DESIGN AND FABRICATION OF POLYNORBORNENE- AND LIQUID CRYSTAL POLYMER-BASED ELECTRODE ARRAYS FOR BIOMEDICAL APPLICATIONS ALLISON ELIZABETH HESS

Transcription

1 DESIGN AND FABRICATION OF POLYNORBORNENE- AND LIQUID CRYSTAL POLYMER-BASED ELECTRODE ARRAYS FOR BIOMEDICAL APPLICATIONS by ALLISON ELIZABETH HESS Submitted in partial fulfillment of the requirements For the degree of Master of Science Thesis Advisor: Professor Christian A. Zorman Department of Electrical Engineering and Computer Science CASE WESTERN RESERVE UNIVERSITY May, 2008

2 CASE WESTERN RESERVE UNIVERSITY SCHOOL OF GRADUATE STUDIES We hereby approve the thesis/dissertation of candidate for the degree *. (signed) (chair of the committee) (date) *We also certify that written approval has been obtained for any proprietary material contained therein.

3 For my grandmother

4 Table of Contents List of Tables... 5 List of Figures... 6 List of Figures... 6 Acknowledgements Introduction Neural Interfaces Conventional Neural Interfaces Materials for flexible microfabricated nerve electrodes Chapter 2 : Evaluation of PNB as a Material for Implantable Electrodes Motivation for using PNB Hermeticity testing using interdigitated electrodes Motivation Method Design Fabrication of all-pnb IDEs Fabrication of hybrid PNB/LCP IDEs Test Setup Discussion PNB Biocompatibility Tensile Testing of PNB Structures Conclusion Chapter 3 : Design #1 -- Microfabricated FINE Design

5 3.1 Motivation Flexible Array Design Silicone Housing Connector Discussion Chapter 4 : Fabrication of All-PNB Design #1 Devices All-PNB Devices with SiO 2 Sacrificial Layer and Ti/Pt Metallization Fabrication process Results All-PNB Devices with SiO 2 Sacrificial Layer and Cr/Pt Metallization All-PNB Devices with Aluminum Sacrificial Layer and Ti/Pt Metallization Base Layer Patterning Metal Layer Patterning Capping layer patterning Device Release Summary Chapter 5 : PNB/LCP Electrode Arrays Rationale Device Fabrication Discussion Wet LCP Etch RIE LCP Patterning

6 5.3.3 Metal/LCP Interface Chapter 6 : Design #1 Evaluation Something Pulse testing in saline Test Setup Results Discussion Finite Element Analysis Flexing-Conductivity Measurements Hybrid PNB/LCP devices tested on Aplysia buccal ganglion Introduction Experimental Stimulation Recording Summary Chapter 7 : Design #2 -- Single-Substrate Peripheral Nerve Electrode Arrays Motivation Design Device Fabrication In vivo experiments Discussion Chapter 8 : Conclusion and Future Work Applendix A LCP Bonding

7 A.1 Liquid Crystal Polymer A.2 LCP Bonding Experiments A.3 Patterning bonded LCP sheets A.4 Patterning Non-bonded LCP with Non-copper etch mask A.6 Bonding Foil to LCP A.7 Discussion of LCP-based devices Appendix B: Alternative PNB Releases B.1 Unity TM as a Sacrificial Release Layer B.1.1 Unity TM Background B.1.2 Unity as a Sacrificial Layer for Avatrel TM -based Devices B.1.3 Experimental Methods B.1.4 Unity Discussion B.2 Alternative PNB-based Device Release Methods B.3 Aluminum Etch Release References

8 List of Tables Table 1-1: Electrical, mechanical, and moisture absorption properties of various polymers typically use in microfabrication processes Table 2-1: Avatrel TM 2585P mechanical and electrical properties Table 2-2: Total number of cells counted on each substrate after 7 days Table 2-3: Contact angle measurements for an oxide layer on a Si wafer and for PNB to determine the degree of hydrophobicity [49] Table 2-4: Comparison of Young s modulus, percent strain at break, and tensile strength of unsoaked PNB strips and PNB strips soaked in PBS for 72 hours at 87ºC Table 4-1: Summary of results from each type of metallization and release Table 6-1: Mean and standard deviation of measured resistance across trace 9 of hybrid PNB/LCP device in different curvature configurations Table 6-2: Calculated values for resistance per unit length in each configuration, based on both the mean measured resistance for each configuration and the first order geometric model developed in Chapter Table A-1: Description of each LCP-LCP bonding experiment in wafer bonder Table B-1: Documentation of non-sacrificial layer alternative release mechanisms attempted for the release of the microfabricated FINE structures

9 List of Figures Figure 1-1: Cross sectional schematic of a FINE closed around a nerve [5] Figure 1-2: Optical photograph of conventional FINE devices Figure 1-3: Schematic of a multichannel recording microelectrode array (Michigan Probe) [12] Figure 1-4: Diagram of three-dimensional microelectrode array made from multiple planar arrays [14] Figure 1-5: Scanning electron micrograph of Utah Intracortical Electrode Array [16] Figure 1-6: Flexible cortical array with probes fabricated on a planar substrate and bent to form 3D array [22] Figure 1-7: Typical polyimide fabrication process utilizing two metal layers [23] Figure 1-8: Flexible parylene-based electrode array for spinal cord simulation and recording [27] Figure 2-1: Keesara s all-pnb device [30] Figure 2-2: Schematic of design of IDE with capping layer. The IDEs used had 25 digits per electrode, each 1 cm long, 100 µm wide, and separated by 50 µm Figure 2-3: Cross-section views showing all-pnb IDE fabrication process Figure 2-4: Cross-section views showing hybrid PNB/LCP IDE fabrication process Figure 2-5: Photographs of all-pnb and PNB/LCP 50-digit IDEs Figure 2-6: Setup used for leakage current measurement of interdigitated electrodes Figure 2-7: Baseline measurement of leakage current in PNB IDE over 8 hours at 5-V bias without saline Figure 2-8: Measurement of leakage current in PNB IDE over 100 hours at 5-V bias

10 Figure 2-9: Leakage current measurement over time exposed to saline for the hybrid PNB/LCP IDEs Figure 2-10: Result of microtensile testing of PNB strips #1 and #2 that had not been soaked in PBS Figure 2-11: Result of microtensile testing of PNB strips #1 and #2 that had been exposed to saline Figure 2-12: Result of microtensile testing of PNB strips #3 and #4 that had not been exposed to saline Figure 2-13: Result of microtensile testing of PNB strips #3 and #4 that had been exposed to PBS Figure 3-1: Plan view of microfabricated FINE design Figure 3-2: Base layer photomask for microfabricated FINE arrays Figure 3-3: Metal layer photomask for microfabricated FINE arrays Figure 3-4: Capping layer photomask for microfabricated FINE arrays Figure 3-5: Electrode end of microfabricated device with vias for each contact. Dimensions are in microns Figure 3-6: Connector end of device with vias for each contact. Dimensions are in microns Figure 3-7: Custom-molded silicone housing for microfabricated FINE Figure 3-8: Jig used for the purpose of holding the housing and insert in place while silicone was applied to the housing and the backside of the flexible electrode array Figure 3-9: Hirose FH-27 flex circuit connector soldered to a printed circuit board

11 Figure 4-1: Process sequence for fabricating PNB-based electrode arrays with HF Irelease Figure 4-2: Microfabricated FINE arrays with PNB base and capping layers Figure 4-3: Two all-pnb microfabricated electrode arrays in a molded silicone housing. The housing is splayed open to demonstrate the flexibility of the structure Figure 4-4: Pt electrode pad with Ti adhesion layer on a HF-released PNB-based electrode array Figure 4-6: SEM micrograph of a Pt electrode pad with Cr adhesion layer on a HFreleased PNB-based electrode array Figure 4-7: Process sequence for fabricating PNB-based electrodes with anodic aluminum release Figure 4-8: Avatrel TM 2585P cure cycle [30] Figure 4-9: Polarization behavior of Cr and Al layers at potentials between -1.5 V and 1 V in 2M NaCl solution [58] Figure 4-10: Schematic of anodic aluminum dissolution used as a release mechanism for the PNB-based devices Figure 4-11: (left) Electrode-end opening and (right) Pt traces from anodic aluminum dissolution-released PNB-based device with cures at 160ºC Figure 5-1: Process sequence used to fabricate hybrid PNB/LCP FINE arrays Figure 5-2: Microfabricated FINE array with LCP base and PNB capping layer. A closer look at the electrode end of the device shows the roughness of the LCP being translated to roughness on the metal surface

12 Figure 5-3: Two microfabricated FINE arrays with LCP base and PNB capping layer embedded in molded silicon housing Figure 6-1: Setup for current pulse testing in saline Figure 6-2: Typical output waveform for pulse testing of PNB-based electrodes Figure 6-3: All-PNB structure that failed upon handling Figure 6-4: Finite element modeling of stresses in microfabricated structure when connector end is anchored and a downward force is applied at the electrode end Figure 6-5: Flexed configuration of a PNB/LCP device Figure 6-6: Geometric model to determine the change in cross-sectional area of the metal: (left) traces are curled outward, away from the tubing; (right), traces are curled inward, facing the tubing Figure 6-7: Plot showing the cross-sectional area of the metal (perpendicular to the length of the trace) as a function of trace curvature and configuration calculated from both the mean measured resistance of the trace as well as the geometric model developed in this chapter Equation 6-6: Layout of pinned down buccal ganglion and nerves [66] Figure 6-8: Recordings at buccal nerve 2 when the ganglion was stimulated with electrode 6 from a microfabricated FINE array Figure 6-9: Stimulation setup with array on the underside of the ganglion Figure 6-10: Close-up photograph of electrode array on the underside of the ganglion.113 9

13 Figure 6-11: Comparison of conventional extracellular recording at buccal nerve 2 with the recording of action potentials on electrode 6 of the flexible, microfabricated array Figure 7-1: Mask layout for peripheral nerve electrode arrays. Devices labeled a are recording electrodes, b and c are stimulating electrodes. Magenta corresponds to the outline of the base layer; green corresponds to the outline of the metal layer; blue corresponds to the outline of the capping layer Figure 7-2: Schematic of single-substrate peripheral nerve electrode stimulation array.120 Figure 7-3: Fabrication sequence for single-substrate peripheral nerve electrode arrays Figure 7-4: Completed dual array peripheral nerve electrode device for stimulation Figure 7-5: Completed PNB/LCP peripheral nerve electrodes in flex circuit connectors soldered to circuit boards Figure 7-6: Microfabricated peripheral nerve electrode blank sutured around a canine nerve Figure 7-7: Schematic showing location of implanted electrode arrays Figure 7-8: No stimulation Figure 7-9: Flexion of fingers upon stimulation of Channel 1 of the proximal array Figure A-1: Results from LCP lamination using hot plate and steel roller. This shows a square cover layer laminated wafer-shaped base layer. Air bubbles are circled in red. [30] Figure A-2 :Cross-sectional view of bonding layers Figure A-3: Photograph of bonded LCP sheets demonstrating flexibility

14 Figure A-4: a) Photograph of completed patterned, bonded, non-metalized LCP structures b) Photograph demonstrating flexibility of bonded structures Figure A-5: (Top Right) traces stemming from shaft to electrodes appearing to be broken at interface. (Top Left) traces stemming from shaft to electrodes enlarged. (Bottom) connector end of the device with traces appearing to be cut at capping layer interface Figure B-1: Avatrel TM base layers after curing on photodefinable Unity TM Figure B-2: Avatrel TM base layers after curing on non-photodefinable Unity TM

15 Acknowledgements Foremost, I would like to express my sincere gratitude to my advisor, Prof. Zorman, for his enthusiasm, support, and wealth of ideas and information. With his help, I have improved as a researcher, expanded my skill set, and was exposed to a variety of interesting projects. I am also thankful for the opportunities, and insight presented by Prof. Tyler, which has encouraged me to keep thinking about ways to improve the devices. I would also like to thank Prof. Mohseni for serving on my thesis committee and for providing valuable suggestions. Additionally, I would like to express sincere thanks to Jeremy Dunning for his suggestions and assistance, training, many of the photographs and illustrations in this thesis, as well as for being a good friend. I would also like to thank past and present members of the research group, who have provided encouragement and help, and from whom I ve learned a great deal. Likewise, I would like to thank Prof. Tyler s group for their efforts in collecting data that supports my thesis. I also want to thank Dave Greer from the EDC for the metal deposition and tips for Avatrel processing. I would also like to thank Ed Jahnke from the MFL, who was always able to help when I had problems with or questions about equipment. I would also like to thank members of the Promerus, Inc. technical staff, who were wonderful in supplying different Avatrel and Unity formulations, as well as in helping with testing, or offering suggestions to improve the processes. Also, I would like to thank Rogers Corp. for supplying the LCP that was used for work toward this thesis. Lastly, I would like to thank my parents for always supporting me, encouraging me, and otherwise helping me wherever possible. Without question, having my parents 12

16 living nearby helped me get through some difficult events of the past couple years, but I have also enjoyed celebrating successes with them. I cannot thank you enough for everything you do. 13

17 DESIGN AND FABRICATION OF POLYNORBORNENE- AND LIQUID CRYSTAL POLYMER-BASED ELECTRODE ARRAYS FOR BIOMEDICAL APPLICATIONS Abstract by ALLISON ELIZABETH HESS Conventional peripheral nerve electrodes are often hand-made devices that incorporate platinum foil contacts and stainless steel wires, requiring significant skill and time to fabricate, and the contact density is limited by the manufacturing method. Microfabrication allows for batch processing, eliminates hand assembly, and permits a much higher contact density. Liquid crystal polymer (LCP) is a good choice as a substrate material for flexible, microfabricated electrode arrays because it is mechanically robust and resistant to moisture absorption. However microfabrication of multilayerd LCP devices is challenging because lamination is required. This thesis describes the development of peripheral nerve electrode prototypes fabricated using LCP, polynorbornene and thin film platinum. The fabrication process involves photolithographic patterning of the platinum and PNB layers, and laser micromachining of LCP. Benchtop testing was performed to evaluate the mechanical, electrical and hermetic properties of the devices. Preliminary in-vivo evaluation showed that the prototypes could be used for stimulation. 14

18 1. Introduction 1.1 Neural Interfaces Conventional Neural Interfaces Functional electrical stimulation (FES) is a technique in which short electrical pulses are used to restore nerve activity to patients experiencing paralysis from such disabilities as spinal cord injury or stroke. The electrical pulses generate action potentials in neurons, which result in muscle contractions [1]. FES has developed and expanded in the past several decades, as demonstrated by the evolution from muscle-based electrodes to the more efficient approach of peripheral nerve stimulation. Stimulating electrodes have progressed accordingly, from surface electrodes to percutaneous electrodes for direct muscle stimulation, then to cuff and intra-neural electrodes for peripheral nerve stimulation. Peripheral nerve stimulation is a more efficient technique than muscle-based stimulation techniques since one peripheral nerve array can stimulate many muscles to produce a variety of motions, while muscle-based electrodes require that electrodes are placed in or on each muscle to be activated. Peripheral nerve stimulation takes advantage of the idea that despite the fact that the neural pathways between the central nervous system and the peripheral nervous system have been interrupted, the peripheral nerves remain intact. The peripheral nerves can be stimulated with electrical pulses to produce useful muscle contractions. There are several challenges in using peripheral nerve stimulation to allow a patient to regain use of upper or lower extremities, including signaling a stimulator to send a current pulse to the correct electrode at the correct time, powering a chronically 15

19 implanted device, and allowing for stimulation of multiple extremities. In this thesis, the focus is on the design of electrodes for nerve interfacing. Peripheral nerve electrodes are designed either as intraneural electrodes or extraneural electrodes. The intraneural electrodes are used to electrically stimulate and record from small numbers of neurons. These penetrating electrodes typically have the form of one or two needles or wires [2, 3], but multi-electrode arrays have also been used in intraneural peripheral nerve stimulation [4]. The selectivity demonstrated by electrodes of this type is high compared to extraneural electrodes, but since these are penetrating electrodes, the potential for short- and long-term nerve damage is more of a concern than with non-penetrating electrodes. The flat interface nerve electrode (FINE) is a peripheral nerve cuff electrode [5]. This particular nerve cuff is unique in that it takes advantage of the oblong cross-section of the nerve to maximize the interfacial area between the electrode array and the nerve. While other cuff electrodes, such as the traditional cuff electrode [6] or the spiral cuff [7, 8], utilize a circular cross-section, the FINE has a rectangular cross section, as shown in Figure 0-1. This rectangular cross-section brings the electrodes into closer proximity to the nerve fascicles than is possible in the electrodes with circular cross-sections. This increases the stimulation selectivity of the nerve fascicles for a given current pulse magnitude, as compared to cuff electrodes with circular cross sections. This design may also reduce the amount of current required for stimulation compared to the required current for cuff electrodes, reducing the amount of power needed for stimulation. 16

20 Figure 0-1: Cross sectional schematic of a FINE closed around a nerve [5]. The FINE devices are made by first molding a silicone elastomer into the correct shape. Eight or twelve platinum foil contacts (0.5 x 0.5 x mm) are embedded in the silicone and spot-welded to Teflon-coated stainless steel wire. A 0.4 mm diameter window is opened through the silicon to the contacts. A photograph of these devices is shown in Figure 0-2. Figure 0-2: Optical photograph of conventional FINE devices 17

21 1.1.2 Microfabricated Neural Interfaces While hand-fabricated nerve electrode arrays like the FINE have proven to be successful in selective stimulation of peripheral nerves, these devices do have significant limitations. First, fabrication of each device requires significant skill and a large amount of time to produce. Hand-fabricated devices inherently suffer from device-to-device variability with respect to electrode size, window size, electrode placement, as well as in aspects such as electrode wrinkling and wire bending from handling, and variations in the electrode-wire weld. The density of contacts is limited by the size of electrodes that can be handled practically, as well as the precision with which they can be positioned. These limitations have resulted in a push toward the development of microfabricated electrode arrays for neural stimulation and recording. Microfabrication techniques, including photolithography, thin film deposition, and wet and dry etching, support batch processing, allowing for many devices to be produced at the same time. The fabrication parameters, such as photoresist thickness, exposure time, metal thickness, and etch parameters, can be closely controlled to reduce the variability from device to device. Further, because microfabrication techniques and tools do not rely on manual manipulation to position each electrode discretely, they enable the fabrication of dense arrays, which is valuable for selectivity in stimulation and recording in some neural applications. Many of the microfabricated neural interfaces are designed as cortical interfaces implanted directly into the brain for recording and stimulation. Two of the most notable microfabricated neural interfaces are the planar Michigan probe and the three-dimensional Utah Electrode Array. Both use silicon substrates and silicon bulk 18

22 micromachining techniques to form the probes. Other microfabricated silicon-based neural electrodes have been developed by other groups, but these are often variations on the Michigan probe and Utah array designs, with variation in shape, fabrication sequence, or materials [9] [10]. An early version the Michigan Probe is shown in Figure 0-3. A typical probe for chronic implants consists of one to four shanks connected to a flexible silicon ribbon cable [11]. The basic fabrication process for such devices begins with selective doping of a silicon substrate with boron through a patterned thermal oxide mask. Alternating layers of SiO 2 and Si 3 N 4 are deposited, then the conductor metal or polysilicon is deposited and patterned. A second set of alternating SiO 2 and Si 3 N 4 layers are then deposited and patterned with a plasma etch to open windows to the recording sites and to the bonding pads. Leaving the photoresist etch mask on the wafer, gold is deposited and patterned using lift-off, leaving only the recording sites and bonding pads gold-coated. The dielectric layers are then patterned in the shape of the probes, and the silicon wafer is thinned from the back in an isotropic wet etch. The individual probes are released in an unmasked etch in ethylene diamine-pyrocatechol, which dissolves the silicon wafer, but not the highly doped probes. 19

23 Figure 0-3: Schematic of a multichannel recording microelectrode array (Michigan Probe) [12]. Probes built on silicon substrates have the advantage of being inherently ICcompatible. As such, on-chip circuitry for signal processing can be placed directly on the silicon probe [12]. The Michigan probe has been further developed to include microfluidics channels embedded into the substrates for drug delivery [13]. A microelectrode array developed by Hoogerwerf and Wise [14] uses an array of the planar Michigan probes to allow for true 3-D recording (Figure 0-4). This array is comprised of a microfabricated silicon platform, two spacer bars, and the planar probes. Each probe is inserted through the slots in the silicon platform to produce an array of spikes. 20

24 Figure 0-4: Diagram of three-dimensional microelectrode array made from multiple planar arrays [14]. The Utah Intracortical Electrode Array [15], shown in Figure 0-5, is a threedimensional silicon-based structure made up of a 10 x 10 array of 1.5 mm long Si needles projecting from a 4.2 mm x 4.2 mm x 0.2 mm thick substrate. The needles are each approximately 80 µm wide at the base and taper to a sharp tip. The tips of the needles are metalized with platinum. Polyimide is used as a passivation layer, insulating all but approximately 50 µm at the tip of the needle. Platinum contact pads are also formed on the back side of the substrate to provide electrical access to each of the electrodes. The electrodes are electrically isolated with glass dielectric in the base around each needle. 21

25 Figure 0-5: Scanning electron micrograph of Utah Intracortical Electrode Array [16]. The Utah Arrays are fabricated from a three inch, 1.7 mm thick n-type silicon wafer. Thermomigration is used to create paths of p + type silicon traversing the thickness of the silicon wafer. These paths will eventually become the needles in the array and are electrically isolated by the back-to-back pn junctions that result from the p + paths in the n-type wafer. Next, a dicing saw is used to cut 1.5 mm into the wafer along the n-type grid to form a 10x10 array of p + -type rectangular columns. A two-step etch process is then used to etch the columns into the needle shapes. Both used isotropic Si etches in 5% HF and 95% H 2 NO 3, but in the first step, a process is used to ensure that the entirety of the columns were etched isotropically, and in the second step, the tips were sharpened and polished. To selectively deposit a layer of gold, then a layer of platinum on the needles and to ensure that each needle remains electrically isolated from the others, the tips of the needles are poked through a thin piece of metal foil, which acts as a mask. Thus, about 1.0 mm of each needle is coated with gold and platinum. The array is insulated with polyimide by placing the array onto a surface with the electrodes pointing upward, then flooding the array with a polyimide solution. This leaves a thick layer of polyimide on the base, and a thinner layer along each electrode. A metal foil mask 22

26 technique is again used to mask all but 0.5 mm at the tip of the electrodes such that the polyimide at the tip of each electrode can be removed in an oxygen plasma etch. Silicon-based devices like the Michigan probe or Utah array have the advantages of being inherently biocompatible, since silicon is recognized as a biocompatible material. Further, these devices have the ability to incorporate on-chip signal processing capabilities using integrated circuit technology developed for silicon substrates. However, the neural applications of these devices are limited partly because of the stiffness of silicon which has a Young s modulus of approximately 190 GPa, as compared with brain, nerve, and other tissue which has a Young s moduli on the order of 1 to 100 kpa. For chronic implantation, it can be expected that the neural tissue, whether is be cortical tissue or nerve tissue, will experience some motion in the body. A stiff implanted device may not move with the tissue, resulting in a relative motion between the electrodes and the tissue. This relative motion may lead to poor recordings or stimulation as the electrode positioning with respect to the nerve or neurons may change. Furthermore, this relative motion may result in tissue damage. Moreover, an unyielding substrate such as silicon cannot conform to the non-planar geometry of the brain or peripheral nerves. As a result, flexible substrates have been investigated for both cortical and peripheral nerve interfaces. 1.2 Materials for flexible microfabricated nerve electrodes Flexible biomedical microdevices for chronic implantation require a biocompatible, compliant substrate that exhibits a high tensile strength and a high moisture resistance. The high tensile strength is required to ensure that the device has 23

27 adequate mechanical robustness to survive the potentially harsh implantation procedure, as well as the forces that may be applied during chronic implantation. An insulation layer, which may or may not be the same as the substrate material, must also be biocompatible, flexible, and be highly resistant to moisture absorption. Moisture absorption of the substrate and insulation layers must be minimized to prevent crosstalk between electrodes and electrode corrosion. Polymers that have been used for microfabricated nerve electrodes include polyimide, parylene, PDMS, and benzocyclobutene (BCB). These, and other polymers commonly processed with microfabrication tools and techniques, are compared in Table 0-1. Dielectric constant Water absorption Young's Modulus Tensile strength % (GPa) (MPa) Polyimide < Parylene C PDMS 3 n/a 7.5 x 10-4 <12.5 SU n/a BCB LCP R/flex Avatrel 2585P Table 0-1: Electrical, mechanical, and moisture absorption properties of various polymers typically use in microfabrication processes. Polyimide is one of the more popular polymers used as substrate and insulation layers in flexible microfabricated nerve electrodes. Rodriguez et al. employed polyimide as the substrate and insulating materials in a tubular, self-sizing peripheral nerve cuff that is capable of both recording and stimulation of peripheral nerves [17]. The same group developed a sieve electrode of the same materials developed to interface with a regenerating peripheral nerve [33]. Boppart, et al. described a flexible, planar microelectrode array to record potentials in brain slices [19]. The innovation in this 24

28 device was the etching of perforations through the polyimide to allow artificial cerebrospinal fluid to better circulate to the recording surface of the tissue and increase the viability of the brain slices used in the recordings. Microfabricated, flexible electrode arrays on polyimide with perforations were also developed by Gonzalez and Rodriguez [20], although these electrodes were designed for use in nerve and muscle tissue recordings. The popularity of polyimide as a substrate and passivation material in microfabricated, flexible electrode arrays is further exemplified by Rousche, et al. [21], who implemented a Michigan probe-like design in polyimide. A flexible cortical electrode array was developed by Takeuchi, et al. [22], shown in Figure 0-6. The probes were fabricated on the planar substrate, but then bent under the influence of a magnetic field to create a three dimensional array. Figure 0-6: Flexible cortical array with probes fabricated on a planar substrate and bent to form 3D array [22]. Polyimide (PI) devices are typically fabricated using a process similar to that shown in Figure 0-7. PI is first spun onto a Si wafer and cured. Next, a photoresist lift-off mask is patterned and the first metal layer is sputter-deposited. A second PI layer that acts as insulation for the first metal layer is then spun on and cured. Next, a second photoresist lift-off mask is patterned and the second metal layer is sputter-deposited and patterned. 25

29 A final PI layer is then spun on to insulate the second metal layer. An Al etch mask is deposited and patterned with a wet etch to define windows to the electrode sites and the outer shape of the devices. The PI is then patterned with reactive ion etching (RIE). The Al etch mask is then stripped with Al etchant. Tweezers are used to mechanically strip the devices from the wafer [23]. Figure 0-7: Typical polyimide fabrication process utilizing two metal layers [23]. The techniques used in the fabrication of polyimide-based nerve electrodes are similar to techniques used in surfacing micromachining of silicon-based MEMS devices. Therefore, special tooling or radically new techniques are not required for the fabrication of this type of device. However, as can be seen from Table 0-1, polyimide has one of the highest rates of moisture absorption in comparison to the other polymers used in microfabrication. At this rate of moisture absorption, it is likely that the device will remain viable for only a couple years, at best. The eventual goal of our group is to 26

30 fabricate a device that will be suitable for implantation for 20 years. One method to achieve this goal is to develop devices from micromachinable polymers that exhibit extremely low moisture absorption characteristics. An alternative flexible substrate material to parylene is polydimethylsiloxane, or PDMS. This is the same material used for the molded structure in the FINE, and can be refined to a medical grade material suitable for implants. PDMS is not photodefinable, instead often being patterned using micromolding. However, PDMS can be spin-coated and patterned using a technique similar to metal lift-off [24] or with a laser [25]. PDMS is attractive for neural interfaces for stimulation and recording because it is very flexible, biocompatible, and chemically inert. As such, PDMS is being used as a substrate for a retinal electrode array [24]. On the other hand, PDMS exhibits a very low tensile strength compared to other biocompatible polymers. Additionally, metal adhesion to PDMS is poor. Therefore, PDMS-based electrode arrays require either complicated processes to promote metal adhesion, or will suffer from metal delamination over time [26]. Further, PDMS shrinks when cured, which can cause difficulties in aligning subsequent layers to the PDMS base structure [26]. Parylene is a third biocompatible polymer that has been investigated for use in neural electrode arrays. A parylene-based spinal array is shown in Figure 1-8 [27]. Parylene exhibits a high resistance to moisture absorption compared to polyimide, but a lower tensile strength. Parylene is not spin castable, and not photodefinable. Instead, parylene is deposited by chemical vapor deposition and patterned using reactive ion etching (RIE) in an oxygen plasma. There are concerns, however, regarding the adhesion properties of parylene. Further, because parylene is vapor deposited, it can be difficult to 27

31 produce high-quality, thick layers [28]. Additionally, there are reports of parylene coatings cracking after long-term implantation [29]. Figure 0-8: Flexible parylene-based electrode array for spinal cord simulation and recording [27]. Based on the properties detailed in Table 0-1, it appears that LCP is a good choice as a polymer for use in a microfabricated peripheral nerve electrode interface due to its relatively high tensile strength and low moisture absorption, yet LCP is not a material that has been widely used in the development of neural interfaces. This may be attributed to LCP being difficult to process using standard microfabrication tools and techniques. In contrast to polymers that can be vapor-deposited or spin-casted, LCP is available as a cross-linked polymer in pressed sheets. Because one of the principal uses of LCP is as a flexible printed circuit board substrate, it can be acquired from commercial vendors replete with copper cladding. A multilayered LCP device with embedded thin film metal electrodes is very difficult to produce in comparison to the same device design made in polyimide. LCP layers are stacked using lamination, which requires high pressure and a temperature near the melting point of LCP (295º C). Patterned copper structures made from the cladding layers (typically 18 µm in thickness) remain viable 28

32 through a lamination process, but this may not be true for thin film, biocompatible metals with thickness on the order of 0.25 µm. As such, the development and fabrication of a multilayered LCP structure with embedded thin film electrodes is a very challenging task. The main hypothesis of this thesis is that a flexible, microfabricated nerve electrode array suitable for acute testing on a mammalian peripheral nerve can be produced using liquid crystal polymer, polynorbornene, and thin film platinum as the main materials of construction. This hypothesis is evaluated through the development of polynorbornene-based and hybrid liquid crystal polymer/polynorbornene device fabrication processes, building upon previous work in Prof. Zorman s group [30]. Two electrode array designs for peripheral nerve stimulation are developed, with devices realized using the aforementioned fabrication processes. These devices are used to evaluate the hypothesis that the thin film electrodes have a adequately small resistance to allow for nerve stimulation. Further, preliminary evaluation of PNB will support the hypothesis that PNB is a suitable material for use in acute testing conditions. Finally, it will be shown that the microfabricated electrodes are able to stimulate neural systems in both in vitro and in vivo environments. 29

33 Chapter 2 : Evaluation of PNB as a Material for Implantable Electrodes 2.1 Motivation for using PNB The difficulties associated with the development and fabrication of a multilayer all-lcp device with embedded thin film electrodes, connector contacts, and interconnects were examined in this thesis and are described in detail in Appendix A. Similar difficulties in processing LCP observed by Keesara [30] prompted him to consider other moisture resistant, mechanically robust, biocompatible materials for the substrate and insulating capping layers of his generic electrode array design. While polyimide has been widely used to in other neural interface designs and is compatible with standard monolithic processing, the relatively high rate of moisture absorption will likely lead to device failure after a couple years of implantation. In fact, a study has shown that polyimide disintegrated after two years in buffered saline [31]. This anecdotal evidence suggests that polyimide lacks the chemical stability to be an unprotected structural material for long term neural implants. Therefore, polyimide was not considered an option as a substrate or as an insulating material for the peripheral nerve interfaces developed in this thesis. A polymer based on polynorbornene (PNB), marketed under the trade name Avatrel TM 2585P, was investigated for use in the fabrication of the microfabricated nerve interfaces developed in this thesis. This polymer has been developed for use in highdensity packaging [32] and for optical interconnects [33]. The Avatrel TM 2585P formulation of PNB is both spin-castable and photodefinable, providing for straightforward microfabrication without the use of special techniques or tools. Films with thickness of up to approximately 100 µm can be produced from a single spin-cast 30

34 layer of the polymer. PNB exhibits good adhesion to sputtered metals without the need of an adhesion promoter [34]. Good metal adhesion is essential in the ability to produce a device fit for chronic implantation. Various mechanical and electrical properties of PNB are shown in Table 2-1. Property Test Condition Avatrel TM 2585P Tensile Strength Room temperature 18 MPa Young s Modulus Room temp. 0.8 GPa Elongation Room temp. 32% T g 10ºC/min 280ºC Coefficient of 10ºC/min 180 ppm Thermal Expansion Dielectric constant 1 MHz / Room temp Dissipation factor 1 MHz / Room temp Volume resistivity Room temp. 2 x Ω cm Water absorption 24 hours/room temp. 0.07% Table 2-1: Avatrel TM 2585P mechanical and electrical properties. Most notable of the properties documented in Table 2-1 is that the water absorption of PNB is almost as low as that of LCP, and much lower than polyimide. Further, it should be noted that Avatrel TM 2585P is but one formulation of PNB offered by the vendor (Promerus, Inc, Brecksville OH), and there are several formulations of PNB available. Modifications to the PNB matrix can be made to tune the mechanical properties, the adhesive properties, and the photosensitivity of the polymer. As such, while Avatrel TM 2585P was used in the fabrication of the devices detailed in this thesis, future formulations of the polymer may be better suited toward neural interfaces. The biocompatibility of PNB has not been thoroughly studied for use in biomedical implant devices. As a result, device testing is limited to acute animal studies until the biocompatibility of PNB is established. 31

35 Keesara [30] was able to successfully fabricate the facsimile of a multilayer, flexible electrode array that incorporated PNB base and capping layers shown in Figure 2-1. This motivated further investigations to determine if PNB is a suitable candidate material for peripheral nerve electrode arrays made from biologically relevant designs. These investigations include a pilot biocompatibility study, a preliminary study on the effects of saline soaking on the mechanical properties of PNB, and a study of the moisture absorption of PNB as a capping layer on both PNB and LCP base layers. Figure 2-1: Keesara s all-pnb device [30]. 2.2 Hermeticity testing using interdigitated electrodes Motivation A hermetic material and seal are required for chronically implanted electronic devices to prevent the permeation of ion-containing fluids from the body through the device. Permeation of an ion-containing moisture can lead to corrosion and degradation of the electrodes, as well as adhesion failure at polymer interfaces. 32

36 Methods by which other implanted medical devices have been made hermetically sealed are not suitable for the peripheral nerve electrodes developed in this thesis. These approaches include sealing the device in a borosilicate glass that can be bonded to the native oxide layer that grows on the surface of many metals [35], using a pressurized gasket to provide a hermetic seal [36], or placing the device into a titanium capsule [37]. All of these require rigid structures, whereas one of the design requirements of the microfabricated FINE is that it be flexible. Therefore, the substrate and capping materials that encapsulate the electrode traces need either to be hermetic themselves, or made hermetic with an additional coating or coatings. Additionally, the interfaces between the various materials used in the devices need to be hermetically sealed so that the fluids cannot permeate the device through these interfaces. One of the reasons for choosing LCP and PNB for the flexible nerve electrode arrays was the high moisture resistance that exhibited by both materials. Based on the moisture resistance parameters of both materials, it was hypothesized that both of these polymers would provide an adequate barrier to ion-containing fluids for acute testing. However, it was not clear that these properties would hold after undergoing unconventional processes used in the fabrication of the microfabricated FINE devices. More specifically, since PNB has not been used in combination with LCP in the manner that it is utilized in the proposed flexible electrode arrays, it is important to determine if the process used to fabricate the devices is allowing for the optimal material properties. 33

37 2.2.2 Method There are several different device structures that have been used to monitor moisture absorption. Total moisture absorption can be determined using specular X-ray reflectivity [38], gravimetric methods [39], or triple-track patterns [40], [41], [42] have been used in combination with a leakage current measurement setup. Interdigitated test structures are a popular option, with a wide range of digit width and digit spacing utilized [43]-[46]. The interdigitated electrode (IDE) structures were chosen as the test structures to monitor moisture absorption of the materials used for the flexible electrode arrays. One advantage to this structure is that moisture absorption through the capping layer polymer can be measured with either capacitance or the leakage current between the electrodes over time. Additionally, this test structure has advantages over the triple track pattern because a fabrication flaw, such as a broken digit, on an IDE does not necessarily mean that a device cannot be used. In the case of the triple-track pattern, if there is any break in the metal, the entire device is sacrificed. The initial intent was to measure the capacitance between the electrodes as a function of time exposed to moisture. As the dielectric constant of water is approximately 81 and the dielectric constant of PNB is approximately 2.42, as moisture is absorbed into the polymer, the dielectric constant of the capping layer polymer is expected to increase as the polymer takes on more of the characteristics of water. Since capacitance is directly proportional to the dielectric constant, the capacitance of the interdigitated electrode structure will increase with the absorption of moisture according to the formula in Eq. 2-1Error! Reference source not found.. 34

38 Equation 2-1 ε γ ε ε ε wet = water dry + dry 3 [41] Leakage current measurements can also be used to indicate moisture absorption. In this case, instead of exposing the capping layer of the device to deionized water, the capping layer is exposed to phosphate buffered saline, which closely mimics the ionic makeup of fluid in the body. As the saline is absorbed into the polymer, conductive pathways are created between the two electrodes. When a DC bias voltage is applied across the electrodes, current will flow along these conductive pathways. Therefore, the leakage current, or the current that flows from one electrode to the other under DC conditions, will increase with increasing moisture absorption Design Initially, the IDEs were designed to have an adequately large capacitance between the two electrodes while using features sizes comparable to those of the flexible electrode array design. Therefore, any dependence on the geometry of the underlying metal on the moisture resistance of the material is accounted for. Further, fabrication processes had already been developed for the flexible electrode arrays. Therefore, using similar dimensions should allow for the use of the same process parameters. Two IDE sizes were designed: one with 200 digits total and one with 50 digits total. The dimensions of the individual digits and the spacing between them were identical for the two different sized devices. Each digit was 1 cm long and 100 µm wide with 50 µm spacing between each digit. A schematic of a 50-digit device is shown in Fig. 2-2Error! Reference source not found.. 35

39 Figure 2-2: Schematic of design of IDE with capping layer. The IDEs used had 25 digits per electrode, each 1 cm long, 100 µm wide, and separated by 50 µm. A custom saline reservoir was designed to sit on top of the electrodes and expose the electrodes to saline only through the capping layer on any device. As such, there was no available pathway for the saline to the electrodes through the PNB or LCP base layer, through the base layer/capping layer interface, or by wicking through the PNB/metal interface. The vessel was secured to the PNB capping layer with the use of RTV silicone Fabrication of all-pnb IDEs Two types of IDEs were fabricated, the all-pnb IDEs with a PNB base layer, Cr/Al metal electrodes, and a PNB capping layer, and the hybrid PNB/LCP IDEs with a LCP base layer, Cr/Al metal electrodes, and a PNB capping layer. The all-pnb IDEs were fabricated using the process shown in Figure

40 1. Spin-on Avatrel base layer 2. Sputter-deposit and pattern metal traces and electrodes Si Cr/Al Avatrel SiO 2 3. Spin-on and pattern Avatrel capping layer to complete device Figure 2-3: Cross-section views showing all-pnb IDE fabrication process. The 50 µm PNB base layer was fabricated by first spin casting the film onto an oxide wafer for 10 s at 800 rpm, then 30 s at 1000 rpm. The film was soft-baked on a hotplate for 5 minutes at 120ºC, then flood exposed using UV radiation through a 365 nm filter for s. Next, the film was baked on a hot plate 4 minutes at 90ºC, then cured for 1 hour at 160ºC. The metal layer was patterned using photolithography and lift-off with AZ nlof 2035 negative-tone photoresist. First, the photoresist was spun onto the PNB base layer to a thickness of 3.5 µm, then soft-baked on a hotplate for 60 s at 110ºC. The resist was patterned with i-line UV radiation for 8 s, then again baked on a hotplate for 60 s at 110ºC. The resist was immersion developed in AZ nlof 300MIF for 2 minutes, then rinsed. Next, a 50 nm Cr adhesion layer and 2500 nm Al were sputter-deposited onto the surface of the base layer and resist. The resist was dissolved in EBR 70/30 using 37

41 ultrasonic agitation, lifting off excess metal and leaving behind the patterned interdigitated electrodes. Next, a 12 µm PNB capping layer was spun onto the wafer for 10 s at 800 rpm, then 30 s at 2000 rpm. The film was then soft-baked on a hotplate for 5 minutes at 120ºC. The capping layer was exposed with i-line UV radiation for 250 s through a 365 nm filter. Next, a post-exposure bake was performed on a hotplate for 4 minutes at 90ºC. The capping layer was then cured in a polymer curing furnace for 1 hour at 160ºC. The wafer was then diced to separate the IDEs Fabrication of hybrid PNB/LCP IDEs The hybrid PNB/LCP IDEs were fabricated using the sequence shown in Figure 2-4. To begin, copper-clad LCP was cut into the shape of wafers using a craft knife. Next, the copper was stripped from the backside of the sheet in a sodium persulfate solution. The sheet was then adhered to a silicon wafer using Shipley 1813 photoresist as an adhesive. 38

42 1. Adhere LCP sheet to Si wafer with photoresist 2. Sputter-deposit and pattern metal traces and electrodes 3. Spin-on and pattern Avatrel capping layer 4. Dissolve photoresist to release device Si LCP Avatrel Cr/Al Photoresist Figure 2-4: Cross-section views showing hybrid PNB/LCP IDE fabrication process. As for the all-pnb IDEs, the metal layer was patterned using photolithography and lift-off with AZ nlof 2035 negative-tone photoresist. First, the photoresist was spun onto the PNB base layer to a thickness of 3.5 µm, then soft-baked on a hotplate for 60 s at 110ºC. The resist was patterned with i-line UV radiation for 8 s, then again baked on a hotplate for 60 s at 110ºC. The resist was immersion developed in AZ nlof 300MIF for 2 minutes, then rinsed. Next, a 50 nm Cr adhesion layer and 2500 nm Al were sputterdeposited onto the surface of the base layer and resist. The resist was dissolved in EBR 70/30 using ultrasonic agitation, lifting off excess metal and leaving behind the patterned interdigitated electrodes. Next, as for the all-pnb IDEs, a 12 µm PNB capping layer was spun onto the wafer for 10 s at 800 rpm, then 30 s at 2000 rpm. The film was then soft-baked on a hotplate for 5 minutes at 120ºC. The capping layer was exposed with i-line UV radiation 39

43 for 250 s through a 365 nm filter. Next, a post-exposure bake was performed on a hotplate for 4 minutes at 90ºC. Instead of dicing the PNB/LCP IDEs, the IDEs were separated by simply cutting them apart using a craft knife along the same lines that the all-pnb IDEs were diced. A photograph of completed all-pnb and hybrid PNB/LCP IDEs are shown in Figure 2-5. Figure 2-5: Photographs of all-pnb and PNB/LCP 50-digit IDEs Test Setup Ultimately, moisture absorption was monitored by measuring the leakage current between the electrodes when exposed to PBS. This is primarily due to ease of test setup. A capacitance meter requires four-point contact to the electrodes. As such, this requires four micromanipulator probes for best accuracy. Further, the PNB devices were built 40

44 upon doped silicon wafers, there was some concern about parasitic capacitance. In contrast, leakage current measurement requires only two micromanipulator probes. The leakage current was measured with a Keithley 6485 picoammeter under the control of a computer using DAQFactory (Azeotech) data acquisition software. This software can be used to automate communication with external instruments used in test setups. In the case of the leakage current measurements, a DAQFactory script was written to communicate with the picoammeter only. Using the picoammeter documentation, a script was written to command the picoammeter to measure the current and return the value to the computer and the DAQFactory program at specified time intervals. A schematic of the test setup is shown in Figure 2-6Error! Reference source not found.. Vacuum micromanipulator probes were mounted on stacked acrylic blocks. These probes were connected to a power supply and the picoammeter. The probes were then lowered to the electrode pads of the IDEs to measure the leakage current between the electrodes over time. A 5 VDC potential was applied between the electrodes. The micromanipulator probes, leads, and device under test were all placed in a Faraday cage. All testing was performed at room temperature. Figure 2-6: Setup used for leakage current measurement of interdigitated electrodes 41

45 A baseline leakage current was determined for each IDE set by measuring the leakage current over time in the as-fabricated configuration. This measurement was generally steady, as expected. The average leakage current of the dry device was fa, which is very close to the minimum current of 20 fa that can be measured with the picoammeter. Next, the custom-built channel was attached to the IDE. This was done by first applying a layer of RTV silicone around the bottom edge of the channel using a syringe. This edge had been filed down to make it flat. The channel was then pressed firmly against the IDE. Care was taken to ensure that the channel edge was completely within the boundaries of the capping layer and to avoid moving the channel once contact had been made with the structure so that silicone would not be spread over the IDEs. The silicone was then allowed to cure overnight at room temperature. The leakage current through the electrodes was first measured every 11.5 seconds for 8 hours without saline. The result is shown in Figure 2-5. The curve is fairly flat, and has a mean value of fa. Again, the picoammeter is capable of measuring currents only as small as 20 fa, and is accurate to 10 fa under the best possible test setup. As such, the negative current under DC bias indicates that the noise is below the measurement range of the instrument. Vibrations and external noise can produce additional error. 42

46 Leakage current (pa) Time (hours) Figure 2-7: Baseline measurement of leakage current in PNB IDE over 8 hours at 5-V bias without saline At this point, the baseline reading with the silicone-attached channel was determined, and the saline was added to the channel using a small rinse bottle. The leakage current was then measured using the Keithley 6485 Picoammeter in 30 second intervals for 100 hours under the control of a computer using DAQFactory software. The results from one device, with each point representing a 10 minute average, are shown in Figure 2-8. The leakage current clearly increases over time. Several hours were required for the leakage current to stabilize. While the current did increase, the current was still fairly small at a maximum of approximately 200 fa. 43

47 0.250 Leakage Current (pa) Time Exposed to Saline (hours) Figure 2-8: Measurement of leakage current in PNB IDE over 100 hours at 5-V bias The results of saline exposure of one hybrid PNB/LCP IDE are shown in Figure 2-7. Again, the leakage current clearly increases over the course of approximately 35 hours. It clearly increases more than the all-pnb IDEs. This may due to incomplete exposure or incomplete curing of the PNB capping layer. The UV exposure time required to fully expose the PNB is a function of the reflective properties of the substrate. The PNB layer on both types of devices was UV exposed for 250 s. A more reflective substrate, such as a Si wafer, reflects the UV radiation back into the photosensitive polymer, increasing the effective exposure. On the other hand, a substrate such as LCP has a low reflectivity and absorbs the UV radiation. As such, for the same UV exposure time and power, the actual exposure of the PNB on the LCP substrate is significantly lower. In order to account for this, the UV exposure time for a PNB capping layer on an LCP substrate should be increased. 44

48 Leakage current [pa] Time exposed to saline [Hours] Figure 2-9: Leakage current measurement over time exposed to saline for the hybrid PNB/LCP IDEs Discussion While it appears that the PBS was able to permeate the PNB capping layer on both the all-pnb and especially the PNB/LCP IDEs, it is also clear that catastrophic failure did not develop for either type of device. The current did not increase to such a level that would prevent the devices from functioning correctly for acute testing. Further, while the test procedures mirrored soak testing of other materials [31, 47],,this does not mimic the way in which the electrode arrays are to be used. In practice, current pulses on the order of several hundred microamps in magnitude are applied to the electrode. However, the leakage current measurement is not meant to be a reflection of the properties of the metal traces or of the passivation layer when current pulses are applied. Instead, the leakage 45

49 current measurement is simply a means of monitoring the absorption of moisture by the insulating PNB capping layer. From the two types of devices, it can be said that the capping layer on the hybrid PNB/LCP IDE absorbed more moisture than the capping layer on the all-pnb device. Gravimetric methods may be helpful to relate the leakage current measurements to actual moisture absorption. The data presented herein indicate that additional leakage current testing of these structures is needed. For the time period and temperature at which these experiments were carried out, it is not clear if the leakage current would continue to increase indefinitely, if the material would saturate and the leakage current would level off, or if at some point there would be catastrophic failure that would essentially indicate a short across the electrodes. If additional long term testing indicates that the electrodes will not survive chronic implantation for approximately 20 years, new solutions such as hermetic coatings on the devices or alternative flexible materials, will have to be investigated to progress this project. 2.3 PNB Biocompatibility In a previous biocompatibility study by Keesara [30], it was found that growth of colorectal cancer cells on PNB was indistinguishable from the growth of colorectal cancer cells on a polycarbonate control after 24 and 48 hours. This preliminary study suggests that cell growth on PNB proceeds normally, and that PNB does not leech any substances that would cause a change in growth kinetics or morphological characteristics. This study was by no means comprehensive, but did support the justification to further 46

50 develop PNB-based nerve electrode devices. In this thesis,, PNB samples were fabricated for a second preliminary biocompatibility trial. The PNB samples were fabricated by first spin casting a 50 µm PNB film onto a 1 µm thermal SiO 2 layer on a Si wafer. The film was then soft-baked on a hotplate for 5 minutes at 120ºC, then flood exposed using i-line UV radiation through a 365 nm filter for s. The film was post-exposure baked for 4 minutes at 90ºC, then cured in a polymer curing furnace for 1 hour at 160ºC. The wafer was then diced into 10x10 mm 2 squares. The PNB squares were then released from the diced wafer by dissolving the thermal SiO 2 layer in HF. The PNB samples were then thoroughly rinsed and dried. Two of the control materials in the study were Si and SiO 2. These substrates were prepared for the study by dicing bare Si wafers and Si wafers with a thermal SiO 2 layer into 10x10 mm 2 squares. In the pilot biocompatibility trial, carried out by Nora Lee, a student in the Department of Biomedical Engineering at CWRU, NIH 3T3 mouse fibroblast cells were cultured on the PNB samples, as well as the control substrates glass, Si, and SiO 2. A calcein AM/ ethidium homodimer-1 (EthD-1) live/dead assay was done on each of the cell cultures, allowing for the comparison of the cell counts from each material. Three samples of each material were incubated for 1, 3 and 7 days post-seeding, respectively. Each sample was cleaned in six-step process involving alternating ultrasonic agitation in organic solvents and rinsing with deionized water. The cells were seeded on each substrate with a cell density of cells/mm 2 in a DMEM culture medium with 10% newborn calf serum, 2% L-Glutamine, and 1% penicillin/streptomycin. After each trial period, the respective samples were rinsed to 47

51 remove any cells that were not adhering well to the substrates. The samples were then incubated with 4µM EthD-1 and 1 µm calcein AM for 50 minutes at 37ºC, rinsed with c- PBS, then mounted on glass microscopy slides for fluorescent imaging with a Zeiss Axio lmager upright fluorescent microscope. Using calcein AM to indicate live cells that were observed with a FITC filter set and EthD-1 to fluoresce red cells that were observed with a Texas Red filter set, the total number of live cells and the total number of dead cells was counted. The percentage of live cells on each of the four substrates was compared. It was found that the highest percentage of live cells in comparison to the total cell count was found on the glass substrate, as expected. The percentage of live cells in comparison to the total cell count on PNB was slightly smaller than that of glass, but comparable to both Si and SiO 2. However, the total number of cells on the PNB substrate was much lower than on any of the other substrates, demonstrated by the total cell count numbers after 7 days, shown in Table 2-2. Substrate Total Number of Cells (after 7 days) Glass 3965 Si 5991 SiO PNB 964 Table 2-2: Total number of cells counted on each substrate after 7 days. The results from this preliminary study suggest that when cultured on PNB, the ratio of live cells-to-dead cells is relatively high. However, more testing is required to determine the cause of the lower number of total cells that adhered to the PNB substrate in comparison with the control substrates. It may be that cells simply do not adhere to 48

52 PNB to the same extent that they adhere to the control substrates, or it may be the case that cell proliferation on PNB is somehow impeded. Contact angle measurements on water droplets were made on a thermally oxidized Si wafer and on PNB to determine the degree of hydrophobicity of the two substrates (Table 2-3). These measurements suggest that PNB is more hydrophobic than SiO 2. This may provide some insight into the reasoning for so few cells adhering to the PNB surface after rinsing. While water droplets are likely to spread on a hydrophilic surface, droplets are likely to bead and roll off of a hydrophobic surface. Similarly, cells do not adhere as well to hydrophobic substrates as they do more hydrophilic substrates [48]. This data suggests that the cause for the lower cell count on the PNB substrate compared to the control substrates may be poor adhesion between cells and the hydrophobic PNB surface. Substrate Contact Angle SiO 2 48º PNB 90º Table 2-3: Contact angle measurements for an oxide layer on a Si wafer and for PNB to determine the degree of hydrophobicity [49]. 2.4 Tensile Testing of PNB Structures Preliminary work to investigate the effects of phosphate buffered saline (PBS) on the mechanical properties of PNB was executed. This work was done to ensure that catastrophic failure would not result from exposure to a physiological environment during acute testing. To execute this experiment, eight test strips of PNB with dimensions 49

53 10 x 60 mm 2 were made prepared. These test strips were sized and shaped for use in the microtensile tool (Instron 5564). Furthermore, these structures are the standard in testing the microtensile properties of PNB, providing a basis of comparison to other microtensile tests. The strips were fabricated by spin-casting a 50 µm thick PNB layer on a Si wafer with a 1 µm thermal oxide layer. The film was then soft-baked on a hotplate for 5 minutes at 120ºC, then patterned using i-line UV radiation through a 365 nm filter for s. The film was post-exposure baked for 4 minutes at 90ºC, then spray-developed using methyl amyl ketone (MAK) and rinsed with PGMEA. The film was then cured in a polymer curing furnace for 1 hour at 160ºC. Next, a 12 µm thick PNB layer was cast on top of the patterned 50 µm strips on the wafer. The film was then soft-baked on a hotplate for 5 minutes at 120ºC, then patterned using i-line UV radiation through a 365 nm filter for 250 s. The film was post-exposure baked for 4 minutes at 90ºC, then spray-developed using methyl amyl ketone (MAK) and rinsed with PGMEA. The film was then cured in a polymer curing furnace for 1 hour at 160ºC. The two-layer PNB test strips were then released from the wafer by dissolving the thermal SiO 2 layer in HF. Upon completion of fabrication, half of the ribbons were soaked in PBS for 72 hours at 87ºC to simulate extended exposure to the PBS under operating conditions at 37ºC. Determination of the acceleration factor would require long-term testing in addition to further accelerated testing. However, Dokmeci, et al. [50] previously used an Arrhenius model to calculate an activation energy of 1.26 ev for soak tested polysilicon. This activation energy can then be used to find the acceleration factor from one temperature to another: 50

54 t37c EA / k( 1/ T37 1/ T87 ) Equation 2-2: AF = = e t [50] 87C Assuming, for estimation purposes, that a similar activation energy is applicable in the soaking of PNB in saline, the 72 hours at 87ºC is equivalent to 5.7 years at 37ºC. This assumption may not be valid because activation energy is related to the material, and the activation energy of 1.26 ev is based on the mean time to failure of a polysilicon coating. Accelerated parylene soak testing by Li, et al. assumed an activation energy of at least 1.05 ev. If the activation energy of PNB is a similar value, then 72 hours at 87ºC is equivalent to 2.1 years at 37ºC. At this rate, to find the effects of saline on PNB at body temperature for 20 years, the PNB test strips would need to be soaked for approximately 32 days. Each soaked and unsoaked PNB test strip was subjected to microtensile testing to failure with an Instron 5564 tool on the Promerus campus. The results of the microtensile testing are shown in Figure 2-10 through Figure 2-13, as well as in Table 2-4. The Young s modulus of the unsoaked strips averaged 887 MPa, while that of the soaked strips averaged 741 MPa. Any difference in percent strain at break or tensile stress was not statistically significant. 51

55 Figure 2-10: Result of microtensile testing of PNB strips #1 and #2 that had not been soaked in PBS. Figure 2-11: Result of microtensile testing of PNB strips #1 and #2 that had been exposed to saline. 52

56 Figure 2-12: Result of microtensile testing of PNB strips #3 and #4 that had not been exposed to saline. Figure 2-13: Result of microtensile testing of PNB strips #3 and #4 that had been exposed to PBS. 53

57 Unsoaked Soaked Device # Young s Modulus (MPa) % Strain at Break Tensile Strength (MPa) Young s Modulus (MPa) % Strain at Break Tensile Strength (MPa) Table 2-4: Comparison of Young s modulus, percent strain at break, and tensile strength of unsoaked PNB strips and PNB strips soaked in PBS for 72 hours at 87ºC. The data shown in Table 2-4Error! Reference source not found. shows that the measured values of the unsoaked strips compare reasonably well with the reported literature values of Avatrel TM 2585P in Table 2-1. The data suggests that soaking in saline for 72 hours at 87ºC is of minimal consequence with respect to the tensile strength of the PNB film structure. This accelerated testing condition is the equivalent of an extended testing condition at 37ºC, which is the temperature at which the devices would normally be operating. A t-test analysis was used to determine if the difference in the means of the tensile strength of the unsoaked strips and the soaked strips is statistically significant. In this test, a t value is calculated using the formula: xt xc Equation 2-3: t = 2 2 σ T σ C + n T n C, 54

58 where x T is the mean tensile strength of the soaked (treated) samples, x C is the mean of the unsoaked (control) samples, soaked samples, 2 σ C 2 σ T is the variance in the tensile strength of the is the variance of the tensile strength of the unsoaked samples, nt is the size of the sample set of soaked sample, and n C is the size of the sample set of the unsoaked samples. The calculated t value, 0.689, was compared to the values in a standard t-table at a significance of This table shows that the critical value of t 0.05 is The associated p-value is with a power of 0.10 assuming a difference of 1 MPa is negligible. To achieve a power of 0.90, 77 samples would be required for both the soaked and unsoaked treatments. Similarly, the effect of soaking in saline on the percent strain at break of the strips is inconclusive. For percent strain at break, p = with a power of To increase the power to 0.95, the sample size of the unsoaked and soaked strips would need to be increased to 42. The soaked strips did exhibit a lower Young s modulus than the unsoaked strips (p = 0.007). This increase in flexibility is actually desirable because it reduces the mechanical mismatch between the microfabrication electrode structure and the nerve tissue, which has a Young s modulus on the order of kpa [51]. From this experiment, we can conclude that the mechanical properties after soaking in saline remain satisfactory for the duration of acute testing at a minimum. If an Arrhenius relation can be assumed, and if an activation energy of 1.05 ev can be assumed, this may hold true for at least a couple years. However, it must again be noted that this work was a preliminary investigation. 55

59 2.5 Conclusion Preliminary evaluations of PNB have been performed to test the hypothesis that PNB is a suitable material for use in acutely implanted microfabricated electrodes for biomedical applications with respect to moisture absorption, biocompatibility, and mechanical properties. It was shown that moisture absorption through PNB will remain small, allowing for leakage current in the range of hundred of femtoamps to several picoamps over the course of several hours under dc applied voltage conditions, the period of time required for acute testing. Further, preliminary biocompatibility data suggests that PNB is not toxic to cultured cells. Lastly, after soaking in phosphate buffered saline for 72 hours at 87ºC, the PNB did not simply disintegrate. In fact, the small number of samples tested did not indicate that there is any appreciative difference in tensile strength of the soaked PNB samples compared to the unsoaked PNB samples. Through these investigations, it has been determined that PNB is a suitable material for acutely implanted biomedical devices and justifies the development of the peripheral neural electrodes described in this thesis. 56

60 Chapter 3 : Design #1 -- Microfabricated FINE Design 3.1 Motivation The first microfabricated electrode array design developed in this thesis was based on the Flat Interface Nerve Electrode [5] described in Chapter 1. This device, shown in Figure 0-2, is comprised of platinum foil electrodes set in a silicone structure such that the electrodes can stimulate or record from a nerve when the silicone structure is closed around the nerve. The rectangular opening in the closed silicon structure allows the device to take advantage of the oblong shape of the nerve and maximize the interfacial area with which the electrode array has contact. While effective [5, 52-54], this array requires a great amount of skill and time to produce. Furthermore, since each array is handmade, they lack in the repeatability that is desired for medical applications. A microfabricated array, on the other hand, takes advantage of batch processing used by IC fabrication techniques, which allows for efficient, repeatable fabrication. Microfabrication tools also permit the production of a dense electrode array, which may allow for a higher degree of specificity when stimulation to or recording from a nerve. The microfabricated version of the FINE was designed to be comprised of two flexible, microfabricated electrode array inserts set in a molded silicone housing. The microfabricated electrode arrays were required to have a flexible substrate, a metal layer, and an insulating capping layer with openings to the electrodes. The microfabricated inserts were designed to have an electrode array at the bio-interface end of the device that fits into the silicone housing, a long shaft determined by the size of the wafer, and connector contacts to provide a means by which the electrodes can be interfaced with 57

61 external electronics. The first microfabricated FINE design was sized for a human femoral nerve. 3.2 Flexible Array Design There are three main features of each flexible electrode device: ten electrode contacts to stimulate the nerve, ten connector contacts for interfacing with external electronics, and the long shaft that connects the two. A plan view of the device is shown in Figure 3-1. Figure 3-1: Plan view of microfabricated FINE design. The flexible electrode arrays were to be fabricated using a three photomask system, shown in Figure 3-2, Figure 3-3, and Figure 3-4. Photomask layouts for each layer were generated using AutoCAD and submitted to a mask manufacturing vendor (The Photoplot Store) for manufacture. The polarity of each photomask was chosen with regard to the tone of the photoresist that would be used in that process step. The electrode end of the device and the connector end of the device are shown in more detail in Figure 3-5 and Figure 3-6, respectively. 58

62 Figure 3-2: Base layer photomask for microfabricated FINE arrays. Figure 3-3: Metal layer photomask for microfabricated FINE arrays. 59

63 Figure 3-4: Capping layer photomask for microfabricated FINE arrays. Figure 3-5: Electrode end of microfabricated device with vias for each contact. Dimensions are in microns. 60

64 Figure 3-6: Connector end of device with vias for each contact. Dimensions are in microns. A long, thin, flexible shaft supports the interconnect traces between the electrode and the connector ends of the microfabricated array. This shank was designed to be thin, and leave approximately 1 mm of base polymer on either side of the long traces, for a total width of 3 mm. The length of the shaft was determined by its position on the 100 mm wafer design. Eight devices were fabricated on each 100 mm wafer, with shaft lengths that vary from 2 cm to 6.5 cm. 3.3 Silicone Housing The overall dimensions of the molded silicone housing were the same as those used in the conventional, hand-fabricated FINEs sized for a human femoral nerve, which is approximately 8mm x mm [54]. The silicone housing, shown in Figure 3-7, Error! Reference source not found.was designed such that two planar microfabricated arrays could be positioned and held in place. This housing consisted of hinged end, a clasp, and a frame to provide the shape of the FINE closing. The frame provided the 61

65 opening cross section that had been used in conventional FINE devices, as well as a lip to help secure the microfabricated array into place. The shape and dimensions of the electrode end of the microfabricated arrays were chosen such that the fit between the substrate and the housing would be tight. Figure 3-7: Custom-molded silicone housing for microfabricated FINE The electrode end of each flexible array was set into the housing with the aid of a custom jig (Figure 3-8Error! Reference source not found.) that was used to affix the device to the housing. To complete one side of the device, the microfabricated structure was set into the housing, electrodes facing downward against a steel block that served as a rigid surface such that when pressure was applied to the backside of the microfabricated structure, the array did not push completely through to the bottom half of the silicone housing. Silicone was applied to the housing and the backside of the electrode to completely backfill the space between the top of the housing and the microfabricated structure, thus anchoring the array. After the first insert was successful secured into position in the housing, the jig is flipped over and the procedure was repeated for the second flexible insert. 62

66 Figure 3-8: Jig used for the purpose of holding the housing and insert in place while silicone was applied to the housing and the backside of the flexible electrode array. 3.4 Connector A number of options have been developed to interface a microfabricated substrate with external electronics. Michigan probes with an integrated silicon ribbon cable used a flexible printed circuit board mounted to a commercially available microconnector [55]. Stieglitz, et al. [56] has developed the MicroFlex Interconnection technique for connection between a stiff substrate such as an IC and a flexible ribbon cable. This technique requires thermosonic ball wedge bonding. In this process, vias in the ribbon cable contacts are aligned over the bonding pads of the rigid substrate, and a gold ball bond is formed between the two substrates to produce an electrical and mechanical contact. For acute testing and applications, the sieve electrode nerve interface used a commercially off-the-shelf zero insertion force connector [57]. As the devices produced from this first design would initially only be used in acute applications, the inserts were designed to interface with a commercial off-the-shelf 63

67 flexible circuit connector. Hirose has several lines of flexible circuit connectors with varying numbers of contacts and varying pitch. The connector chosen for the prototypes developed in this thesis was a 10-contact FH-27, which has the smallest commercially available pitch at 0.4 mm. The contacts at the connector end of the microfabricated device were designed to align with the metal contacts on the connector itself. Further, the shape of the connector end of the microfabricated device was designed to closely fit into the connector. The FH-27 connector was soldered to a patterned circuit board, as shown in Figure 3-9Error! Reference source not found.. Metal traces on the circuit board lead to a second connector into which long, discrete wires can be placed. A cable formed by the discrete wires was connected to a second board with pins onto which leads from external electronics can be clipped. Figure 3-9: Hirose FH-27 flex circuit connector soldered to a printed circuit board. 3.5 Discussion The design described for the flexible microfabricated electrode array inserts for the molded silicone FINE housing emulates other microfabricated neural interface 64

68 designs in that it is comprised of a base substrate layer, a metal layer for the electrodes, contacts, and interconnects, and a top insulating layer. The basic design incorporates an electrode array at one end with interconnects that provide distance between the connector and the electrodes. This ensures that the heaviest part of the device, the connector, can be distanced from the nerve and will not pull on the nerve. This device is unique in that it combines microfabricated, flexible arrays with a molded silicone housing. Further, this device is implemented with unique materials that have not been used for microfabricated peripheral nerve interfaces. One design issue that should be addressed in future flexible electrode designs is the connector scheme that allows for interfacing to external electronics. While the current connection scheme is acceptable for acute testing procedures, it will not be practical for chronic implants. First, the stability of the connection may not suit chronic implantation. A long-term connection scheme would involve a hermetically-sealed connector. This may be achieved by finding a commercial connector that is hermetically sealed, modifying the current connection scheme to hermetically seal it, or developing a new connection scheme that is hermetically sealed. The current connection scheme may be hermetically sealed by encasing the connector in a moisture-resistant coating, such as parylene, silicone, or polynorbornene. The connection scheme described above requires that wires would have to traverse the skin in order to allow for interfacing of the nerves to external electronics. Both stimulation and recording require interfacing with electronics. A long term solution would require a connection scheme that does not require wires or any component of the electrodes to interface with electronics outside of the body. As such, a better option is to 65

69 to package the necessary electronics for complete implantation into the body. In order to monitor nerve recordings, or program nerve stimulation, some kind of wireless capability would be necessary. Microfabrication processes have the ability to produce a dense array of electrodes. However, the current design and connection scheme require that each of these electrodes must lead to a discrete wire or connector to which a stimulator or other external electronics can be connected. This can be cumbersome and heavy as the number of electrodes in the array grows. Further, it places a high demand on the external electronics as in order to take full advantage of the array, the stimulator or other device needs one channel for each of the electrodes. Otherwise, the electronics will have to be disconnected from one electrode and connected to another, eliminating the possibility of stimulating to or recording from both electrodes at the same time. An alternative that should be considered in future designs is the incorporation of a multiplexer that will allow for reducing the number of discrete wires and connections to external electronics needed. This idea has been implemented in an 18-channel polyimide-based spiral cuff nerve electrode [34]. Commercial off-the-shelf ICs were used to successfully reduce the number of channels from 12 to 4. Building upon this work would allow for a more reasonable number of wires. Additional work may focus on the shaft connecting the electrode and connector ends of the device. The shaft length in the devices described in this thesis is currently limited by the width of the 100 mm wafer to approximately 65 mm. This may not be long enough for all applications. As such, once way in which this flexible ribbon cable can be made longer is by using a spiral or serpentine shape. This will require that the substrate 66

70 material is sufficiently mechanically robust to endure the resultant twisting in the shank, but can allow for a much greater distance between the electrode and connector ends of the device. However, it should be noted that this will allow for fewer devices on each wafer. 67

71 Chapter 4 : Fabrication of All-PNB Design #1 Devices 4.1 All-PNB Devices with SiO 2 Sacrificial Layer and Ti/Pt Metallization Fabrication process The first fabrication process used to implement Design #1 using PNB as both the substrate and capping layers is based on the process developed by Keesara [30], and is shown in Figure 4-1. From Keesara s work, it was determined that a 40 μm base layer and a 12 μm capping layer were sufficient to form viable free-standing structures, and thus the all PNB devices fabricated in this thesis utilize these dimensions. This is similar to the process used to fabricate the all-pnb IDEs described in Chapter 2. Figure 4-1: Process sequence for fabricating PNB-based electrode arrays with HF Irelease. The 40 µm PNB base layer was fabricated by first spin casting the film onto the oxide wafer for 10 s at 800 rpm, then 30 s at 1000 rpm. The film was soft-baked on a hotplate for 5 minutes at 120ºC, then flood exposed using UV radiation through a 365 nm 68

72 filter for s. Next, the film was baked on a hot plate 4 minutes at 90ºC, then cured for 1 hour at 160ºC. The metal layer was patterned using photolithography and lift-off with AZ nlof 2035 negative-tone photoresist. First, the photoresist was spun onto the PNB base layer to a thickness of 3.5 µm, then soft-baked on a hotplate for 60 s at 110ºC. The resist was patterned with i-line UV radiation for 8 s, then again baked on a hotplate for 60 s at 110ºC. The resist was immersion developed in AZ nlof 300MIF for 2 minutes, then rinsed. Next, a 20 nm Ti adhesion layer and 2500 nm Pt were sputter-deposited onto the surface of the base layer and resist.the resist was dissolved in EBR 70/30 using ultrasonic agitation, lifting off excess metal and leaving behind the patterned interdigitated electrodes. Next, a 12 µm PNB capping layer was spun onto the wafer for 10 s at 800 rpm, then 30 s at 2000 rpm. The film was then soft-baked on a hotplate for 5 minutes at 120ºC. The capping layer was exposed with i-line UV radiation for 250 s through a 365 nm filter. Next, a post-exposure bake was performed on a hotplate for 4 minutes at 90ºC. The capping layer was then cured in a polymer curing furnace for 1 hour at 160ºC. The devices were released by dissolving the thermal oxide layer in HF. Completed microfabricated arrays are shown in Figure 4-2. Two microfabricated FINE arrays placed in the silicone housing are shown in Figure

73 Figure 4-2: Microfabricated FINE arrays with PNB base and capping layers. Figure 4-3: Two all-pnb microfabricated electrode arrays in a molded silicone housing. The housing is splayed open to demonstrate the flexibility of the structure. 70

74 4.1.2 Results While the metal traces and electrodes appeared flat prior to release, the metal buckled after release, as shown on the electrode pad in Figure 4-4.The effect also appeared along the long traces, which were fully encapsulated by the PNB, and the connector contacts. Keesara reported a similar phenomenon for the Ti/Pt electrodes in his PNB devices, and suggested that this was the result of compressive stress. For one experiment, Ti/Pt electrodes were patterned on a PNB base layer without a capping layer, and were released in HF. Upon release, the Ti/Pt electrode and connector contacts lifted completely from the polymer. Since Ti is etched by HF, it was determined that even though only a very thin edge of Ti was exposed to the acid, it was enough exposure to completely etch the adhesion layer. Since all of the electrode and connector contact edges were encapsulated by PNB, the metal was anchored and prevented from lifting completely from the device. However, the metal was easily sliced with a probe tip, revealing a void between the metal and the PNB base layer. It has not yet been determined if this is the result of delamination from excessive compressive stress, or if HF was able to permeate through the device and etch the Ti to create a void between the metal and the PNB substrate layer. 71

75 Figure 4-4: Pt electrode pad with Ti adhesion layer on a HF-released PNB-based electrode array. 4.2 All-PNB Devices with SiO 2 Sacrificial Layer and Cr/Pt Metallization Since PNB is believed to be very resistant to moisture, it did not seem very likely that the HF would be able to permeate through the PNB and etch away all of the Ti. Since Cr is another widely used metal adhesion layer and will not be etched in HF, devices were fabricated using a Cr adhesion layer with Pt electrodes. This second fabrication process used to implement design #1 in an all-pnb device was nearly identical to the process described in Section 4.1, the only difference being that in place of Ti/Pt metallization, a 30 nm Cr adhesion layer was sputtered with the 250 nm Pt. A photograph of the Cr/Pt electrodes is shown in Figure 4-5Error! Reference source not found.. There are two items to note on this electrode pad. First, is the texture of the metal is clearly different than if Ti is used as an adhesion layer. Unlike the Ti/Pt metal buckling that developed upon releasing the structures from the wafer, the crack-like texture of the Cr/Pt electrodes was visible after metal deposition, before the lift-off step. 72

76 Second, while the electrode pad appears to be cracked, it is a continuous film. An SEM of a Pt/Cr thin film electrode on PNB is shown in Figure 4-6Error! Reference source not found.. Figure 4-5: Pt electrode pad with Cr adhesion layer on a HF-released PNB-based electrode array. Figure 4-6: SEM micrograph of a Pt electrode pad with Cr adhesion layer on a HFreleased PNB-based electrode array. The Cr/Pt metallization on the PNB substrate did not provide any additional insight as far as why the Ti/Pt electrodes were so severely buckled. The Cr appears to 73

77 yield and crack during sputtering, and the Cr/Pt electrodes do not buckle when released with HF. Howeever, the Cr/Pt films are continuous, as demonstrated by the SEM micrograph, as well as the relatively small resistivity of approximately 157 Ω/cm when resistance measurements are made across structures with numerous crack-like defects. This resistivity was determined by measuring the resistance between the electrode and the connector contact with using an Agilent 34401A digital multimeter and micromanipulator probes, then dividing by the length of the trace. Since Ti is known to be a biocompatible metal, while not all varieties of Cr are biocompatible, we could not simply use Cr as the adhesion layer and abandon Ti as the adhesion layer in this process. It was hypothesized that a different release mechanism than the HF release by dissolution of a thermal SiO 2 layer may provide a solution to the problem. Finding an alternative release mechanism proved to be a difficult challenge to overcome, with several failed attempted documented in Appendix B. The large interfacial area between the substrate and the PNB structures prevented several techniques that had been successfully implemented by others for much smaller device structures. Anodic aluminum dissolution was investigated, and as shown in Section 4.3, proved to be a successful method to releasee the PNB-based devices from a Si wafer. Moreover the Ti/Pt metallization did not suffer from the same adhesion issues that were seen with the HF release. 4.3 All-PNB Devices with Aluminum Sacrificial Layer and Ti/Pt Metallization The third fabrication sequence developed for design #1 prototypes utilizes an Al release layer that is dissolved anodically. The process used to fabricated the all-pnb 74

78 devices is shown in Figure 4-7. As is evident in the figure, the fabrication process uses a sequence of processing steps that follows conventional monolithic microfabrication. To realize these significant differences in film thickness, a high viscosity solution (4500 cps) was used to form a 40 μm base layer, and a low viscosity solution (1000 cps) was used to form the 12 μm capping layer. 1. Sputter-deposit Cr and Al release layers 2. Spin-on and pattern Avatrel base layer 3. Sputter-deposit and pattern metal traces and electrodes 4. Spin-on and pattern Avatrel capping layer 5. Anodically dissolve Al sacrificial layer to release devices 6. Completed device Si Avatrel Pt Cr Al Figure 4-7: Process sequence for fabricating PNB-based electrodes with anodic aluminum release Base Layer Patterning 75

79 The PNB fabrication sequence requires a rigid substrate for processing. Silicon wafers are logical substrates to be used in PNB processing as they provide a very smooth surface onto which the PNB can be spin-cast. Further, 100mm Si wafers are sized for the tooling in the microfabrication facilities at Case. The PNB-based devices were built on Si wafers coated with sputter-deposited 100 nm Cr and 500 nm Al layers. The Al film serves as a sacrificial layer that is dissolved at the end of the fabrication process to release the PNB structures from the wafer. The next step in the fabrication process was to spin-cast a layer of PNB onto the Si wafer using the more-viscous PNB solution, a programmable Laurell Spin Coater was used for the spin casting. Approximately 5 ml of the solution was statically dispensed onto the center of the wafer by pouring directly from the bottle. Immediately after the PNB solution was dispensed, a two-step spinner program was started. The first step was to spin the wafer for 10 seconds at 800 rpm to spread the polymer on the wafer. The second step was to spin for 30 s at 1000 rpm to cover the wafer at a uniform thickness. Immediately after starting the spinner, a piece of clean room paper was placed over the opening of the spinner. This was to ensure that the spinner chamber remained saturated with solvent, preventing polymer strings from forming and disturbing the otherwise smooth surface of the polymer film. The base layer was then soft baked on a hot plate for 5 minutes at 120ºC. This was followed by exposure in a Karl Suss MA6 aligner for s with the use of a 365 nm filter. The power was set to 10 mw, but the filter reduced the power to 6.4 mw/cm 2. As such, the total dose provided was 2000 mj. A dark field photomask was used for this exposure because Avatrel TM 2585P acts as a negative-tone photoresist. As such, the UV-exposed polymer begins cross-linking upon exposure to the 76

80 radiation. A post-exposure bake followed for 4 minutes at 90ºC to further cross-link the polymer. The base layer was then spray-developed using methyl n-amyl ketone (MAK), or 2-heptanone, and rinsed with PGMEA. The spray development was done by placing the wafer on the spinner chuck, then setting the spinner to spin at 2000 rpm. During the first 90 s of spinning, the wafer was sprayed with MAK to develop the polymer, then rinsed for the next 5 s by spraying the wafer with PGMEA. The final 30 s of spin-development was used to dry the wafer and ensure that the MAK and PGMEA had been spun away. During the development step, the polymer that had not been exposed to UV radiation, and therefore was not cross-linked, was removed from the wafer by rinsing. After development, the wafer was cured in a polymer curing furnace. The temperature profile for the furnace is shown in Figure 4-8. The wafers enter the furnace at room temperature, after which the temperature was increased at a rate of 5ºC/minute until reaching a temperature of 160ºC. The temperature was held at 160ºC for 1 hour, then the temperature was decreased to room temperature at a rate of 5ºC/minute. The cure ensured that the polymer was fully cross-linked and that all solvents were removed from the polymer film. 77

81 Temperature 25 C 5 C/minute 160 C 1 hour 5 C/minute 25 C Time Figure 4-8: Avatrel TM 2585P cure cycle [30] Metal Layer Patterning After the base layer patterning was complete, the PNB base layer surface was prepared for metal deposition and patterning. First, AZ nlof 2035 negative tone photoresist was spin-cast onto the wafer to produce a 3.5µm film. Approximately 5 ml of resist was statically dispensed from a 50-mL beaker onto the center of the wafer. The resist was then spread at 500 rpm for 10 seconds, and then spun at 3500 rpm for 30 s to evenly distribute the resist on the wafer. The resist was soft-baked on a hot plate at 110ºC for 60 s. Next, the clear-field metal layer photomask was aligned to the PNB base layer, and the resist was exposed for 8.0 s with a power of 10.0mW. Next, a post-exposure bake was done on a hot plate at 110ºC for 60 s. The resist was then immersion developed for 120 s in AZ MIF 300. For the development, each wafer was placed in a wafer boat on the side furthest from the H-bar. The H-bar was then used as a handle to gently agitate the developer in a larger container during the development. The wafer was then immediately rinsed in deionized water in a dump rinse bath. The wafer was then dried using a nitrogen gun. After drying, the wafer was ready for metal deposition. 78

82 For these devices, the thin film metal electrodes, traces, and contact pads were sputter-deposited in a Discovery 18 sputter system. The wafer surface was first prepared with an RF preclean for 90 seconds at 125W to expose the surface of the polymer to an argon plasma for improvement of the adhesion of the metal to the PNB base layer. Next, 20 nm Ti and 250 nm Pt were sputter-deposited under vacuum on the PNB base layer and negative-tone resist. The metal patterning was completed in a liftoff step. Ultrasonic agitation was used to dissolve the resist in the solvent AZ EBR 70/30. Approximately 1 L of the solvent was poured into a 4-L beaker. The wafer was then placed in a wafer boat at the end of the boat far from the H-bar. The H-bar was then used as a handle to submerge the wafer into the beaker with the support of the wafer boat. The beaker was then placed into an ultrasonic bath that was partially filled with water. The liftoff step took 3 to 4 hours to completely remove the excess metal. The wafer was then rinsed in a dump rinse bath to remove all of the solvent from the surface of the wafer Capping layer patterning Upon completion of the metal patterning step, the wafer was prepared for patterning of the insulating capping layer. This procedure was very similar to that used for patterning the base layer. In this case, the less viscous PNB solution was used. Approximately 5 ml of polymer solution was statically deposited onto the center of the wafer by pouring directly from the bottle. Immediately after pouring the solution onto the wafer, the spinner program was started. The first step in the program was to spread the polymer solution for 10 s at 800 rpm, and the second step was to evenly coat the polymer 79

83 over the wafer for 30 s at 2000 rpm. Cleanroom paper was again placed on top of the spinner opening to ensure that the spinner chamber was saturated with solvent. The capping layer was then soft baked for 5 minutes at 120ºC on a hot plate. Next, the dark field capping layer mask was aligned to the metal layer on the wafer. The capping layer was exposed to UV radiation through a UV filter for 250 s. Again, the power was set to 10.0 mw, but 3.6 mw of the power was lost through the filter, allowing 6.4mW to pass through to the film. After UV exposure, the post-exposure bake was done on a hotplate at 90ºC for 4 minutes. Next, the capping layer was spray developed with MAK and PGMEA using the procedures developed for the base layer. At 2000 rpm, MAK was sprayed for 15 s, and then the PGMEA was sprayed for 5 s, followed by a 30 s spin-dry step. Finally, the wafer was again placed in the polymer curing furnace to cure the capping layer of PNB. The same furnace recipe was used as for the base layer cure. The program started with room temperature, ramped up to 160ºC for 1 hour, then ramped back down to room temperature Device Release The devices were released using anodic aluminum dissolution. This was investigated as an alternative to the HF release to remove the PNB structures from a rigid wafer. This method is used on an industrial scale for electrochemical machining and electropolishing [58]. Anodic aluminum dissolution occurs in an electrolytic cell (a beaker) filled with a 2M NaCl solution. The sample from which metal is being removed is the anode in the cell. A piece of metal foil, such as stainless steel foil, acts as the cathode in the configuration. When a potential is applied across the cell, the 80

84 electrochemical reactions noted in Error! Reference source not found. and Error! Reference source not found. occur [58]. Equation 4-1 Equation Al Al + 3e - + 2Al 3H2O Al2O3 6H 6e [58] [58] A similar reaction will occur with many metals used in microfabrication, but at different applied potentials. We have taken advantage of the selectivity of the metal dissolution with the use of the data presented in Figure 4-9Error! Reference source not found.. Here, layers of Cr or Al with a Ag/AgCl cathode were polarized at potentials ranging from -1.5 V to 1 V in 2M NaCl solution and the current density was measured with a potentiostat/galvanostat. From this plot, we can conclude that Al dissolution begins to occur at a potential of approximately -0.5 V, while Cr dissolution does not begin until a potential of approximately 0.75 V is applied. As such, applied potentials between -0.5 V and 0.75 V will selectively dissolve Al over Cr [58]. This method has been used to successfully to release and partially detach planar and nonplanar microdevices fabricated using polyimide, SU-8, and platinum [58]. 81

85 Figure 4-9: Polarization behavior of Cr and Al layers at potentials between -1.5 V and 1 V in 2M NaCl solution [58]. Anodic aluminum dissolution has been used to release other biomedical polymer devices from rigid wafers, and it was reasonable to believe that this would resolve the PNB device release issues. This method does not require any toxic chemicals and would not harm the electrodes embedded between the polymer layers. To begin this process, 100 nm Cr and 500 nm Al were sputtered onto a Si wafer. PNB base layer structures were patterned onto the sputtered metal as described in Section Error! Reference source not found.. In the first run, the anodic Al dissolution method was used to release only the base layer structures, without the metal or capping layers. After curing the base layer structures on the Si wafer with the Cr/Al layers, the wafer was placed in a beaker that was filled with 2 L of 2M NaCl in DI water. A power supply was connected to both the wafer and a piece of stainless steel foil with alligator clips, and both the wafer and foil were submerged in the solution. The potential between the wafer and the foil was set to 0.70 V. 82

86 As the Al dissolved, the Cr layer remained intact, providing a continuous conductive film across the wafer (see Figure 4-10). After approximately 2 hours, the structures fully released from the wafer and floated to the top of the salt water solution. Figure 4-10: Schematic of anodic aluminum dissolution used as a release mechanism for the PNB-based devices. The metal on the released all-pnb devices is shown in Figure 4-11Error! Reference source not found.. This metal on the electrode pad does not resemble that from the devices released by dissolving an SiO 2 sacrificial layer in HF. Instead, the metal on the devices released with anodic aluminum dissolution is rippled throughout, which can be attributed to compressive stress in the polymer. Unlike the metal on the HFreleased devices, the metal on anodic Al dissolution-released devices has not lifted from the base polymer layer at all, but resembles the compressed Au on PDMS samples fabricated by Lacour, et al. [59]. 83

87 Figure 4-11: (left) Electrode-end opening and (right) Pt traces from anodic aluminum dissolution-released PNB-based device with cures at 160ºC. 4.4 Summary A summary of the results of each release is shown in Table 4-1. As electrodes in the PNB devices released with anodic aluminum dissolution appears to have superior adhesion to those released with HF, and because release in 2M NaCl is easier and safer than release in HF, future devices should utilize the anodic aluminum release process in place of the HF release. 84

88 Metallization Release Result Ti/Pt HF Buckling, void between metal and PNB Ti/Pt Cr/Pt Anodic Al Dissolution HF 160ºC cure: rippling indicative of compressive stress, adhesion not compromised Cracking during sputter-deposition step, but continuous Pt film Table 4-1: Summary of results from each type of metallization and release. 85

89 Chapter 5 : PNB/LCP Electrode Arrays 5.1 Rationale The microfabricated PNB-based electrode arrays are not sufficiently mechanically robust to survive handling during implantation. After release, many of the PNB-based devices were easily broken and required gentle handling. The base layer mechanical properties must be improved so that the devices are sufficiently robust to survive aggressive handling. Additional options included a more viscous formulation of Avatrel TM 2585P that would allow for thicker films to be spin-cast onto the wafer, or a completely different Avatrel TM formulation that produces films with a higher tensile strength. Along these lines, PNB layers were patterned on wafers using a more viscous solution of PNB, Avatrel TM 2580P. This PNB formulation has both a higher Young s modulus (1.56 GPa) and a higher tensile strength (38.1 MPa) than Avatrel TM 2585P. Furthermore, the viscosity of the Avatrel TM 2580P solution was higher than either of the two Avatrel TM 2585P solutions previously used. This solution allowed a thicker film to be cast. While the target thickness for this film was 100 µm, the actual thickness of the cured film was measured to be approximately 60 µm. The metal and capping layers were not processed in this set of devices. When the base structures were released from the wafer in HF, it was found that these devices were more brittle than the 40µm Avatrel TM 2585P structures. Additionally, the tensile strength was not large enough to pursue this material further. Another option was to fabricate a hybrid PNB/LCP device utilizing a LCP base layer and a PNB capping layer. This approach takes advantage of the superior mechanical 86

90 properties of LCP and the convenience of PNB fabrication processes without significantly sacrificing moisture resistance. As such, efforts were made to fabricate a hybrid device with a LCP base layer and a 12 µm thick PNB insulating capping layer. The mechanical properties of the device would be dominated by the more mechanically robust, 50 µm thick LCP base layer. Both LCP and PNB exhibit moisture resistance superior to that of other polymers, such as polyimide. As such, a device employing both materials should exhibit a high moisture resistance as well. Since PNB is spin-castable, photodefinable, and translucent, it would be significantly less difficult to pattern a PNB capping layer on top of a LCP sheet with patterned metal than fabricating a multilayer LCP electrode array device. Further, the risk of damaging patterned thin film metal electrodes on a LCP substrate by spin casting and patterning a PNB capping layer was small since PNB had been spin-cast and patterned on top of the same patterned thin film electrodes on a PNB substrate. The patterning of the LCP base layer was more complicated than the patterning of the PNB base layer, but was possible by taking advantage of the copper cladding on the backside of the LCP sheet and the front-to-back alignment feature of the Karl Suss MA6 aligner. 5.2 Device Fabrication The hybrid PNB/LCP arrays were fabricated by first patterning the metal layer on the LCP, then spin casting and patterning the PNB capping layer, and finally patterning the LCP base layer with RIE, as shown in the fabrication process shown in Figure 5-1. This is in contrast to the all-pnb devices, where the base layer was patterned first. 87

91 Patterning the LCP base layer last facilitated alignment of the copper cladding with the metal and capping layers. 1. Adhere LCP/Cu-laminated sheet to Si wafer with photoresist 2. Sputter-deposit and pattern metal traces and electrodes 3. Spin-on and pattern PNB capping layer 4. Flip and adhere to glass wafer with photoresist 5. Pattern and etch Cu to form RIE etch mask 6. RIE Pattern and remove Cu to complete device Wafer Pt/Ti LCP Cu PNB PR Figure 5-1: Process sequence used to fabricate hybrid PNB/LCP FINE arrays. LCP is available in sheets of varying thicknesses, with or without 18 µm-thick copper cladding. In this project, single-sided copper clad sheets of R/flex 3600 LCP (Rogers Corp.) with an LCP thickness of 50-µm were used. Wafer-shaped sheets of LCP were produced by cutting the LCP sheet with a craft knife and a Si wafer stencil. Standard microfabrication tooling requires a rigid substrate. Keesara s method of mounting the wafer-shaped LCP sheets to rigid Si wafers using photoresist as an adhesive was used here [30]. Shipley 1813 positive-tone photoresist was statically dispensed onto 88

92 a Si wafer, then allowed to spread at 500 rpm for 2 seconds. The LCP sheet was then applied to the resist-coated wafer copper-side down. A nitrogen gun was used to hold down the LCP sheet as the spinner spin speed was incrementally increased from 500 rpm to 4500 rpm. These wafers were then placed into an oven for 20 minutes at 90ºC. Thin film platinum electrodes were patterned using a procedure similar to the procedure used to pattern the metal electrodes on the all-pnb FINE arrays. First, AZ nlof 2035 negative-tone photoresist was patterned on top of the LCP to serve as a liftoff mask for the metal. Approximately 5 ml of resist was statically dispensed onto the surface of the LCP. The resist was allowed to spread for 10 seconds at 500 rpm, then the spin speed was increased to 3500 rpm for 30 seconds to evenly coat the LCP with resist. The resist was soft-baked for 60 seconds at 110ºC on a hot plate, then exposed for 8 second at 10 mw with i-line UV radiation. Next, a post-exposure bake for 60 seconds at 110ºC on the hotplate was used to complete the reaction that makes the resist insoluble in its developer. The resist was immersion-developed using AZ 300 MIF developer for 120 seconds. The LCP sheet was immediately rinsed with DI water in a dump rinse basin at the completion of the development. Next, 50 nm Ti and 250 nm Pt were sputter-deposited onto the LCP and photoresist using a Discovery 18 sputtering system. The Ti acts as an adhesion layer between the Pt and LCP. It has been suggested that the mechanism involves the formation of a titanium carbide layer at the polymer/ti interface [60]. Lift-off was performed in an ultrasonic bath by dissolving the photoresist underneath the sputtered metal in AZ EBR 70/30. This liftoff step also dissolved the photoresist adhesive between 89

93 the LCP sheet and the wafer, requiring that the sheet be re-adhered to the wafer before further processing could be done. Next, the 10 μm PNB capping layer was patterned on top of the LCP and the platinum electrodes. This was done by statically dispensing approximately 5 ml PNB directly from the bottle onto the surface of the LCP and patterned electrodes. The polymer solution was spread for 10 seconds at 800 rpm, and then spun for 30 seconds at 2000 rpm. The PNB was soft-baked on a hotplate for 5 minutes at 120ºC. The dark-field capping layer mask was aligned to the metal layer. The PNB was then exposed with i-line UV radiation through a 365 nm filter for 250 seconds. As in the fabrication of the all- LCP devices, the power was set to 10.0 mw, but since the 365 nm UV filter was used, the actual power was 6.4 mw. The polymer was post-exposure baked on a hotplate for 4 minutes at 90ºC. The capping layer was then spray-developed at 2000 rpm with MAK for 15 seconds, rinsed with PGMEA for 5 seconds, and spin-dried for 30 seconds. Prior to curing the capping layer, the LCP was removed from the wafer and the photoresist on the copper side was removed with acetone. The LCP was then taped to the Si wafer with Kapton tape, which can easily withstand the heat of the cure. The PNB capping layer was cured in a polymer curing furnace for 1 hour at 160ºC to complete the process on the front-side of the LCP. After curing the PNB capping layer, the LCP sheet was removed from the silicon wafer and adhered copper-side up to a transparent Pyrex wafer with Shipley 1813 photoresist as was done to adhere the LCP to Si wafers. The transparent wafer acts as a rigid mechanical substrate, but also allows for a procedure known as front-to-back alignment. AZ nlof 2035 negative-tone photoresist was spun onto the copper cladding, 90

94 then soft-baked on the hotplate for 60 seconds at 110ºC. In the aligner tool, the metal layer and capping layer features are visible to cameras mounted to view the bottom of the wafer. The base layer mask was aligned to these features, and the photoresist was exposed for 8 seconds. A post-exposure bake was done on a hotplate at 110ºC for 60 seconds, and then the resist was immersion-developed in AZ 300 MIF. The photoresist serves as a wet etch mask for copper patterning. A sodium persulfate solution (250g sodium persulfate: 1 L deionized wafer) served as a copper etchant to pattern the copper. LCP can be patterned using either a wet etch in a heated KOH/ethanolamine solution or with a dry oxygen plasma etch. The wet etch was more readily available because it only requires a heated bath, the KOH/ETA solution, and the single wafer holder, all of which were already available in the laboratory. However, adhesion between the LCP and PNB was compromised during exposure to the heated KOH/ETA. As such, devices using these two polymers were not fabricated successfully while using the KOH/ETA solution to pattern the LCP. The second option was to use an oxygen plasma etch to pattern the LCP. New substrates were prepared with patterned Ti/Pt electrodes and a PNB capping layer as before. The copper served as an etch mask in the oxygen plasma such that only the field area of the LCP would be etched. The LCP was plasma-etched in an RIE system using a gas mixture of 15% CF 4 and 85% O 2, with an etch rate of approximately 0.6 μm/min. The LCP sheet was placed on top of a silicon wafer inside of the chamber to protect the stage from any damage that might occur. It was important to ensure that the LCP sheet was secured flat against the wafer surface to prevent energy from being trapped between the LCP sheet and the wafer, which could cause the LCP to burn. To prevent any 91

95 burning, a layer of tape secured the LCP to the silicon wafer around the edge. Kapton (polyimide) tape was used initially, but the etch recipe was very efficient in etching this tape. A better solution was Teflon tape, which did not etch very quickly in the oxygen plasma. The etch rate was higher along the edges of the wafer and the edges of the Teflon tape. Finally, the copper was stripped from the devices using the copper etchant. A completed device is shown in Figure 5-2, and a completed hybrid PNB/LCP microfabricated FINE is shown in Figure 5-3. Figure 5-2: Microfabricated FINE array with LCP base and PNB capping layer. A closer look at the electrode end of the device shows the roughness of the LCP being translated to roughness on the metal surface. 92

96 Figure 5-3: Two microfabricated FINE arrays with LCP base and PNB capping layer embedded in molded silicon housing. 5.3 Discussion Wet LCP Etch A wet etch for patterning the LCP in PNB/LCP devices was not successfully implemented. The viability of a PNB structure in the heated KOH/ETA solution was investigated prior to placing the LCP sheet with metal and PNB capping layers by placing a broken piece of an all-pnb device in the solution for two hours. This was the estimated time that would be required to complete the LCP wet etch. The results did not demonstrate catastrophic failure, though there was evidence of discoloration. While the coloring of typical PNB-based devices tends to be brown, the part of the PNB device that had been soaked in the heated KOH/ETA bath was closer to yellow in color. Even so, there was no evidence that the adhesion between the two polymer layers was failing. The LCP sheets with the patterned metal and PNB capping layer did not survive in the heated KOH/ETA solution. The single wafer holder prevented the exposure of the PNB/LCP interface to the KOH/ETA solution until the solution was able to break through the etched LCP. At this point, the adhesion at the PNB/LCP interface failed with the PNB layer completely detaching from the LCP sheet. The metal layer remained 93

97 adhered to the LCP in some cases, to the PNB in others, and to neither in the remainder of the devices. It is possible that an adjustment in the processing parameters of the PNB capping layer may have led to better adhesion between the PNB and LCP. Two parameters to take into account are PNB exposure time and PNB cure temperature. While the UV exposure time was the same as was used in the all-pnb devices, the time required for exposure is highly dependent upon the reflective properties of the substrate. An oxidized Si wafer, even with a 50-µm base PNB layer patterned on it, is much more reflective than the R/flex 3600 LCP sheet that was being used for the fabrication of these devices. As a result, it is possible that increasing the UV exposure time would allow for a better interface between the two polymers. Secondly, there is a range of temperatures at which the PNB can be cured. 160ºC is at the low end of this range, with 250ºC at the high end. Increasing this temperature may increase the amount of cross-linking in the PNB, making it less vulnerable to the KOH/ETA solution. In any case, neither of these strategies was used to produce a device with PNB and LCP. Instead, use of an oxygen plasma dry etch was pursued, as was discussed in the previous section RIE LCP Patterning The LCP in the PNB/LCP hybrid nerve electrode arrays was patterned using a plasma etch in a reactive ion etch (RIE) system using a gas mixture of 15% CF 4 and 85% O 2 at a power of 325 W and a pressure of 300 mtorr in a March CS-1701 tool. This recipe exhibited an etch rate of approximately 0.6 μm/minute. By patterning the LCP, 94

98 complete flexible electrode arrays with a LCP base and PNB capping layer were fabricated. Previously, this type of device had not been fabricated to completion as Keesara did not pattern the LCP base layer in his version of the PNB/LCP hybrid array [30], and the combination of PNB and LCP has not been used elsewhere. Upon completion of the oxygen plasma etch, the copper was stripped from the backside of the device, providing a completed device that would then be secured into the molded silicone housing Metal/LCP Interface One concern in the LCP based devices is the quality of the metal on the LCP surface. Because the LCP used is in the form of pressed sheets, it is not as smooth as a polymer such as PNB or polyimide that can be spin-cast. While surface roughness can promote adhesion, roughness with undulations larger than the thickness of the metal itself can be highly demanding of the thin film metal. This is another drawback of LCP. One way to lessen this problem may be to planarize the LCP. There are several ways in which this might be done. One might be to spin coat a thin layer of PNB on top of the LCP. The PNB solution would then fill the crevices at the LCP interface, but provide a planar layer onto which electrodes can be built. Wang et al. suggested two techniques for reducing the surface roughness of the LCP sheet. The first was to spin coat the LCP with a 3μm thick layer of PI 2610 polyimide. The second method involved polishing the LCP using a lap master polisher with a 0.06 μm alumina slurry for 2 hours, which can reduce the roughness from a root-mean-square of approximately 190 nm [61] to approximately 10 nm [62],[63]. Similar techniques may be used using the tooling at Case. 95

99 Chapter 6 : Design #1 Evaluation 6.1 Something Implementation of design #1 with the all-pnb and the PNB/LCP hybrid fabrication process allowed for evaluating the devices to determine if they are suitable for acute implantation and testing. These tests include current pulse testing in phosphate buffered saline, finite element analysis of the geometry of the structure, trace resistance measurement as a function of flexed curvature, and in vitro testing on an invertebrate neural system. 6.2 Pulse testing in saline Test Setup The devices were also tested using a current pulse technique that is standard for testing biomedical electrode devices. This testing determines the impedance of the electrodes, which is important for the determination the amount of current that can be delivered to the nerve. Furthermore, it determines the quality of the metal of the device. The test setup for this set of experiments is shown in Figure 6-1Error! Reference source not found.. 96

100 Figure 6-1: Setup for current pulse testing in saline. A beaker was filled with phosphate-buffered saline. The solution was prepared using 1 L deionized water, 8.77 g sodium chloride, 8.04 sodium biphosphate, and 1.00 g disodium biphosphate in a beaker. A piece of stainless steel foil was partially submerged into the saline. The foil had multiple-wire wire spot welded to it to allow for connection to a digital multimeter. Also submerged into the saline was the electrode end of the microfabricated arrays. The connector end of the device was inserted into the flex-circuit connector and the connector interfacing system. A constant-current pulse generator was connected to one of the pins of the interfacing system, which was in turn connected to the corresponding electrode contact. A current pulse of duration 1 ms and magnitude 1 ma was applied to the electrode. An oscilloscope was used to measure the voltage between the stainless steel foil and the electrode to which the current pulse was applied. In doing 97

101 so, the resistance of the setup could be determined by dividing the magnitude of the voltage waveform by the magnitude of the current pulse that was applied to the electrode Results A typical output waveform measured using the oscilloscope is shown in Figure 6-2Error! Reference source not found.. For the PNB base devices, the output voltage waveform had a magnitude of approximately 3V. Using Ohm s Law, the resistance of the setup is 3V/1mA=3kΩ. For the hybrid PNB/LCP devices, the output voltage waveform had a magnitude of approximately 1V, from which a system resistance of 3kΩ can be inferred. Figure 6-2: Typical output waveform for pulse testing of PNB-based electrodes Discussion The experimental setup is more like what the device would experience in vivo. The ph of the saline is near to the ph of fluids in the human body. Furthermore, in this setup, electric current is converted to ionic current through the saline, and back to electric current at the stainless steel foil. This is similar to what happens as the stimulus current 98

102 through the electrode is converted to an ionic current through bodily fluids, and through the nerve to the muscle. Most importantly, this set of experiments shows that the resistance of these electrodes is sufficiently small that the stimulator is able to supply enough current to stimulate. If the stimulator has a 16 V rating, and the intention is to provide 1 ma current pulses, the maximum resistance is 16V / 1mA, or 16kΩ. The resistance of these electrodes is similar to that of standard platinum foil/stainless steel electrodes, which have a resistance of approximately Ω. Furthermore, these electrodes have a much smaller resistance than those reported by Lee, et al. [63], which have an impedance of 23 kω at 1 khz. It should also be noted that the resistance measured with this method does not match with the resistance measured using a multimeter and probes. There are several possible explanations for this. First, the measurement methods are very different. In the case of the pulse testing, the resistance of the saline is also taken into account. In the case of the probes, there is a very high resistance at the probe point where the probe meets the electrode contact. Furthermore, the capacitance that comes into play in the pulse testing is not the same as capacitance that comes into play in the probe measurements. The capacitance in the pulse testing is proportional to the entire exposed area of the electrode contact, whereas the capacitance in the probe measurements is proportional to the area of the probe tip, which is significantly smaller than the electrode contact. Future pulse testing should be performed over a long period of time to simulate the number of pulses that a typical electrode would see over a 20-year implant lifetime. 99

103 The change in pulse shape or magnitude over time should be analyzed to determine how well the thin film electrodes are enduring the pulses in a physiological environment 6.3 Finite Element Analysis Finite element analysis (FEA) was performed using ANSYS to study the mechanical properties of the FINE design and to evaluate the hypothesis that when handled, the stress will be concentrated at the corners where the connector meets the shaft. If stress is localized at these points and the substrate is not sufficiently robust, devices may break at these points, as shown in Figure 6-3. This device failure resulted from handling while the connector end was placed in the connector, and the electrode end was being manipulated for electrical testing. Figure 6-3: All-PNB structure that failed upon handling. The FEA utilized the dimensions of the base layer as a model for the microfabricated array. This simulation was concerned with only the geometry of the base structure, and therefore did not take any material properties, such as tensile strength or Young s modulus, into account. As a result, this model cannot predict, for example, a force required to initiate device failure. In the simulation, a displacement was applied to 100

104 the top edge of the electrode end of the device while the connector end was anchored. This simulation is reasonable because the board with the flexible circuit connector is the heaviest part of the device. Therefore, in benchtop testing, the connector end of the device can be considered to be anchored, while the electrode end is free. From this simulation, it was seen that the highest stress was located at the corner where the shaft expands to fit the FH-27 flex circuit connector, as shown in Figure 6-4. Figure 6-4: Finite element modeling of stresses in microfabricated structure when connector end is anchored and a downward force is applied at the electrode end. The results of this simulation were not surprising as the point of the highest stress in the model and the point of failure in the real device was a corner. As such, it was likely that stresses would be localized to that sharp point. In future designs, this issue may be addressed by rounding the edges of all corners in the design. However, while the geometry of the devices may be improved in future designs, it was very apparent from the device in Figure 6-1 and others that the PNB used in these devices was simply not sufficiently robust for handling. These devices are easily broken, even with gentle 101

105 handling. Since it cannot be expected that these devices will be handled with such care during an implant, it was determined that all future devices would be fabricated using the hybrid PNB/LCP process until a suitable formulation of PNB can be identified. 6.4 Flexing-Conductivity Measurements One of the advantages of microfabrication technology is that it allows for batch fabrication, allowing many devices to be fabricated at once. For high yield, microfabricated devices must be processed in a clean room facility as particles in the air that settle on the surface of a substrate may cause device failure. This may still occur in a clean room environment, but the concentration of particles that can settle on a wafer is very small. Typical microfabricated devices are small in area and high in density such that the device yield may still be high even if some devices fail due to defects in the process. In effect, the fraction of the effective area lost to flaws in the fabrication process is small. In contrast to devices typically fabricated photolithographically, the microfabricated FINE arrays are very large, allowing for only eight arrays per wafer or four complete FINE devices. As such, a small flaw during the fabrication process may result in the failure of one of the eighty traces, causing the effective failure of one-eighth of the device area of the wafer. A viable array has the requirements that all ten traces are conductive, and all ten traces are electrically insulated from the others. This further requires that the base polymer layer have a texture adequate for supporting thin-film metal traces across the width of the wafer, that the metal layer photomask pattern is transferred in detail to 102

106 the photoresist lift-off mask, and that there are no flaws in the metal patterning process that would cause shorting between traces. To determine if each trace was continuous, the resistance between the center of the contact pad and the electrode pad was measured with a digital multimeter and micromanipulator probes. A measurable resistance between the contact and electrode pads indicated a continuous interconnect, while an overload resistance that is too large to be measured with the Agilent 34401A multimeter indicated a broken trace. On a single wafer of all-pnb devices with Cr/Pt electrodes, 4 traces were broken. Two of these traces were on a single array, producing five complete electrode arrays, or a 62.5% yield. However, those arrays with one or two broken traces may still be useful, allowing for a 95% yield of viable electrodes. On a wafer of all-pnb devices with Ti/Pt electrodes, 3 traces were broken on a single array, producing seven complete electrode arrays, or a 87.5% yield. The yield of viable electrodes was 96.25%. When normalized to length, the resistivity of the Cr/Pt electrodes measured Ω cm -1, and the Ti/Pt electrodes were measured to have a resistivity of Ω cm -1. The variation in resistivity can be attributed to variation in the metal thickness across the wafer, variation in trace width from one trace to another, as well as variation in probe placement for the measurement. The electrical properties of the electrodes as a function of the mechanical condition of the substrate are an important consideration in the determination of whether the fabricated flexible electrode structures are viable. After being subjected to flexing, the metal must remain continuous. Otherwise, the benefits behind using a flexible substrate are negated. After releasing devices, the resistance across the traces was measured to 103

107 ensure that the traces were continuous. Traces were selected to compare the unflexed resistance of the traces to the flexed resistances on the hybrid LCP-based devices. Resistance data were collected on a conventional probe station stage using an Agilent 34401A multimeter in a 4 wire configuration to measure resistance between the electrode pad and the corresponding pin of the connector. Two probes were placed on the electrode pad and two connections were made to the connector. Five resistance measurements were made in each of five configurations. Between each resistance measurement, both probes on the electrode were picked up, moved slightly, and placed back down to make contact with the electrode. A photograph of a curled device is shown in Figure 6-5. Figure 6-5: Flexed configuration of a PNB/LCP device. Five sets of data were collected: flat shaft, shaft curled inward around 0.25 inchdiameter cylinder, shaft curled inward around 0.5 inch-diameter cylinder, shaft curled outward around 0.25 inch-diameter, and shaft curled outward around 0.5 inch-diameter cylinder. A two-sample t-test was performed between the set of measurements from each curled configuration and the measurements from the flat configuration to determine if the difference in means statistically significant. In each case, it was determined that the 104

108 difference in resistance in configurations was statistically significant, as all p-values were much less than Keesara performed a similar experiment when he made 1200 measurements of his electrode resistance while flat and 1200 resistance measurements of a curled shaft, without picking up the probes. He found that the resistance difference between the curled and flat shafts was statistically significant, but did not suggest a reason for the difference in resistance. I suggest that as the thin film metal traces are wrapped around a cylindrical object, the metal is either stretched or compressed, depending upon the direction of the flexing of the LCP, causing a change in the cross-sectional area of the metal trace. If the metal is stretched, the cross-sectional area will decrease compared to the resistance of the flat trace, and if the metal is compressed, the cross-sectional area will increase compared to the resistance of the flat trace. As the measured resistance is inversely proportional to the cross-sectional area of the trace, the measured resistance of a stretched trace is expected to increase and the measured resistance of a compressed trace is expected to increase. To test this hypothesis, the resistance of a trace on a hybrid PNB/LCP device was tested in five configurations: flat shaft, shaft curled inward around 1/8 inch-diameter cylinder, shaft curled inward around 1/4 inch-diameter cylinder, shaft curled outward around 1/8 inch-diameter, and shaft curled outward around 1/4 inch-diameter cylinder. For each configuration, 5 resistance measurements were made using a digital multimeter and a conventional probe station. The probes were picked up the probe and placed back down between each measurement. The resistance was measured using a 4-wire 105

109 configuration between electrode 9 and the corresponding connector contact for each configuration. The results are shown in Table 6-1. Measurement INWARD 1/8 INWARD 1/4 FLAT OUTWARD 1/4 OUTWARD 1/ Average St. Dev Table 6-1: Mean and standard deviation of measured resistance across trace 9 of hybrid PNB/LCP device in different curvature configurations. This data can be analyzed using a simple model of the curled configurations described in the following equations: Equation 6-1: R = R + R R total flat curled L flat total = ρ + Aflat L A curled curled In this model, the total resistance of the curled trace is treated as a sum of the curled resistance and the flat resistance. The curled segment is modeled as exactly one wrap around the circumference of a cylinder, without taking into account any twisting or offset in the setup. The length of the flat segment is taken to be the difference between the total length of the trace and the length of the curled segment. Using this data, the resistance per unit length can be calculated for the flat configuration, which can then be used to calculate the resistance per unit length and cross-sectional area for each configuration, shown in Table 6-2. Based on these calculations, the resistivity of the sputtered Ti/Pt films is 20.8 µω-cm. This result is reasonable as the resistivity for a 30nm Ti/ 300nm Pt film has been reported to be between about 20 and 30 µω-cm [65]. 106

110 Further, geometrical models can be used to provide a first order approximation of the expected cross-sectional area for each configuration to determine if the calculated cross-sectional areas based on the measured resistance values are reasonable. Several assumptions are made in these models. First, that the thickness of the LCP is 50 µm, the thickness of the PNB is 12 µm, and the thickness of the Ti/Pt metal is 300 nm. Second, because the mechanical properties of the device are dominated by LCP, it is assumed that the device conforms to the curvature determined by the backside of the LCP. As such, to determine the change in cross-sectional area of the metal layer, the thickness of the metal and the thickness of the LCP are determined from building upon the radius from the center of the tubing to the backside of the LCP. This is illustrated in Figure 6-6. Figure 6-6: Geometric model to determine the change in cross-sectional area of the metal: (left) traces are curled outward, away from the tubing; (right), traces are curled inward, facing the tubing. In the case of the traces curling inward, the outside radius of the LCP is taken to be the sum of the tubing radius and the thicknesses of the PNB and LCP layers. The area the flat cross sections of LCP and metal wrapped around the cylinder shown in Figure

111 are then calculated using A = 2π R t, where R LCPouter is the outer radius of the LCP LCPouter LCP0 LCP and t LCP0 is the initial thickness of the LCP (50 µm) and A = 2π R t. metal LCPouter metal0 When wrapped around the cylinder, the outer radius of the LCP is fixed, but the inner surface of the LCP must compress to accommodate the inner radius. The new thickness of the LCP is determined by making a third assumption, that the area of the cross section normal to the direction of curling is a constant. Therefore, the inner radius of LCP can be determined using. Equation 6-2 A = πr πr R 2 2 LCP LCPouter LCPinner LCPinner = π R 2 LCPouter A π LCP The inner radius of the LCP is also the outer radius of the metal, and the inner radius of the metal can be calculated using the same method as in Equation 6-3: Equation 6-3 A = πr πr R 2 2 metal M, outer M, inner Minner, = π R 2 M, outer π A Given this new thickness of the metal, and making a fourth assumption, that the width of the trace is constant at 50 µm, the cross-sectional area perpendicular to the trace can be calculated as: metal A = R R w. Equation 6-4,, (,, ) metal t w M outer M inner trace The expected resistance per unit length can therefore be calculated using:. Equation 6-5 R L curled = ρ A metal metal,, t w. 108

112 The calculated resistance per unit length based on the resistance measurements and the resistance per unit length based on this model for each configuration is compared in Table 6-2. Configuration R/L [Ω/cm] from measurement from model Inward 1/8" Inward 1/4" Flat Outward 1/4" Outward 1/8" Table 6-2: Calculated values for resistance per unit length in each configuration, based on both the mean measured resistance for each configuration and the first order geometric model developed in Chapter 6. This data suggests that the change in cross-sectional area of the trace metal is, in fact, contributing to the change in resistance/length of the Ti/Pt traces. The calculations based on measurement and the calculations based on the model trend in the same direction. Further, the Δ(R/L) between the each curled configuration and the flat trace is well within a factor of 2 for the sample comparison based on the model. Some of the factors that may have caused the discrepancy between the model and the measurement are the assumptions made by the model, as well as a variation in tension between the electrode end and the connector end. The corresponding cross-sectional area, both based on measurements and the model (Figure 6-7) supports the hypothesis that the crosssectional area of the trace will increase when flexed inward and will decrease when flexed outward. In all cases, the resistance of the trace still falls well within limits that allow for current stimulation with a 16-V stimulator. 109

113 16.00 Cross-sectional Area [um 2 ] Inward 1/8" Inward 1/4" Flat Outward 1/4" Outward 1/8" Configuration Measurement Model Figure 6-7: Plot showing the cross-sectional area of the metal (perpendicular to the length of the trace) as a function of trace curvature and configuration calculated from both the mean measured resistance of the trace as well as the geometric model developed in this chapter Hybrid PNB/LCP devices tested on Aplysia buccal ganglion Introduction The flexible electrode arrays were sized to fit the human femoral nerve, one of the larger nerves in the human body. However, as the biocompatibility of PNB has not yet been established, these devices cannot be tested on a human femoral nerve. Efforts have been made to make new molded silicon housings to fit the nerves of other mammals. The electrode arrays can be trimmed to fit into these smaller housings. However, access to mammals for device testing is still limited due to cost and required protocols. 110

114 An invertebrate neural system, the buccal ganglion and nerves from Aplysia californica, has been made available by Prof. Hillel Chiel in the Department of Biology at Case Western Reserve University for in vitro testing of the arrays. This system was convenient due to its availability, as well as due to the fact that this system has been highly studied and is well-known. As the human femoral nerve and the flexible electrode arrays are much larger than the Aplysia buccal nerves and buccal ganglion, it was not practical use the flexible electrode arrays with the silicone housing on the buccal ganglion or nerves. Still, by placing the array, if only one electrode, on top of the ganglion, it was hypothesized that the ganglion could be stimulated to or recorded from. As the microfabricated arrays had not otherwise been tested on a real neural system, the Aplysia buccal ganglion may provide information as to the viability of the devices for a neural system, and provide direction as to where improvements to the design might be made Experimental The Aplysia buccal ganglion was removed from the invertebrate and prepared as follows. A live Aplysia with a targeted mass off approximately g was selected, removed from the tank, placed into a tray filled with Aplysia saline, and weighed. The animal was then anesthetized using MgCl 2 of volume equal to 2/3 of the weight of the Aplysia, with portions of the MgCl 2 going into the head, mid-section, and foot-ends, in that order. Next, incisions were made into the head area of the animal to remove the buccal mass. The buccal mass was placed in a small beaker filled with Aplysia saline. 111

115 Next, the buccal ganglion and associated nerves were separated from the buccal mass and pinned down in a dish. Equation 6-6: Layout of pinned down buccal ganglion and nerves [66] Stimulation Stimulation was to be provided to the Aplysia buccal ganglion, and the response to be recorded using conventional methods of extracellular recording, as described in [67]. The electrode array was interfaced with the ganglion in two ways. First, the array was pressed down onto the ganglion. A square pulse stimulus was applied through each of the microfabricated electrodes sequentially as the signal from buccal nerve 2 was recorded. Correlation between stimulation and traditional extracellular recording may be seen in the plots in Figure 6-8Error! Reference source not found.. 112

116 0.015 Signal voltage at BN2 [V] Pulse timing from Electrode Time (s) Figure 6-8: Recordings at buccal nerve 2 when the ganglion was stimulated with electrode 6 from a microfabricated FINE array. While this first configuration was able to stimulate the ganglion as determined through nerve recordings, there were some issues in the performing the experiment. First, it was difficult to hold the array in place such that the interface between the array and the ganglion was snug. Second, as the LCP is opaque, it was not possible to see exactly which electrode(s) was making contact with the ganglion. Attempts were made to stimulate after flipping the array over and stimulate the underside of the ganglion so that the placement of the ganglion with respect to the electrode geometry could be determined (Figure 6-9Error! Reference source not found. and Figure 6-10). Further, it was hypothesized that using the tendency of the array to curl along with the tension of the 113

117 pinned-down buccal nerves, the contact between the electrodes and the buccal ganglion would be sufficient to allow for stimulation without having to use external force. However, the stimulation was not successful in this configuration. This is likely due to a poor electrode/ganglion interface. Figure 6-9: Stimulation setup with array on the underside of the ganglion. Figure 6-10: Close-up photograph of electrode array on the underside of the ganglion. 114

118 6.5.4 Recording A separate set of devices was fabricated for use in ganglion recording. These devices are for capacitive recording and differ from the original devices in that windows in the capping layer down to the electrodes were omitted. This was done by using the base layer photomask to pattern the PNB capping layer, but the connected contact end was taped off so that there would be an opening for the flex circuit connector. The purpose of this was to ensure that the electrodes are properly insulated in a saline environment. Electrodes submerged in the conductive saline environment will be shorted to each other if there is not a proper seal. Further, it was hypothesized that the neural signals would be able to be sensed capacitively. In essence, this was already occurring on some level as there is an insulating sheath around the buccal nerves and buccal ganglion that acts as a dielectric for any recording. Further, there is a layer of saline between the electrode and the ganglion. As such, adding a thin layer of polymer was simply adding another dielectric layer. This layer has a dielectric constant of 2.42, and its thickness can be tuned. To begin with, the standard 12-µm capping layer thickness was used, which results in an estimated capacitance of 5 pf. As in the stimulation experiments, these devices were most effectively used when applying a downward force onto the electrode array onto the top of the buccal ganglion. The result of a recording is shown in Figure Ideally, correlation could be made between each action potential recorded from buccal nerve 2 with a set of action potentials recording from the buccal ganglion with the microfabricated electrodes. However, this would require careful placement of the electrode on the ganglion to ensure that buccal 115

119 nerve 2 was connected to the neurons in the ganglion in closest proximity to the microfabricated electrode. This device does not allow for the precision placement that is required. However, the recording from buccal nerve 2 suggests that the ganglion and nerves are active and should be producing signals that can be recorded. The recording from the ganglion through the microfabricated electrode is from some collection of neurons in the ganglion. A design comprised of densely packed electrodes may be able to record from a smaller collection of neurons, or possibly even from individual neurons. 5 x 10-5 Signal voltage at BN2 [V] Signal voltage on Electrode 6 [V] 5 x Time (s) Figure 6-11: Comparison of conventional extracellular recording at buccal nerve 2 with the recording of action potentials on electrode 6 of the flexible, microfabricated array. Further, it was determined that one of the major issues in the development of a flexible microfabricated electrode array is determining a method by which the array can 116

120 be clamped onto the ganglion. This clamp would ideally be automatically closed, and would remain closed reliably and with enough force that the array stays in place for a relatively long period of time. This set of experiments demonstrated for the first time that the hybrid PNB/LCP devices with thin film platinum electrodes are capable of neural stimulation and recording. While the microfabricated electrodes used in chapter were not designed for Aplysia, it was possible to stimulate and record by simply pressing the array down on the ganglion. This is not suitable as a long-term solution for ganglion stimulation or recording, but does provide proof-of-concept. A dense, 50-contact array has been designed toward use on the Aplysia buccal ganglion. However, at this time, this design has not been successfully implemented in the hybrid PNB/LCP system. 6.6 Summary This set of experiments evaluated the hypothesis that the all-pnb and the hybrid PNB/LCP devices based on design #1 are suitable for acute peripheral nerve stimulation studies. This hypothesis was supported for both types of devices by pulse testing in saline, which mimics the in vivo environment and sends a relevant current signal down the trace, by showing that the resistance of the thin film electrodes and the saline is sufficiently small to support stimulation with a 16-V stimulator. However, finite element analysis and anecdotal evidence showed that the all-pnb devices would simply not be sufficiently mechanically robust to survive implantation and use. Therefore, all further experiments used only the hybrid PNB/LCP devices. Trace resistivity as a function of flexing was analyzed, showing that the resistance of the traces remains within an 117

121 acceptable range even when curled with a small radius of curvature. Finally, the devices were tested on an in vitro invertebrate neural system for which they were not designed. However, these experiments showed that the microfabricated electrodes are capable of both stimulation and recording. This evaluation points to continuing to use the hybrid PNB/LCP type devices. 118

122 Chapter 7 : Design #2 -- Single-Substrate Peripheral Nerve Electrode Arrays 7.1 Motivation While several complete versions of the microfabricated FINE were successfully fabricated using design #1, these were all designed for use on a human femoral nerve. As this technology is so new, and is therefore untested, in vivo testing on a human femoral nerve was not a realistic goal. As a result, new peripheral electrode arrays were designed with sizing geared toward nerves of varying diameter. These would allow for testing on non-human mammalian nerves. While one of the attractive features of the microfabricated FINE is the ease with which the electrode arrays can be secured around the nerve with a clasp, the lip that was built in to the molded silicone housing to secure the microfabricated electrode array on the inside of the housing also prevents the electrodes from being flush with the nerve. Since the proximity of the electrode to the individual fascicles in the nerve is one of the advantages of the FINE configuration, increasing the distance between the electrode and the fascicles with a silicone lip negates this benefit. A second design issue with the microfabricated FINE is that it did not incorporate the strain relief that had been developed for the conventional hand-fabricated FINE. Any tugging at the connector of the microfabricated FINE will be directly translated into tugging at the nerve. This issue was mitigated with a built-in strain relief in the conventional FINE. When the leads would be pulled, the stress would be concentrated on a section of bent tubing instead of at the nerve. 119

123 7.2 Design These issues described in Section 7.1 were considered in the design of a new set of flexible, thin-film, peripheral nerve electrode arrays. This design incorporates a strain relief, does away with the silicone housing, and introduces recording capability to the peripheral nerve electrodes. The layout of the new photomask set is shown in Figure 7-1Error! Reference source not found., with a schematic of a stimulation device Figure 7-2. Figure 7-1: Mask layout for peripheral nerve electrode arrays. Devices labeled a are recording electrodes, b and c are stimulating electrodes. Magenta corresponds to the outline of the base layer; green corresponds to the outline of the metal layer; blue corresponds to the outline of the capping layer. 120

124 Figure 7-2: Schematic of single-substrate peripheral nerve electrode stimulation array. The electrode arrays labeled a or b consist of two electrode arrays fabricated on a single substrate. To implant and secure around a nerve, the substrate is folded down a centerline between the two arrays, positioned and closed around the nerve, then secured with sutures. This eliminates the need for a molded silicone housing. The two long shafts on each side of the electrode arrays support interconnects between the connector contacts and the electrodes. The U-shaped sections of the shafts act as built-in strain relief and prevent pulling on the connector from translating into pulling on the nerve. The electrodes labeled a are recording electrodes. Each array has a grounding strip on either side. The devices labeled c are similar to the microfabricated FINE inserts, but with varying numbers of electrodes to suit varying nerve diameters. All electrodes in this design have the same dimensions. The connector contacts were designed to interface with a 10-contact Hirose FH-19 flex-circuit connector.. 121

125 7.3 Device Fabrication These electrode arrays were fabricated using the only the hybrid PNB capping layer on an LCP substrate approach only as these materials provided a mechanically robust platform that was fairly straightforward to fabricate. However, while the oxygen plasma etch was successfully used to pattern the LCP in the microfabricated FINE design, this approach was cumbersome. To simplify this process, a CO 2 laser cutting tool was acquired to pattern the LCP. This tool allows for cutting the periphery of the base layer into the LCP without cutting through the copper cladding using laser parameters of 15% power (3 W), 50% speed, and 500 pulses per inch. As such, this tool was used to cut the base layer pattern into the LCP as the first step in the fabrication process. Next, the metal and capping layers were patterned as in Chapter 5, both aligned to the base layer markings. Devices were released by stripping the copper from the backside of the LCP in sodium persulfate. Cross-sectional schematics of the fabrication process are shown in Figure 7-3Error! Reference source not found.. Completed devices are shown in Figure 7-4Error! Reference source not found. and Figure 7-5Error! Reference source not found.. 122

126 Figure 7-3: Fabrication sequence for single-substrate peripheral nerve electrode arrays. Figure 7-4: Completed dual array peripheral nerve electrode device for stimulation. Figure 7-5: Completed PNB/LCP peripheral nerve electrodes in flex circuit connectors soldered to circuit boards. 123

127 7.4 In vivo experiments These devices have been used in two sets of non-human mammal experiments. In the first trial, blank devices without metal or capping layers were placed around a canine nerve. The purpose of this experiment was to determine the feasibility of using sutures to secure the device in place. A photograph is shown in Figure 7-6Error! Reference source not found.. The results of this procedure were promising as the substrate was able to be sutured in place around the nerve without any major issues. Figure 7-6: Microfabricated peripheral nerve electrode blank sutured around a canine nerve. The next trial involved using completed electrodes in non-human primates (macaca mulatta) performed by members of Prof. Tyler s research group and his collaborators at Northwestern University. The purpose of this experiment was assess the performance of these electrode arrays with the objective of demonstrating that the stimulation of different channels on the same nerve would produce different actions in 124

128 the hand by selective stimulation of the nerve fascicles. For this experiment, two devices were implanted on the median nerve in the upper extremity of a subject, with the sites chosen to target different muscles (Figure 7-7). The arrays were folded along the centerline between the arrays, then half of the device was slipped under the nerve and positioned such that the nerve rested in the fold of the array. The two halves of the array were closed and secured with a suture. Figure 7-7: Schematic showing location of implanted electrode arrays. A multi-channel, square wave, charge-balanced current generator was used to stimulate each electrode. Channels 1 and 4 were stimulated on each device. On the distal array, stimulation of channel 1 resulted in adduction of the thumb without movement of other digits. On the same device, when channel 4 was stimulated, the long finger was independently activated. On the proximal array, stimulation of channel 1 resulted in the fingers flexing into a fist. On the same device, when channel 4 was stimulated, the forearm rotated. Photographs from this experiment are shown in Figure 7-8Error! Reference source not found. and Figure 7-9Error! Reference source not found.. These 125

129 results indicate that different muscles can be activated by stimulating different electrodes on the same nerve. Figure 7-8: No stimulation. Figure 7-9: Flexion of fingers upon stimulation of Channel 1 of the proximal array. 7.5 Discussion These electrodes were able to produce different results when different electrodes were stimulated, which was the objective of the experiment. This set of experiments marked the first time that any of the peripheral nerve electrode designs fabricated using LCP or PNB and thin film Pt electrodes had been used in mammals. These results show 126

130 that the microfabricated electrode platform developed through work by Keesara, as well as through this thesis, is able to selectively stimulate in vivo for acute studies, supporting the primary hypothesis of this project. While benchtop testing provided evidence that these devices were fit to be used in acute studies on mammalian nerves, this set of experiments proved that these devices are viable for such a purpose. One point to be addressed in future designs is the lengths of the shafts. The circuit board onto which the flex circuit connector is soldered is relatively heavy and much heavier than the rest of the device. Due to the constraints of the diameter of the Si wafer, the circuit board ends up lying too close to the nerve, making experimentation awkward and difficult. A means by which the circuit board can be distanced from the nerve is desirable. 127

131 Chapter 8 : Conclusion and Future Work This research project has shown that microfabricated peripheral electrode nerve electrode arrays with thin film platinum electrodes on a liquid crystal polymer substrate and capped with a polynorbornene layer are viable in acute mammalian peripheral nerve studies. Investigation of PNB as a candidate material for microfabricated, flexible nerve electrode arrays were carried out further than had been established by Keesara. These investigations involved suggest that PNB is sufficiently moisture resistant as a capping layer for a period of time sufficient for acute in vivo studies. Accelerated testing in PBS showed that the mechanical properties of PNB structures are not negatively impacted by soaking in PNB, which mimics the physiological environment of the body. As a biocompatibility study more thorough than reported in Keesara s thesis [30] supported the hypothesis that PNB is fit for further investigation for biomedical implants. An alternative release mechanism, anodic aluminum dissolution, was demonstrated for use to release the large interfacial area PNB-based arrays from a wafer in place of an HF release. A release with this technique does not evoke the metal adhesion issues that resulted from an HF release. Furthermore, the anodic release does not involve the safety issues that a release with the toxic acid HF. It was determined that all-pnb devices made with the Avatrel TM 2585P formulation of PNB are not sufficiently mechanically robust to withstand the stresses applied to devices in preparing devices for implantation and during the implantation procedure itself. The results of the PNB investigations point toward this material being a good candidate material for nerve electrode arrays, with the exception of the overall 128

132 mechanical strength. Therefore, this material was used as an insulating capping layer on LCP-based devices. This allowed for taking advantage of the superior mechanical properties of LCP, while circumventing the processing difficulties in fabricating a multilayered LCP device. Advancement of the PNB and LCP processes leading toward testing the electrode arrays in vivo included process development utilizing a front-to-back alignment feature and an oxygen plasma etch in RIE to pattern the LCP base layer, then process development utilizing a CO 2 laser engraving system to pattern the shape of the base layer into the LCP. Three electrode array designs have been produced in support of this thesis. The first was designed to be a microfabricated version of the flat interface nerve electrode with two microfabricated arrays incorporated into a molded silicone housing. This design fastened to a flexible circuit connector, which allowed for interfacing with external electronics. Keesara s generic electrode array design did not incorporate a means by which external electronics could be easily interfaced with the flexible electrode array, so this was large advancement toward using these microfabricated electrode arrays for acute in vivo testing. The microfabricated electrodes were used to stimulate to and record from Aplysia buccal ganglia, showing that the thin film Pt electrodes on LCP with a PNB capping layer provide a viable platform for use in nerve stimulation and recording applications. The microfabricated FINE was never tested on the human femoral nerve for which it was designed. However, this design allowed for the identification of improvements that could be made in future designs. These improvements were 129

133 implemented in a second design, which incorporated both electrode arrays on a single LCP substrate that could be folded and sutured closed around the nerve, eliminating the need for a molded silicone housing. Elimination of the silicone housing allows for closer contact between the electrodes and the nerve fascicles, an important consideration in extraneural electrode arrays. A built-in strain relief was also incorporated into the singlesubstrate electrode arrays, preventing a force applied at the connector from being transferred completely to the electrode array around the nerve. This also reduces the likelihood that a stress applied at the connector end of the device will cause the shaft to tear off at the electrode end of the device. A third electrode array was designed to show that a high density of contacts can be fabricated on a flexible polymer substrate. This electrode array was targeted toward use in stimulation and recording from the Aplysia buccal ganglia and nerves. To increase the yield of the devices, and the yield of viable traces on each device, a PNB layer should be cast onto a LCP base layer before spin-casting and patterning the phororesist metal lift-off mask. As PNB on LCP provides a smoother surface than the LCP alone, the demand on the metal traces to conform to the surface is lowered, which may increase device yield. Further, other materials, such as parylene, may be investigated to fabricate these structures as a translucent material is desirable to be able to see where the electrodes are placed in relation to the neurons in the buccal ganglion. In the future, the flexible, thin-film microfabricated electrode arrays designed for mammalian peripheral nerves will need to be subjected to additional testing to determine how long these devices will be viable as an implant. The devices will then need to be developed further to increase the lifetime to allow for chronic implantation for a time 130

134 period of approximately twenty years. The electrical and mechanical robustness of the electrodes should be evaluated. One set of experiments that should be done is to subject the thin film electrodes to many current pulses in saline over time to mimic the number of pulses that will be seen in the lifetime of a chronically implanted electrode. A more electrically or mechanically robust electrode material than thin film Ti/Pt electrodes, such as Pt foil, electroplated electrodes, or conductive diamond, may be necessary for chronic implantation. A noise analysis should be performed to determine the maximum resolution of a signal through the devices. The entire flexible structure should be subjected to mechanical testing involving repeated twisting and bending to determine the robustness of the polymer, metal, and device as a whole against fatigue. Further, the alternative pathways of moisture leakage should be evaluated with alternative interdigitated electrode setups to gain more insight into how long the signal leakage between traces will remain sufficiently small to allow for a high degree of selectivity in recording from and stimulation to nerve fascicles. Alternative, or additional, passivation layers should be investigated to produce a device that absorbs less water than the current devices. Options may include an insulating silicon carbide layer or thin parylene coating the polymer layers. A means by which not only the flexible device can be made hermetically sealed, but also a means by which any implanted ICs that control the electrodes can also be hermetically sealed is necessary. 131

135 Applendix A LCP Bonding A.1 Liquid Crystal Polymer Liquid crystal polymer (LCP) is a good candidate material for a flexible electrode structure because LCP has a very high tensile strength at 120 MPa, and is known for being very chemically inert, biocompatible, and highly resistant to moisture absorption with a water absorption of 0.04 wt. %. An adequate seal between two layers of LCP with a metal layer between them may allow for a device that meets the specifications of an implantable electrode array, and is reliable as a long term chronic implant. A multi-layered LCP structure with a LCP base layer, patterned thin film metal, and a LCP capping layer, is difficult to process using standard microfabrication methods because LCP cannot be spin-cast on a wafer, vapor deposited with a reactor, and it is not photodefinable. Instead, it is available in pressed sheets of polymer with typical thicknesses of 1-,2-, or 4- mil. It is also available with a copper cladding laminate with a thickness of approximately 18μm on one or both sides of the sheet. Multi-layered LCP devices are produced by lamination with the application of heat and high pressure. Patterning a capping layer and base layer of laminated LCP sheets with embedded thin film electrodes is challenging because LCP is opaque. Therefore, patterning of the LCP must be aligned to the metal, which requires a complicated fabrication process. LCP has most commonly been used as a substrate in applications where a flexible printed circuit board is desired. It is a good choice as a high frequency packaging material due to its low thermal expansion characteristics. Further, it is a good choice in antenna 132

136 applications due to its constant dielectric constant of 3.1 over the RF range up to 110 GHz, and also has a very low loss tangest at This makes LCP a good material to use in the design of millimeter wave applications [68]. It has also been used in yarns to produce a fiber with very high strength and rigidity [69] LCP has been used in several biomedical applications, including use as the substrate for a microelectrode array designed to record from a rat sciatic nerve. However, a second layer of LCP was not used as a cover layer, and it is not clear what material was used for a cover layer, if any [70]. Thompson et al. were able to achieve multilayered LCP structures for the packaging of monolithic microwave integrated circuits (MMIC). Their objective was to fabricate homogeneous packaging tall enough to enclose the MMIC and a capacitor. To achieve the desired height, it was required that several sheets of LCP be bonded together. To ensure that the MMIC operated properly, it was required that the structure be homogeneous and not be bonded using any kind of adhesive or epoxy or any other non- LCP material. As such, it was required that the LCP sheets be laminated to one another. In order to do so, the process suggested by Rogers Corp. was followed. This requires that the LCP be placed in a wafer bonder chamber at a pressure of 300 PSI and a temperature of 285ºC for 45 minutes. This was successful in bonding the LCP [71]. Multi-layered LCP MEMS pressure sensor arrays have been fabricated by Palasagaram and Ramadoss [72]. LCP-based pressure sensors are attractive because LCP is low-cost, is mechanically flexible, and has low moisture absorption. These pressure sensor arrays are fabricated by first photolithographically patterning the copper to form the bottom and top electrodes. Next, holes that will serve as cavities in the completed 133

137 structure are machined in a third sheet of LCP with deep reactive ion etching. A thermocompression bonding process is used to laminate the three sheets [72]. Keesara used a lamination approach that involved heating the LCP to its glass transition temperature T g. He was unable to access a vacuum hot press, and therefore used an ordinary hot press and a hot plate to laminate two LCP sheets. He was unable to achieve success using the hot press because it did not have the temperature or pressure capabilities required. However, using a steel roller, he was able to achieve a pressure greater than 200 psi. This was successful in laminating some areas of the LCP sheets, but there were still air bubbles present, as shown in Figure A-1 [30]. By his estimate, the device yield could be expected to be approximately 25% using this method of LCP lamination. Figure A-1: Results from LCP lamination using hot plate and steel roller. This shows a square cover layer laminated wafer-shaped base layer. Air bubbles are circled in red. [30] 134

138 A.2 LCP Bonding Experiments Building on previous work, a wafer bonder tool that is traditionally used to bond silicon wafers was used to laminate two sheets of liquid crystal polymer together. This was first done without patterned metal between the sheets so that it could be determined if this method was viable. Determination of the optimal wafer bonding parameters to bond two sheets of LCP required several experimental runs. For each run, two pieces of single-sided copper-clad, 2-mil thick LCP (R/Flex 3600) were sandwiched between two silicon wafers, in the arrangement shown in Figure A-2. Graphite chuck Si Wafer #2 Cu cladding #2 LCP sheet #2 LCP sheet #1 Cu cladding #1 Si Wafer #1 Chamber stage Figure A-2 :Cross-sectional view of bonding layers. With the flags in the EV501 wafer bonder pulled out, the first silicon wafer was placed on the stage, aligning the wafer flat with the pins in the chamber. The first LCP sheet, which had been cut into the shape of a Si wafer, was placed on top of this wafer, copper-side down. A flat cut on the LCP sheet was aligned with the primary wafer flat on the Si wafer. Separately, the other sheet of LCP was placed on top of the other silicon wafer with the copper side of the LCP toward the wafer, and the primary flats aligned to each other. After pushing the flags inward such that they were on top of the wafer and 135

139 LCP sheet in the chamber, the second silicon wafer/lcp sheet combination was placed on top of the flags in the chamber, with the LCP sheet facing downward and both the sheet and the wafer resting on the flags. Ideally, the two LCP sheets would not yet be in physical contact, but in reality they made contact in the center of the sheets due to the flexibility of the sheets. The graphite chuck was then set on top of the top silicon wafer, and the chamber was closed, being sure to tighten the pins evenly such that pressure was applied evenly. After setting the stack height, the data recorder was started and the chamber was pumped down to 1x10-3 mbar, followed by a nitrogen purge for 30 seconds. Next, the chamber was pumped down to 3x10-4 mbar. Upon reaching this chamber pressure, the wafer bow button was pressed such that the pin in the center of the chuck would push down onto the wafer and LCP sheet, making forced contact at the center of each LCP sheet. Next, the flags were then pulled out, removing the barriers at the perimeter of the LCP sheets. The piston was pressed downward to apply pressure evenly to the entire wafer. At this point, heat was applied in order to reach a temperature of 300ºC. The bonding time is the time that the setup was allowed to sit at the specified temperature upon reaching that specified temperature. After the bonding time had been reached, the heat and vacuum pump were turned off and the cooling water was turned on. The piston was lifted and the chamber was purged. The chamber was then vented, opened, and allowed to cool. 136

140 Run Temperature (ºC) Pressure (mbar) Force (N) Bonding Time (min) x x x Table A-1: Description of each LCP-LCP bonding experiment in wafer bonder. It was seen in the first run (Table A-1) that the bond was good around the perimeter of the wafer, but there was clearly an air-bubble in the center. To improve the bond, the bonding time was increased to 30 minutes in the second run. After being exposed to the bonding conditions for 30 minutes in Run 2, the LCP had begun to flow from the between the edges of the copper cladding, indicating that the bonding time was too long. In Run 3, the bonding time was reduced to 15 minutes, but this was still long enough to cause the LCP to flow between the layers of copper cladding. As such, the optimal bonding time was determined to be 5 minutes. After bonding, the copper cladding was stripped from both of the bonded LCP sheets in a sodium persulfate solution. The resulting structure is shown in Figure A-3. The structure was an improvement upon Keesara s method of LCP lamination as there were no trapped air bubbles. However, there was a section along the periphery of the structure where the lamination was incomplete. The critical structures fabricated from Keesara s mask design lie in the center of the wafer, not around the periphery. As such, the bonding flaw on the outside edge of the structure would not lower the device yield if devices were produced from this wafer. 137

141 Figure A-3: Photograph of bonded LCP sheets demonstrating flexibility. A.3 Patterning bonded LCP sheets One LCP bonded structure was processed further to investigate the viability of the bond in additional process steps. For these experiments, two new sheets of LCP were laminated together and the copper cladding was not stripped. The copper cladding was to serve as an etch mask for the LCP in a LCP wet etchant. As such, the Cu on one side of the structure was patterned using photolithography and wet etched using sodium persulfate. The LCP was then etched in a solution of 20/40/40 % by wt of Ethanolamine/ KOH/Water at 85ºC. A single wafer holder was used to assist in the wet etch of the LCP. This holder secured the LCP structure in place, as well as protected the edges and backside of the structure from being etched. Two o-rings seal the holder such that the backside of the wafer will not be exposed to the etchant. As such, the LCP field area and the vias that correspond to where metal contacts would have been patterned were etched in the heated solution over a period of 2 hours. After the LCP etch was complete, the patterned, bonded structures were submerged in a sodium persulfate solution to strip any 138

142 remaining copper from the surface of the LCP. The resulting devices are shown in Figure A-4. Figure A-4: a) Photograph of completed patterned, bonded, non-metalized LCP structures b) Photograph demonstrating flexibility of bonded structures. Patterning the bonded LCP structures established that these additional process steps did not result in catastrophic damage to the bonded structure. The patterned devices appear to have a good bond between them, and there was no evidence of the integrity of the bond between the two sheets being compromised. These results suggest that if the fabrication process can be augmented with patterned electrodes, a hermetic electrode array using LCP as both a substrate and capping layer may be able to be fabricated. The superior mechanical and moisture resistance properties of LCP make this highly desirable. 139

143 A.4 Patterning Non-bonded LCP with Non-copper etch mask In an effort to remove copper, a toxin in the human body, from the LCP device fabrication sequence, thin film metals were explored as alternatives for use as masks for the LCP wet etch. First, wafer-shaped LCP sheets were cut from a roll of copper-clad LCP using a craft knife and a Si wafer as a template. Next, the copper cladding was stripped from each sheet in a sodium persulfate solution. Prior to metallization, each sheet was secured to the sputterer chuck with Kapton tape. The next step was to sputter deposit the 250 nm Al layer onto the same surface of the LCP. Next, the LCP sheet was adhered to a wafer Al side up, and the Al was patterned using photolithography and Al etchant. The single LCP sheet with the patterned Al was then placed into the single wafer holder and the heated KOH/ETA solution. The thin film Al was not sufficient to act as an etch mask for the LCP in the heated KOH/ETA solution. Instead, the Al pattern was stripped from the LCP, resulting in failure. It was therefore determined that thin film Al is not an appropriate etch mask for LCP in the wet etch solution. The above process was repeated using Ni with a Ti adhesion layer, but the result was the same. It was clear that a thin film metal was not going to be sufficient to act as an etch mask for the LCP in the wet etch solution. This led to further concern over whether the wet etch could be used for devices that have patterned thin film metal traces between the two sheets of LCP. As such, new solutions were sought that did not require the use of the LCP wet etch solution 140

144 The LCP wet etch is very difficult as it requires such harsh chemicals, as well as elevated temperatures. However, it has the advantage that it is readily available and only requires the correct chemicals, a container to do the process, a means to heat the chemicals, and a single wafer holder for the LCP. Once all of these are obtained, it is relatively simple to prepare the setup. On the other hand, other methods that may be used to etch or cut LCP requires expensive equipment such as a reactive ion etching system or a laser cutting system. The use of these tools is discussed in subsequent chapters. A.5 Bonding to metalized LCP Thin film Ti/Pt electrodes and traces were patterned on a sheet of LCP to investigate the effects of bonding two sheets of LCP on embedded thin film metal traces. On a second sheet of LCP, holes were cut using an X-acto knife such that when grossly aligned to the bottom LCP sheet, the connector contact pads and the electrode pads were exposed. Next, the two LCP sheets were bonded together using the optimal recipe determined in Section 2.2. Upon completion of the bond, the conductivity between the electrode pads and the connector contact pads was measured using a digital multimeter. A relatively low resistance measurement (< 10kΩ) would be indicative of a continuous trace that had remained viable through the bonding procedure. Photographs of the resultant devices are shown in Figure A

145 Figure A-5: (Top Right) traces stemming from shaft to electrodes appearing to be broken at interface. (Top Left) traces stemming from shaft to electrodes enlarged. (Bottom) connector end of the device with traces appearing to be cut at capping layer interface. The electrodes were examined under a microscope after the bonding was complete. It became apparent that at the edge where the window of the top sheet of LCP crossed the metal traces, the sharp edge of the LCP sliced through the metal traces. None of the traces were viable as they were all broken at the window edge. As such, the effect of bonding on the traces between the two sheets of LCP was unable to be determined. It is not clear if those traces were left intact, or if they were broken during the bonding process as there were no exposed bonding pads that could be used to measure the resistance of the traces to determine continuity. Additional attempts to determine the effect of bonding on the thin film traces between two sheets of LCP have not been done. Since it is suspected that the sharp edges 142