). (f ) An optrode consisting of a silica ber and a tungsten electrode[START_REF] Gradinaru | Targeting and readout strategies for fast optical neural control in vitro and in vivo[END_REF] (g) Metal microwire array equipped with a silica ber[START_REF] Kravitz | Optogenetic identication of striatal projection neuron subtypes during in vivo recordings[END_REF] (h) Optetrode combining four tetrode bundles and an optical ber[START_REF] Anikeeva | Optetrode: a multichannel readout for optogenetic control in freely moving mice[END_REF]. The inset shows the cross-section schematic. (i) Utah array incorporating a tapered silica ber[START_REF] Zhang | Integrated device for optical stimulation and spatiotemporal electrical recording of neural activity in light-sensitized brain tissue[END_REF](j) Silicon multitrode probes integrated with tapered silica

Figure 5 . 4 :

 54 Figure 5.4: Examples of optogenetic engineered tools featuring µLEDs. (a) Injectable, cellularscale thin (~8µm) neural probe with ILED integration [307]. (b) Fabricated Opto-μECoG array with µLED mounted on a Parylene substrate [308]. (c) Conceptual diagram of a Parylene neural probe with OLED integration [155].

Figure 5 . 5 :

 55 Figure 5.5: Representation of our Parylene C tetrode featuring a blue OLED

Figure 5 . 6 :

 56 Figure 5.6: Fabrication steps of an OLED on a Parylene substrate. (a) 20 µm Parylene C deposition on a silicon substrate (b) ITO conformal deposition (c) ITO etching (d) Organic semiconductors evaporation (e) Aluminium evaporation (f ) Encapsulation layer (g) Device release. When a voltage is applied between the anode (ITO) and the cathode (aluminium), a blue light is emitted between these layers and is visible from the backside of the device.

Figure of meritΦ T C (kΩ - 1 ) @ 470 nm 22 1 Table 5 . 1 : 10 R

 115110 Figure of meritΦ T C (kΩ -1 ) @ 470 nm

Figure 5 . 7 :

 57 Figure 5.7: (a) Photograph of a shining OLED on a Parylene C substrate (b) Electrical characterization of dierent OLEDs throughout the wafer, showing rather satisfying process homogeneity.

  Photoresist AZ nLOF 2535 5µm

Figure C. 2 :

 2 Figure C.2: Représentation du placement de dispositifs d'enregistrement et des types de signaux enregistrés (adapté depuis [5, 6])

Figure C. 3 :

 3 Figure C.3: Électrodes intracorticales communes. (a) Le cône neurotrophique (ou pipette de verre) (b) Un l de tungstène passivé avec de la gélatine (c) Une tétrode (quatre ls de tungstène passivés enrobés ensemble) (d) la Utah array (10x10 électrodes de silicium avec une pointe métallique) (e) La Michigan probe (implant en silicium avec sur chaque branche, plusieurs pistes métalliques d'enregistrement).

Figure C. 4 :

 4 Figure C.4: Schéma d'une Interface Cerveau Machine

Figure C. 5 :

 5 Figure C.5: Schéma de la réaction gliale sur le long terme

Figure C. 6 :

 6 Figure C.6: Implants à base de polyimide[START_REF] Mercanzini | Demonstration of cortical recording using novel exible polymer neural probes[END_REF], SU-8[START_REF] Altuna | SU-8 based microprobes for simultaneous neural depth recording and drug delivery in the brain[END_REF] et Parylène C[START_REF] Kim | 3D Parylene sheath neural probe for chronic recordings[END_REF]

Figure C. 7 :Figure C. 8 :

 78 Figure C.7: (a) L'eet de ambement d'un implant en polyimide sur la surface d'un cerveau de poisson, puis l'insertion réussie grâce à un lm de soie biodégradable intégré à l'implant [114] (b) Insertion d'un implant en Parylène C grâce à une tige rétractable[START_REF] Kim | 3D Parylene sheath neural probe for chronic recordings[END_REF] (c) Un implant semi-rigide avec des zones en silicium et des zones en Parylène C, an que l'implant possède susamment de rigidité pour pénétrer dans l'organisme.[START_REF] Kim | A hybrid silicon-parylene neural probe with locally exible regions[END_REF]

Figure C. 10 :

 10 Figure C.10: (a) Première génération d'implant (b) Deuxième génération d'implants (c) Schéma de packaging développé pour être compatible avec le protocole d'expérimentation in vivo.

Figure C. 11 :

 11 Figure C.11: Résultats des tests MTT et en uorescence, montrant la viabilité des cellules SH-SY5Ysur les implants Parylène.

  Ainsi, un implant en Parylène de faible dimensions a du mal à pénétrer dans le cerveau. Comme nous l'avons vu au Chapitre 1, plusieurs méthodes ont été mises en ÷uvre pour pallier à ce problème. Dans les prochaines sections, on se propose d'étudier deux stratégies d'insertion d'implants exibles : -l'insertion de notre implant développé en Chapitre 2 dans l'hippocampe d'une souris à l'aide d'un lm de soie biorésorbable -l'insertion d'un implant en Parylène avec plusieurs tiges dans l'hippocampe d'un rat

Figure C. 14 :

 14 Figure C.14: (a) Schéma du protocole de dépôt de broïne de soie dans un moule PDMS (b) Schéma de l'implant avec un lm de soie intégré à l'arrière (c) Image au microscope de la pointe de l'implant en Parylène avec un lm de soie (teint en bleu) (d) Test de compression standard d'un implant avec et sans lm de soie en face arrière (e) Test de dissolution de la soie dans un environnement protéolytique (PXIV) selon le temps de water-annealing (f ) Images d'insertion d'un implant Parylène avec et sans soie dans un gel agarose à 0.6% représentant le cerveau (g) Insertion réussi de l'implant dans un cerveau de souris.

Figure C. 15

 15 Figure C.15: (a) Neuroprothèse hippocampale en Parylène développée dans l'équipe d'Ellis Meng (b) Modèles d'implants 2D avec resp. 1, 4 et 8 tiges (c) Modèle d'implant 3D avec 3x8 tiges (d) Mesure des forces de cisaillement pendant la descente des implants dans des gels agarose en fonction du nombre de tiges (e) Idem avec les forces d'insertion dans les gels.

Figure C. 16 :

 16 Figure C.16: Modes d'échec à l'échelle cellulaire et tissulaire des implants intracorticaux sur le long terme ([6])

Figure C. 17

 17 Figure C.17: (a) Schéma du protocole expérimental (b) Résultats d'enregistrements spontanés et stimulés, pour une électrode en or et une électrode en PEDOT.

Figure C. 18 :

 18 Figure C.18: (a) Schéma expérimental des tests de vieillissement (b) Évolution de l'impédance des électrodes à 1kHz au cours du temps, et images MEB associées

Figure C. 19 :

 19 Figure C.19: (a) Schéma représentatif du protocole d'implantation in vivo sur souris et un échantillon de signal typique enregistré pendant les séances (b) Évolution de l'amplitude des ondes Theta enregistrées en fonction du temps, pour une électrode stable et une instable (c) EIS avant et après implantation pendant 30 semaines

  , D. Bourrier, P. Dubreuil, E. Descamps and C. Bergaud, Plasma etching of thick Parylene C for fabrication of biocompatible electrodes, accepted for oral presentation. neuroprothèses sont un domaine de recherche visant à restaurer les fonctions de personnes atteintes de déciences sensorielles ou motrices. Les implants neuraux assurent une communication bidirectionnelle entre le cerveau et les ordinateurs. Ils permettent par exemple de favoriser la communication et la mobilité des personnes présentant une décience motrice grave, rétablir la perception sensorielle (vision,

	List of acronyms Résumé
	ACF Conception et caractérisation de microélectrodes exibles pour le développe-Anisotropic Conductive Film NGF Nerve Growth Factor ment de neuroprothèses implantables
	ACSF	Articial CerebroSpinal Fluid	NI	Neural Interfaces
	AFM	Atomic Force Miroscopy	OECT	Organic ElectroChemical Transistors
	ALS	Amytrophic Lateral Sclerosis	OLED	Organic Light-Emitting Diode
	AP	Action Potentials	PaC	Parylene C
	BCB	Benzocyclobutene	PBS	Phosphate Buer Saline
	BCI	Brain Computer Interface	PDMS	PolyDiMethylSiloxane
	BDNF	Brain-Derived Neurotrophic Factor	PEDOT	Poly(3,4-EthyleneDiOxyThiophene)
	CA	Cornu Ammonis	PEEK	PolyEther Ether Ketone
	ChABC	Chondrointinase ABC	PEG	PolyEthylene Glycol
	CNS	Central Nervous System	PLGA	Poly(Lactic-co-Glycolic Acid)
	CRCA	Research Center of Animal Cognition	PR	Photoresist
	CVD	Chemical Vapor Deposition	PSS	PolyStyrene Sulfonate
	DBS	Deep Brain Stimulation	PVAc	Poly(Vinyl Acetate) composite
	DEX	Dexamethasone	PXIV	Protease XIV
	DG	Dentate Gyrus	RIE	Reactive Ion Etching
	DMEM	Dulbecco's Modied Eagle Medium	SC	Semiconductor
	DMSO	DiMethyl SulfOxide	SCP	Slow Cortical Potentials
	ECM	Extracellular Matrix	SEM	Scanning Electron Microscopy
	ECoG	Electrocorticography	SF	Silk Fibroin
	EEG	Electroencephalography	SNR	Signal-to-Noise Ratio
	ERP	Event-Related Potentials	UEA	Utah Electrode Array
	FBR	Foreign Body Response	USC	University of Southern California
	FDA	Food and Drug Administration	USP	United States Pharmacopeia
	FET	Field Eect Transistors
	FP	Field Potentials
	GFAP	Glial Fibrillary Acidic Protein
	HS	Headstage
	ICP	Inductive Coupling Plasma
	ISO	International Organization for
		Standardization
	ITO	Indium Tin Oxide
	LFP	Local Field Potentials
	LIS	Locked-In Syndrome
	MEA	MicroElectrode Array
	MEMS	MicroElectroMechanical Systems
	MTT	3-(4,5-dimethylthiazol-2-yl)-2,5-
		diphenyltetrazolium
		bromide

Les audition) et réduire des symptômes neurodégénératifs (Parkinson).

Les dernières avancées technologiques et la meilleure compréhension des facteurs déclenchant les réactions inammatoires permettent d'envisager des implants corticaux chroniques ables. Les implants traditionnels, basés sur des matériaux rigides comme le silicium ou le tungstène, sont souvent associés à une réaction immunitaire importante, du fait de leur pauvre biocompatibilité et du stress qu'ils induisent sur les tissus environnants. En ce sens, les implants exibles, basés sur des biomatériaux souples, sont de plus en plus étudiés. Le substrat s'adapte aux micromouvements du cerveau (respiration, pulsation cardiaque) et de se fait promouvoit un meilleur contact tout en diminuant la réaction inammatoire.

  1.2.3.3 Silicon-based Multielectrode arrays (MEAs)

  3.1. INSERTION STRATEGIES FOR FLEXIBLE NEURAL PROBES IN THE LITTERATURE 73

	Biotic eect and drug	integration	PLGA hydrogel:	integrable NGF in PEG	microspheres -DEX	integrated in PLGA	nanoparticles
	Degradation	rate /	Sti-to-soft rate	Weeks to	months	(tunable)[214,	215, 217]
	Main biomedical applications	other than neural interfaces		Resorbable sutures, bone	repair, vascular graft, drug	delivery[216]
	Biocompa-	tibility				++ [215]
	Young's	modulus			1.3-2 GPa	[215, 214]
		Polymer			PLGA	[213, 214, 83]	(=PLA+PGA)
								Bioresorbable	polymers

Table 3 . 3

 33

					Single shaft or
				Silicon Glass	1x10 array (pitch 550/1100 µm)	100 × 120 175 × 500	17 49.3	3 -6
				Polyimide	Single shaft	10 × 460	66
				Tungstene		Ø135	15 -20
	Andrei [201, 200]	10	100	Silicon	Single shafts	[200 -400] ×[50 -150]	10 -50	6
	Fekete [270]	20	175	Silicon	Single shaft	[200 -400] ×400	0 -90	∼ 2 -3
	This work	100		Parylene	Single shaft and up to 3x8 array (pitch 250 µm)	20× [70 -100]	47	2.8 -5

: Non-exhaustive comparison table of in vitro or in vivo neural probe insertion parameters inventoried from the literature Besides these theoretical considerations, implant type (single shank or 3D arrays) and material (silicon or soft polymer), insertion depth (surface or deep) and procedure (animal type, age or brain region

Table 4 .

 4 1: ACSF composition used in this study 4.3.2 Evolution of electrical properties

Au cours de cette thèse, nous avons conçu et fabriqué un implant exible à base de Parylène C, polymère souple de plus haute classe de biocompatibilité atteinte par la législation américaine (USP Classe VI), sur lequel des électrodes en or sont positionnées. Divers procédés de la microélectronique, comme la photolithographie et la gravure plasma, utilisés communément pour le développement de microsystèmes en métal ou semi-conducteurs, ont été adaptés à la structuration d'implants en Parylène C. Par le biais de la culture cellulaire in vitro, nous avons montré que des cellules neuronales dérivées se diérenciaient correctement sur les implants, validant ainsi la biocompatibilité des dispositifs.Cependant, ces nouveaux implants ont tendance à se courber à la surface du cerveau lors de leur insertion, empêchant le bon déroulement de l'implantation. Nous proposons ici une méthode basée sur l'intégration d'un lm biorésorbable à l'arrière de l'implant. Ce lm rigide permet d'assurer la pénétration de l'implant dans les tissus cérébraux, avant de se dissoudre de façon inoensive dans l'organisme. Le lm est réalisé en broïne de soie, extrait des cocons de vers à soie. Ce matériau, plus résistant que le Kevlar, est utilisé depuis des siècles comme ls de suture biodégradable. La mise au point de l'extraction de la broïne de soie et sa structuration sur l'implant à l'aide d'un moule en polymère, ainsi que l'optimisation de la méthode de dépôt permet l'obtention d'une couche de soie en forme de gouttière, ce qui facilite l'insertion tout en limitant les contraintes et pressions indésirables lors de l'insertion. Nous avons montré à travers une série de test in vitro dans des gels et in vivo sur souris, que la soie augmentait par 100 la rigidité de l'implant et pouvait se résorber à taux accordable dans l'organisme.Un aspect primordial des implants neuraux concerne leur tenue et leur abilité sur le long terme.Si les implants traditionnels en silicium sont matière à de nombreuses études sur le sujet, les implants en polymères souples ne se sont développés que récemment et ne bénécient pas encore du même recul.Nous proposons une étude préliminaire in vitro dans du liquide cérébro-spinal articiel et in vivo sur souris permettant de mettre en évidence l'augmentation de la durée de vie de nos implants. Les résultats ont montré qu'au bout de six mois, les dispositifs ne présentent pas de signe de délamination, corrosion ou gonement, ce qui se caractérise par la stabilité des propriétés électriques des électrodes.En conclusion, les implants conçus au cours de cette thèse présentent des caractéristiques prometteuses pour le développement de neuroprothèses implantables exibles ables sur le long terme.

Acknowledgements

First of all, I would like to thank the CNRS, for both nancing my PhD and providing me with the greatest lab someone could hope for.

I am very grateful to Emeline Descamps and Christian Bergaud, who supervised me during this PhD.

They have been great mentors to me, pushing me forward and allowing me to be independent. Emeline, you've been there side by side with me from the internship, shared small successes and disappointments, thank you for your optimism and your every-day commitment. Christian, you've taught me to look at the bigger picture and gain hindsight, thank you for challenging me, and for always keeping an open door whenever I needed it. I have learned from both your skills, your expertise, your perspective, and I am glad we managed to work together. Thank you also, for your amazing support during my pregnancy.

Awards and distinctions

L'Oréal-UNESCO Award "For Women in Science" This 15,000¿ prize was awarded in 2016 to 30 young female researchers among 1000 applicants for their research during a doctorate or post-doctorate position. It is dedicated to assist laureates in their research and their career. Besides, laureates

Ethical considerations

Today, the exploration of cerebral mecanisms involve an intrusion in memory, thoughts, emotions and behaviours. The possible interventions on the nervous system are multiplying, and some questions need to be asked: what do we read, what do we track down, what do we heal? Are we risking to modify the human race? We need to be aware that these new brain technologies are bearers of hope, but also of fear of manipulation, of free-will breach and violation of intimacy.

The emergence of neurotechnologies gives rise to questionings, concerns and a real need to debate on the impact of these technologies on our society.

An intrusion in our minds

The brain is a cradle of thoughts, perception and behaviour, as well as the expression of our personal identity. The new possibilities in terms of behaviour modication, such as appetite, sleep, sexuality, mood or memory are expanding rapidly. These modications can be happen whether or not they were intended. For instance, almost 350,000 patients are currently beneting from a Deep Brain Stimulation device, which highly improves their quality of life. However, behavioural side eets recount hallucinations, compulsive gambling, hypersexuality and depression. The presence of an intruder into our brains is not and shall not be treated lightly.

If our behaviours and decisions can be described as biological mecanisms, if these emerging neurotechnologies are able to predict cognitive deciencies and record electrical pathways from willing or unwilling behaviours, what will be their purpose and their limits? Imaging a functioning brain is becoming more and more accurate, but what are these images, and what do they represent ? If all these technologies are allowing us to push forward high-resoluted therapies and diagnostics, their abuse or misappropriations are to be carefully foreseen and evaluated. Control or alteration over a region of the brain could be achieve through electrical or chemical means, intimacy might be breached through the recording or imaging of brain areas. Finally, it appears obvious that there should be an enlightened consent for a treatment or an image, but this consent is not enough, for public power is responsible for the integrity and inviolability of the human body.

Enhancement of normal functions

There is a real technological greed that arises in this century, where new technologies are fancied not for their use, but for what they represent : a jewel of technological performance. Combined with the wish of self-control and self-knowledge, neurotechnologies are becoming more and more attractive.

Performance enhancement does not seem blameworthy if its purpose is to assist communication and individual autonomy, for instance help amputees to walk again. However, assisting technologies are more and more becoming enhancement technologies, where the primary function of a neurotechnology is diverted into an advanced function. Amputees athletes are nowadays faster than two-legged runners, because at some tilting point, a leg prosthesis was able to outperform human function. In the military eld, neural implants might be used to control an exoskeleton with extraordinary strength or to communicate speechlessly via the emission/reception of brain waves. Incidentally, the Army research fundings