7
Biomaterials for Brain Implants
7.1 Introduction
Damage to the central nervous system (CNS) affects at least two million people per year. This may cause physical or sensory disabilities that partially or completely remove functioning. The life of these patients undergoes a drastic decline in quality, so much so that they are no longer able to look after themselves. Neuroengineering strives to provide improved quality of life to these people [1]. A brain (or neural) implant is a technological system that enables communication between the brain and electronic devices, thus permitting brain activity to be modified, recorded, and/or translated for the manipulation of devices [2].
These devices can improve our understanding of the organization and operation of the nervous system and may lead to improving the current state‐of‐the‐art neural technologies for tackling of some mankind’s most debilitating disorders, including deafness, paralysis, blindness, epilepsy, and Parkinson’s disease (PD) [3].
Brain–machine interfaces (BMIs) have been used since the second half of the twentieth century to register and stimulate neural tissue in humans and animals. Today, neural treatments are being used more and more to treat neurological disorders. In fact, electrical stimulation of the brain can alter the brain function by injecting electrical signals into neurons (Deep Brain Stimulation – DBS) [4, 5].
Much progress has been made in the development of intracortical systems for recording and translating brain signals to control external devices. These systems are called BMIs [6]. Such systems are potentially valuable for restoring lost neuronal function associated with neurological diseases and injuries.
Neural interfaces communicate with the nervous system via implantable electrodes that transduce electric signals to and from bioelectric signals [3].
To allow these devices to achieve the desired result, the electrodes must have a high signal‐to‐noise ratio (SNR) for a period of time sufficient for recording and must be able to record as many neurons as possible.
This translates toward the need of electrode materials that allow to develop high‐density neural probes that are biologically transparent and biocompatible, support seamless integration with neurons, and remain functional for a long period of time. The performance of electrode–tissue interface ultimately rests on the quality of the martial substrate, which enables a long‐lasting functional neural device.
Existing neural electrodes use conventional electronic materials that are often not intrinsically compatible with biological systems and do not conduct integration with neural tissue. They are usually made of metals that have biocompatibility characteristics, but the oxides produced greatly limit their use.
The challenge for materials science is to apply nanotechnology strategies and develop innovative biocompatible nanomaterials that mimic neural tissue characteristics, cause minimal inflammation and neuronal cell loss, and are functional for a long period of time.
Nanomaterials are particularly suitable for these applications because the neural tissue has complex nanoscale structural features that require an interface with nano‐sized components. Glial cells and extracellular matrix (ECM) have nanoscale dimensions; thus, the unique intrinsic properties of nanomaterials offer a great promise to seamlessly integrate with neural tissue and simulate features and functions of cells and ECM [7].
Carbon nanotubes (CNTs), silicon nanowires, gallium phosphide nanowires, and conducting polymer (CP) nanotubes are conductive nanomaterials able to provide a more‐effective surface area than regular metals for signal transduction at the electrode–tissue interface, thus enhancing the electrical characteristics of neural recordings and stimulations, such as SNR [8, 9]. Furthermore, the incorporation of drugs and bioactive molecules into EANs can improve the biocompatibility of neural electrodes, reduce reactive tissue response, and promote neural process outgrowth [3].
Advances in biocompatibility technology have increased the lifespan of brain implants to several years as opposed to the months, weeks, or days of their predecessors [1]. Most of the advances that increase biocompatibility can be attributed to novel materials and coatings [10]. The main focus of this review is to provide an overview of the most recent human in vivo materials for implementing brain implants.
7.2 Brain Implants Classification
7.2.1 Brain Implants for Recording
The brain implants that base their operating principle on the recording of brain activity are the so‐called brain–computer interfaces. In a brain computer interface (BCI), electrical activity is measured from specific parts of the brain, via penetrating microelectrode arrays, electrode matrices resting on the brain surface, or electrodes attached to the scalp. The recorded signals are amplified, digitized, and resolved into components by a computer processor [11].
Brain activity can be recorded, thanks to invasive and noninvasive techniques. Some of the most common noninvasive techniques for measuring brain activity are electroencephalography (EEG), functional magnetic resonance imaging (fMRI), magnetoencephalography (MEG), and functional near‐infrared spectroscopy (fNIRS) [12].
They acquire data from changes in magnetic fields, electrical current, and oxygen consumption. These techniques are not ideal for real‐world applications due to the size of the equipment and the poor spatial resolution. The noninvasive methodologies just mentioned and the materials that constitute them will not be dealt with in this review as they are not part of a proper brain implant.
Invasive methods achieve a high SNR by eliminating the volume conduction problems caused by tissue and bone. Electrodes are directly placed onto the cortical matter. This makes excellent detection of high‐frequency oscillatory activity possible. However, surgery is required for the transcutaneous wire implantation and signals from thousands to millions of cells may be superimposed [13]. Electrocorticography (ECoG) and single unit recording are the most used techniques.
7.2.1.1 Electrocorticography (ECoG)
ECoG electrodes are typically composed of platinum–iridium disks that are embedded in silastic sheets that are surgically implanted directly beneath the dura (subduraly), though they may also be implanted above it (epidurally). Like EEG, ECoG measures the summed local synaptic field potentials (LFPs) of underlying neuronal populations, a direct measure of neural activity inaccessible to fMRI. Because the electrodes are in direct contact with the cortical surface, the signal quality is greatly enhanced, especially for high‐frequency activity [14]. ECoG electrodes can also feasibly be combined into grids and strips consisting of dozens or hundreds of electrodes to sparsely cover large and/or distributed regions of the brain. The ECoG arrays that are used have 2.3 mm diameter electrodes and 18 mm electrode spacing, but can be manufactured with much smaller dimensions. They can even reach a diameter and a spacing of 1 mm. With a technology of this type, it is possible to obtain a measurement of neural activity with a very high spatial and temporal resolution [15].
However, ECoG also has drawbacks. Its biggest drawback is its relative invasiveness, requiring burr holes or craniotomies to place the electrodes [14]. For this reason, ECoG has most often been used in drug‐resistant epilepsy patients, as well as in tumor patients, undergoing resection.
7.2.1.2 Multielectrode Arrays (MEAs)
MEAs record activity from single neurons with excellent temporal resolution but sample this activity from a very small patch of cortex [16]. These two arrays are capable of recording and conducting neural activity with high SNRs because they interface closely with neuronal populations at depths not reachable by other neural interfaces.
However, a trade‐off of intracortical recording interfaces is the manifestation of acute and chronic complications due to tissue damage and neuronal cell death resulting from their invasive presence in the brain parenchyma. The active lifespan of these interfaces is largely dependent on its histocompatibility, which is dictated by a combination of factors such as the material, size, shape, and implantation and tethering modalities [17]. With current devices, the recordings degrade within a year or so [18].
In the most recent research, the signals recorded from the brain have been used to generate trains of electrical stimuli delivered to muscle nerves via surface electrodes, eliciting movements of the user’s own arm. In this case, the system is a combination of a BCI and a NS and has been called a “neural bypass” [19, 20] (Figures 7.1 and 7.2).
Figure 7.1 ECoG grid and activation thresholds. (a) Optical image of a typical high‐density ECoG grid (4 cm × 4 cm in size, 1 mm thick), (b) radiograph of a skull with the ECoG grid implanted, (c) the position of the ECoG grid on brain, and (d) activation thresholds were determined using G1 and G2 electrodes (depicted in black), serving as return electrodes for Electrodes 1–32 and Electrodes 33–64, respectively. Threshold values are expressed in terms of charge exchange per second (μC•Hz) necessary to induce a response. A gray line demarcates electrodes positioned above the motor and somatosensory cortices. Stimulation parameters included a pulse amplitude ranging from 1 to 7 mA, pulse width from 50 to 400 μs, and frequency from 50 to 500 Hz. The color bar illustrates the range of charge exchange values across the electrode.
(From Ref. [21], 2017, PLOS ONE, CC0 1.0).
Figure 7.2 The Utah Intracortical Electrode Array is depicted in this scanning electron micrograph, showcasing an electrode length of 1.5 mm. Each electrode tip is metallized with platinum, and a glass dielectric insulates each electrode from its neighboring ones.
(From Ref. [22], 1998, Elsevier).
7.2.2 Brain Stimulator Implants
Neurostimulators have been used to treat and relieve symptoms for neurological disorders, including pain, PD, epilepsy, tremor, dystonia, tinnitus, stroke, incontinence, gastroparesis (GP), and obesity [23].
From a theoretical point of view, the excitation of whole nerves or individual nerve axons and cell bodies with electrical stimulation is well understood. In practice, the main problem is to activate just those whole nerves or groups of axons within nerves that elicit the desired action and not others. Selectivity is a crucial issue in implants targeting structures within the CNS, where ensembles of neurons with opposite functions may be located less than a millimeter apart [11]. Overall, the neurostimulators can be classified into the following:
Noninvasive stimulation, through skin. Nerves innervating limb muscles are relatively accessible, allowing pulses of electrical current that activate them to be applied through the skin via self‐adhesive, conductive gel electrodes, conductive rubber electrodes coated with gel, or metal plate electrodes with an intervening layer of spongy material that is moistened with water.
Transcutaneous stimulation is commonly used in surface stimulators such as TENS (transcutaneous electrical nerve stimulation) and NMES (neuromuscular electrical stimulation) units.
Implantable stimulators: Only implantable brain stimulators will be covered in this review. The most common and studied implantable brain devices for stimulation are DBS and cortical visual implants.
Deep brain stimulator [24], for disorders characterized as oscillopathies, where patients’ symptom severity is correlated with excessive rhythmic neural activity at the DBS target region and projection targets (e.g. PD, essential tremor, and dystonia), high‐frequency electrical stimulation has been shown to suppress the rhythmic neural activity and concurrently alleviate patients’ symptoms [25]. DBS involves implanting a battery‐operated, electrical pulse generator underneath the clavicle, tunneling a lead wire to the top of the skull, and inserting, through a craniotomy, a linear electrode array deep into the brain. Typical DBS targets are the subthalamic nucleus and internal globus pallidus [26] (Figure 7.3).
Figure 7.3 The illustration portrays the configuration of a deep brain stimulation system, featuring the deep brain stimulation lead, lead extension, and their infraclavicular placement connected to the implanted pulse generator.
(From Ref. [27], 2011, Elsevier).
Visual cortex implant [28, 29] in this electrical stimulation can target different areas of the visual pathway, such as the retina, optic nerve, lateral geniculate nucleus (LGN), optic radiations, and the visual cortex, depending on the impaired region of the visual pathway. V1 stimulation is usually applied to patients with nonfunctional retinae or optic nerves, such as when complete blindness results from age‐related macular degeneration (AMD), retinitis pigmentosa (RP), glaucoma, or diabetic retinopathy. There are many reasons to use V1 as a site for implanting: It has uniform thickness and density of cells for central vision and peripheral vision and also has a well‐organized mapping of visual space onto neurons (visuotopic map). In addition, it has enough area for implantation of electrodes and is easy to access, compared to the retina, optic nerve, or LGN. The primary visual cortex also has a large magnification factor which means that more cortical tissue is devoted to a given visual field angle, and thus, the receptive field mapping on the cortex can be covered with more electrodes (if required) and higher resolution images can be evoked.
7.2.3 Brain Regenerative Medicine
One aspect that has not yet been covered in this review is regenerative medicine brain implants. This type of implant bases its operation mostly on stem cells. Advances in research on neural stem cells (NSCs) and pluripotent stem cells (PSCs) are expected to achieve wide clinical applications in the field of neural regenerative medicine [30]. Clinical trials have already started for several diseases including AMD [31], spinal cord injury (SCI) [32], PD, and Alzheimer’s disease (AD) using an embryonic stem cells (ESCs).
In this review, we will focus only on the implants located in the brain, so only for PD and AD therapies. Current medications and treatments for AD and PD that are designed to treat clinical symptoms are seldom successful. For example, dopamine and dopamine agonists administered to patients with PD and cholinesterase inhibitors are drug therapies for patients with AD. Though these drug interventions temporarily help with symptoms, they do not cure the damage to the neural tissue itself. Stem cell therapeutic approach is relatively new and emerging for both diseases [33].
There have been some clinical trials with stem cells for the treatment of AD. For example, mesenchymal stem cells (MSCs) have had much preclinical success and have supported the approval for human trials in the AD.
In a recent clinical trial, intracranially injected MSCs evaluated for safety and tolerability. No slowing of cognitive decline found at the next 24 months, and no decrease of AB pathology observed. None of the patients showed adverse side effects from the surgery and transplantation [34]. Another AD clinical study had more success using Nerve Growth Factor (NGF) producing fibroblasts. This study used gene therapy to program fibroblasts to produce NGF, which prevents the death and stimulates the growth of cholinergic neurons in the basal forebrain. All patients showed a response to the NGF through the growth of axons toward the site of transplantation. No adverse effects observed, and the study is currently moving forward to a phase 2 trial [35].
Also, regarding clinical trials using stem cell therapies for PD patients, many studies have been done. The tissue transplant to replace dopamine neurons existed for years. Buckland et al. performed one of the first to surgically transplant autologous medulla tissue into the caudate nucleus of two Parkinson’s patients in the year 1985. Neither of the patients significantly benefitted from the procedure. However, recent developments with autologous stem cells have proven to help restore the motor function in PD patients. One recent study [36] showed good results about intra‐arterial autologous stem cell implantation. The cells were directly implanted into an artery that feeds the substantia nigra, allowing the natural environment of the cells to determine a differentiation. No adverse effects observed, and an improvement in the quality‐of‐life scores was observed.
Using stem cells and growth factors, a common limitation seen in regeneration techniques is the method of delivery and the retention in the transplantation site. This problem is being addressed in recent developments using encapsulation with hydrogels. These are 3D scaffolds made of hydrophilic polymers which provide mechanical support in delivering growth factors [37]. Similarly, another type of scaffold using different materials and 3D printing was developed. This scaffold has the potential to deliver NSCs, which can then be cued to differentiate using low‐level light therapy (LLLT). LLLT used on 3D‐printed scaffolds could inhibit cell apoptosis and enhances migration and proliferation [38].
Also, nano‐neurology is a rapidly growing research field due to its versatile uses in medicine. Nanoparticles are shown to be useful drug delivery systems because they can pass through the blood–brain barrier (BBB) [39]. For Parkinson’s treatment, nanoparticles can be injected directly in precise brain regions without damaging surrounding areas.
7.3 Causes of Failure
It is anticipated that the progress in the neurological implant field has been hampered by a combination of technological and biological factors, such as the limited understanding of the long‐term behavior of implants, unreliability of devices, and biocompatibility of the implants, among others. This chapter aims to give an overview of the factors by which the implants interact acutely and chronically with the tissue.
The implantable brain devices described above consist mainly of electrodes implanted in the brain. For these reasons, the causes of failure will focus mainly on electrodes.
The nervous tissue response to implantable microelectrodes is a complex process characterized by a cascade of biochemical alterations and chemical reactions occurring at the level of the tissue material interface [40]. These biochemical and chemical alterations may lead in an undesired foreign body response. Additionally, the electrode must maintain its properties constant over time because in a long‐term implant, the loss of these could lead to a worsening of compatibility and durability.
Body fluids and components that make up body tissues prune the human body to be an environment extremely prone to corrosion. The presence of oxygen, macromolecules, and dissolved ions can lead to the detachment of the electrode surface.
For a better description of the causes of failure, it is necessary to distinguish when the failure is due to the implantation process itself or to a phenomenon that occurred after the implantation.
Once surgically implanted, microelectrodes must remain intact for several years to ensure the efficacy of the therapy and device functionality. To provide successful integration, reliability, and durability once implanted in the brain tissue, microelectrodes must fulfill the following requirements:
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Biocompatibility: This is an essential requirement for any implantable device. In the case of the electrodes, it concerns the nontoxicity of the electrode surface toward the cells. It must therefore not cause any adverse reactions with the surrounding tissues.
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Biomimicry
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