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![2: Mouse brain neurons stained with brainbow technique (adapted from [7, 8]).](profile/Valentina-Castagnola/publication/278827467/figure/fig2/AS:613909576822785@1523378857029/Mouse-brain-neurons-stained-with-brainbow-technique-adapted-from-7-8.png)
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Implantable neural prosthetics devices offer, nowadays, a promising opportunity for the restoration of lost functions in patients affected by brain or spinal cord injury, by providing the brain with a non-muscular channel able to link machines to the nervous system. The long-term reliability of these devices constituted by implantable electrodes ha...
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... wafer surface was previously patterned with AZ-nLOF 2035 photoresist (Fig.2.12.g), then a layer of 100 nm of Nickel (Ni) was deposited by electron beam physical vapour deposition on top of it and a second lift-off procedure was performed in order to use the Ni layer as a hard mask for the second etching process ...
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... wafer surface was previously patterned with AZ-nLOF 2035 photoresist (Fig.2.12.g), then a layer of 100 nm of Nickel (Ni) was deposited by electron beam physical vapour deposition on top of it and a second lift-off procedure was performed in order to use the Ni layer as a hard mask for the second etching process ...
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... smallest electrochemically active area dimensions (10 µm in diameter) were also verified With ultramicroelectrode, a steady state is readily obtained, even without stirring. The same LSV was performed in the electrolyte in order to subtract this capacitive component (red line in Fig.3.12) to the measured curve and obtain the faradic component (black line in Fig.3.12). The radius of the electrode is determined using the current intensity limit, that corresponds to the diffusion plateau (see eq. 11 in Appendix B [41], and knowing the initial concentration (10 mM), the Faraday constant and the number of exchanged electrons (1 electron for this reaction), it was possible to estimate a value of the radius of 4.5±1 µm. ...
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... smallest electrochemically active area dimensions (10 µm in diameter) were also verified With ultramicroelectrode, a steady state is readily obtained, even without stirring. The same LSV was performed in the electrolyte in order to subtract this capacitive component (red line in Fig.3.12) to the measured curve and obtain the faradic component (black line in Fig.3.12). The radius of the electrode is determined using the current intensity limit, that corresponds to the diffusion plateau (see eq. 11 in Appendix B [41], and knowing the initial concentration (10 mM), the Faraday constant and the number of exchanged electrons (1 electron for this reaction), it was possible to estimate a value of the radius of 4.5±1 µm. ...
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... the whole process can be divided in four main steps. A standard silicon wafer was used as a sub- strate for the whole process. Parylene-C (Fig.2.12.a) was deposited through ther- mally activated CVD at the pyrolysis tem- perature of 700 • C. The thickness of the deposition was measured each time using a standard Profilometer (KLA Tencor) and it was found to be proportional to the mass of precursor dimer charged in the sublimation chamber (see ...
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... the Au electrode patterning, a thin layer of PXC (about 1 µm obtained from 1.6 g of precursor) is deposited on top of the wafer as a passivation layer ...
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... etching rate of PXC with these parameters was found to be 350 nm/min, whereas the etching rate of AZ-ECI photoresist is about 450 nm/min ...
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... electrode surfaces and the contacts were opened by dry etching of the PXC using a photoresist mask (AZ-ECI 3027, MicroChemicals) with a thickness of about 2 µm ...
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... of the first silicon-based multielectrode arrays was made by Wise, Starr and Angell [45,46]. These structures have then evolved in what we know nowadays as Michigan array (Fig.1.12.a). In this array several microelectrode sites are patterned on each shank of the structure, providing higher density of sensors while reducing the displaced tissue when compared to a microwire bundle implant [47]. [43]). b) The Utah array and the insulation coated electrodes with exposed platinum tips (from ...
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... to the process carried out on polyimide, Au circular electrodes were patterned on the PXC surface thanks to a metallization followed by a lift-off process ...
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... well-known example is represented by the Utah array ( Fig.1.12.b). This three-dimensional electrode array consists in 100 conductive sharpened silicon needles, each of them electrically isolated from its neighbours in which the tip is coated with platinum and the shanks are insulated with polyimide. Electrical contact is made from the back side of the structure using insulated gold wire [49,50]. Researchers now routinely employ multiple single microelectrodes aligned into arrays to provide ever-increasing numbers of electrode sites in one device. Some of these devices have positionable electrodes, while others have modified single electrodes (with larger site sizes and/or reduced impedances) which are capable of recording neural activity without precise positioning. These devices can remain functional upon implant for few months, but the same individual neurons cannot be "tracked" longer than about six weeks. Like the first generation of devices, these intracortical interfaces can remain into the brain for extended time periods, but recording quality and electrode yield typically diminish with time. The surface area substrate/encapsulation material is much larger compared to the areas of the active electrode sites; therefore, these materials have an important role in long-term stability and functioning of neural implants and they need to be biocompatible, biostable and to exhibit good dielectric ...
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... impedance spectroscopy Normally the elctrode/solution interface is represented through an equivalent circuit called "Randless circuit" (Fig.12), which is basically constituted by a capacitor, C ...
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... plasma oxygen etching was performed as described above during a time duration (6 cycles of 15 min) long enough to allow the cutting of the 24 µm-thick layer of PXC. During the physical etching it is important to perform short cycles in order to avoid overheating of the sample and prevent Nickel cracking. The Ni layer was then chemically etched ( Fig.2.12.i). Finally the PXC on the edge of the Si wafer was scratched with tweezers and the electrode structures were easily peeled off from the Si surface, or released in a water bath ( Fig.2.13), keeping their planar shape. The details of the whole process are given in Table 2.2, whit the corresponding steps described in ...
Citations
... A wide range of conductive materials have been used as electrode interface materials including gold [8], platinum [9], iridium [10], titanium nitride [11], platinum-iridium alloys [12]. Recently conductive polymers such as PEDOT [13], poly(ethylenedioxythiophene):poly (styrenesulfonate) (PEDOT:PSS) [6], [14], polypyrrole (PPY)-modified CNTs (CNTs-PPY) [15], PEDOT/carbon nanotubes [16]- [17], polypyrrole/graphene oxide [18] were also used. ...
... The application of microelectrodes provides better spatial resolution with high density of recording sites but as the site diameter decrease the impedance increases resulting in low signal to noise ratio and poor recording quality. Recently, several groups developed or used various deposition techniques to obtain high surface area electrode 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60 A c c e p t e d M a n u s c r i p t 15 interfaces. PEDOT:PSS coatings [43]- [46] was utilized due to its mixed electronic and ionic conductivity and high ionic mobility. ...
Objective:
Intracranial EEG (iEEG) or micro-electrocorticography (µECoG) microelectrodes offer high spatial resolution in recordings of neuronal activity from the exposed brain surface. Reliability of dielectric substrates and conductive materials of these devices are under intensive research in terms functional stability in biological environments.
Approach. The aim of our study is to investigate the stability of electroplated platinum recording sites on 16-channel, 8 micron thick, polyimide based, flexible µECoG arrays implanted underneath the skull of rats. Scanning electron microscopy and electrochemical impedance spectroscopy was used to reveal changes in either surface morphology or interfacial characteristics. The effect of improved surface area (roughness factor = 23±0.12) on in vivo recording capability was characterized in both acute and chronic experiments.
Main results. Besides the expected reduction in thermal noise and enhancement in signal-to-noise ratio (up to 39.8), a slight increase in the electrical impedance of individual sites was observed, as a result of changes in the measured interfacial capacitance. In this paper, we also present technology processes and protocols in details to use such implants without crack formation of the porous platinum surfaces.
Significance. Our findings imply that black-platinum coating deposited on the recording sites of flexible microelectrodes (20 microns in diameter) provides a stable interface between tissue and device.
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... However, implantation of soft and flexible structures into the brain is challenging as the precision and depth of implantation is compromised. Attempts to overcome this problem are done with the construction of various insertion aids in the forms of removable stiff-backbone stiffeners, additional layers of dissolvable materials or by piercing the tissue with other instruments prior to the implant placement (Takeuchi et al., 2005;Felix et al., 2013;Castagnola, 2014;Barz et al., 2015). ...
... Difficulties accounted during deposition of metals onto polymers are poor adhesion and mismatch in thermal and mechanical properties, causing metal films to crack when probes are subjected to high temperatures or significant strain (McClain et al., 2011). Both can be improved with the use of adhesion promoting layers and pre-roughening of polymer surface (Nakamura et al., 1996;Mercanzini et al., 2008;Castagnola, 2014). Rather than using physical deposition methods, metals can be applied a foils, which are integrated within a polymer sandwich by lamination. ...
Implantable neural interfaces for central nervous system research have been designed with wire, polymer, or micromachining technologies over the past 70 years. Research on biocompatible materials, ideal probe shapes, and insertion methods has resulted in building more and more capable neural interfaces. Although the trend is promising, the long-term reliability of such devices has not yet met the required criteria for chronic human application. The performance of neural interfaces in chronic settings often degrades due to foreign body response to the implant that is initiated by the surgical procedure, and related to the probe structure, and material properties used in fabricating the neural interface. In this review, we identify the key requirements for neural interfaces for intracortical recording, describe the three different types of probes—microwire, micromachined, and polymer-based probes; their materials, fabrication methods, and discuss their characteristics and related challenges.
The long-term reliability of the neural electrode is closely related to its implantation behavior. In order to realize the quantitative research of the implantation behavior in a low-cost and accurate way, a refined brain model containing meninges is proposed. First, the expected simulation material was selected through measuring the elastic modulus based on the method of atomic force microscope indentation technique. As a result, the 2% (mass fraction) agarose gel simulated the gray and white matter, the 7: 1 (volume ratio) polydimethylsiloxane (PDMS) sheet simulated the pia mater, and the polyvinyl chloride (PVC) film simulated the dura mater. Second, based on designing a three-layer structure mold, the brain model was prepared by inverted pouring to realize a flat implantation surface. Finally, the simulation behavior of the brain model was investigated with the rat brain as a reference. For mechanical behavior of implantation, the implantation force experienced two peaks both in the brain model and the rat brain, maximum values of which were 10.17 mN and 7.69 mN respectively. The larger implantation force in the brain model will increase the strength requirement for the electrode, but reduce the risk of buckling of that in practical application. For humoral dissolution behavior, the dissolution rates of the polyethylene glycol (PEG) coating of the electrode in the brain model and rat brain were 7 000 µm3/s and 5 600 µm3/s, respectively. The faster dissolution rate in the brain model will cause the larger thickness of the coating design but provide sufficient implantable time in practical application. The establishment of the brain model and the research of its simulated behavior are beneficial to the size design of the electrode substrate and coating, and research of the implantation mechanism, and further increase the functional life of the electrode.
