Atomic layer deposited aluminum oxide and parylene C bi-layer encapsulation for biomedical implantable devices

Biomedical implantable devices have been developed for both research and clinical applications, to stimulate and record physiological signals in vivo. Chronic use of biomedical devices with thin-film-based encapsulation in large scale is impeded by their lack of long-term functionality and stability. Biostable, biocompatible, conformal, and electrically insulating coatings that sustain chronic implantation are essential for chip-scale implantable electronic systems. Even though many materials have been studied to for this purpose, to date, no encapsulation method has been thoroughly characterized or qualified as a broadly applicable long-term hermetic encapsulation for biomedical implantable devices. In this work, atomic layer deposited Al2O3 and Parylene C bi-layer was investigated as encapsulation for biomedical devices. The combination of ALD Al2O3 and CVD Parylene C encapsulation extended the lifetime of coated interdigitated electrodes (IDEs) to up to 72 months (to date) with low leakage current of ~ 15 pA. The long lifetime was achieved by significantly reducing moisture permeation due to the ALD Al2O3 layer. Moreover, the bi-layer encapsulation separates the permeated moisture (mostly at the Al2O3 and Parylene interface) from the surface contaminants (mostly at the device and Al2O3 interface), preventing the formation of localized electrolyte through condensation. Al2O3 works as an inner moisture barrier and Parylene works as an external ion barrier, preventing contact of AI2O3 with liquid water, and slowing the kinetics of alumina corrosion. Selective removal of encapsulation materials is required to expose the active sites for interacting with physiological environment. A self-aligned mask process with three steps was developed to expose active sites, composed of laser ablation, oxygen plasma etching, and BOE etching. Al2O3 layer was found to prevent the formation of microcracks in the iridium oxide film during laser ablation. Bi-layer encapsulated iridium oxide had higher charge injection capacity and similar electrochemical impedance compared with Parylene C coated iridium oxide film after deinsulation. The Al2O3 and Parylene C bi-layer encapsulation was applied to Utah electrode array (UEA)-based neural interfaces to study its long-term performance. The median tip impedance of the bi-layer encapsulated wired Utah electrode array increased slowly during the 960 days of equivalent soak testing at 37 °C. Impedance for Parylene coated UEA dropped 50% to 75% within 6 months. In addition, bi-layer coated fully integrated Utah array-based wireless neural interfaces had stable power-up frequencies at ~910 MHz and constant RF signal strength of -50 dBm during the 1044 days of equivalent soaking time at 37 °C. This is much longer than lifetime achieved with Parylene C coating, which was about one year at room temperature.

ATOMIC LAYER DEPOSITED ALUMINUM OXIDE AND PARYLENE C BI-LAYER ENCAPSULATION FOR BIOMEDICAL IMPLANTABLE DEVICES by Xianzong Xie A dissertation submitted to the faculty of The University of Utah in partial fulfillment of the requirements for the degree of Doctor of Philosophy Department of Electrical and Computing Engineering The University of Utah December 2013 Copyright © Xianzong Xie 2013 All Rights Reserved The Unive r si t y of Utah Gradua t e School STATEMENT OF DISSERTATION APPROVAL The dissertation of __________________ Xianzong Xie________________ has been approved by the following supervisory committee members: Florian Solzbacher , Chair 09/03/2013 Date Approved Loren Rieth , Member 08/26/2013 Date Approved Carlos Mastrangelo , Member 08/26/2013 Date Approved Richard Normann , Member 08/28/2013 Date Approved Richard Cohen , Member 08/30/2013 Date Approved and by _________________Gianluca Lazzi_________________ , Chair/Dean of the Department/College/School o f ______ Electrical and Computer Engineering and by David B. Kieda, Dean of The Graduate School. ABSTRACT Biomedical implantable devices have been developed for both research and clinical applications, to stimulate and record physiological signals in vivo. Chronic use of biomedical devices with thin-film-based encapsulation in large scale is impeded by their lack of long-term functionality and stability. Biostable, biocompatible, conformal, and electrically insulating coatings that sustain chronic implantation are essential for chip-scale implantable electronic systems. Even though many materials have been studied to for this purpose, to date, no encapsulation method has been thoroughly characterized or qualified as a broadly applicable long-term hermetic encapsulation for biomedical implantable devices. In this work, atomic layer deposited Al2O3 and Parylene C bi-layer was investigated as encapsulation for biomedical devices. The combination of ALD Al2O3 and CVD Parylene C encapsulation extended the lifetime of coated interdigitated electrodes (IDEs) to up to 72 months (to date) with low leakage current of ~ 15 pA. The long lifetime was achieved by significantly reducing moisture permeation due to the ALD Al2O3 layer. Moreover, the bi-layer encapsulation separates the permeated moisture (mostly at the Al2O3 and Parylene interface) from the surface contaminants (mostly at the device and Al2O3 interface), preventing the formation of localized electrolyte through condensation. Al2O3 works as an inner moisture barrier and Parylene works as an external ion barrier, preventing contact of AI2O3 with liquid water, and slowing the kinetics of alumina corrosion. Selective removal of encapsulation materials is required to expose the active sites for interacting with physiological environment. A self-aligned mask process with three steps was developed to expose active sites, composed of laser ablation, oxygen plasma etching, and BOE etching. Al2O3 layer was found to prevent the formation of microcracks in the iridium oxide film during laser ablation. Bi-layer encapsulated iridium oxide had higher charge injection capacity and similar electrochemical impedance compared with Parylene C coated iridium oxide film after deinsulation. The Al2O3 and Parylene C bi-layer encapsulation was applied to Utah electrode array (UEA)-based neural interfaces to study its long-term performance. The median tip impedance of the bi-layer encapsulated wired Utah electrode array increased slowly during the 960 days of equivalent soak testing at 37 °C. Impedance for Parylene coated UEA dropped 50% to 75% within 6 months. In addition, bi-layer coated fully integrated Utah array-based wireless neural interfaces had stable power-up frequencies at ~910 MHz and constant RF signal strength of -50 dBm during the 1044 days of equivalent soaking time at 37 °C. This is much longer than lifetime achieved with Parylene C coating, which was about one year at room temperature. iv To My wife Yi Li and my dear family, who made all this possible TABLE OF CONTENTS ABSTRACT...........................................................................................................................iii ACKNOWLEDGEMENTS...................................................................................................ix CHAPTER 1. INTRODUCTION............................................................................................................... 1 1.1 Implantable Devices.......................................................................................................2 1.2 Electrode Arrays............................................................................................................. 3 1.3 Encapsulation of Implantable Devices..........................................................................5 1.4 Failure of Implantable Devices..................................................................................... 7 1.5 Hypothesis, Approaches and Specific Aims................................................................9 1.6 References.....................................................................................................................14 2. STATE OF THE ART: NEURAL ELECTRODE ARRAYS, ENCAPSULATION MATERIALS, AND SELECTIVE DEINSULATION.......................................................22 2.1 Introduction...................................................................................................................22 2.2 Neural Electrode Arrays...............................................................................................22 2.2.1 Microwire Arrays..................................................................................................23 2.2.2 Silicon-based Microelectrode Arrays.................................................................. 24 2.2.2.1 The Michigan Array....................................................................................... 24 2.2.2.2 The Utah Electrode Array..............................................................................25 2.3 Hermetic and Thin-film-based Encapsulation............................................................26 2.3.1 Hermetic Encapsulation........................................................................................ 27 2.3.2 Thin-film-based Encapsulation.............................................................................28 2.4 Nonpolymeric Materials for Encapsulating Implantable Devices........................... 29 2.4.1 Silicon Oxide and Silicon Nitride.........................................................................29 2.4.2 Ultrananocrystalline Diamond and Diamond-like Carbon.................................30 2.4.3 Silicon Carbide......................................................................................................31 2.5 Polymeric Materials for Encapsulating Implantable Devices...................................32 2.6 Atomic Layer Deposited Al2O3 ...................................................................................34 2.6.1 The Chemistry of ALD Al2O3 ..............................................................................35 2.6.2 The Growth Rate of ALD Al2O3 ..........................................................................37 2.6.3 Plasma-enhanced ALD......................................................................................... 38 2.7 Parylene........................................................................................................................38 2.7.1 Parylene Variants...................................................................................................39 2.7.2 Parylene Deposition...............................................................................................39 2.7.2.1 Vaporization....................................................................................................40 2.7.2.2 Pyrolysis Process........................................................................................... 41 2.7.2.3 Polymerization................................................................................................41 2.7.3 Parylene Adhesion.................................................................................................42 2.8 Tip Deinsulation........................................................................................................... 43 2.9 References.....................................................................................................................53 3. PLASMA-ASSISTED ATOMIC LAYER DEPOSITION OF AL2O3 AND PARYLENE C BI-LAYER ENCAPSULATION FOR CHRONIC IMPLANTABLE ELECTRONICS....................................................................................................................68 4. LONG-TERM BI-LAYER ENCAPSULATION PERFORMANCE OF ATOMIC LAYER DEPOSITED AL2O3 AND PARYLENE C FOR BIOMEDICAL IMPLANTABLE DEVICES.................................................................................................74 5. SELF-ALIGNED TIP DEINSULATION OF ATOMIC LAYER DEPOSITED AL2O3 AND PARYLENE C COATED UTAH ELECTRODE ARRAY-BASED NEURAL INTERFACES ........................................................................................................................84 5.1 Abstract........................................................................................................................84 5.2 Introduction...................................................................................................................85 5.3 Materials and Methods.................................................................................................88 5.3.1 Fabrication of SIROF Test Structures and UEAs............................................... 88 5.3.2 Deinsulation Process for Alumina and Parylene Coating..................................89 5.3.3 Experiments........................................................................................................... 90 5.4 Results and Discussion.................................................................................................92 5.5 Conclusion.....................................................................................................................97 5.6 References...................................................................................................................107 6 . LONG-TERM RELIABILITY OF AL2O3 AND PARYLENE C BI-LAYER ENCAPSULATED UTAH ELECTRODE ARRAY-BASED NEURAL INTERFACES FOR CHRONIC IMPLANTATION..................................................................................111 6.1 Abstract.......................................................................................................................111 6.2 Introduction................................................................................................................ 112 6.3 Experimental Details..................................................................................................115 6.3.1 Integrated Neural Interfaces...............................................................................115 6.3.2 Alumina and Parylene C Deposition................................................................. 117 6.3.3 Tip Deinsulation..................................................................................................117 6.3.4 Testing Setup.......................................................................................................118 6.4 Results and Discussion...............................................................................................119 6.5 Conclusion...................................................................................................................124 6.6 References...................................................................................................................132 7. CONCLUSIONS AND FUTURE WORK.................................................................... 137 vii 7.1 Conclusions.................................................................................................................137 7.1.1 Long-term Performance of ALD Al2O3.............................................................138 7.1.2 Selective Etching of ALD Al2O3 and Parylene C............................................. 139 7.1.3 Long-term Reliability of Al2O3 and Parylene C............................................... 140 7.2 Future Work............................................................................................................... 142 7.2.1 Long-term In Vivo Experiment..........................................................................142 7.2.2 Hydrogen Reduction or Elimination in Al2O3 Film..........................................142 7.2.3 Cap Layer for Preventing Al2O3 Dissolution.................................................... 143 7.2.4 Multilayer Configuration....................................................................................143 7.2.5 Nucleation on Neural Interface Surfaces...........................................................144 7.2.6 Biocompatibility Improvement...........................................................................144 7.2.7 Improving Substrate Stability.............................................................................145 7.3 References...................................................................................................................146 viii ACKNOWLEDGEMENTS Many people deserve credit for assisting me during the past five years of my PhD study. My deepest gratitude goes to my advisor Prof. Florian Solzbacher for his continuous support of my PhD research, for his insight, motivation, and immense knowledge. I would like to sincerely thank my committee members for their guidance: Dr. Loren Rieth, Prof. Richard A. Normann, Prof. Carlos Mastrangelo, and Prof. Richard Cohen. Special thanks to Dr. Loren Rieth for his insight on scientific problems and article proofreading. Appreciation is extended to Dr. Prashant Tathireddy for invaluable discussion. I also thank Rohit Sharma, Ryan Caldwell, Dr. Mohit Diwekar, Mahender Avula, Tanya Abaya, Je-Min Yoo, Dr. Xiaoxin Chen, Dr. Sandeep Negi, Dr. Rajmohan Bhandari, Dr. Asha Sharma, and Dr. Layne Williams for their assistance during my PhD research. Gratitude is extended to Microfab staff for their help on equipment maintenance and training. Finally, I would like to thank my dear wife Yi Li, and my family for their unconditional support and love, which made this long journey possible. CHAPTER 1 INTRODUCTION Implantable electronic systems and devices have undergone significant development over the past few decades for both research and clinic applications, to monitor, stimulate, and record physiological responses in vivo. The progress in implantable devices is made possible by both the accumulating knowledge of human neuron-motor systems, and technology advances in semiconductor industry and microelectromechanic systems (MEMS). Neural interfaces are implantable devices developed for applications such as neuroprosthetics and neuroscience to diagnose and treat neuron-related disorders and diseases. Lack of long-term functionality and stability of these devices has prevented them from widely chronic usage. Various factors could contribute to the failure of implantable devices, including device corrosion and decreased encapsulation impedance caused by coating degradation, connector problems, and/or foreign body responses. Fully integrated, wireless, silicon-based neural interfaces have been developed to eliminate connector problems and remove the risk of infection associated with percutaneous wired connectors. Long-term stable, conformal, biocompatible, and highly insulating coating materials and methods have been investigated to address the failure modes due to encapsulation failure for chronic implantation. Even through a large number of materials have been proposed for encapsulating biomedical implantable devices, they all have 2 unique drawbacks and limitations. In this work, an atomic layer deposited (ALD) Al2O3 and chemical vapor deposited (CVD) Parylene C bi-layer encapsulation is studied as a candidate for encapsulating chronic neural interface implants. The ALD Al2O3 works as a water vapor barrier due to its extremely low water vapor transmission rate; Parylene C thin film serves as an ion barrier. Moreover, Parylene prevents the direct contact of liquid water with ALD Al2O3, thus stopping the ALD Al2O3 dissolution. This bi-layer encapsulation is used to significantly reduce water vapor permeation and separate the substrate surface contaminants (ions, metal particles, etc.) from the penetrated water moisture at the interface between Al2O3 and Parylene. The introduction chapter is composed of five sections, starting with various implantable devices and their applications, followed by the requirements for encapsulating implantable devices. Then encapsulation failure modes are discussed for different materials and methods in section 4. The approaches, aims, and results of this work are introduced in the last section. 1.1 Implantable Devices Implantable devices, such as bio-sensors, cardiac peacemakers, implantable cardioverter defibrillators, cochlear implants, deep brain stimulators, and neural interfaces, are being implanted into patients worldwide [1 -8] for different research and clinic purposes. Pacemakers utilize implanted electrodes to deliver electrical pulses to control the heart rate. Cochlear implants are electronic implantable devices that directly stimulate the cochlea to enable hearing in profoundly deaf patients. According to the Food and Drug Administration, approximately 219,000 people have received cochlear 3 implants as of December 2010. Deep brain stimulators are used to treat movement and affective disorders such as chronic pain, Parkinson's disease, epilepsy, tremor, and dystonia, by sending electrical stimulating pulses to the specific parts of the brain [9]. A deep brain stimulator typically consists of three components: the neurostimulator, the lead, and the extension. The extension, essentially an insulated wire, connects the lead, and the neurostimulator. Deep brain stimulation has demonstrated therapeutic benefits for otherwise treatment-resistant diseases [10-13]. Neural interfaces have been developed for neuroprosthetics to restore functions for patients with communication issues between the central and peripheral nervous system or the muscles. The potential of regaining functions using neural prosthesis has been pursued for decades for paralyzed patients [14-18]. Clinical trials of neural interfaces were made possible by major advances in developing implantable systems, which demonstrated the potential efficacy of this technology [8, 19-21]. Combination of neural interfaces with prosthetic devices as therapies for neuronal disorders is very promising. 1.2 Electrode Arrays Two major types of electrodes have been developed for neural interface devices to record or stimulate neural signals: surface electrodes and penetrating electrodes. Surface electrodes are mostly noninvasive or less invasive, thus causing less tissue damage and foreign body response. This is at the cost of low selectivity and sensitivity. They usually measure localized field potentials (LFPs) from relatively large populations of neurons. Also, they lack the ability to access neural signal from deeper in the tissue. Penetrating electrodes can detect smaller signals from a single neuron unit due to high selectivity and 4 sensitivity, at the cost of tissue damage and foreign body response. Examples of penetrating electrodes for chronic implantation include the iridium wire array [22, 23], the floating microelectrode array (FMA) [24], the Michigan array [15], and the Utah electrode array [25, 26]. The iridium wire arrays have been investigated to have long-term stability for chronic implantation [22, 23] . The good recording performance of the iridium wire array was achieved by the findings that there was negligible connective tissue encapsulation or edema at the active electrode tips and large neurons presented around the active electrode tips. The issue with this handcrafted iridium wire array is the lack of quality control and repeatability. Therefore, they are not suitable for mass production. Floating microelectrode arrays (FMAs) with electrodes made of platinum/iridium 70%/30% have showed the potential for chronic implantation [24]. The advantages of FMAs are the flexibility of electrode length and potential random distribution of individual electrodes. However, the fabrication process is expensive, time-consuming, and also lacks control. Development of silicon-based micromachined electrode arrays with long-term stability, repeatability, and potential of mass production for commercialization was made possible by the advances in MEMS technology. Even through large numbers of fabrication methods and configurations of microelectrode arrays can be found in the literature for neural recording and stimulation, two major silicon-based electrodes have been commercialized and widely used: the Michigan array and the Utah electrode array (UEA). Active recording sites of the Michigan array are positioned along the silicon electrode shanks. This design enables the Michigan array to be able to record neural signals from variable depths of tissue on each electrode shank. However, tissue damage during the implantation process decreases the quality of recorded signals. Also, the glial scar formation after insertion can isolate the active sites from adjacent neurons and impair the recording capabilities. The UEA consists of 100 microelectrodes with typical lengths of 1.0 and 1.5 mm and pitch of 400 |im, as shown in Fig 1.1. The encapsulation of the electrode tip was removed to expose the active metal (iridium oxide) electrode sites for recording/stimulation purposes, as shown in Fig 1.2. In contrast with the Michigan array, active sites of the UEA are only available at the electrode tips, where tissue damage is typically minimal. The UEA is also the only FDA-cleared neural interfaces, which has an investigation device exemption (IDE). Clinical research usage of UEA has been report in recent years [8, 21, 27], demonstrated the efficacy of the UEA-based neural interfaces for neuroprosthetics. 1.3 Encapsulation of Implantable Devices Implantable devices integrated with active electronics need to be protected from the physiological environment in order to perform their designated functions, which is a particular challenge for chronically implanted devices. Encapsulation needs to meet specific requirements for individual applications, but there are some basic requirements that apply to most of the implantable devices, including biocompatibility, biostability, sufficient mechanical strength, high electrical resistance, low dielectric constant, conformal and pin-hole free coating, low process temperatures, and compatibility with sterilization process(es). 5 (a) Biocompatibility: The encapsulation materials must be nontoxic, and should have minimal or no contribution to the acute and chronic foreign body responses due to the implantation of the devices, which can be bioinert or bioactive. (b) Biostability: There must be no discernible dissolution or degradation of the material and no material property changes in the physiological environment for the intended lifetime of the device. (c) Mechanical strength: Sufficient mechanical strength is required to maintain the coating integrity during the handling (surgical and fabrication) and implantation process. (d) High insulation resistance and low dielectric constant: Coating with high insulation resistance and low dielectric constant can reduce the signal loss through shunting and capacitive cross-talk between channels, and maximize signal to noise ratio. (e) Conformal and pin-hole free coating: Conformal coating helps to maintain the original geometry of the devices, which can affect the surgical process and foreign body response after implantation. (f) Low process temperature: Implantable medical systems usually contain multiple components, composed of various materials. Polymers, solders, metal contacts, and integrated circuits are susceptible to high temperature (over 200 °C). The lowest temperature tolerance among all materials in the whole device sets the limit for encapsulation process temperature. 6 7 (g) Sterilization: The coating has to be able to withstand one or more sterilization processes, which is essential for an implantable device before implantation. The common sterilization procedures are steam and ethylene oxide gas. Other than the aforementioned requirements, selective deinsulation of encapsulation without affecting the overall coating performance to expose the localized active sites is necessary for information exchange between the implantable device and the physiological environment. The proper selective etching process has to be developed for the encapsulation. 1.4 Failure of Implantable Devices There are three main failure mechanisms for implantable devices: connection failure, failure due to foreign body response, and encapsulation failure. Connection failure is ascribed to mechanical stress, handling forces, etc. This failure mode can be solved by developing wireless implantable devices [28-33]. The elimination of tethering forces can also reduce foreign body response [34]. The absence of wires for connection also reduces infection likelihood [35]. Another major failure mode of an implantable device results from the foreign body response. The initial tissue damage due to the implantation process evokes inflammatory response to protect the body from potential hazards. The mechanisms behind this are not fully understood. The foreign body response can be partially alleviated by minimizing surgical trauma. Also, implantation procedures and surgical techniques can be optimized to reduce the foreign body response induced by the implants [36, 37]. The geometry of implanted devices is reported to also have impact on acute immune response [38]. The presence of foreign materials provoke the foreign body 8 response, leading to formation of scar tissue encapsulating the implants [39]. As the scar tissue grows, it isolates the implant from its surrounding tissue and neurons, thus attenuating the electrical signals. The eventual complete isolation of signals leads to a loss of function for implantable devices. For UEAs, the typical progression of degradation is first decreasing single unit signal intensity, then loss of single unit but continued LFPs and multi-unit recording, with gradually degrading loss of signals. The foreign body response is affected by mechanical flexibility [40] and surface properties of the implants [41]. Coating implantable devices with a noninteractive or antifouling surface [42] to reduce the protein absorption was used to reduce scar tissue formation and enhance biocompatibility. Use of conductive polymers (such as PEDOT and Polypyrrole) combined with agents aimed at promoting neuron growth around recording sites is also utilized to improve the performance of implantable devices [43, 44]. Encapsulation failure is another major failure mode. The primary failure points of implantable devices are the interfaces of various components and the coating layer, where water vapor and ions permeate and accumulate [45]. Moisture ingress can lead to failures such as open circuits [46], short circuits [47-49], corrosion/dissolution of different materials [50], electrical leakage [47], and delamination of coating materials. The consequences are catastrophic and can lead to complete device failure. Tremendous efforts have been devoted to address this issue using different materials and approaches, including silicon carbide, diamond-like carbon (DLC), silicon nitride, urethanes, polyimide, Teflon, silicone, Parylene, etc. [51-58]. Most of those materials and methods have their own drawbacks, which make them nonideal candidates for encapsulation of implantable devices. For example, silicon carbide typically requires high deposition 9 temperature and is prone to have pinholes [59], silicon nitride slowly dissolves in PBS solution [57], and DLC coating has adhesion problems and also delaminates over time [56, 60-62]. Polymeric materials exhibit relative high water vapor transmission rate (WVTR) and poor adhesion [57, 58]. This work is focusing on the encapsulation failure and trying to address this issue by combing atomic layer deposited (ALD) Al2O3 and Parylene C. The ALD Al2O3 is used to block water vapor permeation; the Parylene C layer is an ion barrier and also prevents liquid water from contacting with ALD Al2O3 and dissolving it. 1.5 Hypothesis, Approaches and Specific Aims Encapsulation with low water vapor transmission rate (WVTR) is important to reduce the water vapor permeation and slow down the corrosion process. Typically WVTR for polymers is in the order of 10"2 gmm/m2 day, which is too high for moisture sensitive applications [63]. Atomic layer deposited (ALD) Al2O3 (alumina) has demonstrated extremely low WVTR in the order of 10"10 gmm/m2 day [64-67]. The biocompatibility of bulk Al2O3 is comparable to that of corrosion-resistant metals like titanium [6 8 ]. It has been reported that ALD alumina coated glass slides had slight better biocompatibility compared with uncoated glass slides in terms of cell proliferation and cell activity [69]. Also bulk alumina was used as substrate for floating microelectrode arrays for neural recording, suggesting it is reasonable for use with neural tissue, at least if encapsulated [24]. Liquid water is known to corrode ALD Al2O3 thin films [70] mostly due to the high concentration of hydrogen in the form of hydroxyls in the film [71, 72]; therefore, ALD Al2O3 alone is not suitable for encapsulation of biomedical implants directly exposed to 10 physiological environment. A thin-film encapsulation layer that drastically reduces the dissolution kinetics of the alumina film could be highly effective at preserving the integrity of this layer, and allowing it to maintain its ultra-low permeation characteristics. We investigate the use of Parylene-C as the overlayer, due to its demonstrated effectiveness for implantable devices, and in particular the UEA. Parylene C has been widely utilized in various electronic and biomedical devices to protect them from the harsh environment. Biomedical applications of Parylene include blood pressure sensors, stent coating, bone pins, bio-MEMS, and neural recording/stimulating electrodes [53, 73-78]. Among polymeric materials, Parylene C has a relatively low water absorption (0.1%) [79]. Parylene C has demonstrated thermal and chemical stability [79]. Parylene C is an excellent ion barrier to Na+, K+, Cl", etc. [80], which is critical for devices exposed to physiological environment. Parylene is also believed to be nontoxic after being used in medical devices for many decades with very few negative reports [8, 81-83]. Although Parylene C has relatively low WVTR of 0.4 g mm/m2 day [84] among polymeric materials, better moisture barrier is needed to significantly reduce the water vapor permeation rate and slow down the condensation of moisture around ion contaminants to form electrolyte, in order to further extend the lifetime of implantable devices [85]. Our hypothesis is that the combination of ALD Al2O3 and CVD Parylene C can address the encapsulation failure mode by significantly reducing moisture permeation. Moreover, the bi-layer encapsulation separates the permeated moisture (mostly at the Al2O3 and Parylene interface) from the surface contaminants like ions and metal particles (mostly at the device and Al2O3 interface), preventing the formation of electrolyte through moisture condensation. Al2O3 works as an inner moisture barrier and Parylene works as an external ion barrier, and slows down the kinetics of alumina corrosion. The specific aims of this study are as follows: (a) Develop deposition process and optimize process parameters of ALD Al2O3 at low temperature using plasma-assisted ALD. (b) Characterize ALD Al2O3 thin film and evaluate the Al2O3 and Parylene C bi-layer encapsulation performance based on interdigitated electrode (IDE) test structures. (c) Investigate and compare the effect of temperature, topography, and bias voltage on Al2O3 and Parylene C bi-layer encapsulation and Parylene C encapsulation on lifetime based on IDE test structures. (d) Develop and optimize selective etching process for Al2O3 and Parylene C bi-layer encapsulated implantable devices to expose active sites for neural recording/stimulation. (e) Study the effect of deinsulation process on charge injection capacity (CIC), charge storage capacity (CSC), and electrochemical impedance of the tip metal iridium oxide for neural interface applications. (f) Evaluate the long-term impedance stability, device reliability, RF power-up frequency, and signal strength constancy of Al2O3 and Parylene C bi-layer coated Utah electrode array (UEA)-based neural interfaces. Chapter 2 reviews the state-of-the-art of coating materials and related deposition techniques. Research background and literature review of different coating approaches are also covered in this chapter. 11 Chapter 3 is reprinted form the article published in Applied Physics Letter [86 ]. It includes the deposition, characterization of Al2O3 and Parylene C bi-layer encapsulation. It also reports the in vitro soak testing performance of the bi-layer encapsulation based on IDE test structures. Chapter 4 covers the long-term performance of Al2O3 and Parylene C bi-layer encapsulation for chronic implantable devices [87]. The effects of temperature, topography, and bias voltage on encapsulation performance were studied and comparison between Al2O3 and Parylene C bi-layer encapsulation and Parylene C encapsulation were performed. Chapter 5 reports a self-masked deinsulation process for Al2O3 and Parylene C bi­layer encapsulated neural interfaces. The effects of deinsulation process on CIC, CSC, and electrochemical impedance of iridium oxide were evaluated. Chapter 6 assessed the long-term performance of Al2O3 and Parylene C bi-layer encapsulation on neural interfaces. Long-term impedance stability of wired UEAs and RF power-up frequency and signal strength consistency of wireless integrated neural interfaces were assessed. Chapter 7 concludes the work of this dissertation and proposes future work. 12 13 Fig 1.1 Scanning electron micrograph of the UEA with 100 (10 by10) silicon electrodes. The electrode length is 1.5 mm and space between electrodes is 400 |im. Fig 2.2 Scanning electron micrograph of single exposed electrode of the UEA with exposed active tip. 14 1.6 References [1] D. C. Klonoff, "Technological advances in the treatment of diabetes mellitus: Better bioengineering begets benefits in glucose measurement, the artificial pancreas, and insulin delivery," Pediatric Endocrinology Reviews, vol. 1, pp. 94-100, 2003. [2] T. 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Sharma, et al., "Long­term bi-layer encapsulation performance of atomic layer deposited Al2O3 and Parylene c for biomedical implantable devices," Biomedical Engineering, IEEE Transactions on, vol. 60, 2013. CHAPTER 2 STATE OF THE ART: NEURAL ELECTRODE ARRAYS, ENCAPSULATION MATERIALS, AND SELECTIVE DEINSULATION 2.1 Introduction This chapter conveys the state of the art of neural electrode arrays and advances in wireless neural recording and stimulation by integrating active electronics, and existing encapsulation methods for implantable devices. Wireless neural interfaces proposed new challenges for encapsulation, especially for devices under continuous bias voltage in the physiological environment. The methods of traditional hermetic encapsulation and emerging thin-film-based encapsulation are compared. Various encapsulation material candidates that are widely used in both research and industry for biomedical implantable devices and their corresponding application techniques are reviewed. The advantages and drawbacks of each material/deposition method are covered based on the requirements of implantable devices. The last section of the chapter discusses the deinsulation processes for selectively exposing active electrical sites for biomedical applications. 2.2 Neural Electrode Arrays The technology developed in this work is a platform technology that can likely be applied to many biomedical implantable devices. This work focused on developing 23 encapsulations for neural interfaces, due to the recognized challenges like sizes, materials compatibility, and complex geometries in these systems. Neural interfaces based on the Utah electrode array (UEA) have been developed for decades as implantable devices for recording/stimulating neurons at the University of Utah [1-3]. Neural interfaces were chosen to evaluate the performance of atomic layer deposited Al2O3 and Parylene C bi­layer encapsulation because they are widely used in both research and clinical trials and lack of long-term effective encapsulation for chronic implantation. Moreover, UEA-based neural interfaces are very representative of implantable devices because of their complex topography, combination of different materials, and complicated multiple fabrication processes. 2.2.1 Microwire Arrays Microwire arrays for neural recording in cortex was pioneered by Salcman using glass insulated Pt/Ir wire [4] or Parylene insulated pure Ir wire [5]. The use of single microwire was designed to reduce the displacement of microwire after implantation due to the cortical movement. However, single wire can only record from a small area. Therefore, microwire arrays were developed to increase the available recording sites and investigate the arrangement of neural circuits. Microwire arrays use commercially available corrosion-resistive wires that have enough mechanical strength for fabrication and insertion. Typical materials include tungsten, iridium, or platinum/iridium alloy wires. Williams et al. reported a 35-^m tungsten wire-based microelectrode array [6 ]. The array consists of 2 rows of11 microwires, electrically connected to a back-end connector, as shown in Fig 2.1. Liu et al. fabricated another type of microelectrode array by inserting 16 Ir wires into epoxy backplate to form a 4 by 4 array [7, 8]. Commercially available microwire arrays have been developed similar to aforementioned structures. Microprobe Inc. provides up to 64 channels of stainless steel or Pt/Ir fine wire-based microelectrode arrays with Teflon or polyimide as insulation material [9]. Tucker Davis Technologies uses polyimide insulated tungsten microwires to make arrays up to 64 channels [10]. The assembly process and size constraints limit the mass production of microwire arrays. 2.2.2 Silicon-Based Microelectrode Arrays Compared with microwire arrays, silicon-based micromachined electrodes have many attractive properties. Advantages include precise control of electrode geometry, the elimination of time-consuming assembly processes, potential of mass production with high repeatability, and compatibility with integration of active electronics for wireless implantable systems. Two major and commercially available designs of silicon-based microelectrode arrays are the Michigan array and the Utah electrode array (UEA). 2.2.2.1 The Michigan Array The University of Michigan has been developing and improving the silicon-based Michigan array for the last four decades [11-13]. A micromachined planar array on a tapered silicon beam was reported by Wise et al. [13], shown in Fig 2.2. Silicon oxide was used as insulation materials for the array tip. A revised version of the Michigan multielectrode array was developed by Najafi in 1985 [14]. Advances include the achievement of multiple recording sites on a single shank with 3-mm in length, 50-^m in 24 25 width, and 15-^m in thickness. Highly boron-doped silicon by diffusion was used as a etch stop during the wet etching. The fabrication process was compatible with metal-oxide semiconductor (MOS) integrated circuit fabrication allowing the addition of circuitry for signal amplification and multiplexing [15, 16]. The development of silicon ribbon cable provided high flexibility between the microelectrode array and back-end connectors [17]. This was further improved by the later-developed Parylene cable [11]. The three-dimensional multielectrode Michigan array was designed and fabricated by Hoogerwerf in 1994 [12], as shown in Fig 2.3. Other major modifications include the change of recording-site metal from gold and platinum or iridium [18], and adding wireless capabilities for eliminating connecting cables [19]. 2.2.2.2 The Utah Electrode Array Other than the aforementioned commercially available Michigan array, the Utah Electrode Array (UEA) is the other popular option for neural interfaces, with commercial availability and food and drug administration (FDA) clearances. Richard Normann first designed and fabricated the three-dimensional UEA for intracortical stimulation [2]. A dicing saw was used to cut silicon wafers and create columns with dimension of 150 |im square, 1.5 mm tall, and pitch of 400 |im. The columns were first thinned and then tapered by wet etching. Platinum was used as electrode active site metallization and the UEA was insulated by polyimide using dip coating. Later versions of the UEA were improved by the utilization of Pt/Ti/W/Pt as active tip metal [1]. Glass was used as insulation material between individual electrodes to improve the electrical isolation of electrodes and reduce cross-talk [1]. The Utah Slant 26 Electrode Array (USEA) was later developed by varying the length of electrodes in one direction, which gave the ability of accessing the nerve fascicle across the cross-section of peripheral nerves [20]. The insulation of the UEA was further improved by adopting chemical vapor-deposited Parylene C to replace dip coated polyimide, in order to obtain a conformal and pin-hole free insulation layer [21]. Trials with human patients have been performed and the long-term effectiveness has been demonstrated with UEA-based neural interfaces [3, 22, 23]. Wireless integrated neural interfaces based on UEA have been reported recently to eliminate the wire bundles for electrical connection [24, 25], as shown in Fig 2.4. Compared with microwire arrays and Michigan arrays, the UEA has the advantages of batch process reproducibility and tip recording from undamaged tissue. This work uses UEA-based neural interfaces to evaluate the performance of proposed bi­layer encapsulation. 2.3 Hermetic and Thin-film-based Encapsulation Implantable devices with active electronics must be able to perform their intended functions effectively and stably in the physiological environment over a long period of time. Encapsulation is critical to the success of implantable devices by protecting them from the body fluids. There are two major categories of encapsulation: hermetic and near-hermetic encapsulation. Hermetic encapsulation typically utilizes metal capsules, biocompatible ceramics, and glasses to build up an airtight and waterproof environment and keep the implanted device from being corroded by body fluids [26]. Near-hermetic encapsulation, on the other hand, use thin film layers to slow down or prevent the ingress of body fluids into the devices. 27 2.3.1 Hermetic Encapsulation Typical materials that provide a hermetic barrier for implantable devices are metals, ceramics, and glasses. Metallic packaging generally uses a biocompatible metal capsule such as titanium. Metal capsule-based hermetic encapsulation has been successfully applied to pacemakers [27], cardioverter defibrillators [27], neuromuscular stimulators [27], and cochlear implants [28]. However, power-receiving coil and communication antenna need to be placed outside of the metallic hermetic packaging to avoid the interference of radio-frequency signal and loss of power through eddy-current formation in the packaging. Biocompatible ceramics and glass have the advantage of radio-frequency transparency over metallic hermetic packaging. The application includes neuromuscular microstimulators [29], cochlear implants [30], and artificial retina implants [31]. Biocompatible ceramics used for hermetic coating include alumina [32-34], zirconia [35], and ceria stabilized zirconia poly-crystal [36, 37]. Biocompatible glasses, such as borosilicate glass, have been used for encapsulating neuromuscular microstimulators [38]. One of the challenges for hermetic encapsulation is feedthroughs. Conducting wires are necessary for signal entering and exiting the hermetic packaging. Fusion welding methods are usually used to form a hermetic seal between the packaging components and conductive wires, which could potentially go through a high-temperature process and damage the encapsulated device. Additionally, hermetic capsules tend to take more space compared with thin-film-based encapsulation, which conflicts with miniaturization for space-limited applications, such as neural implants. 2.3.2 Thin-film-based Encapsulation There are mainly two different kinds of materials used for thin-film-based near-hermetic encapsulation for implantable devices: nonpolymeric and polymeric materials [26, 39]. Nonpolymeric materials mainly include silicon oxide, silicon nitride, silicon carbide, ultrananocrystalline diamond (UNCD), and diamond-like carbon (DLC); polymeric materials include polytetrafluoroethylene (PTFE), silicone elastomer, polyimide, and Parylene, etc. Nonpolymeric materials like silicon oxide, silicon nitride, silicon carbide, UNCD, and DLC usually require chemical vapor deposition (CVD) with a relatively high temperature to obtain a uniform coating layer, which makes them incompatible with implantable devices incorporated with active electronics and various polymeric materials. Other drawbacks include dissolution in phosphate saline solution (PBS) (silicon nitride and silicon oxide) [40, 41], structural defects (pin-holes)[42], and nonconformal deposition [43]. On the other hand, polymeric materials are typically flexible and inexpensive. They also tend to have low process temperature. The challenges for polymer encapsulation of chronic implantable devices are relative high water vapor transmission rate (WVTR) (typically > 1 g/m2 day)[44], degradation of the material itself, and difficulties of forming conformal and pin-hole free thin films. The high WVTR makes moisture penetrating through the coated devices easier. A large volume of permeated moisture condenses around ion contaminants to from electrolyte, which corrodes the coated substrate and thus reduces the lifetime of the device. Degradation of materials includes hydrolytic, oxidative, and enzymatic mechanisms that deteriorate the chemical structures of the 28 29 polymer [45]. Polymers applied by dip coating or spin coating usually are lack of control of the film quality, such as pin-holes, adhesion, thickness, and its variations. 2.4 Nonpolymeric Materials for Encapsulating Implantable Devices 2.4.1 Silicon Oxide and Silicon Nitride Silicon oxide can be easily grown by oxidation process or deposited by CVD process. Silicon oxide has been used as insulation layer for the Michigan array back to 1970 [13]. The silicon oxide was formed by a thermal oxidation process, which requires high temperature (~ 1000 °C). Later CVD silicon oxide was used for protecting the Michigan array against the physiological environment [14]. SiO2 passivation has also been applied to retina implants [46]. The major issue with SiO2 passivation is that it dissolves in vivo over time. It has been showed that SiO2 coating exhibited dissolution and underneath electronics (with 500 nm SiO2 passivation) were exposed for retina implants after 10 months in vivo [40, 46]. Silicon nitride has also been utilized as an encapsulation material for implantable devices. It was used as part of the passivation for the Michigan array through CVD process as an addition to SiO2 passivation [14]. Also, it was used to encapsulate photodiode in retina implants [42]. Silicon nitride often exhibits pin-holes and chemical reaction with saline solution with a dissolution rate of 1 - 2 nm per day, which lead to the loss of insulating functions to the underneath photodiode for retina implants [42]. Similar to silicon oxide, silicon nitride dissolves in PBS solution at 37 °C [41], indicating it is not an ideal candidate for encapsulating chronic implantable devices. 30 2.4.2 Ultrananocrystalline Diamond and Diamond-like Carbon Ultrananocrystalline diamond (UNCD) films were reported to be relatively chemically inert [47], biocompatible, and bio-inactive [48-50], which make them a promising candidate for encapsulating implantable devices. In addition, the unique nanostructure of UNCD leads to a high wear-resistance and low friction surface [51], which facilitate maintaining the integrity of the coating during and after implantation. UNCD film was developed first by Argonne National Laboratory with microwave plasma-enhanced chemical vapor deposition (MPECVD) [52, 53]. Diamond films were traditionally synthesized by conventional chemical vapor deposition (CVD) using H2/CH4, requiring high process temperature of above 700 °C [54]. MPECVD of UNCD was achieved at 400 °C because of both the change of chemistry from H2/CH4 to Ar-rich/ CH4 plasma and reduction of activation energy from 20 kcal/mol to 6 kcal/mol [50, 54, 55]. The crystal orientation and surface morphology can be tuned by controlling nucleation and MPECVD parameters [53]. The usage of MPECVD greatly reduced the synthesis temperature from 700 °C to 400 °C, resulting in a reasonable temperature range that is compatible with most silicon-based devices. However, most of implantable devices contain polymeric materials other than silicon. The ideal process temperature would not exceed 200 °C. Diamond-like carbon is an alternative with relatively low deposition compared with UNCD. It has similar properties compared with UNCD, including low friction, low wear, chemical inertness, and high biocompatibility [56-58]. DLC have been extensively used for coating biomedical implants [59-61], including orthopedic (e.g., in artificial hips, 31 knees, and joints) applications, and vascular applications (e.g., in stents, heart pumps, and heart valves) [62, 63]. DLC can be deposited at room temperature or higher [59, 64], depending on the deposition method. Methods that have been used to deposit DLC includes direct ion beam deposition [65], pulse laser deposition [6 6 ], filtered cathodic deposition, sputtering [67], and plasma enhance chemical vapor deposition (PECVD) [60]. Different deposition methods and parameters can greatly affect the quality and properties of the DLC film. Despite wide applications in implantable devices, one of the major issues for DLC is its adhesion to the substrate [63]. Adhesion-promoting interlayers have been developed to address this issue, including Ti, Si, Cr, Si-DLC, etc. [63]. However, dissolution of the interlayer was observed, which caused the delamination of DLC and therefore significantly undermined the coating performance [6 8 ]. 2.4.3 Silicon Carbide Silicon carbide can be grown as single crystalline c-SiC, polycrystalline p-SiC, and amorphous a-SiC. Both c-SiC and p-SiC require a process temperature over 600 °C [69], which is not compatible with most implantable devices. Amorphous SiC has low dielectric constant (4.2-4.9) and low water intake [70, 71]. It also exhibits chemical inertness due to the strong Si-C chemical bonds [72]. A-SiCx:H has been applied in fields such as optoelectronics [73], solar cell technology [74], and surface passivation [74-76]. In addition, the usage of a- SiCx:H as biomedical encapsulation material has been reported to improve the hemocompatibility of implanted 32 stents [77, 78], and reduce thrombosis and restenosis rate after angioplasty [79]. It has also been investigated as insulation for neural interfaces [41, 43]. A few low-temperature deposition techniques have been developed to grow a-SiC in order to meet the temperature requirements for a variety of applications. Those techniques include pulsed laser deposition (PLD) [80, 81], sputter deposition [82-84], and plasma-enhanced chemical vapor deposition [41, 43, 85, 86 ]. PLD and sputtering deposition both have flux directionalities during the deposition process, leading to nonuniform thickness of the coated film. Moreover, sputtered films tend to have structural defects (pin-holes), which is not acceptable for coating of implantable devices. High compressive stress induced by PLD requires high annealing temperature (~ 600 °C) to relieve the stress [87]. PECVD is the most commonly used method for depositing SiC at low temperature [43, 8 8 ]. Both the plasma and thermal energy are used to dissociate gas precursors and create free radicals for chemical reaction. Therefore, the temperature of PECVD required for dissociating precursors can be significantly lowered than for conventional CVD using thermal energy only [8 8 ]. Incorporation of hydrogen into SiC during the growth is almost unavoidable due to the high hydrogen concentration in the precursor. 2.5 Polymeric Materials for Encapsulating Implantable Devices Polytetrafluoroethylene (PTFE), best known as Teflon, a trademark of DuPont, is a bioinert, biostable, and low friction material, and has a chemical formula of (C2F4)n. PTFE does not dissolve in vivo [89]. The good biocompatibility of PTFE may be partial 33 due to its bioinertness and hydrophobic surfaces that minimize foreign-body recognition in vivo [89]. PTFE has also been used as a material for vascular grafts [90]. Plasma treatment has been used to improve the cell adhesion of endothelial cell onto PTFE vascular grafts by adding amide functional groups [90]. The surface hydrophilicity, which is favorable for cell adhesion, was also improved by this technique. PTFE is generally applied by either spraying or dipping coating [91]. It is almost impossible to obtain a uniform coating film using liquid-phase techniques for implantable devices with complex 3-D geometries. Alternative deposition techniques have been investigated to gain uniform PTFE films. Plasma-enhanced chemical vapor deposition (PECVD) [92] and hot-wire chemical vapor deposition (HWCVD) [93] have been proposed as potential techniques to conformally deposit fluorocarbons that have similar chemical formulas to PTFE. The PECVD deposited PTFE has two major drawbacks. First, it contains byproducts like CF3 and CF other than CF2 groups. Additionally, free radicals were present in PECVD PTFE films, which can potentially react with oxygen and water over time and thus may undermine the coating performance of the film [92, 94]. HWCVD is proposed to overcome the aforementioned drawbacks by minimizing the incorporation of byproducts and free radicals in the deposited films during PECVD [94]. Initiated chemical vapor deposition (iCVD) is one type of HWCVD, which utilizes chemical initiators to initiate polymerization reaction. The usage of initiator enables further decreasing the filament temperature required to dissociate the precursors. Consequently, iCVD requires lower power and has higher deposition rate compared with conventional HWCVD. HWCVD or iCVD fluorocarbon films have better fluorine to 34 carbon ratio (close to 2) than films deposited by PECVD due to higher percentage of CF2 and lower byproduct incorporation. Silicone has been widely used for biomedical implantable devices [95]. The traditional challenge of applying silicone conformally was overcome by the iCVD technique. Researchers have demonstrated the possibility of conformally depositing siloxane by iCVD [96, 97]. In addition, iCVD siloxane has been used as effective encapsulation for neural probes [98]. One major drawback of silicone coating is that it has relatively higher water vapor transmission rate (WVTR) than polymers like PTFE and Parylene [99]. The large volume of penetrated moisture due to high WVTR can nucleate around interface contaminants to form electrolytes, which corrode the device and significantly reduce its lifetime [10 0 ]. Polyimide is another widely used polymeric material for implantable devices. Polyimide has demonstrated its biocompatibility in neural interface devices and retina implants [101-104]. However, it is very challenging to obtain a conformal polyimide coating by the commonly used spin-casting technique. Additionally, polyimide has high water absorption and water vapor transmission rate (WVTR) [49, 99], which can undermine the long-term insulation performance of polyimide. Table 2.1 shows WVTR of commonly used polymers for encapsulating biomedical implantable devices and atomic layer deposited Al2O3. 2.6 Atomic Layer Deposited Al?O3 Atomic layer deposited (ALD) Al2O3 is one of the two materials used in this work. Alumina (Al2O3) is known for its high hardness, high abrasion resistance, and bioinertness [105]. It has demonstrated good biocompatibility in vivo in forms of both bulk alumina and ALD alumina thin film [106, 107]. Alumina has been widely used as a bioceramic for dental and bone implants [34, 108-110]. It is also used as a substrate material for floating microelectrode arrays for neural recording [111]. Retina implants utilized alumina as a coating material to insulate the device from physiological environment [112]. What differentiates alumina from other materials is its extremely low water vapor transmission rate (WVTR). Atomic layer deposited (ALD) alumina has been demonstrated as an excellent moisture barrier with WVTR as low as in the order of ~ 10' 10 gmm/m2 day [113-116], to prevent the degradation of extremely moisture-sensitive organic light emitting diodes (OLEDs). Another major application of alumina is the passivation of solar cells for higher efficiency [117, 118]. The success of applying alumina for solar cell and OLED passivation is mostly attributed to the ALD process, which generates conformal, dense, and pin-hole free films. Different from bulk Al2O3, ALD Al2O3 can be easily corroded by liquid water [119], due to the incorporation of hydrogen in the form of OH groups in the film [120, 121]. 2.6.1 The Chemistry of ALD Al2O3 ALD is able to achieve atomic layer control and conformal deposition using sequential, self-limiting surface chemical reactions [122]. ALD was proposed back to the 1960s and started to gain popularity in the beginning of the 1990s, driving its potential application in scaling down microelectronic devices. A typically ALD cycle is composed of four steps [123]: 35 (a) A self-limiting chemical reaction of the first precursor (precursor A) with the absorbed second precursor from the previous cycle. (b) A purge of inert gas to remove excess precursors and byproducts from the chamber. (c) A self-limiting reaction of the second precursor (precursor B) with the absorbed precursor A. (d) A purge of inert gas. The growth of material is achieved by repeating of the aforementioned four steps. The characteristics of self-limiting surface reaction lead to the ALD as a surface-controlled process, where parameters other than precursors and temperature have little or no effect on the growth of the film. When sufficient time between precursors is present, gas phase transport of precursors into tight gaps with high aspect ratio is possible, contributing to the very high degree of conformality. Therefore, films grown by ALD are extremely conformal, uniform, and reproducible. The first report of thermal ALD Al2O3 using trimethylaluminum (TMA) and H2O as precursors was back in the late 1980s and early 1990s [124, 125]. The surface chemistry of ALD Al2O3 using H2O as oxidant can be described as [122, 126, 127] (A) AlOH* + Al(CH3)3 ^ AlOAl(CH3)2*+ CH4 (B) AlCH3* + 2H2O ^ AlOH* + CH4 where the asterisks denote the surface species. The growth of Al2O3 happens during the alternating exposures to TMA and H2O. The overall chemical reaction of ALD Al2O3 is [1 2 2 ] 2Al(CH3)3 + 3H2O ^ Al2O3 + 3CH4 AH = -376 kcal 36 The formation of a strong Al-O covalent bond makes the reaction very efficient. ALD is very similar to chemical vapor deposition (CVD) based on binary reaction format of A + B ^ C + byproduct. The distinction between ALD and binary reaction-based CVD is the time of the precursor presence. For CVD, reactants A and B are present simultaneously on the surface and it grows films continuously. On the other hand, reactants A and B are timely separated in the ALD process and only one reactant is present at a time in the form of a monolayer. The growth of the film is more stepwise. It is the separation of precursors in time domain which ensures that the ALD only happens at the surface. 2.6.2 The Growth Rate of ALD Al2O3 The plasma-enhanced ALD reaction for Al2O3 used in this work is illustrated schematically in Fig 2.5. The growth of ALD Al2O3 is extremely linear with a rate of about 1.1 A per cycle due to the highly repeatable self-limiting surface chemistry. The deposition rate of ALD Al2O3 is distinct from the monolayer thickness of Al2O3, which is estimated to be 3.8 A [1 2 2 ]. The surface chemistry, surface species, and precursor coverage limits the possibility of depositing a full monolayer of Al2O3. The deposition rate of ALD Al2O3 is temperature dependent because the chemical reaction during the ALD growth requires thermal energy to generate free radicals. Higher process temperature leads to the reduction of time needed for each cycle. However, studies have shown that the growth per cycle decreases progressively as temperature increases from 177 to 300 °C, due to less coverage of AlOH and AlCH3 surface species at 37 38 higher temperatures [128, 129]. A typical measurement of ALD Al2O3 growth with quartz crystal microbalance (QCM) is shown in Fig 2.6. 2.6.3 Plasma-enhanced ALD Other than the widely used H2O oxidant during ALD process, other oxidants like ozone and oxygen have also started to gain popularity to obtain better dielectric properties and lower leakage current [130-133]. Oxidants like ozone and oxygen make chemical reaction much less likely or even impossible by using only thermal energy due to the lack of free radicals. By creating free-radicals using plasma, plasma-enhanced ALD can deposit Al2O3 using TMA and O2 at temperature as low as room temperature [133] . The low temperature is extremely help for coating of thermal fragile substrate like polymers or microsystems with underfiller materials. The plasma-enhanced ALD Al2O3 films exhibit higher electrical breakdown voltage and higher dielectric constant, which are ascribed to the higher density of the deposited films than thermal ALD Al2O3 films [134]. The improved electrical properties lead to better passivation of silicon substrate [135, 136]. In addition, plasma-enhanced ALD reduces the hydrogen incorporation in Al2O3 films compared with thermal ALD Al2O3 films [136, 137], thus improving the film quality and reduce leakage current [138]. 2.7 Parylene Poly-para-xylene (Parylene) was first discovered by Szwarc in 1947 using chemical vapor deposition (CVD). Parylene was deposited using para-xylene solvent as precursor. This method had very low yield and high impurities. The deposition process was 39 improved by Gorham from Union Carbide using vacuum pyrolysis of di-para-xylene precursors [139, 140]. The room temperature deposition process and good chemical and physical properties make Parylene very attractive for many electronics and biomedical applications [21, 141-144]. 2.7.1 Parylene Variants The chemical structure of the Parylene monomers is composed of an aromatic group with methylene groups attached at the para positions. Parylene variants are created by replacing the aromatic or aliphatic hydrogen atoms with other side groups. There are three most common variants: Parylene C, Parylene D, and Parylene N. Fig 2.7 shows the chemical structure of those three major Parylene variants and Table 2.2 presents the their properties. Parylene N and Parylene C have been categorized as USP class VI polymers [145], which requires demonstration of biocompatibility and indiscernible toxicity in the systemic injection test, intracutaneous test, and implantation test. In addition, Parylene C has lower water vapor transmission rate compared with Parylene N (0.4 vs 5.4 gmm/m2 day) [99]. Parylene C is widely used for biomedical applications. 2.7.2 Parylene Deposition Parylene is deposited by chemical vapor deposition (CVD). The deposition process is composed of three major steps: vaporization, pyrolysis, and polymerization. The chemical reaction of each step is described in Fig 2.8. Di-para-xylene dimer is used as 40 precursor and is vaporized and then pyrolyzed into free radical monomers, which then undergo polymerization to form poly-para-xylene (Parylene) at lower temperature. The Parylene deposition system is designed according to the aforementioned three-step polymerization process, composed of vaporization furnace, pyrolysis furnace, and deposition chamber. Additionally, a cold trap and a vacuum pump are included in the system to absorb the unreacted monomer and maintain the required low pressure (around 10 mTorr for base pressure) during the deposition, respectively. Fig 2.9 is a schematic view of the deposition system. 2.7.2.1 Vaporization The dimer sublimation rate varies with vaporization temperature, affecting the morphology and crystallinity of deposited Parylene [146, 147]. The threshold temperature for dimer sublimation is about 60 °C and sublimation rate increases as the sublimation temperature rises. A typical vaporization temperature is around 130 °C to generate a sublimation rate sustaining sufficient vapor pressure and deposition rate. Other than temperature, the sublimation rate is also dependent on the exposed surface of the dimer (i.e., kinetics factors). The surface area of exposed dimer decreases as a function of sublimation time; therefore, a slight increase in sublimation temperature is needed to compensate the sublimation rate drop and maintain a constant vapor pressure. A constant vapor pressure helps to obtain Parylene films with consistent properties during the whole deposition process. 2.7.2.2 Pyrolysis Process Pyrolysis is the process in which heat is used to dissociate dimers into reactive monomers, which participate in the polymerization process at lower deposition temperature. The required temperature for fully converting dimers into monomers has been reported to be 565 °C [148]. The suggested pyrolysis temperature setting is 670 °C from the vendor and various reports [140]. Complete dissociation of dimers into monomers is needed to reduce the presence of dimer in the deposited Parylene film. High sublimation rate has been reported to reduce the residence time of dimer in the pyrolysis chamber, leading to incomplete dissociation of dimers. The "un-cracked" dimers contained in the reactive monomers presents in the deposited Parylene film without any chemical polymerization process [147]. The short residence time of dimer vapor in the pyrolysis furnace can be compensated by increase the length of the pyrolysis furnace. 2.7.2.3 Polymerization The polymerization process is achieved through free radicals and it happens at temperature lower than 80 °C and typically uses room temperature. The low-temperature deposition process makes it very attractive to coat materials and devices that require a low thermal budget. The essence of the growth of Parylene film is a free radical chemical reaction process. A vapor-deposition polymerization model and surface roughening kinetics have been proposed [149, 150]. The growth of the Parylene film is achieved by two chemical reactions: initiation, in which new carbon chains are formed, and propagation, in which existing carbon chains are extended into higher molecular weight. 41 42 Unlike physical vapor deposition, reaction only occurs at the end of the polymer chain instead of many available sites during the vapor deposition polymerization. Surface diffusion, intermolecular interaction, and chain relaxation can occur during the deposition [149]. Pressure, temperature, and deposition rate can affect the polymerization process and thus the film properties. The deposition pressure has been suggested to be lower than 100 mTorr to obtain high-quality films [151, 152]. It is also suggested that low-temperature polymerization process increases the Parylene growth rate and chain length [151]. Longer chain length leads to better thermal stability of the Parylene film. 2.7.3 Parylene Adhesion Parylene is known to have poor adhesion to substrates like polyimide, silicon, glass, and metallic materials. Therefore, surface modification is necessary before Parylene deposition to improve the adhesion. There are mainly three approaches have been proposed to enhance the adhesion: (a) Using plasma to remove contaminants and clean the surface [153, 154]. (b) Using plasma-enhanced chemical vapor deposition to deposit a thin layer of polymer as adhesion promoter [154, 155]. (c) Using gas or liquid phase silanization process to add an adhesion layer with functional groups to form chemical bonds with both substrate and Parylene [156]. The thin film deposition method has been reported to be more effective in improving adhesion than the plasma cleaning process for a variety of substrates [154]. However, the requirement for additional PECVD system adds more complications. Silanization process is less complicated and inexpensive compared with the plasma-based techniques for improving adhesion. It utilizes functional groups from organic-silane to provide covalent linkage between substrate and Parylene. Among many potential coupling agents [157], Silquest A-174® silane (Gamma- Methacryloxypropyltrimethoxysilane) is the most common option for improving Parylene adhesion. Silquest A-174® silane, patented by Union Carbide, has a simplified form of X-R-SiY3. Fig 2.10 shows the chemical structure of Silquest A-174® silane. The Y group is a hydrolysable functional group which can form silanol intermediates when reacting with water, which further form covalent bond with substrate. The X group is a vinyl group that can form covalent bond with reactive monomers during the Parylene deposition process. 2.8 Tip Deinsulation Selective removing of encapsulation materials is required to expose the active sites for interacting with physiological environment. Primary methods of selective etching include reactive ion etching and laser ablation in the manner that keeps the encapsulation of the rest of the device intact. Removal of Parylene by wet etching is challenging due to its chemical inertness. Early era removal of Parylene includes burning out of Parylene by high temperature and electrical breakdown by high-voltage arc [143]. The high temperature introduced by the excessive heat could break down the nearby insulation material and damage the active electronics. The high-voltage arc method led to formation of fractures in Parylene near the deinsulation site. In addition, it is difficult to precisely control the deinsulated area, 43 44 which is directly related to the impedance and therefore critical to the performance of the device. Reactive ion etching (RIE) was developed as an alternative dry etching method to remove Parylene. An oxygen plasma is used to selectively remove Parylene C on the tip of Utah Electrode Array (UEA) [21, 158]. Other groups have also used reactive ion etching to etch Parylene [159, 160]. Anisotropic etching is often preferred in MEMS and integrated circuit (IC) fabrication for precise control; isotropic etching is more suitable for substrate cleaning. For deinsulation of biomedical devices like neural interfaces with complex three-dimensional geometry [21], isotropic plasma etching is desired to identically remove the encapsulation film from all directions while maintaining the original geometries. Therefore, inductively couple plasma is preferred over capacitively coupled plasma due to its more isotropic characteristic. Masking is required to define the etching area during the plasma etching. Thin aluminum foil has been adopted as a mask layer for Parylene etching [21]. The major drawback of this masking method is the lack of control in exposure area, which leads to big variation in tip impedance. Alternatively, photoresist has been reported as an etching mask for three-dimensional electrodes [161, 162]. However, the usage of photoresist could affect the surface hydrophobicity [163] and therefore the biocompatibility of Parylene film. In addition, applying photoresist to biomedical devices with complex geometry would be very challenging. In addition to plasma etching, laser ablation has also been demonstrated to effectively remove Parylene [164-166]. The biggest advantage of laser ablation is that no mask is required, which is extremely beneficial for biomedical devices with complex geometries making them difficulty to mask. One of the concerns with laser ablation is the damage to the film/material underneath Parylene [166, 167]. This effect can often be eliminated or minimized by manipulate the power and pulse number of the laser [166]. Another drawback is the carbon redeposition around the ablation spot [165, 168]. A brief oxygen plasma treatment can be used to clean the carbon residual and generate electrode cites that have performance similar to RIE deinsulated cites. 45 Fig 2.1 Microwire arrays made by Williams et al., 22 tungsten microwires were connected to the back-end connector [reprinted with permission from Elsevier, © 1999, Elsevier]. 46 OUTPUT LEADS/ RIBBON CABLE Fig 2.2 Examples of the 2-D Michigan microelectrode array with multiple recording/stimulation sites [reprinted with permission from IEEE © 2008 IEEE]. Fig 2.3 3-D Michigan array. At the top, a 64-site 8-channel Michigan array with CMOS electronics; at the bottom, four probes are assembled onto a platform to form a 256-site 32-channel array [reprinted with permission from IEEE © 2008 IEEE]. 47 Fig 2.4 Fully integrated wireless neural interfaces based on UEA. An ASIC chip was flip-chip bonded at the backside of UEA. Gold coil was wire-bonded and SMD capacitors were soldered for inductive powering and wireless communication. Fig 2.5 Schematic representation of ALD Al2O3 using self-limiting surface chemistry and an A (TMA) +B (oxygen plasma) binary reaction sequence. 48 7W ■C pea f *P0 m m m a a goo m m m ™ Thi*fs) Fig 2.6 QCM measurements for Al2O3 ALD at 58 °C showing the linear growth of the Al2O3 ALD film over many reaction cycles. The average Al2O3 mass gain per ALD cycle is 30 ng/cm2. [Reprinted with permission from Chemistry of Materials M. Groner, F. Fabreguette, J. Elam, and S. George, "Low-temperature Al2O3 atomic layer deposition," Chemistry of Materials, vol. 16, pp. 639-645.© 2004, American Chemical Society]. 49 Fig 2.7 Chemical structures of Parylene-N,-C and -D. [reprinted from J. B. Fortin and T. M. Lu, Chemical vapor deposition polymerization: the growth and properties o f Parylene thin films, Springer, 2004, with permission from Springer]. 50 Fig 2.8 The chemical vapor deposition process for Parylene C. The dimer is first vaporized and then dissociated into monomers with free radicals. The monomers undergo polymerization process when cooling down. Fig 2.9 Schematic view of a Parylene deposition system. The system is consisted of five major components: a vaporizer, a pyrolysis furnace, a deposition chamber, a cold trap, and a vacuum pump. 51 Fig 2.1Q The chemical structure of Silquest A-174 ® silane. The functional group X can form covalent bonds with Parylene monomer, and Y (right) can be hydrolyzed and form covalent bonds with substrates. Table 2.1 Water vapor transmission rate (WVTR) for polymers and atomic layer deposited Al2O3. Material WVTR (gmm/m2 day) Epoxy 0.5-1 Polyimide (DuPont) 1.4 Silicone, RTV 46.8 PTFE (DuPont) 0.1 Parylene C (Specialty Coating Systems) 0.4 Parylene N (Specialty Coating Systems) 5.4 Atomic Layer Deposited Alumina ~10"1Q 52 Table 2.2 Material properties for Parylene N, C and D [reprinted with permission from J. B. Fortin and T. M. Lu, Chemical vapor deposition polymerization, © 2004 Springer]. Property Parylene N Parylene C Parylene D Electrical property Dielectric Constant 1 MHz 2.65 2.95 2.8 1 KHz 2.65 3.10 -- 60 Hz 2.65 3.15 -- Dissipation factor 1 MHz 0.0006 0.013 0.002 1 KHz 0.0002 0.019 -- 60 Hz 0.0002 0.020 -- Dielectric strength (MV/cm) 300 185-220 215 Volume resistivity (23°C, 50% RH) 1.4x1017 8.8x1016 2x1016 Surface resistivity (23°C, 50% RH) 1x1013 1x1014 5x1016 Physical Property Melting point (°C) 420 290 380 Glass transition (°C) 13-80 35-80 110 Linear coefficient of expansion (25°CX10'5, K-1) 6.9 3.5 Heat capacity (25°C, J/gK) 1.3 3.5 -- Thermal conductivity (25°C, kW/mK) 1.3 1.0 Density (g/cm3) 1.110 1.289 1.418 Refractive index 1.661 1.639 1.669 Mechanical Property Tensile modulus (Gpa) 2.4 3.2 2.8 Tensile strength (Mpa) 45 70 75 Yield strength (Mpa) 32 55 60 Elongation to break (%) 30 200 10 Yield elongation (%) 2.5 2.9 -- Static coefficient of friction 0.25 0.29 0.35 Dynamic coefficient of friction 0.25 0.29 0.31 Hardness (Gpa) 0.6 (least) Moderate -- Moisture resistant Water absorption (%) after 24 hrs 0.1 0.1 0.1 Coating performance Crevice penetration Best Good Least Molecular activity Highest Good Least Coating uniformity Best Good -- Thickness control Good Best -- Coating speed Lowest Moderate Highest 53 2.9 References [1] P. 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