A novel neural prosthesis providing long-term electrocorticography recording and cortical stimulation for epilepsy and brain-computer interface

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  • 1 AB MEDICA S.p.a., Cerro Maggiore(MI), Milan;
  • 2 Computer Science Division, School of Science and Technology, University of Camerino, Italy; and
  • 3 Biomedical Research Center, Polygone Scientifique Grenoble (CLINATEC Campus), University of Grenoble Alpes, Grenoble, France
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OBJECTIVE

Wireless technology is a novel tool for the transmission of cortical signals. Wireless electrocorticography (ECoG) aims to improve the safety and diagnostic gain of procedures requiring invasive localization of seizure foci and also to provide long-term recording of brain activity for brain-computer interfaces (BCIs). However, no wireless devices aimed at these clinical applications are currently available. The authors present the application of a fully implantable and externally rechargeable neural prosthesis providing wireless ECoG recording and direct cortical stimulation (DCS). Prolonged wireless ECoG monitoring was tested in nonhuman primates by using a custom-made device (the ECoG implantable wireless 16-electrode [ECOGIW-16E] device) containing a 16-contact subdural grid. This is a preliminary step toward large-scale, long-term wireless ECoG recording in humans.

METHODS

The authors implanted the ECOGIW-16E device over the left sensorimotor cortex of a nonhuman primate (Macaca fascicularis), recording ECoG signals over a time span of 6 months. Daily electrode impedances were measured, aiming to maintain the impedance values below a threshold of 100 KΩ. Brain mapping was obtained through wireless cortical stimulation at fixed intervals (1, 3, and 6 months). After 6 months, the device was removed. The authors analyzed cortical tissues by using conventional histological and immunohistological investigation to assess whether there was evidence of damage after the long-term implantation of the grid.

RESULTS

The implant was well tolerated; no neurological or behavioral consequences were reported in the monkey, which resumed his normal activities within a few hours of the procedure. The signal quality of wireless ECoG remained excellent over the 6-month observation period. Impedance values remained well below the threshold value; the average impedance per contact remains approximately 40 KΩ. Wireless cortical stimulation induced movements of the upper and lower limbs, and elicited fine movements of the digits as well. After the monkey was euthanized, the grid was found to be encapsulated by a newly formed dural sheet. The grid removal was performed easily, and no direct adhesions of the grid to the cortex were found. Conventional histological studies showed no cortical damage in the brain region covered by the grid, except for a single microscopic spot of cortical necrosis (not visible to the naked eye) in a region that had undergone repeated procedures of electrical stimulation. Immunohistological studies of the cortex underlying the grid showed a mild inflammatory process.

CONCLUSIONS

This preliminary experience in a nonhuman primate shows that a wireless neuroprosthesis, with related long-term ECoG recording (up to 6 months) and multiple DCSs, was tolerated without sequelae. The authors predict that epilepsy surgery could realize great benefit from this novel prosthesis, providing an extended time span for ECoG recording.

ABBREVIATIONS BCI = brain-computer interface; DCS = direct cortical stimulation; ECoG = electrocorticography; ECOGIW-16E = ECoG implantable wireless 16-electrode; EEG = electroencephalography; GFAP = glial fibrillary acidic protein; Iba-1 = ionized calcium-binding adapter molecule–1; ISO = International Organization for Standardization; MICS = Medical Implant Communication Service; NT = neoformed tissue; SSEP = somatosensory evoked potential; Vim = vimentin.

OBJECTIVE

Wireless technology is a novel tool for the transmission of cortical signals. Wireless electrocorticography (ECoG) aims to improve the safety and diagnostic gain of procedures requiring invasive localization of seizure foci and also to provide long-term recording of brain activity for brain-computer interfaces (BCIs). However, no wireless devices aimed at these clinical applications are currently available. The authors present the application of a fully implantable and externally rechargeable neural prosthesis providing wireless ECoG recording and direct cortical stimulation (DCS). Prolonged wireless ECoG monitoring was tested in nonhuman primates by using a custom-made device (the ECoG implantable wireless 16-electrode [ECOGIW-16E] device) containing a 16-contact subdural grid. This is a preliminary step toward large-scale, long-term wireless ECoG recording in humans.

METHODS

The authors implanted the ECOGIW-16E device over the left sensorimotor cortex of a nonhuman primate (Macaca fascicularis), recording ECoG signals over a time span of 6 months. Daily electrode impedances were measured, aiming to maintain the impedance values below a threshold of 100 KΩ. Brain mapping was obtained through wireless cortical stimulation at fixed intervals (1, 3, and 6 months). After 6 months, the device was removed. The authors analyzed cortical tissues by using conventional histological and immunohistological investigation to assess whether there was evidence of damage after the long-term implantation of the grid.

RESULTS

The implant was well tolerated; no neurological or behavioral consequences were reported in the monkey, which resumed his normal activities within a few hours of the procedure. The signal quality of wireless ECoG remained excellent over the 6-month observation period. Impedance values remained well below the threshold value; the average impedance per contact remains approximately 40 KΩ. Wireless cortical stimulation induced movements of the upper and lower limbs, and elicited fine movements of the digits as well. After the monkey was euthanized, the grid was found to be encapsulated by a newly formed dural sheet. The grid removal was performed easily, and no direct adhesions of the grid to the cortex were found. Conventional histological studies showed no cortical damage in the brain region covered by the grid, except for a single microscopic spot of cortical necrosis (not visible to the naked eye) in a region that had undergone repeated procedures of electrical stimulation. Immunohistological studies of the cortex underlying the grid showed a mild inflammatory process.

CONCLUSIONS

This preliminary experience in a nonhuman primate shows that a wireless neuroprosthesis, with related long-term ECoG recording (up to 6 months) and multiple DCSs, was tolerated without sequelae. The authors predict that epilepsy surgery could realize great benefit from this novel prosthesis, providing an extended time span for ECoG recording.

ABBREVIATIONS BCI = brain-computer interface; DCS = direct cortical stimulation; ECoG = electrocorticography; ECOGIW-16E = ECoG implantable wireless 16-electrode; EEG = electroencephalography; GFAP = glial fibrillary acidic protein; Iba-1 = ionized calcium-binding adapter molecule–1; ISO = International Organization for Standardization; MICS = Medical Implant Communication Service; NT = neoformed tissue; SSEP = somatosensory evoked potential; Vim = vimentin.

In Brief

This study was done to prove the feasibility of long-term ECoG recording. The results are very promising for the use of wireless long-term ECoG recording in humans that is aimed toward restorative neurosurgery and responsive stimulation treatments for epilepsy.

In patients with refractory epilepsy or brain tumors, electrocorticography (ECoG) and direct cortical stimulation (DCS) are the gold standard intraoperative techniques used to identify the tissue for resection, particularly with regard to neighboring eloquent cortex.6,7,9,16,19,30,38 DCS consists of application of an electrical stimulus directly to the cortex, to assess the contralateral muscle contraction or the related electromyography discharge, in anesthetized patients. Moreover, it can be used to generate transient language and behavioral effects in awake patients while they perform motor or cognitive tasks.19,30 DCS allows precise mapping of cortical organization in patients undergoing a resective procedure, providing a valuable instrument to spare the resection of eloquent cortex and to avoid catastrophic neurological sequelae.16,17,30 Intraoperative DCS is especially useful in neurosurgical procedures involving the resection of gliomas close to or involving eloquent cortex, because it allows the surgeon to maximize the extent of the resection and to prevent neurological injury.17

Electrocorticography offers the additional opportunity to record the remote effect of electrocortical stimulation without distortion within a limited distance of a few millimeters, and can provide further details about the functional reorganization caused by the individual’s brain pathology.16,19,26 Moreover, given the high spatial (approximately 1 mm) and temporal (within the timescale of neural activity) resolution, ECoG recording is essential to identify an ictal focus generating drug-refractory seizures and to guide the resection. ECoG signal analysis is a valuable emerging tool for both brain mapping10,14,15 and BCI,5,18 due to its high signal-to-noise ratio. It allows the examination of high-frequency bands (unavailable for scalp electroencephalography [EEG] recordings) and allows the use of spectral analysis ECoG recording from sensorimotor cortex.14 At the current state of art, this method provides the most effective tool for brain-computer interface (BCI) applications, i.e., to drive robotic prostheses and to enhance neurorehabilitation.21,27,28,35,42

The available commercial ECoG systems require cables connecting the subdural grid with an external recording system. The cables connecting the electrodes placed on the cortex with the external apparatus leave the skull through a subcutaneous channel, thus providing a path for CSF leakage and consequent meningeal infection.4 The patient undergoing invasive monitoring for epilepsy needs to be confined to a special recording room under careful but necessarily short observation. In cases in which no seizures are recorded, the wound is reopened and the grid removed. In essence, the use of cables connecting the epicortical grid with an external device is the most significant shortcoming of current ECoG techniques because it allows only short-term recordings, with the risk of meningeal and cerebral infection growing steadily day by day.20,31 Here we describe our second experience with a long-term implantable wireless ECoG device allowing prolonged (up to 6 months) wireless recording. A first pilot procedure in which a preliminary device was used is described in Piangerelli et al.,32 and medium-term electrophysiological testing performed 3 months after implantation is also described in Zippo et al.47

Methods

Surgical Procedure

In this study, we used a male macaque monkey (Macaca fascicularis) weighing 6.95 kg. The experimental protocol was approved by the regional committee (Cometh [Committee on Ethics] Grenoble) and registered to the national committee under the number 12/136 Clinatec-NTM-01. The experiment complied with the EU (European Union) directive approved on September 22, 2010 (2010/63/EU), on the care and use of laboratory animals. A 4.7-T MRI study was performed before the craniotomy (BioSpec; Bruker BioSpin) to provide image guidance for the placement of the grid above the sensorimotor cortex. The animal was anesthetized using a loading dose of 5 mg/kg xylazine and 20 mg/kg ketamine hydrochloride administered intramuscularly, and then a maintenance dose of 1.25 mg/kg and 5 mg/kg xylazine and ketamine, respectively. Physiological parameters were monitored by the veterinarian staff during the surgical procedure: heart rate, blood pressure, respiratory depth, and body temperature. Standard aseptic conditions were guaranteed during the surgical procedure. When deep anesthesia was achieved, the animal was secured to a stereotactic frame and a 3 × 2.5–cm square craniotomy was performed over the left sensorimotor cortex, aiming to center the grid placement over Brodmann area 4. The dura mater was cut in a Y fashion, and the flaps were retracted and sutured on the sides to expose the central sulcus and the surfaces of the primary motor (M1) and sensory (S1) cortex. Radiographic images were acquired during surgery to guide device placement. Electrophysiological confirmation of motor cortex localization was obtained through cortical stimulation with bipolar wand intraoperative neural monitoring (ISIS; INOMED Medizintechnik GmbH).

The grid was centered above the hand knob of the left motor cortex (Fig. 1), and then the dura was extended over the grid without suturing it, to avoid excessive pressure over the stem in the subdural penetration point. A silicone adhesive (KWIK-SIL—a translucent silicone elastomer with medium viscosity) was applied over the dura to facilitate the closure and to avoid contact between bone cement and cortex. We used KWIK-SIL elastomer because it produces only a small amount of hydrogen gas during condensation (the traditional room temperature–vulcanizing silicone systems produce toxic chemicals).3

FIG. 1.
FIG. 1.

Intraoperative photograph showing device placement over the motor cortex: the central sulcus is visible under the grid. L = left; R = right. Figure is available in color online only.

The bone removed during the craniotomy was replaced and affixed with screws and plates. We avoided damage to the grid stem by excessive pressure or overbending. Finally, bone cement was applied around the grid case to smooth the edges and provide better protection to the case. Once the cement became solid, the wound was washed and closed with sutures. Antibiotics (Augmentin, 14 mg/kg) were given intramuscularly, and the animal was placed in the observation cage with heating pads.

The 16-Electrode Device

The device called the ECoG implantable wireless 16-electrode (ECOGIW-16E) monitor (patent numbers US9031657 B2, EP2699145 B1, AU2012245942 B2, JP2014514944 A, and CN103648367 A) was designed specifically for monkeys.32 It consists of 2 parts (Fig. 2): the grid and the body. The grid is a single sheet of flexible polyimide support that integrates 16 platinum electrodes. The body, cased in polyetheretherketone, includes a microcontroller handling local processing and a transceiver module for implantable medical applications within the Medical Implant Communication Service (MICS) band (402–405 MHz), with an 800/400/200 kbps raw data rate. In addition, it includes a triaxial accelerometer, a stimulus generator, a sensor of temperature and load current, and a lithium-ion battery (3.6 V, 150 mA/hr; International Organization for Standardization [ISO] 13485). A polyetheretherketone case was chosen for its high and well-documented biocompatibility in prostheses in different fields of medicine.34 Finally, we have adopted a charging apparatus to provide an induction charger (250 mW, 70 mA at 3.7 V) for a wireless rechargeable battery. The interface consumes 58 mA (16 channels at 500 samples per second [SPS] + radio link), 30 mA (16 channels at 500 SPS), and 7 mA in standby mode.

FIG. 2.
FIG. 2.

The scheme of the ECOGIW-16E. Measurements are in millimeters (mm). Figure is available in color online only.

The device was positioned orthogonally and the grid was centered above the hand knob of the left motor cortex (Fig. 1), also providing coverage of the central sulcus and part of S1.

Wireless Recharging Cage

The ECOGIW-16E device is wirelessly rechargeable using a special cage for nonhuman primates that was developed by Aethra Telecommunication Co. to recharge the device. The cage (patent numbers EP2755469 B1, US2015180267, AU2013278927 B2, and CN104427864 A) is a nylon structure containing coils for electromagnetically induced high-frequency recharge in the x, y, and z directions of the space. The coils, generating a constant magnetic field inside the volume they encompass, provide wireless recharging of the ECOGIW-16E device, allowing freedom of movement and avoiding any constraint to the monkey. Cage details are available in Table 1.

TABLE 1.

Details of the recharging cage characteristics

CharacteristicMeasure
Dimensions
 Weight450 kg
 External dimensions, in mm920 (width) × 1700 (length) × 2000 (height)
 CharacteristicsETS 123
Power supply
 Voltage230 VAC
 Absorbed power1200 W
HF power unit
 Frequency1000 kHz
 Horizontal coils (no.)3
 Vertical coils (no.)3 + 2
 Max magnetic field163 A/m IEEE C95.1

ETS = European Treaty Series (European Convention for the Protection of Vertebrate Animals Used for Experimental and Other Scientific Purposes); HF = high frequency; IEEE = Institute of Electrical and Electronics Engineers; max = maximum; VAC = volts of alternating current.

Both the ECOGIW-16E device and the recharging cage are wireless systems. This novelty introduces several improvements in experimental settings researching sensorimotor cortex functions. In typical wired systems, the cables need to be protected from the hands of the primates because they typically try to strip them away. To avoid this problem, behavioral experiments require that the monkey sit in a special chair that does not allow the animal’s arms to reach the cables. Our cable-free system therefore opens a new window of opportunity to observe the neural activities in unrestrained, freely moving animals. Using the smart wireless recharge cage, it is possible to recharge the implanted device during the physiological rest period while recording the ECoG signals at the same time.

Electrophysiological Measures

In this study, we used commercial software for ECoG recording (Micromed), and we tested software specifically developed for our hardware (Aethra Telecommunication Co.) for DCS and impedance measurement. The ECoG recording is done using the previously described customized cage with the animal fully awake and in freely moving condition. A video recording was synchronized with ECoG. Impedance checks of each contact preceded the recording and consisted of electrochemical impedance spectroscopy of each electrode performed using AB MEDICA internally developed software. It is possible to set up electrochemical impedance spectroscopy from 1 to 250 Hz for each electrode. Moreover, the motor evoked potential was obtained by DCS of cortical electrodes coupled to a video recording of the motor response, with the animal anesthetized. We achieved cortical stimulation using embedded capabilities of the implant by sending electrical pulses to each electrode at constant voltage (3 V at 1–2 mA). We registered median and tibial nerve somatosensory evoked potentials (SSEPs) once a week after induction of general anesthesia outside the recording cage, inside a dedicated mobile ECoG unit.

We briefly describe the SSEP protocol: the nonhuman primates were anesthetized using a mixture of ketamine (7 mg/kg) and xylazine (0.6 mg/kg) administered intramuscularly, and then put onto a comfortable table that was covered with an absorbent sheet. The median nerve area at the level of the wrist and the tibial nerve area at the level of the ankle were shaved, and 2 needle electrodes were placed following the nerve trajectory. A pulse generator (Energy Light; Micromed) provided a bipolar electrical stimulation to the nerve (stimulation parameters: 1 Hz, anode proximal, cathode distal, pulse width 125–250 μsec, amplitude variable). The generator was synchronized to the ECoG recording software. When the nerve stimulation was performed, it was possible to record the related cerebral signal in correspondence to the somatosensory area. This procedure was done several times, to identify exactly which electrodes more frequently record the resulting signal to the stimulation. We identified electrodes 4, 5, 7, and 13. We used data derived from this study phase to develop the DCS stimulation protocols (Table 2).

TABLE 2.

Description of the 3 stimulation protocols

Stimulation ProtocolValue
No. 1
 Electrode no.4
 Impedance4 KW
 ShapeBipolar
 Frequency1 Hz
 V positive3 V
 V negative3 V
 Pulse width (tp = tn)100 μsec
 Tpn5 μsec
No. 2
 Electrode nos.4, 5, 7, 13
 Impedance4, 3.8, 5, 5 KW
 ShapeBipolar
 Frequency7 Hz
 V positive3 V
 V negative3 V
 Pulse width (tp = tn)500 μsec
 Tpn5 μsec
No. 3
 Electrode nos.4, 5, 7, 13
 Impedance4, 3.8, 5, 5 KW
 ShapeBipolar
 Frequency30 Hz
 V positive3 V
 V negative3 V
 Pulse width (tp = tn)500 μsec
 Tpn5 μsec

Tn = negative time; tp = positive time; tpn = time between positive and negative component.

Tissue Preparation and Histological Analysis of Brain Reactivity

Thirty weeks after implantation, the monkey was perfused transcardially with 0.9% NaCl followed by 10% formalin, as previously described by our team.28 To avoid damaging the brain tissue during histological procedures, the monkey brain—including the device and the skull—was fixed in 4% formalin. After fixation, we removed the device carefully to test its adhesion to the brain cortex. Then, the brain was removed from the skull and postfixed overnight in the same fixative. Next, we placed the brain in phosphate-buffered saline with the addition of 30% sucrose until the brain sank. Before freezing, we cut the brain to form a block, and both cerebral tissue and fibrotic tissue covered the implant. Furthermore, we investigated the control tissue. Subsequently, the brain was serially coronally sectioned using a freezing microtome. Sections were collected and processed for Nissl and Perls iron staining (RAL Diagnostics), and immunohistochemistry analysis for glial fibrillary acidic protein (GFAP), vimentin (Vim), and ionized calcium-binding adapter molecule–1 (Iba-1). We incubated sections with the following primary antibody solutions overnight at 4°C, including GFAP (1:500, polyclonal rabbit IgG; Dakocytomation) to identify astrocytes; Vim (1:500, monoclonal mouse IgG; Dakocytomation) to identify meningeal-derived fibroblasts; and Iba-1 (1:500, polyclonal rabbit IgG; Wako Chemicals GmbH) to identify macrophage and/or microglia. Secondary antibodies were diluted, including goat anti–rabbit IgG (Alexa 488; Molecular Probes) for GFAP and Iba-1. All sections were counterstained by incubation with the nuclear dye propidium iodide (Sigma). Sections treated only with secondary antibody but with no primary antibody were used to determine nonspecific binding. Tissue sections were mounted with Fluorsave, and bound primary antibodies were visualized on a confocal microscope.

Results

Prolonged Wireless ECoG Recording

The procedure and postoperative course were uneventful. The monkey recovered immediately and was able to resume all normal motor activities (walking and climbing into the cage and feeding unassisted) within a few hours. The ECoG signals of the 16 electrodes were recorded at a 512-Hz sampling rate and a software-imposed band-pass filter from 0.008 to 400 Hz. We first removed all frequencies below 0.5 Hz from the ECoG signals by using a high-pass filter. The ECoG signals were acquired every day for 6 months and remained stable during the study. Figure 3 shows 3 screenshots of EcoG signals: a few hours after the implant, after 3 months, and after 6 months. Electrode impedance values remained at 40 KΩ, below the acceptable threshold of 100 KΩ. We also computed the fast Fourier transform of the ECoG traces. The power spectral estimate is based on the Welch method applied to a centered signal, 1-second window analysis, with an overlapping 750 msec, and each section is provided with a Hamming window.43 The frequency spectrum shows the characteristic decrease in amplitude at higher frequencies and all the characteristic frequency components of an ECoG signal. Spectral analysis of the averaged signal showed the expected power reduction in all frequencies between the first and the last month of observation, due most probably to the thickening of fibrotic tissue surrounding the implant.8,13

FIG. 3.
FIG. 3.

ECoG recordings in awake monkey at 24 hours, 3 months, and 6 months after implantation. Electrophysiological recordings were obtained using wireless transmission in the plastic cage.

In addition, we performed a spectral analysis in each channel and evaluated the amount of line noise in our ECoG signal. For each electrode, we calculated the relative power of the 50-Hz line noise band (P49,50) over the total power (P0,250), giving rise to the percentage of electronic noise in each channel (P49,50/P0,250). The percentage of electrical interference in total signal for each electrode was stable at < 5% during the 6 months of the study. Only at the final recording did electrode 10 increase to more than 10% of electrical noise, probably due to a dysfunction of the specific electrode.

We measured values of impedance at different frequencies (from 16 Hz to 250 Hz) in all channels. In Fig. 4 we report impedance values for each electrode (1 kHz), and the inset graphic shows the relative mean value. During the entire duration of the study, impedance values had remained well below the threshold of 100 KΩ (red line in Fig. 4), representing the limit of the impedance over which signal quality becomes poor. Our device has impedance values consistent with the ones found in the literature.21

FIG. 4.
FIG. 4.

Graphs showing electrode impedance at 1 kHz over time. The mean impedance for each electrode during the total study duration is shown. Figure is available in color online only.

Overall, the signal quality remained excellent for the entire 6 months of recording, except for a modest degradation during the 1st week after the procedure (probably due to postoperative debris), as well as during the last 3 weeks, when the implant duration over the cortex drew close to 6 months.

We also performed throughput tests of the antennae in the device. After the surgery, we measured the electromagnetic field in a not-anechoic room without known interferences. The electromagnetic field is almost symmetrical, even in the presence of unknown disturbances, and it decreases with the distance from the device.

Somatosensory Evoked Potentials

We analyzed the evolution of the first component (N1 or P1) of the SSEPs for the median and tibial nerve in the animal. We carefully inserted stimulator needles in the same spot at each acquisition and used constant parameters capable of eliciting muscle contraction. We were able to follow the evolution of the SSEP over time. Also, by signal inspection it was possible to localize the central sulcus and the motor strip projection to the grid. This motor projection, represented by polarity changes in the first inflection curve, was present during the 6-month period of observation (Fig. 5).

FIG. 5.
FIG. 5.

A: Example of median nerve SSEP response of the 15.9 × 11.5–mm ECoG grid. The diameter of the recording sites was 1.09 mm and the electrode pitch was 2.99 mm. Change in polarity can be perceived between electrodes 4 and 5, where dipolar deflections at 13 msec were seen (horizontal black arrowheads). These changes signal the presence of the central sulcus between those electrodes. (Data averaged 156 repetitions, and the vertical black arrowhead designates the trigger mark.) B: Electrode grid position showing the central sulcus (white line). The blue area corresponds to the premotor cortex (positive early deflections), and the red area corresponds to the sensitive strip. C: Cortical stimulation before grid implantation: the position of maximal hand contraction is shown below the fork probes (Inomed Cortical Probes). Figure is available in color online only.

Time delays between the nerve stimulation and the related recording of the signal remained constant throughout the 6-month experiment (median nerve: N1/P1 was 16.64 ± 3.16 msec; tibial nerve: N1/P1 was 22.43 ± 1.07 msec). In contrast, curve amplitude (in μV) was significantly reduced for tibial nerve SSEP (from 135.42 to 56.50 μV, p < 0.0001) but not for median nerve SSEP (from 19.60 to 13.82 μV, p = 0.0739; Wilcoxon matched-pairs signed-rank test) between 7 and 181 days of observation (Fig. 6).

FIG. 6.
FIG. 6.

Bar graph showing the mean variation in amplitude of the first component (N1/P1) of SSEP in the tibial and median nerves. The ECoG microelectrodes record a significant reduction in amplitude in tibial SSEP but not in median SSEP. Encapsulation could attenuate potential amplitude more when the source is far from the recording electrode (ankle stimulation) than when the source is below the electrode grid (hand). ***p < 0.0001.

Direct Cortical Stimulation

Finally, we performed DCS to test the device stimulation features and functionality. We used 3 different stimulation protocols. After careful stimulation in each electrode, we found that it was possible to elicit contralateral arm movements by using the first and third stimulation protocols (Table 2). Protocol 1 in electrode number 4 produced thumb flexion, and protocol 2 in the same electrode provoked not only right thumb but also elbow complete flexion. Protocol 3 was able to reproduce finger flexion in electrode number 5.

Long-Term Biocompatibility Evaluation

Histological examinations were performed postmortem, 30 weeks after implantation. The cortex was inspected carefully to assess macroscopic signs of damage caused by the prolonged contact of the grid; visual inspection failed to show macroscopic signs of tissue defect. Observation of the implantation site showed a slight brown pigmentation of the brain surface under the grid, probably due to resorbed bleeding. The brain site underlying the electrode array showed its integration into a newly formed dural layer. This neoformed tissue (NT) appeared to be continuous to the constitutive dura mater. We observed no adhesion between the newly formed connective tissue covering the grid and the cortical surface; the grid was easily removed without damaging the cortex.

Immediately after explantation, the grid was extracted from the surrounding dural layers. Some of the exposed electrodes revealed cracks and wrinkles on the surface. We analyzed the dural layers and grid under the stereomicroscope to assess histological features of reactive dural tissue and to observe the electrodes’ mechanical condition (Fig. 7).

FIG. 7.
FIG. 7.

A: The monkey brain after 30 weeks of contact with the electrode array. The dotted lines delimit the contact area between the brain cortex and the electrode array. B: Electrode array explanted after 30 weeks, embedded in a newly formed dural layer. Ant. = anterior; Lat. = lateral; Med. = medial; Post. = posterior. Figure is available in color online only.

Observation of a capsule transversal section revealed a connective tissue formation on two sides of the electrode array (Fig. 8A). The thickness of the reactive tissue ranged between 450 μm (inner layer) and 800 μm (outer layer) (Fig. 8B–D). We observed in Fig. 8B and C that the dural scar (reactive dura mater) was interfacing with the outer layer of NT. As shown in Fig. 8C and D, Perls iron staining revealed hemosiderin in the reactive tissue, suggesting an old hemorrhage in these areas. Note that the thickness of the constitutive dura mater is approximately100 μm after tissue preparation.

FIG. 8.
FIG. 8.

Macroscopic and microscopic views of the encapsulated electrode array (EA). A: Photograph of the electrode array embedded between NTs on top and below. B: Histological slice showing constitutive dura mater (CDM) and reactive dura mater (RDM) covering the NT. Nissl staining. C: Histological slice showing hemosiderin deposits into the NT. Perls iron staining. D: Histological slice showing hemosiderin deposits into the NT. Perls iron staining. Figure is available in color online only.

We performed immunohistological analysis to better assess features of the reactive tissue encapsulating the grid. Figure 9 shows representative GFAP, Iba-1, and Vim expression patterns in the capsule. As shown in Fig. 9B, the reactive tissue surrounding the electrode array did not contain any GFAP-positive cells, indicating the absence of astrogliosis in the capsule. Adjacent sections used Iba-1 and Vim antibodies to label microglial cells and meningeal fibroblasts, respectively, and this showed that reactive tissue contains homogeneously activated microglial cells with an increased density in hemosiderin-containing areas (Fig. 9D), suggesting a recruitment of microglial cells for hemosiderin elimination. As shown in Fig. 9E, we observed Vim-positive cells in the NT with a higher protein level in the outer layer, outside of hemosiderin deposits. This result suggests that meningeal-derived fibroblasts migrate into the reactive tissue from the meningeal space and contribute to the foreign body response. These results clearly show the well-known characteristics of encapsulation for subdural implantation of electrode arrays, with microglial and fibroblastic components.

FIG. 9.
FIG. 9.

A: Photograph of the electrode array (EA) embedded between NTs on top and below. B: Representative GFAP expression pattern (green) in the reactive tissue around the electrode array. The reactive tissue showed no GFAP antibody staining. C: Representative Iba-1 expression pattern (green) in the reactive tissue around the electrode array. Microglia is positively marked in all the NT, with a higher density of microglial cells in areas containing hemosiderin deposits. D: Representative Vim expression pattern (green) in the reactive tissue around the electrode array. The outer layer of reactive tissue (in contact with the reactive dura mater [RDM]) shows a higher Vim level. Bar = 200 µm (B–D). Figure is available in color online only.

The last part of the histological investigation was conducted to evaluate brain cortex reaction beneath the electrode array and to detect signs of inflammation. We performed Perls iron staining to detect signs of alteration and old hemorrhage. No sign of cortical damage was found over the region covered by the grid, except in a single point where repeated stimulation was performed; this was not representative of the stimulated area. As shown in Fig. 10C and E, a small region of cortical damage (diameter < 1 mm) can be identified under electrode number 4, which was used for multiple mapping procedures.47

FIG. 10.
FIG. 10.

A: Photographs of incomplete brain coronal sections showing contact areas of the electrode array (EA) with the cortical surface and the control side. B: Representative histological slice demonstrating the absence of cortical defect on the control side. C: Representative histological slice demonstrating alteration of the brain cortex under the electrode array. Hemosiderin deposits are stained in blue (arrow). D: Representative histological slice demonstrating the absence of cortical defect on the control side. E: Representative histological slice demonstrating alteration of the brain cortex under the electrode array. Hemosiderin deposits are stained in blue (arrow). Perls iron staining (B–E). Bar = 5 mm (B, C); 1 mm (D, E). Figure is available in color online only.

During the electrical stimulation, we adopted all precautions to limit or avoid damage to the brain tissue. Shannon’s equation has proved the empirical relationship between charge delivered to the tissue and tissue damage.37 It has been considered that K = 1.7 with a phase of 500 μsec, the electrode has a 1-mm pad made of platinum, and the maximum current that can be delivered while still avoiding damage to brain tissue is 3.4 mA.11 As shown in Fig. 4, the mean impedance value on electrode number 4 is 10 KΩ and, considering the battery voltage of 3 V, the maximum current delivered on electrode number 4 is less than 1 mA—well under the maximum value of 3.4 mA.

We performed an analysis of cortical GFAP expression in electrode array–covered and electrode array–uncovered brain cortex areas and noted that expression of the astrocyte marker was highly concentrated in the glia limitans. In electrode array–covered brain cortex, we observed a loss of continuity for the glia limitans and reactive astrocytes in the cortex in the spot where the cortical surface was altered. In some areas, the thickness of electrode array–covered glia limitans was increased because of the proliferation of reactive astrocytes. GFAP staining revealed a mild astrogliosis in the cortex under the electrode array. The Iba-1 immunostaining (Iba-1 expressed by microglia and macrophages) showed the absence of activated microglial cells into the control brain cortex after 30 weeks (Fig. 11). Observation of the electrode array–covered brain cortex showed the presence of activated microglial cells into the reactive glia limitans and into the brain cortex under the electrode array after 30 weeks. Note that expression of the microglial marker is highly concentrated in the reactive glia limitans and in the first cortical layer. These results suggest that the electrode array induced localized and mild brain tissue response after a 30-week period.

FIG. 11.
FIG. 11.

Representative Iba-1 expression patterns in electrode array–uncovered brain cortex (A) and electrode array–covered brain cortex (B and C). Bar = 20 µm. GL = glia limitans (arrows). Figure is available in color online only.

Discussion

A detailed prospective analysis of the reliability of a long-term wireless ECoG device implanted on the sensorimotor cortex of a nonhuman primate is reported. The evaluation tested not only the wireless transmission and the quality of the signal, but also the daily impedance of each electrode, as well as the elicitation of SSEPs and the motor effects of DCS. This neuroprosthesis also provides wireless recharging within a dedicated primate cage, allowing a constant battery charge in freely moving animals. Furthermore, we evaluated the biocompatibility of the entire neuroprosthesis, including not only the subdural grid but also the wireless case implanted on the skull near the craniotomy.

The implantable wireless device presented here allowed us to perform prolonged ECoG recording and DCS for 30 weeks. Long-term wireless ECoG recording was possible in a noisy, nonisolated environment, in a freely moving animal. The ECoG signals recorded a power reduction at 6 months, which was explained by local dura mater neoformation. The ECoG signal quality remained excellent throughout the experiment, with low impedances observed. The device was removed easily at the end of the 30 weeks of observation. Encapsulation was partially responsible for attenuation in the tibial SSEP signal. The median SSEP signal was stable, probably due to the distance between the source of SSEP and the microelectrodes (the grid, after sensorimotor cortex mapping, was implanted directly over the cortical hand representation). Line noise was maintained below 5% of the total signal throughout the experiment, giving an indirect measure of stable impedance changes of the electrode-tissue interface. We found no adhesion between the device and the cortex. The cortex underlying the implanted grid showed no macroscopic signs of damage. Immunohistochemical analysis revealed mild superficial astrogliosis, moderate microglial activation, and mild resorbed bleeding under the electrode array. Overall, the neuroprosthesis was well tolerated, considering the extended implantation time. The dural encapsulation of the grid is, in our view, not a matter of concern in humans: the dural neoformation seen here is peculiar to nonhuman primates, which typically show very intense neodurogenesis following dural opening. A similar phenomenon is very unlikely to be seen in a clinical context.

The possibility shown here to greatly extend the implantation time of a subdural grid paves the way for prolonged ECoG recording during epilepsy surgery, thus improving the chances of localizing the epileptogenic focus and extending the range of patients to whom invasive ECoG recording can be offered. This device, which provides a bidirectional interface well tolerated over a long time, also lends itself to further developments toward closed-loop seizure control and BCI applications.

Epilepsy has been addressed in recent years by brain-machine interfaces, with the main objective of detecting seizures and transmitting an alert to patients.12 One category of epilepsy alert devices is noninvasive wearable devices, providing a fast response to ongoing seizures of the tonic-clonic type, based on accelerometers (Epilert: https://bio-lert.com; Smartwatch: https://smart-monitor.com/about-us/news-events/). More closely related to our applications are the invasive seizure detection systems, integrating an ECoG signal. They can detect seizures and transmit an alert to the patients before secondary generalization occurs.12 Some of them can be coupled to an electrical stimulator, delivering electrical pulses to therapeutic targets (i.e., anterior thalamus) when seizures are detected.40 However, these devices are not suitable for presurgical evaluation of epileptic patients: they have very low spatial resolution due to electrode characteristics (low number, pitch, and long diameter). There is no recording and stimulating device for presurgical epilepsy evaluation that is similar to the prototype presented here. Furthermore, the device proposed here is a preliminary proof of concept developed for primate surgery. Another device providing similar long-term recording and wireless transmission and recharging is currently being tested in animal models, in anticipation of clinical application.

Patients with drug-refractory epilepsy could greatly benefit from prolonged wireless recording, allowing them to reduce the risk of intracranial infections proceeding along the intracranial cables. Prolonged wireless ECoG recording could be offered to a substantial population of patients affected by nondaily or episodic seizures. As of today, this patient sample is not a surgical target, because the short implantation time of conventional wired intracranial recordings added to the surgical stress and effects of anesthesia does not allow time enough to record and localize the epileptogenic focus. In such cases, a longer period of observation could help to identify the epileptogenic focus and increase the accuracy of the resection.

Some potential barriers preventing the clinical use of a device implanted over a long period are signal attenuation over time, foreign body rejection, and development of subdural hematomas. Moreover, a much larger recording surface than that offered by this proof-of-concept device will be necessary for clinical applications in epilepsy surgery. We are currently engaged in the preclinical testing of a device providing 128 channels spread over a thin-layer silicone grid (thus aiming to prevent the development of subdural hematomas while providing a large recording surface that can be tolerated over multiple weeks). The results of these preclinical studies are rather encouraging and will be reported soon.

We have developed and tested a 128-channel device because this represents the best compromise between sample rate and radio band used. Our device transmits to 400 MHz according to the MICS specification. The MICS band was created by the Federal Communications Commission specifically to allow the safe use of implanted wireless devices for diagnostic and therapeutic applications. The frequency used reduces the risk of interference with other devices and requires a transmission power limited to 25 μW. Doubling the number of channels (256) would imply reducing by half the sample rate and the quality of the received signal. Therefore, it is technically possible to increase the number of channels, but the drawback is the detriment to the quality of the acquired signal, and therefore to the possibility of making an adequate diagnosis.

A wireless long-term system can provide recordings of previously undetectable seizures because of the longer observation period. Current systems of invasive EEG (i.e., stereo-EEG) are limited to 7–15 days of exploration, with cable and extensions passing through the skin. The risks of infection are reduced in a completely closed system. One can also imagine a home system, which does not require hospitalization and dedicated medical personnel.

Currently, most ECoG BCIs have an electrical connector that passes through the skull and skin.22,36,45 This is not a practical solution for long-term implantation. The challenge has been to improve the next generation of implantable devices by making them wireless, with a large number of electrodes and long duration of functionality. Wireless interfaces are beginning to emerge; e.g., Vansteensel et al.41 have described a first wireless ECoG implantation in a locked-in syndrome secondary to amyotrophic lateral sclerosis. The fully implanted BCI consisted of 4 commercially available subdural electrodes (4 electrodes on each strip; each electrode was 4 mm in diameter, and there was a 1-cm interelectrode distance) placed over the motor cortex and a transmitter (Activa PS) placed subcutaneously in the left side of the thorax. Only 2 strips are connected to the implantable pulse generator, making accessible 8 electrodes to wireless bidirectional transmission. This is a very promising approach that has the merit of being almost ready for clinical use (the Activa PS transmitter is not commercially available yet), but interelectrode distance and only strip geometry for electrodes may be an issue if used for epilepsy surveillance and cartography.

Su et al.39 introduced a compact implantable wireless 32-channel bidirectional brain-machine interface (BBMI) to be used with freely moving primates. It is a more dedicated device with stimulation and sensing characteristics, but it still uses intracortical electrodes (Utah array) as the neural interface, which is not suitable for epilepsy exploration. The Nicolelis group has also published some wireless interface data obtained in monkeys,33 allowing the animal to pilot a wheelchair. The device used a Utah array, and the wireless portion of the implant is exposed in a case not covered by skin. This makes the device not suitable for long-term implantation.

Other BCI implants, some of them shown in Table 3, are at different technological levels of readiness and are not suitable for presurgical epilepsy evaluation. The Neurochip-2 is a device conceived to be placed in an external chamber in the primate’s head and to have only offline recording and stimulation.46 The WIMAGINE (Wireless Implantable Multichannel Acquisition System for Generic Interface with Neurons) is a fully implantable ECoG device that is in the early stages of clinical testing for use as a BCI in tetraplegic patients, but lacks an electrical stimulation module.27 The PennBMBI (brain-machine-brain interface) is a device with 4 recording and 2 stimulating electrodes with a high sampling rate, designed for closed-loop BCI, not suitable for the large-area investigation necessary in epilepsy surgery.24 The W-HERBS (Wireless Human ECoG-based Real-time BMI System) device has a multielectrode matrix (64 electrodes) for recording large cortical areas, but has no stimulation module. Nguyen et al.29 have advanced a recording system that combines a 32-electrode matrix with optical stimulation, but it is not wireless, requiring a wired connection for signal transmission. Angotzi et al.2 have developed an interesting device for small animals that is completely wireless, with 8 recording channels and 8 stimulating electrodes based on nonrechargeable batteries.

TABLE 3.

Literature review of recently developed wireless implantable devices and comparison of their features

FeatureSu et al., 2016Zanos et al., 2011; Neurochip-2Sauter et al., 2015; WIMAGINELiu et al., 2015; PennBMBIHirata et al., 2012; W-HERBSNguyen et al., 2014Angotzi et al., 2014
Recording channel32 unipolar/bipolar3 unipolar/bipolar64464 × 2328
ADC resolution16 bits, 800 SPS/channel, up to 30 kSPS/channel8 bits, 256 SPS12 bits, 1 kSPS12 bits, 21 kSPS12 bits, 1 kSPS16 bits, 400 kSPS (12.5 kSPS/channel)8 bits, 15 kSPS/channel
Sampling rate0.1–20 kHz10 Hz–7.5 kHz0.5–400 Hz0.05 Hz–6 kHz0.1 Hz–1 kHz0.2 Hz–5 kHz1 Hz–10 kHz
Bandwidth24-GHz RF communication link w/ mESB protocolSerial cable/infrared data linkProprietary UHF link in MICS bandCommunication link, 2.4-GHz RFBluetoothWired2.4-GHz ISM band
CommunicationRechargeable1 or 2 rechargeableInductive linkRechargeablePolymer lithium5-V USB3.7 V, 700 mA/hr
Power supply3.6-V battery3.7-V batteryNA3.7-V battery3.7-V ion batteryNANA
Power consumption4.22–15.4 mA284–420 mV75 mW w/o charging, 350 mW w/ charging7.3-mA transmit for sensor node only4.9 mW (AFE), 300 mW (wireless)
Battery chargingWireless ultrasonic chargingNANANAWireless charging, 4-W coil at 38-mm distanceNA
Size35-mm diam63 × 63 × 30 mm50-mm diam, antenna 10 cm2, 12.54-mm thickness31 × 13 × 8 mm, 43 × 27 × 8 mm, 56 × 36 × 13 mm8-mm thickness, 40-mm diam, 60 × 60 × 8 mm, 20 × 30 × 2.5 mm29.5 × 43.4 mmNA
Stimulation channel4 unipolar/bipolar3 unipolar/bipolar21 optical stimulus8 bipolar
Stimulation voltage20 V15 V (normal), 50 V (high V)12 V9 V
Current density40-µA pulse10–200 µA current or 0.5–5 mA0–1 mA300 µA
Pulse width1 msec minimum0.2, 0.6 msecNA200 µsecNANANA
Pulse frequency250 Hz1 min or 1 min 10 secNANANANANA

ADC = analog-to-digital converter; AFE = analog front end; diam = diameter; ISM = Industrial, Scientific, and Medical; NA = not applicable; RF = radiofrequency; UHF = ultrahigh frequency; mESB = micro Enhanced ShockBurst (a wireless protocol, Nordic Semiconductor); — = not done or no information.

Considering the size of the primate brain, we developed a proof-of-concept device that allowed long-term ECoG recording over 6 months. On the basis of this experience, we then developed a 128-channel wireless subdural grid, which has been tested on suitable animal models (pigs) to proceed to human implantation. The recharging process in human applications is meant to develop a recharging modality that permits patient movements. For this reason, we created a portable wireless recharging system, which works through an external coil. The implanted device is supplied by medical-grade–certified batteries, in a titanium case. By using a cap with a large area covered in Velcro that is worn only during the recharging phase, the external coil could be coupled with the area corresponding to the implanted device. Attaching the wireless recharger to the implantation area allows electromagnetic coupling between the coil integrated into the implanted device and the one integrated into the wireless recharger; in this way, power transfer can start. The external coil could be fixed by a headset cover, to guarantee its stability. The wireless recharger provides visual and auditory signals: a bicolor light-emitting diode becomes red to signal some errors in the recharge phase. If no errors occur, the light-emitting diode becomes green to give feedback about the state of recharge. An auditory signal indicates the joining between the two coils, so that the recharge phase can start.

The management of charge current requires a scientific rationale and adherence to ISO guidelines (ISO 14708–1 specifies requirements for establishing control and protection mechanisms), to prevent harm to the patient caused by heat. The ICNIRP (International Commission on Non-Ionizing Radiation Protection) guidelines for limiting exposure to time-varying electric, magnetic, and electromagnetic fields1,23 report that exposure to magnetic and electrical fields and the relative energy absorbed by the human body (i.e., the SAR [specific absorption rate]) have to be controlled by considering the tissues’ heat elevation. A heat elevation of 1°C–2°C from the baseline does not involve tissue damage. Some authors25,44 propose thermal considerations regarding heat’s effects on neuroimplanted devices. They confirm the limit of +2°C (approximately 39°C), beyond which tissue damage and cell necrosis could occur. A 2°C increase occurs within approximately 1 hour of charging, so they suggest caution in the management of charging time. Because of this, we created a control mechanism consisting of 1) an internal microprocessor measuring temperature, 2) an external thermal protection, and 3) a timer. The temperature and time of the recharge phase can be set by using the application software. The internal microprocessor interrupts the recharge phase in the event that the temperature exceeds the determined value, or in case the recharge time exceeds the determined value. The external temperature sensor measures degrees by the skin of the skull, providing a further temperature protection tool, which is independent of the internal microprocessor.

Conclusions

The novel wireless ECoG device described in this work as a proof of concept provides an excellent platform for prolonged ECoG recording. Potential clinical applications include BCI and epilepsy surgery. In the latter case, the absence of subcutaneous cables allows prolonged monitoring, thus enhancing the chances of achieving seizure focus localization. Moreover, it offers a prospect to treat a large subset of patients with drug-refractory epilepsy who have nondaily seizures and rarely undergo invasive monitoring due to the low likelihood of registering and localizing an ictal focus over a short time. Based on this experience, we are now testing a 128-contact wireless device suited for human use.

Acknowledgments

We express our gratitude to Dr. Chiara Fornoni for the valuable editorial assistance.

Disclosures

AB MEDICA S.p.a. manufactured the ECoG device prototype and also funded the entire project, including materials and structures, with the declared aim to test the in vivo operation and the biocompatibility and device reliability during implantation. Pantaleo Romanelli is a consultant for and Cosimo Puttilli and Mauro Picciafuoco are employees of AB MEDICA S.p.a. Also, ATLC S.r.l., a company of the AB MEDICA group, partially cofunded (in the percentage of one-third) Marco Piangerelli’s doctoral scholarship.

Author Contributions

Conception and design: Romanelli, Benabid, Torres. Acquisition of data: Romanelli, Ratel, Gaude, Costecalde, Puttilli, Picciafuoco, Benabid, Torres. Analysis and interpretation of data: Piangerelli. Drafting the article: Romanelli, Ratel, Gaude, Costecalde, Puttilli, Benabid, Torres. Critically revising the article: Piangerelli, Gaude, Costecalde. Reviewed submitted version of manuscript: Romanelli, Ratel, Puttilli, Picciafuoco, Benabid, Torres. Approved the final version of the manuscript on behalf of all authors: Romanelli. Statistical analysis: Piangerelli, Picciafuoco.

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Contributor Notes

Correspondence Pantaleo Romanelli: AB MEDICA S.p.a., Milan, Italy. radiosurgery2000@yahoo.com.

INCLUDE WHEN CITING Published online May 11, 2018; DOI: 10.3171/2017.10.JNS17400.

Disclosures AB MEDICA S.p.a. manufactured the ECoG device prototype and also funded the entire project, including materials and structures, with the declared aim to test the in vivo operation and the biocompatibility and device reliability during implantation. Pantaleo Romanelli is a consultant for and Cosimo Puttilli and Mauro Picciafuoco are employees of AB MEDICA S.p.a. Also, ATLC S.r.l., a company of the AB MEDICA group, partially cofunded (in the percentage of one-third) Marco Piangerelli’s doctoral scholarship.

  • View in gallery

    Intraoperative photograph showing device placement over the motor cortex: the central sulcus is visible under the grid. L = left; R = right. Figure is available in color online only.

  • View in gallery

    The scheme of the ECOGIW-16E. Measurements are in millimeters (mm). Figure is available in color online only.

  • View in gallery

    ECoG recordings in awake monkey at 24 hours, 3 months, and 6 months after implantation. Electrophysiological recordings were obtained using wireless transmission in the plastic cage.

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    Graphs showing electrode impedance at 1 kHz over time. The mean impedance for each electrode during the total study duration is shown. Figure is available in color online only.

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    A: Example of median nerve SSEP response of the 15.9 × 11.5–mm ECoG grid. The diameter of the recording sites was 1.09 mm and the electrode pitch was 2.99 mm. Change in polarity can be perceived between electrodes 4 and 5, where dipolar deflections at 13 msec were seen (horizontal black arrowheads). These changes signal the presence of the central sulcus between those electrodes. (Data averaged 156 repetitions, and the vertical black arrowhead designates the trigger mark.) B: Electrode grid position showing the central sulcus (white line). The blue area corresponds to the premotor cortex (positive early deflections), and the red area corresponds to the sensitive strip. C: Cortical stimulation before grid implantation: the position of maximal hand contraction is shown below the fork probes (Inomed Cortical Probes). Figure is available in color online only.

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    Bar graph showing the mean variation in amplitude of the first component (N1/P1) of SSEP in the tibial and median nerves. The ECoG microelectrodes record a significant reduction in amplitude in tibial SSEP but not in median SSEP. Encapsulation could attenuate potential amplitude more when the source is far from the recording electrode (ankle stimulation) than when the source is below the electrode grid (hand). ***p < 0.0001.

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    A: The monkey brain after 30 weeks of contact with the electrode array. The dotted lines delimit the contact area between the brain cortex and the electrode array. B: Electrode array explanted after 30 weeks, embedded in a newly formed dural layer. Ant. = anterior; Lat. = lateral; Med. = medial; Post. = posterior. Figure is available in color online only.

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    Macroscopic and microscopic views of the encapsulated electrode array (EA). A: Photograph of the electrode array embedded between NTs on top and below. B: Histological slice showing constitutive dura mater (CDM) and reactive dura mater (RDM) covering the NT. Nissl staining. C: Histological slice showing hemosiderin deposits into the NT. Perls iron staining. D: Histological slice showing hemosiderin deposits into the NT. Perls iron staining. Figure is available in color online only.

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    A: Photograph of the electrode array (EA) embedded between NTs on top and below. B: Representative GFAP expression pattern (green) in the reactive tissue around the electrode array. The reactive tissue showed no GFAP antibody staining. C: Representative Iba-1 expression pattern (green) in the reactive tissue around the electrode array. Microglia is positively marked in all the NT, with a higher density of microglial cells in areas containing hemosiderin deposits. D: Representative Vim expression pattern (green) in the reactive tissue around the electrode array. The outer layer of reactive tissue (in contact with the reactive dura mater [RDM]) shows a higher Vim level. Bar = 200 µm (B–D). Figure is available in color online only.

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    A: Photographs of incomplete brain coronal sections showing contact areas of the electrode array (EA) with the cortical surface and the control side. B: Representative histological slice demonstrating the absence of cortical defect on the control side. C: Representative histological slice demonstrating alteration of the brain cortex under the electrode array. Hemosiderin deposits are stained in blue (arrow). D: Representative histological slice demonstrating the absence of cortical defect on the control side. E: Representative histological slice demonstrating alteration of the brain cortex under the electrode array. Hemosiderin deposits are stained in blue (arrow). Perls iron staining (B–E). Bar = 5 mm (B, C); 1 mm (D, E). Figure is available in color online only.

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    Representative Iba-1 expression patterns in electrode array–uncovered brain cortex (A) and electrode array–covered brain cortex (B and C). Bar = 20 µm. GL = glia limitans (arrows). Figure is available in color online only.

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