Chapter 2 Skin Surface-Mounted MEMS Dry Electrode
2.6 Summary
In this chapter, MEMS-based dry electrode (MDE) is presented for EEG recording. Related
experiments demonstrate its superior low electrode-skin interface impedance property without using any skin preparation. Next, a DS-MDE is also developed for better self-stabilized capability when being applied onto skin tissue. Practical application of drowsiness level monitoring in driving tasks shows reliable feasibility and competitiveness in driving error estimation examination comparing with standard wet electrodes. An alternative approach, transparent MDE, in photodynamic therapy application is realized by utilizing advanced hot-embossing method. The duplicated polymer-based micro-needle array can significantly reduce the required irradiation power and therefore avoid additional damage to surrounding skin tissue.
Chapter 3
Flexible Brain Surface Grid Electrode Array
3.1 Electrocorticography in Brain
Many neuroscience techniques have been used to study the functional relationships in the brain. In addition to the brain imaging techniques such as fMRI and PET, electrical recording techniques play an important role in brain mapping as well. ECoG recorded with non-penetrating grid electrodes are now one of a clinically applied methods to record the electrical activity in the brain. Comparing with the penetrating electrode probes, grid electrode arrays placed on the cortical surface can be used as a less invasive method in some clinical cases, e.g. in epilepsy surgery [42].
ECoG signals are composed of synchronized postsynaptic potentials (local field potentials), recorded directly from the exposed surface of the cortex. The potentials occur primarily in cortical pyramidal cells, and thus must be conducted through several layers of the cerebral cortex, cerebrospinal fluid (CSF), pia mater, and arachnoid mater before reaching subdural recording electrodes placed just below or on the dura mater (outer cranial membrane).
However, to reach the scalp electrodes of EEG, electrical signals must also be conducted through the skull, where potentials rapidly attenuate due to the low conductivity of bone.
For this reason, the signal magnitude of ECoG is much higher than EEG [43].
Because of the short distance between electrode and electrical source, neural activities on-set zone localization can be achieved with subdural electrodes, which is a critical imaging advantage for pre-surgical planning. ECoG offers a temporal resolution of approximately 5 ms and a spatial resolution of 1 cm [44]. It has been claimed that the spatial resolution of the present macro-size subdural electrodes usually used in clinic surgery could be substantially improved by using modern micro fabrication techniques [42].
Results in this chapter were co-worked with Yung-Feng Wang in Microsystem Control Laboratory, National Chiao-Tung University.
3.2 Design Criterion
To date, many types of grid electrode arrays have been reported for ECoG recording [45]-[49]. The materials used for electrode array structure include Polyimide, SU-8 and
Silicone, while the contact materials for sensing electrical activity usually are Gold, Platinum and Titanium Nitrite. According to the previous works, existing problems include:
(1) Different level of height of the passivation layer and sensing electrode due to the thin-film process, which can affect the recording efficiency. Post processing for electrode may need. (2) Low spatial resolution due to the large electrode size/distance between electrodes. (3) Un-predictable structure distortion due to that the thickness of Polyimide varies after required high curing temperature (200-400ºC).
Additionally, ECoG recording is an invasive process. Damage occurs while implantation, which consists of (1) Mechanical Damage, abrasion between sensor structure and cell tissue in setup. (2) Micromotion between skull and cortex surface, which cause abrasion in long-term implantation. (3) Immunity response cause by the implanted device [50]. Each of the damage described above can activate the glial scar in the neural tissue. The hyperplasias of the glial scar will then cover around the implanted device, provide attenuation property to the electrode and disable the functionality of the electrode eventually.
To overcome the disadvantage of the prior art, design key-point should includes
(a) Great flexibility and softness to reduce the damage while implantation and fit to brain curvature.
(b) Biocompatibility to reduce the immunity response.
(c) Reasonable electrode size and density to enhance spatial resolution (d) Low gap of height between passivation layer and sensing electrode
In this chapter, a flexible grid electrode array is designed and fabricated by parylene-C as substrate with platinum as sensing electrode material. A three-layered array structure is presented, which formed by 10um thick parylene-C, 60nm/350nm Titanium/Platinum and 1um thick parylene-C as substrate, electrode and passivation layer. Small micro-structure thickness allows minimal implantation damage [48]. Benefit to the thin and flexible property, the grid electrode array is expected to perfectly fit the brain surface curvature for electrical activities collection.
Poly-para-xylylene (Parylene) is a macromolecule polymer. Today, over 20 types of Parylene has been developed, but for numerous reasons, only 3 were deemed commercially viable: they are Parylene-C, Parylene-N and Parylene-D. The advantages of the use of parylene as the bulk-material compared with technologies based on the use of other materials such as polyimide [51] and silicon (developed for other purposes) [52] include parylene’s pinhole-free conformality, its low water permeability when compared with polyimide [53], its proven intraocular [54] and United States Pharmacopoeia (USP) Class VI biocompatibility, biostability, low cytotoxicity, its transparency, and its flexibility and mechanical strength (Young’s modulus ~4 GPa) [55].
3.3 Fabrication and Characterization
Fig. 3.1 illustrates the proposed fabrication process flow: (A) Sacrificial layer deposition on the substrate. (B) Parylene-C deposition using CVD technique. (C)-(D) Platinum electrode and interconnection wire definition using lift-off technique. (E) 2nd parylene-C deposition.
(F) Hard mask patterning for grid electrode array shape define. (G) Dry etching Parylene-C (H) Release structure by washing the sacrificial layer material.
Fig. 3.1 Fabrication process flow of the presented implantable grid electrode array Fig. 3.2 shows the photographs of the fabricated grid electrode array. (A) Top view. Right hand side denotes the recording area while left hand side of the structure is the bond pad area. (B) Packaged grid electrode array with a one cent coin. (C) Close-view of the recording electrode array. Parylene-C shows its transparent property. (D) The grid electrode array is packaged on a pre-designed PCB with connector (white) to be link to a cable wire for signal transmission. The grid array is 31.2mm in length, 4.3mm in width, and approximately 12um in thickness. Total 16 Platinum electrodes are fabricated with 500um in diameter, 750um in pitch. The width of routing wire is 100um.
To evaluate the impedance performance, the fabricated grid electrode array was characterized in physiologic saline solution (0.9% NaCl) under room temperature. The use of physiologic saline solution is because its electrolytic property is closed to implantation condition. As shown in Fig. 3.3, resulted impedance ranges between 12.6kΩ, -61.4° and 785.36kΩ, -12.32° when the testing input is set as 500mV in amplitude with scanning frequency from 40 Hz to 100 kHz. At 1kHz, test result shows that the impedance is about 1.64k Ohm and phase is -34.95°. Low impedance property induces less signal attenuation during recording therefore suit for the biopotential monitoring [56].
Fig. 3.2 Fabricated grid electrode array Fig. 3.3 Impedance characterization of single electrode
3.4 In-vivo Recording of Auditory Response in Rat
To demonstrate the functionality of fabricated gird electrode array, in-vivo experiment on auditory stimulation response recording is induced. When an authorized rat receives specialized frequency/magnitude sound stimulation, electrical response can be observed on a localized auditory area of cortex. It can be seen as neural activity happens due to the ears receive the sound and transfer into neural signals into brain and appear in auditory sensory brain area. By using the flexible grid electrode array, the interested area of brain is covered and recorded. Therefore, localized electrical response spot could be detected and characterized for functional mapping and event evaluation.
Sprague Dawley rats, aged 10–12 week old and weighting 280–350 g body weight, were used in the experiments. The animals were kept in a room under a 12:12-hr light-dark cycle with food and water provided ad libitum. All surgical and experimental procedures were reviewed and approved by the Animal Ethics Committee of the National Chiao-Tung University. The rats were anesthetized with urethane 2.0 g/kg body-weight (b.w.) intraperitoneal (i.p.). Subsequently, it was placed in a standard stereotaxic apparatus.
Proposed grid electrode array are implanted over the area of auditory cortex. Fig. 3.4 (A) shows the implanted location on the rat brain [57], and the practical implantation optical photograph is presented in Fig. 3.4 (B), which shows that the presented grid electrode displayed a flexible property therefore greatly fit to the exposed brain cortex. Fig. 3.4 (C) indicates the close view of the electrode array contact area. Note that only the electrode covered area is exposed to dura layer, the rest area is skull. Fig. 3.4 (D) illustrates the channel number with respect to the location in electrode array, where ch-14, ch-15, ch-16 acts as ground and ch-1 is used as the reference channel.
The audio stimulator use signal generator, amplifier, audio amplifier and programmable attenuator (TDT-RZ5, TDT-PA5, TDT-RP2.1, TDT-SA1). The stimulate signals are 1kHz, 2kHz, 4kHz and 9kHz in frequency (in the auditory range of rat, 250-60kHz) with 25ms
Fig. 3.4 Electrode implantation location, rat brain anatomy is adopted from reference [58]
duration, 2.5ms linear rise/fall time and 75dB sound pressure level (SPL). Every test sound is repeated for 100 times with 1 second period. Additionally, the speaker output is calibrated by a precise microphone (B&K 4149) before experiment.
Fig. 3.5 shows an example of recorded 16 channel signal in 1 second time frame. The upper green line in the figure illustrates the stimulate signal. The auditory response is then recorded and averaged to explicate the even-related evidence by software.
Fig. 3.5 16-channel recoding in 1 second time frame
Fig. 3.6 shows the time-magnitude plot of the averaged result of the 16 channel auditory ECoG response: (A) 1 kHz 75 dB SPL (B) 2 kHz 75 dB SPL (C) 4 kHz 75 dB SPL (D) 9 kHz 75 dB SPL. The 4 by 4 matrix boxes indicates the related electrode location on the brain. Vertical line in each box denotes the stimulation trigger marker. Four electrodes maker G (ground) and R (reference) are used as ground electrodes and reference electrodes.
Fig. 3.7 presents the time-Frequency plot: (A) 1 kHz 75 dB SPL (B) 2 kHz 75 dB SPL (C) 4 kHz 75 dB SPL (D) 9 kHz 75 dB SPL. Related frequency distribution versus different electrode location can be observed in the results.
Fig. 3.6 Time-magnitude plot of the averaged 16 channel auditory ECoG response (A) 1 kHz 75 dB SPL (B) 2 kHz 75 dB SPL (C) 4 kHz 75 dB SPL (D) 9 kHz 75 dB SPL.
Fig. 3.7 Time-frequency plot of the averaged 16 channel auditory ECoG response (A) 1 kHz 75 dB SPL (B) 2 kHz 75 dB SPL (C) 4 kHz 75 dB SPL (D) 9 kHz 75 dB SPL.
By comparing the recorded ECoG intensity, 4 kHz stimulation shows strongest response, while 2 kHz shows less response and 1 kHz, 9 kHz have weakest response signal. Measured results shows that localized neurons discharge phenomenon (gamma activities, 30 – 200Hz) are recorded by the presented grid electrode array. According to Fig. 3.6 and Fig. 3.7, three different auditory areas are covered by the grid electrode due to the different pattern observation. Group 1 includes 13, 11, 5, 3, Group 2 includes 8, 6, 10, 12, 2, 4, and 9, 7
belongs to the other group. Experiment results shows that recorded evoked potential on auditory cortex is about ±70uV, and the minimal effective sound stimulation magnitude is 20 dB SPL.
3.5 Summary
In this chapter, the implementation and characterization of a MEMS-based flexible grid electrode array utilizing parylene-C as substrate was presented for ECoG measurement applications. An un-symmetric sandwich-type structure consists of two Parylene layer (10um substrate and 1um isolation layer) and one Platinum layer was successfully fabricated. Comparing to previous works that using symmetric sandwich-type with thick isolation structure (polyimide, >20um), proposed method enhanced the adhesion property on brain cortex surface with great flexibility. Additionally, previous works suffer from the level difference between sensing electrode and passivation surface may need another electroplating post process high curing temperature (200-400ºC). In this work, simplified process under room temperature and superior properties were achieved without complex and high-temperature processes.
In-vivo experiments demonstrate the recording capability of the proposed grid electrode
array. Recorded auditory evoked potentials (AEPs) is ±70μV with 20dB SPL minimum sound level under general anesthesia, as well as the localized neurons discharge phenomenon (GAMA Activity, 30Hz~500Hz). The measurement result shows that the presented electrode array can cover the most of brain auditory cortex surface area, and distinguish specific signal characteristics between different electrode sites. Proposed grid array is looking forward to be used in awake animals in future works, which helps the studies on neurons degeneration, the mechanism of brain disease, and the development of the ECoG controlled BCI interface as well.Chapter 4
Three Dimensional Neural Probe Array
4.1 3-Dimensional Electrophysiology Recording of Biological Cells
In recent years, advance micromachined/assembled micro probe arrays with electrical recording ability have come to play an essential role in exploring central neural systems.
Simultaneous observation of a larger number of cell activities has become the general requirement to understand the nervous system [58]. Advances in neuroscience and neuroprosthetics now require microelectrode arrays that are able to access numerous neurons simultaneously with high spatial resolution [59]. Recording of the extracellular action potentials has been accomplished by surgically implanting neural probes into the target neurons of interest, which resulted from neural activities. Probes that could insert a large number of recording sites into neural tissues with minimal tissue damage are therefore needed. Also, the design of the probe arrays should be optimized for an experimental purpose that an electrode diameter of a few micrometers could support single-unit recording [60].
The traditional micro probes, which are made from insulated metal wires and glass micropipettes, cannot provide simultaneously multi-channel recording. The main reason is that the traditional devices function as only a single site on a single probe shaft. Some previous studies have improved the problem by thin-film lithography-based micromachining techniques since 1960s.
High-density probe arrays yielded insights into the organization and function of the neural system [61]. Silicon [62], glass [63], polymer [64] and sapphire [65] substrates have been employed as thin-film electrode probe planks. The thin-film silicon micro probe was developed many years ago for neuroscience and neural prostheses [66]. It has also been widely characterized electrically [67] and mechanically [68] for probe scaling [69], insertion force [70], tissue strain [71] and chronic brain responses [72]. The studies mentioned above provide detailed multi-channel recordings along a single plane, but lacked of full cell activity information in 3-D space [56].
To access the full cell activity that originates in the target tissue, three dimensional microprobe arrays are strongly required with precisely controlled dimensions and front-end circuitry compatibility. In other words, to achieve detailed studies of neural networks and
implementation of neural prostheses, we need to access three-dimensional volumes of tissue with three dimensional distributed recording sites. In modern neural system researches, 3D microprobe array allows the recording and mapping of the neural signal network and interconnections among the 3D brain structure. The recording and mapping would be impossible to achieve by using 2-D planar arrays [56].
Creating 3-D arrays by the assembly of 2-D parts is now the most popular method to construct a 3-D structure [59] [73-77]. The 2-D parts usually include 2-D arrays, vertical spacers and supporting platform. The supporting platform acts as a substrate, and the vertical spacers are erected on the supporting platform by tethers, joints and snap fasteners.
The spacers fixed the 2-D arrays vertically on the supporting platform, and made the probe shafts pass through the holes of the supporting platform. The full 3-D structure is therefore like a PC motherboard. Additionally, active circuitry for signal processing can be designed and fabricated in the back-end of the 2-D arrays to achieve system integration. A unique handling method was developed in [76] for a dual-side, ultra-thin silicon substrate process to fabricate thin probe shafts without using doping etching stop technique. Moreover, stacked probes and PCBs by anisotropic conductive film create the connection for the dual-side wire routing and 3D structure. Therefore, each side of probe can be wired out separately. An alternative solution provided in [77] integrated the silicon probe with flexible ribbon cables by using thermosonic bonded gold bump. Also, a platform with bays and gold clips is designed to connect with probes, which results in an impressive 3D device. The comparison of three-dimensional microprobe arrays with some major design parameters is shown in Table 4.1. However, the studies mentioned above neglect the importance of smaller opening for surgery implantation. Smaller opening of skull can reduce the implantation damage to the subject, prevent the rise of brain pressure, and decrease the infection probability of the wound.
Although previous work creating 3-D arrays by assembly of 2-D arrays successfully achieves high electrode density by packaging active probes onto the supporting platform with some micromechanical packaging technique, some problems still exist. First, previous approaches that use 2-D silicon probes to form full 3-D arrays required complex schemes for assembling submillimeter parts [74]. The main problem of such techniques is that the parts (spacers and supporting platform) were all assembled in orthogonal planes. Thus, perpendicular connectors for interconnections between orthogonal planes were required for signal transmission. Ultrasonic bonding [74] and vertical snap fasteners [59] have been proposed for perpendicular transfer pads, but they suffered from complex assembly steps and precise alignment equipment for 3-D assembly. For example, precise alignment was required to make probe shafts pass through holes of the supporting platform and steady the probe onto the vertical spacers without damage during the assembly process. Second, the probe arrays were fixed only by the perpendicular bonding pads and the tenons. Low structure strength can cause stability problem in implantation. Third, the rooms between the
spacers and the 2-D probes were wasted. The volume of a 3-D structure increases rapidly when increasing the number of 2-D probes. Notably, the supporting platform also defines the minimal surgery opening requirement for implantation. Signal processing chips can either be made in the backend of the 2D arrays or mounted onto the supporting platform to achieve system integration. The former currently provides only low-end circuit processes, and the latter results in tremendously increased volume size and surgical opening for implantation. In fact, fabrication of probe array structures with advanced CMOS process in the same wafer can be a large cost because the probe area may usually much larger then the area of the backend.
Table 4.1 Comparison of 3D microprobe arrays with some major design parameters
Reference [9-10] [77-80] [81] [59, 73-75] [77] [76]
Substrate Si Epoxy,
Polyimide Polyimide/Nickel Si Si Si
Dimension 3D 3D 3D 3D 3D 3D
To improve the problems described above, this work reports a new stacking method for fabricating 3-D neural probe arrays. In this study, the 3-D orthogonal interconnection was replaced with 2-D wire bonding by the present stacking method, and the perpendicular bonding and snap fasteners which were used in previous work were no longer needed.
Compared to previous work, this new stacking method can also provide reliable structure strength. ASIC chips can be substituted for spacers to increase the system integration and volume usage efficiency as well. Additionally, an anti-overflow design based on the capillary principle was exploited to avoid gel overflow onto proximate bonding pad during 3-D array assembly.
4.2 Stacked Multichip Structure Design
A new stacking method to produce three-dimensional neural probe arrays is presented in this work. This method creates 3-D probe arrays by assembling 2-D arrays and spacers layer by layer, as shown in Fig. 4.1. For a 4 × 4 3-D array, four 2-D arrays (gray color) with four probes in each array and three spacers (yellow color) were required. Compared to exising three-dimensional neural probe designs, the present stacking method improved the inconvenient assembly steps which include orthogonal assembly and perpendicular connection techniques. In the stacking method, the shapes of each 2-D arrays were carefully designed so they can be wire-bonded individually with different height levels. Spacers with an anti-overflow mechanism were also proposed in this paper. The present anti-overflow mechanism can also be realized on 2-D arrays if active circuit chips are used as spacers.
Also, the thickness of the spacer determined the spacing between two 2-D arrays. Each planar 2-D array, electrode sites, inter-connection routing and bonding pads were located in the same plane. The bonding pads were arranged on the different sides of four 2-D probe arrays for wire bonding. Therefore, each 2-D array can be wire-bonded individually and the
Also, the thickness of the spacer determined the spacing between two 2-D arrays. Each planar 2-D array, electrode sites, inter-connection routing and bonding pads were located in the same plane. The bonding pads were arranged on the different sides of four 2-D probe arrays for wire bonding. Therefore, each 2-D array can be wire-bonded individually and the