Gallium core-shell microstructure for thermal responsive implantable neural electrode

A critical challenge for ensuring a long-term brain-electrode interface is the staggering mismatch between the mechanical properties of the silicon/metal microelectrode arrays and brain tissue. There has been significant research on the development of more flexible implantable microelectrodes than silicon or metal microwire arrays. However, flexible electrodes buckle during insertion and require either a rigid outer shuttle system or a thicker coating to penetrate cortical tissue, creating a larger lesion than is necessary with stiffer implants. The next generation of microelectrode arrays may need to be mechanically rigid during the insertion procedure, then adapt to a more flexible implant with mechanical properties. Gallium could be the ideal material for next-generation mechanically adaptable microelectrode arrays. Gallium has a unique melting point of 29.36 °C. This indicates Gallium is a rigid solid at room temperature and a liquid (no mechanical strength) at body temperature. This offers an ideal mechanical transition (rigid to soft) between 25°C to 37°C for implantable microelectrode arrays. However, there is limited knowledge on the biocompatibility, biostability, and electrochemical performance of Gallium under physiological conditions. This dissertation represents a comprehensive investigation of Gallium chemistry under physiological conditions, improving Gallium material property for neural interface applications, and in vivo evaluation of Gallium-based microelectrodes. iv A thermal responsive Gallium/PEBAX core/shell structure is first fabricated to demonstrate the temperature-dependent mechanical change of the Gallium-based microelectrodes. However, various spectroscopic studies demonstrated Gallium itself is unstable under physiological conditions due to oxidation, biodegradation in aqueous solution, and biofouling by the interaction between ionized Gallium surface and the plethora of biomolecules in physiologic fluids. Various encapsulation strategies were performed to improve the biostability and electrochemical performance of Gallium under physiological conditions. Optimal electrochemical deposition conditions were confirmed for Au, CNT, and PEDOT on Gallium surfaces. Finally, in vivo physiological signals were recorded from soft encapsulation (PEDOT:BF4) and a Gallium-based liquid metal platform. Single-unit action potential recording using the PEDOT functionalized liquid metal electrodes was performed from nonhuman primates and confirmed the long-term recording stability from an invertebrate model. The in vivo results show our strategies can open numerous design opportunities for next-generation Gallium-based bioelectronic devices.

GALLIUM CORE-SHELL MICROSTRUCTURE FOR THERMAL RESPONSIVE IMPLANTABLE NEURAL ELECTRODE by Taehwan Lim 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 Chemical Engineering The University of Utah December 2021 Copyright © Taehwan Lim 2021 All Rights Reserved The University of Utah Graduate School STATEMENT OF DISSERTATION APPROVAL Taehwan Lim The dissertation of has been approved by the following supervisory committee members: Huanan Zhang , Chair 11/29/2021 Swomitra Kumar Mohanty , Member 11/27/2021 Terry Arthur Ring , Member 11/27/2021 Patrick A. Tresco , Member 11/26/2021 Yunshan Wang , Member 11/27/2021 and by the Department/College/School of Eric G. Eddings Date Approved Date Approved Date Approved Date Approved Date Approved , Chair/Dean of Chemical Engineering and by David B. Kieda, Dean of The Graduate School. ABSTRACT A critical challenge for ensuring a long-term brain-electrode interface is the staggering mismatch between the mechanical properties of the silicon/metal microelectrode arrays and brain tissue. There has been significant research on the development of more flexible implantable microelectrodes than silicon or metal microwire arrays. However, flexible electrodes buckle during insertion and require either a rigid outer shuttle system or a thicker coating to penetrate cortical tissue, creating a larger lesion than is necessary with stiffer implants. The next generation of microelectrode arrays may need to be mechanically rigid during the insertion procedure, then adapt to a more flexible implant with mechanical properties. Gallium could be the ideal material for next-generation mechanically adaptable microelectrode arrays. Gallium has a unique melting point of 29.36 °C. This indicates Gallium is a rigid solid at room temperature and a liquid (no mechanical strength) at body temperature. This offers an ideal mechanical transition (rigid to soft) between 25°C to 37°C for implantable microelectrode arrays. However, there is limited knowledge on the biocompatibility, biostability, and electrochemical performance of Gallium under physiological conditions. This dissertation represents a comprehensive investigation of Gallium chemistry under physiological conditions, improving Gallium material property for neural interface applications, and in vivo evaluation of Gallium-based microelectrodes. A thermal responsive Gallium/PEBAX core/shell structure is first fabricated to demonstrate the temperature-dependent mechanical change of the Gallium-based microelectrodes. However, various spectroscopic studies demonstrated Gallium itself is unstable under physiological conditions due to oxidation, biodegradation in aqueous solution, and biofouling by the interaction between ionized Gallium surface and the plethora of biomolecules in physiologic fluids. Various encapsulation strategies were performed to improve the biostability and electrochemical performance of Gallium under physiological conditions. Optimal electrochemical deposition conditions were confirmed for Au, CNT, and PEDOT on Gallium surfaces. Finally, in vivo physiological signals were recorded from soft encapsulation (PEDOT:BF4) and a Gallium-based liquid metal platform. Single-unit action potential recording using the PEDOT functionalized liquid metal electrodes was performed from nonhuman primates and confirmed the long-term recording stability from an invertebrate model. The in vivo results show our strategies can open numerous design opportunities for next-generation Gallium-based bioelectronic devices. iv To my mentors, and my loving family and friends. TABLE OF CONTENTS ABSTRACT...................................................................................................................... iii LIST OF TABLES ........................................................................................................... viii LIST OF FIGURES ........................................................................................................... ix ACKNOWLEDGEMENTS ............................................................................................... xi Chapters 1. INTRODUCTION .......................................................................................................... 1 1.1 Motivation ................................................................................................................. 1 1.2 Research History ....................................................................................................... 2 1.3 Gallium-based Biomedical Devices ......................................................................... 5 1.4 Organization of the Dissertation .............................................................................. 9 1.5 References ............................................................................................................... 12 2. GALLIUM/POLYMER CORE/SHELL STRUCTURE PREPARATION .................. 24 2.1 Introduction............................................................................................................. 24 2.2 Materials and Methods............................................................................................ 26 2.3 Results and Discussions .......................................................................................... 28 2.4 Conclusion .............................................................................................................. 31 2.5 References ............................................................................................................... 32 3. SURFACE CHEMISTRY OF GALLIUM WITH INCUBATION IN PHYSIOLOGIC CONDITION .................................................................................................................... 39 3.1 Introduction............................................................................................................. 39 3.2 Materials and Methods............................................................................................ 42 3.3 Results and Discussions .......................................................................................... 45 3.4 Conclusion .............................................................................................................. 53 3.5 References ............................................................................................................... 55 4. GALLIUM SURFACE MODIFICATION BY ELECTROCHEMICAL DEPOSITION ................................................................................................................... 72 4.1 Introduction............................................................................................................. 72 4.2 Materials and Methods............................................................................................ 75 4.3 Results and Discussions .......................................................................................... 79 4.4 Conclusion .............................................................................................................. 93 4.5 References ............................................................................................................... 96 5. GALLIUM BASED-BIOMEDICAL APPLICATIONS AND THEIR PERFORMANCE IN VIVO ............................................................................................ 120 5.1 Introduction........................................................................................................... 120 5.2 Materials and Methods.......................................................................................... 121 5.3 Results and Discussion ......................................................................................... 125 5.4 Conclusion ............................................................................................................ 130 5.5 References ............................................................................................................. 131 6. CONCLUSIONS AND FUTURE DIRECTIONS...................................................... 141 6.1 Conclusions ........................................................................................................... 141 6.2 Future Directions .................................................................................................. 145 6.3 References ............................................................................................................. 148 APPENDIX ..................................................................................................................... 152 vii LIST OF TABLES Tables 3.1. EDAX data of the Ga surface for 45 days ................................................................. 69 3.2. Summary of the XPS parameters from C 1s core level spectra ................................. 70 3.3. Nyquist plot parameters of Ga with physiologic incubation in DMEM (from Figure 3.7) .................................................................................................................................... 71 4.1. Fitting results from Randles and Corrosion equivalent circuits in Nyquist plot (Figure 4.6) .................................................................................................................................. 119 LIST OF FIGURES Figures 1.1. Conventional neural prosthetic devices ..................................................................... 19 1.2. Advantages of traditional ultrasoft bioelectrodes ...................................................... 20 1.3. Advantages of traditional stimuli-responsive polymer substrate bioelectrodes ........ 21 1.4. Recent liquid metal applications ................................................................................ 22 1.5. Gallium surface oxidation studies .............................................................................. 23 2.1. Preparation of Ga-based liquid metal wire ................................................................ 36 2.2. Mechanical property of PEBAX by stretching ratio. ................................................. 37 2.3. Penetration and thermal responsive property of Ga/PEBAX structure ..................... 38 3.1. Ga ion mass transfer in water and physiologic buffer at 37oC .................................. 62 3.2. Image validation of Ga with incubation..................................................................... 63 3.3. FT-IR spectra of Ga droplet in DMEM with incubation ........................................... 64 3.4. XPS survey scans of Ga surface in DMEM for 5 days .............................................. 65 3.5. XPS survey scans of Ga surface for 45 days ............................................................. 66 3.6. Electrochemical performance degradation in DMEM with incubation ..................... 67 3.7. Nyquist plot change in DMEM for 3 months ............................................................ 68 4.1. Preliminary SEM and electrochemical studies ........................................................ 105 4.2. PDDA wrapping on the CNT surface ...................................................................... 106 4.3. Optimization of the CNT and polymer ratios .......................................................... 107 4.4. CNT/PDDA composite preparation ......................................................................... 108 4.5. Optical confirmation of Ga surface oxidation after CNT/polymer deposition ........ 109 4.6. Nyquist impedance after CNT/polymer electrochemical deposition ....................... 110 4.7. CNT/Au NPs multilayer nanocomposite preparation .............................................. 111 4.8. Electrochemical properties of nanostructure functionalized Ga electrodes ............. 112 4.9. PEDOT:PSS on the Ga surface ................................................................................ 113 4.10. PEDOT:BF4 nanostructure preparation ................................................................. 114 4.11. Physiologic stabilities of PEDOT:BF4 structure on the Ga surface ...................... 115 4.12. Dopamine (DA) monitoring by surface functionalized Ga wires .......................... 116 4.13. Dopamine (DA) monitoring for selectivity under physiologic hindrance ............. 117 4.14. Fast scan CV (FSCV) test results to confirm in vivo test availability ................... 118 5.1. Single-unit action potential recording from nonhuman primates. ........................... 134 5.2. Electrochemical properties of PEDOT LMEs after implantation (red) ................... 135 5.3. Schematic illustrations showing the method to obtain action potentials (APs) ....... 136 5.4. Repeated action potentials (APs) recordings from invertebrate model ................... 137 5.5. Action potential recording stability from PEDOT LMEs ........................................ 138 5.6. Post-implantation surface images ............................................................................ 139 5.7. Validation of the initial recording performance attenuation .................................... 140 6.1. Schematic illustration showing Young`s modulus of bioelectrode materials. ......... 149 6.2. PEDOT functionalized microelectrodes array fabrication step. .............................. 150 6.3. PEDOT functionalized microelectrodes array performance .................................... 151 A.1. Electrochemical analysis setup in DMEM buffer. .................................................. 155 A.2. 120o bending deformation test setup. ...................................................................... 156 A.3. Cannula assisted PEDOT LMEs setup.................................................................... 157 A.4. Action potential recording setup from invertebrate model ..................................... 158 A.5. Eelectrical signals including stimulation artifacts and corresponded potentials..... 159 x ACKNOWLEDGEMENTS The dissertation reflects my research interest in the material designing field and my life in Utah, which was a significant challenge in my life. Although the biomedical application field was unfamiliar to me at first, I have developed a great understanding of biomaterial selection, design optimization, and the fundamentals of complex material characterization in biological systems. I would like to thank all those who supported me throughout. First and foremost, I`m honored to work with my advisor, Dr. Huanan Zhang, who introduced me to the biomedical field, provided me with valuable direction, support in my work, and helped me adapt to different surroundings in my life. Without his assistance, I would have had a hard time achieving my research goal and graduating here. That`s why I like to especially thank Dr. Zhang throughout our time working together. I would also like to thank all of the members of my supervisory committee. I would like to thank Dr. Terry Ring for his deep and professional insight into the material field to complete my fundamental liquid metal study. I`m grateful to Dr. Yunshan Wang for helping me analyze the deeper and more accurate based on her professional spectroscopic knowledge. Dr. Bobby Mohanty was willing to give his time, whether it was to provide me germinate new ideas or to allow my applicable device fabrication. I also owe a huge thank you to Dr. Patrick Tresco, an expert in the biomedical field and provided me with valuable insights for my research and life. I would like to thank my labmates Frank Curry, Tasmia Tasnim, and Alexandra Boyadzhiev for their warm friendship and help. They enabled me to work easier and settle in our research group well. They also provided me with many and new inspirations to living here as my first international friends. I hope they all achieve their success in life. Lastly, I would like to thank my parents and brother for their love and support. Especially during a challenging season, they always helped me keep calm and carry on my research without dropping out from this hard and solitary work. xii CHAPTER 1 INTRODUCTION 1.1 Motivation Bioelectronic devices have the potential to revolutionize the treatment and diagnosis of many diseases. The devices enable continuous monitoring of biochemical signals in the human body for individual health management purposes, including personalized medication, surgical preparation, postsurgical care, etc.1–3 Already, we have seen the real-time use of biological signals for the operation of several human-machine interfaces,4–7 and the stimulations of excitable tissues.8–10 To sustain chronic biocompatibility, research from many different groups indicates that such bioelectronic devices require materials with tissue-like mechanical properties, high electrical properties, and long-term biostability.11–18 One of the major challenges to monitoring biochemicals is preparing soft and flexible bioelectronics that can minimize the mechanical mismatch between the electronic devices (typically over 10 GPa) and soft biological tissue environments (Young`s modulus = 0.1 kPa to 1 MPa).19–21 This mechanical mismatch can cause tissue damage and chronic inflammation/scar tissue formation.22,23 The scar tissue can isolate electronic devices and surrounding neural tissue and limit the functionality of the devices. Consequently, current researches have been implemented to the development of 2 novel soft electronic materials for bioelectronic interfacing with human tissues. The mechanical softness of the devices can provide critical remedies to cellular degeneration, adverse biological inflammatory responses to implants, and poor chronic material stability that degrades the performance of the interface over time (Figure 1.1 A). Here, material-based approaches for the synthesis and fabrication of soft and electrically conductive materials have been reviewed, focusing on the integration of biostable and biocompatible properties. The softness and bioelectronics` performance can be evaluated by histologic change after implantation in the animals` brain model and by physiologic signal recording such as local field potential (LFP) and possible single-unit action potential deconvolution, respectively. 1.2 Research History Advances in lithography techniques enabled the production of metal wire or silicon-based electrode arrays, allowing for the fabrication of smaller sizes, more recording sites, and rigid structures to prevent buckling.24–26 Michigan electrodes were fabricated consisting of a shank with silicon substrate equipped several electrode sites (Figure 1.1 B).27–29 Utah electrode array was fabricated in the University of Utah with sharpened silicon needles with pitches of 200-400 µm (Figure 1.1 C).30,31 Both the Michigan and Utah probes are insulated up to their tips with biocompatible polymers such as Parylene-C and polyimide, and their bare tips were coated with conductive metals, including Pt or Ir. Both electrodes are still widely used in neuroscience and biomedical research today. However, these silicon and metal electrode arrays (Young`s modulus of GPa 3 range) are significantly more rigid than the host brain tissue (Young`s modulus of kPa range), which can create significant damages to the brain tissues. The damages to the brain tissue generated scar tissues (astrocytes and microglia) around the conventional stiff implants, such as. (Figure 1.1 D). The scar formation can isolate the bioelectrodes from the targeted neurons and block communication, eventually leading to device failure. Hence, extensive research had been performed to reduce scar tissue formation in the past decade. First, ultra-small electrodes were prepared by carbon fiber, which served as the conductive core and provided the mechanical platform of the electrode under 10 µm diameter.32,33 Nanoscale poly(ethylene glycol methacrylate) (PEGMA) coated surround the carbon fiber as an insulator and poly(ethylene 3,4-dioxythiophene):poly(styrene sulfonate) (PEDOT:PSS) deposited on the tip to reduce impedance, resulting in upgraded physiologic signal recording performance.32 Although 12-week implantation and singleneuron recording tests in the motor cortex of rats demonstrated less neuroinflammation response than conventional bulk silicon electrodes. Ultrasoft electrode materials have been developed by using conductive hydrogels or mesh-type polymers platforms.11,13,34–37 Since the soft materials cannot penetrate the extracellular part of the brain themselves, the materials received assistance from the syringe needle (cannula) stiff shuttle to be injected into the brain. PEDOT hydrogel crosslinked by poly(ethylene glycol) achieved intrinsic softness (Young`s modulus of 1 MPa), the conductive gel reduced therefore inflammatory response and scar tissue formation around the hydrogel (Figure 1.2 A and B).11 However, low electrical properties compared to typical metals still challenge obtaining high fidelity signals. Flexible free-standing 4 mesh bioelectrodes from the SU-8 backbone can be injected by syringe needle in the brain (Figure 1.2 C).34,36 The robust mesh structure enabled flexible and small-footprint devices that can improve foreign body responses after implantation. These mesh bioelectrodes were enabled to record single-unit action potentials through low- and highpass filters (Figure 1.2 D). Stimuli-responsive polymer substrates have received attention as the mechanical transition that the substrates sufficiently stiff to facilitate implantation into the brain, but then soften reducing mechanical mismatch with the brain tissue. Taking a biomimetic approach from sea cucumber dermis, stimuli-responsive and mechanically adaptive polymer composites were prepared based on a poly(vinyl acetate) (PVAc)/cellulose nanocrystals (CNCs) (Figure 1.3 A and B).38 The composite showed Young`s modulus of 5.1 GPa when dried; however, Young`s modulus drastically dropped to 10 MPa after absorbing biofluids, which mimics the exposure to physiologic conditions. This polymer composite has been utilized to fabricate bioelectrodes, consisting of metal (Au, Pt, Ti) electrodes and Parylene insulators on the mechanically adaptive substrates (Figure 1.3 C).39,40 Yang et al. reported the cellulose-based substrate equipped Young`s modulus of 120 kPa with the wet condition and then designed gold layers for conductive electrodes.41 The composite electrode displayed good biocompatibility and the ability to obtain neural signals, including single-unit recordings on the brain surface of a rodent model (Figure 1.3 D). Although the in vivo histological evaluations of mechanically adaptive bioelectronics demonstrated that compliant implants exhibited less brain tissue damages than conventional rigid ones, the biostability between the adaptable substrate and conductive electrode can limit chronic use for these electrodes. 5 1.3 Gallium-based Biomedical Devices The ideal material needs to have a low electrochemical impedance (under 106 ohm @ 1 kHz) and an intrinsic soft property similar to biological tissues (Young`s modulus of kPa scale). Ga is regarded as an ideal conductive material because of its unique melting point of 29.36oC, which is greater than room temperature (23oC) but lower than human body temperature (37oC). The property indicates that Ga is a rigid solid at room temperature and a liquid (no mechanical property) at body temperature. Liquid metal from Gallium (Ga) and its alloys have emerged as a promising material for soft bioelectronics due to their excellent electrical conductivity and no stiffness at body temperature.42,43 This offers an ideal sharp mechanical property transition from rigid to liquid-like soft with surgery conditions. Also, Ga is regarded as an alternative mercury material due to its low toxicity and negligible vapor pressure. Kim et al. demonstrated the cytotoxicity of Ga from Ga-based liquid metal in an aqueous environment (Figure 1.4 A).44 The study confirmed Ga ions are released by ionization reaction from Ga oxide monohydroxides (GaOOH). The ionized Ga ions of 101 µM did not affect the cytotoxicity of HeLa cells, ADSCs, and NDFs, which is immortal human cancer cell line, stem/progenitor cells, and human skin tissue cells, respectively. However, live/dead staining images displayed less than 30% proliferation of the cells under over 5x101 µM Ga ions existence, which concentrations were generated by sonication or mechanical agitation of Ga droplets. Another study from Guo et al. supported the biocompatibility of Ga in a saline solution that Ga ions up to 2x101 µM had no severe impact on Mouse 3R3 fibroblasts.45 Although both studies suggested Ga has biocompatible benefit in natural conditions, the dynamic environment 6 of the physiologic condition can lead to more release of Ga ions. It might ultimately cause poor viability of cell cultures, restricting the Ga use for biomedical applications. Due to their softness, high electrical conductivity, and biocompatibility, many studies recently have developed methods to fabricate stretchable, soft, and flexible electronics utilizing liquid metals with the combination of polymer substrates, textiles, hydrogel, and rubber.46–49 Ma et al. reported stretchable electronic devices using Gabased liquid metal for wearable and on-skin electronics.46 The liquid metal was deposited on the poly(styrene-block-butadiene-block-styrene) (SBS) film, which showed a Young`s of 0.8 MPa similar to that of human skin. The resistance of the structure increased only 4% when stretched up to 1,800% that is the maximum stretchability to date (Figure 1.4 B and C). The stretchable liquid metal structure was used in the electrocardiogram (ECG) and sweat sensor application in this study. Since the Ga-based liquid metal has flow and viscous behavior, the linear patterned liquid metal shows a self-healing advantage to implement resilient circuit lines. Veeraphandian et al. confirmed that the liquid metal pattern had high conductivity of 25,000 S/cm and difficult-to-cut (they called “die-hard”) electrical connection, which property was similar to the self-healing mechanism of the disconnected line (Figure 1.4 D).47 The study demonstrated liquid metal printed line allowed to use of stretchable inductive strain sensors. Self-healing electronics have significant attraction; thus, the combination between hydrogels and liquid metals also provides wide scope to achieve practical advances due to their superior elasticity and viscosity.49 The hydrogel/liquid metal structure can maintain high conductivity when stretched as well as compressed. The slight change of resistance when stretched and compressed enabled to use of 7 electromyographic (EMG) monitoring without positional restriction. All in all, the Gabased liquid metals have great potential to substitute typical wearable and on-skin sensor materials due to their stretchability, high conductivity, and self-healing benefit. Unlike the applications for wearable and on-skin use, further studies are needed to understand Ga-based liquid metals for implantable bioelectronics. Despite the low mechanical property and reasonable biocompatibility, the Ga-based liquid metals are prone to oxidation at a ppm level oxygen concentration that deteriorates the electrical properties of liquid metals.50,51 The naturally grown Ga oxide layer consisted of Ga2O3 and Ga2O, leading to a solidified thin layer. Although the oxide layer is less conductive than pure Ga (crystalline Ga oxide is generally used as a wide bandgap semiconductor material) and instantaneous growth in air, the oxidation is negligible due to the selflimitation of the growth approximately 10 nm scale in the air (Figure 1.5 A).52–54 The high adhesion of the solid layer can provide unique flow characteristics in nano- and micro-channel, which enabled the 1D transformation of Ga-based liquid metal.55,56 Also, bandgap change from various Ga surface modifications such as UV treatment, O3 etching, and anodic solution etching is useful to support electrochemical study and semiconductor applications.53 However, further oxidation in an aqueous condition would degrade the electrical and electrochemical properties of Ga for implantable electronic applications. Previous literature provides the further oxidation of Ga in aqueous solution to demonstrate surface chemistry change of Ga-based liquid metals. Aqueous condition introduced another Ga oxide member, GaOOH, produced by the hydrolytic reaction between unstable Ga oxides and ionized water ions.57 There are still arguments to verify the exact mechanism of Ga surface change in water; many studies suggested similar 8 pathways from Ga oxides to GaOOH below57–61 𝐺𝐺𝐺𝐺(𝑂𝑂𝑂𝑂)4 − → 𝐺𝐺𝐺𝐺𝐺𝐺(𝑂𝑂𝑂𝑂) + 𝑂𝑂𝑂𝑂 − + 𝐻𝐻2 𝑂𝑂 (1) 𝐺𝐺𝐺𝐺2 𝑂𝑂 + 4(𝑂𝑂𝑂𝑂)− → 2𝐺𝐺𝐺𝐺𝐺𝐺(𝑂𝑂𝑂𝑂) + 𝐻𝐻2 𝑂𝑂 (2) 𝐺𝐺𝐺𝐺(𝑂𝑂𝑂𝑂)3 → 𝐺𝐺𝐺𝐺𝐺𝐺(𝑂𝑂𝑂𝑂) + 𝐻𝐻2 𝑂𝑂 (4) 𝐺𝐺𝐺𝐺2 𝑂𝑂3 + 𝐻𝐻2 𝑂𝑂 → 2𝐺𝐺𝐺𝐺𝐺𝐺(𝑂𝑂𝑂𝑂) (3) All the postulations suggested the GaOOH formation in water and aqueous solutions and determined further oxidized Ga surface by SEM images (Figure 1.5 B).57,58 External driving forces such as temperature, pressure, and electrical potential generated rapid and further GaOOH crystallite formation on the Ga surface; hence, we conclude the limitation of oxide layer growth varies considerably by incubation condition. Chemical and electrochemical methods can remove this oxide layer on the Ga surface. The simple method is HCl vapor treatment introduced by Kim et al., which is useful to recover nonwetting characteristics of Ga surface.50 The simple method was effective in removing Ga oxides derived from the air (Ga2O3 and Ga2O), and the effect was confirmed by X-ray photoelectron spectroscopy (XPS) and low-energy ion scattering spectroscopy (LEIS). Another method is using the reduction process using the electrochemical technique. Previous literature demonstrates that negative potentials under the lower limit of water hydrolysis window (-1.0 V) were effective in reducing oxidic characteristics of Ga surface.62–64 Morales et al. determined that acidic condition and water-solution salts helped perfectly remove the oxide layer on the Ga surface with a negative potential of -1.2 V (Figure 1.5 C and D).63 The study provided the oxide removal mechanism below 2𝐻𝐻2 𝑂𝑂 (𝑙𝑙) → 𝑂𝑂2 (𝑔𝑔) + 4𝐻𝐻 + (𝑎𝑎𝑎𝑎) + 4𝑒𝑒 − (𝑒𝑒𝑒𝑒𝑒𝑒𝑒𝑒𝑒𝑒𝑒𝑒𝑒𝑒𝑒𝑒𝑒𝑒𝑒𝑒𝑒𝑒) (5) 9 𝐺𝐺𝐺𝐺3+ + 3𝑒𝑒 − → 𝐺𝐺𝐺𝐺 (Ga working) (6) Other studies also suggested at least -1.3 V of negative potential was required to obtain oxide-free Ga droplet structure by the contact angle, interfacial tension, and capacitance 64. Despite the presence of the oxide layer removal mechanism, the method has not been tried in human tissue that might cause tissue degeneration due to strong acid and high electrical currents. Overall, previous studies have revealed many aspects of Ga material chemistry; there are still knowledge gaps for understanding the interaction between Ga and the physiologic environment. First, the physiologic environment is an aqueous mixture of cells, biomolecules, soluble gases, and ions, etc. It is crucial to understand the interaction between Ga and biological molecules. Second, previous Ga ion measurements in cell culture have indicated the release of Ga ions into an aqueous environment, the mechanism of this chemical process is however not yet clearly understood. Also, previous literature recognized that the Ga oxidation is self-limited in air and forms a nano-scale oxide layer on the Ga surface. However, the self-limitation mechanism in the physiologic condition is not yet investigated. Lastly, the long-term chemical dynamics between Ga and physiologic buffer have not yet been reported in the literature. To address the shortcomings of Ga studies, we have conducted a thorough investigation of Ga surface chemistry in physiologic buffer. 1.4 Organization of the Dissertation This body of work has been divided into four chapters. Chapter 2 presents the development of a fabrication method for Gallium-polymer core/shell structure as 10 implantable electrodes. Ga-based liquid metals were encapsulated from the thermal drawing method by a biocompatible and soft rubber, polyether block amide (PEBAX). Tensile tests confirm mechanical property and ultimate Young`s modulus of the Ga/PEBAX core/shell structures (Ga wires) with the thickness and the ratio of core and shell thickness. The recording sites of the core-shell structure are exposed to a physiologic environment to monitor neural signals. It is crucial to understand the chemical changes of Gallium in physiological conditions. Chapter 3 presents the surface chemistry change of Ga in physiologic conditions with incubation time. Various chemical and material characterization techniques were used to understand the surface oxidation and corrosion mechanism of Gallium in the physiological buffer. Then impedance measurement and cyclic voltammetry (CV) scan were accompanied to confirm electrochemical property change with surface oxidation and corrosion. The spectroscopic and electrochemical studies in Chapter 3 demonstrated that bare Ga is unsuitable for bioelectronics due to surface oxidation and contamination. Hence, in Chapter 4, surface modification methods were investigated to improve biostability and biocompatibility, as well as the existing electrical property and softness of the Ga-based bioelectronics. We introduce electrochemical deposition of various functional materials to further improve the advantage of Ga. Gold, single-wall carbon nanotubes, and conductive polymer (polyethylene 3,4-dioxythiophene, PEDOT) were selected to improve the electrochemical performance of the bare Ga surface in this study. And each deposition method was optimized for biosensors or implantable microelectrodes. Electrochemical impedance spectroscopy (EIS) and CV were utilized to 11 evaluate the modified electrochemical properties, thereby improving the performance of bioelectronics. In Chapter 5, in vivo performance was characterized by Ga-based bioelectronics. Single-unit action potentials (APs) recording was performed from nonhuman primates and invertebrate models to analyze neural activities and action potentials, respectively. PEDOT:BF4 deposited Ga wires were prepared for the implantation work and the recording performance was confirmed by acute (from rhesus monkey) and repeated recording (from L. terrestris) for 4 weeks. Chapter 6 provides the future direction of this research. It includes the experiment steps for electrode array fabrication and the plan for evaluating chronic histological change in the brain in vivo. 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Natl. Acad. Sci. 2014, 111 (39), 14047–14051. https://doi.org/10.1073/pnas.1412227111. 19 Figure 1.1. Conventional neural prosthetic devices. A) High-magnification confocal microscope images to confirm scar tissue formation by 200x130x60 µm dimension silicon shaft implantation after (a) 2, (b) 4, (c) 6, and (d) 12 weeks.23 B) Michigan probes by (a) optical and (b) SEM measurements.29 C) Utah microelectrodes array.30 D) Immunohistochemical analysis of astrocytic scar (GRAP, green), Microglia/Microphages (red), and IgG (yellow) for 16 weeks. The images are reproduced and modified from the reference.22 20 Figure 1.2. Advantages of traditional ultrasoft bioelectrodes. Histology analysis to compare PEDOT/PEG soft wire and standard stiff tungsten microelectrode; A) NF-200, Iba-1, GFAP stain for axons, microglia and astrocytes, respectively, B) chronic BBB leakage and potential neural regeneration around implants.11 C) Syringe-injectable meshtype electronics; (a) mesh electronics in glass needle, (b) injection to rat`s brain, (c) the mesh electronics injected and folded in 1xPBS solution as a preliminary study, and (d) Bright-field microscopy image of a coronal slice of the HIP region five weeks postinjection of the mesh electronics. D) Acute in vivo 16-channel recording using mesh electronics injected into a mouse brain and superimposed single-unit neural recordings from on channel after 300-6,000 Hz band-pass filtering. C,D reproduced by 34. 21 Figure 1.3. Advantages of traditional stimuli-responsive polymer substrate bioelectrodes. A) Natural model (sea cucumber) and bioinspired design of chemomechanical nanocomposites, cellulose whiskers, the EO-EPI, and PVAc matrix polymers used. B) (a) Schematic illustration of the architecture and switching mechanism in the artificial nano composites with dynamic mechanical properties and (b) its stimuli (water and temperature) responsive mechanical property measurement. A, B reproduced by 38. C) (a) Micromachined PVAc-TW cortical probe with a lithographically defined Ti/Au electrode and (b) stress-strain plot in dry and wet samples.39 D) Images of bacterial cellulose-gold electrode arrays (left) and physiologic signal (by epileptiform activity) recording from the electrode. A representative single spike can be seen in right.41 22 Figure 1.4. Recent liquid metal applications. A) (a) Released Ga ions in water with sonication. (b) Cytotoxicity tests (left: HeLa, middle: ADSCs, right: NDFs) using Ga ion releases with different concentrations.44 B) Resistance change of the liquid metal/SBS as a function of tensile strain. C) ECG signals from liquid metal/SBS. B,C reproduced by 46. D) Resistance change of the liquid metal microparticles-PEVA composite line while simultaneously repeating the stretch (strain: 100%, liquid metal microparticles from 10 to 20 µm) and repeated blunt cuts with a large tweezer (inset: SEM image after the test, scale bar: 250 µm).47 23 Figure 1.5. Gallium surface oxidation studies. A) (a) Oxide layer creation on the Galliumbased bare liquid metal surface,52 (b) self-limited oxide layer thickness estimation in air.54 B) SEM images of Gallium microparticles in water; (a) initial and (b) after 24 h.57 C) Gallium oxide layer removal steps by electrochemical treatment.63 D) Gallium oxide layer formation and removal by electrochemical potential change.64 CHAPTER 2 GALLIUM/POLYMER CORE/SHELL STRUCTURE PREPARATION 2.1 Introduction Implantable microelectrodes are essential tools for recording and stimulating excitable tissues that have attracted great interest due to their potential to revolutionize the treatments of many diseases such as motor dysfunction, congestive heart failure, epilepsy, depression, Parkinson`s disease, and heart arrhythmias.1–6 Although mechanical rigidity is essential to assist the device implantation process, high stiffness often causes tissue damage and ultimately induces a chronic inflammatory response, which generates gliosis around the implanted device.7,8 Therefore, the ideal implantable electrode needs to be mechanically rigid during the tissue insertion procedure, then adapt to flexible and soft mechanical properties in the physiologic condition to comply with fragile brain tissue. Recently, stimuli-responsive materials, such as cellulose whisker and polymer hydrogel, are researched to generate sharp mechanical property transition with water-swollen.9–12 However, these mechanical transitions are confined to the substrate, not for the conductive electrode material such as Au or Ti. The mechanical discrepancy after softening the substrate solely leads to delamination of stiff electrode material from the substrate, resulting in implantable microelectrode failure. Here, Gallium (Ga) is introduced as an electrode material with a unique melting 25 point of 29.36 oC. This indicates Ga-based implantable electrode will be a rigid solid at room temperature and a liquid (no mechanical stiffness) at body temperature (37oC). This provides an ideal sharp mechanical transition, rigid to liquid-like soft, between the room and body temperature for satisfying both surgical steps and patients. Although Ga is an ideal material to use as an implantable microelectrode, Ga should be encapsulated and designed by another outermost layer to use as an electrode practically. Without encapsulation, the liquefied Ga would dissipate in the brain tissue, resulting in electrode malfunction. This study developed a novel manufacturing process to design a Ga/polymer core/shell structure. The shell polymer can be used not only for the container to prevent leaking when Ga changed to a liquid state but also for the insulator to monitor physiological signals without interference. Polyether block amide (PEBAX) was selected as a shell material due to its flexibility and biocompatibility. The PEBAX is a thermoplastic; thus, the material can be stretched by the melting process to create micronscale structures, which is the general dimension of implantable microelectrodes for recording neural signals.13–16 In this chapter, we have fabricated different core-shell structures by this thermal stretching process. By following the tensile stretching test, Young`s modulus was calculated to confirm the mechanical strength of the microstructure. Different Ga/PEBAX core/shell structures were prepared to compare Young`s modulus with the thickness and the ratio between core and shell. The heating cycle of differential scanning calorimetry (DSC) scan is useful to confirm the melting point of the polymer material and the enthalpy of the melting. The mechanical property of the PEBAX depends on the hard 26 segment of the polymer chain, which can be determined by the enthalpy of the melting. Here we include only a brief mechanical property analysis to prepare an optimal soft Gabased platform for future work. 2.2 Materials and Methods 2.2.1 Materials Gallium (99.995% purity) and its alloy (EGaIn, ≥99.99%, trace metals basis) were purchased from Sigma-Aldrich. Different polyether block amide (PEBAX) tubes (0.0340.085 inches of inner diameter, each tubing has 0.005 inches of wall thickness) were obtained from Zeus Industrial Products, SC, USA. The tubing showed oxygen and CO2 permeabilities of 4 and 32 Barrer (3.35x10-16 𝑚𝑚𝑚𝑚𝑚𝑚 ∙ 𝑚𝑚⁄𝑚𝑚2 ∙ 𝑠𝑠 ∙ 𝑃𝑃𝑃𝑃) respectively. These values are 10-500 times lower than other typical polymers such as polyethylene (PE), polypropylene (PP), and polyethylene terephthalate (PET). In addition, the permeabilities would be linearly decreased after being stretched due to the structural deformation.17 2.2.2 Methods 2.2.2.1 Gallium/Polymer core/shell structure preparation Liquefied Ga was injected into the PEBAX tubing, then stretched when melted by heating (200oC), which is above 30-40oC of the melting point of PEBAX. The stretching % determines the thickness of the core/shell structure. And the initial inner diameter/wall thickness of the tubing dominates the ratio between the final core/shell ratio of the Ga wires. The stretching process is similar to the melt spinning process to fabricate typical multi filaments.18 27 2.2.2.2 Characterization of mechanical property The mechanical force with stretching was measured using a tensile tester (MARK-10, NY, USA). The stress and strain were calculated applying cross-section area and initial length of the Ga wires, respectively. To figure out Young`s modulus of the Ga wires, we set the cross-section as a combination of Ga core and PEBAX shell. The stretching was conducted at the rate of 5 mm/min. 2.2.2.3 Characterization of thermal property The crystallinity and melting point of stretched and pristine PEBAX tubes were analyzed by DSC (Q20, TA Instruments, USA). The heating and cooling scan rate are 10 and -20 oC/min, respectively. The enthalpy at the melting point of the hard segment of PEBAX helps to understand the softness of the PEBAX with stretched %. The enthalpy of melting was obtained by TA official software, integrating the melting peak. 2.2.2.4 Characterization of cross-section The cross-section of the Ga wires was characterized by scanning electron microscopy (SEM, Hitachi S-4800, Japan) at 20 kV accelerating voltage. Backscattering electrons (BSE) imaging was utilized to divide the Ga and PEBAX portion clear. 2.2.2.5 Penetration testing In vitro and ex vivo models were used to determine a minimal thickness of the Ga wires. For in vitro model, 0.6% agarose gel was prepared to mimic the hardness of the rat`s brain. High gelling temperature agarose was purchased from Sigma-Aldrich, and the 28 agarose was dissolved in DI water with heating and stirring, ensuring uniform hardness. Rat brain was harnessed for ex vivo model to support in vitro agarose gel test result. 2.3 Results and Discussions In this study, our objective is to prepare a Ga/soft polymer core/shell structure that can transit from the rigid to soft mechanical properties for the implantable electrodes. Typical implantable microelectrodes have consisted of conductive core and biocompatible insulators to minimize noise recording.13–16,18–22 Chemical vapor deposition (CVD) of parylene on metal structures is often used to create conductorinsulator core-shell structures However, the CVD method is unsuitable for fabricating Ga/polymer core/shell structure. First, it is difficult to fabricate microscale Ga wires due to the low melting point and brittle nature of solid Ga. And, the CVD process requires temperatures well above the melting point of Ga. Therefore, we utilized the thermal drawing process to fabricate the microscale Ga/polymer core/shell structure. The hollow polymer shell can protect Ga leaking when liquefied more effectively and provide perfect insulating property. PEBAX, a thermoplastic elastomer, was selected in this study due to its biocompatibility and softness (Young`s modulus of 20 MPa). The PEBAX consists of a soft polyether backbone block and robust polyamide, and a low melting point (160oC) enabled the melt drawing process easily. Since previous literature verified the biocompatibility and low water absorption of the PEBAX, the material is allowed to use as a polymer shell to encapsulate liquid Ga and contact brain tissue without inflammatory 29 damage.23–25 To fabricate PEBAX encapsulated Gallium core/shell structure, liquefied Ga has injected millimeter-scale PEBAX tubing first. The tubing was heated at 200oC, which is above 40oC of the melting point of PEBAX. The heated PEBAX tubing including liquefied Ga was thermally stretched 500% when the PEBAX tubing was melted (Figure 2.1). The stretched structure was solidified under liquid N2 to obtain a clear cross-section by cutting process. The solid structure (Figure 2.1) can penetrate the brain tissue directly and the exposed conductive Ga surface can communicate with neurons in the brain. The melting and stretching process led to Young`s modulus degradation of the PEBAX here from 20 MPa (non-stretched) to 1 MPa (500% stretched) due to the breakdown of the hard segments that might be seen in Figure 2.2 A (Chemical structure can be seen in Figure 2.2 B). The thermal drawing process typically leads to higher Young`s modulus because the amorphous part of the polymer chain can be aligned to form orientation-inducing crystallization.17,26,27 However, the results in Figure 2.1A indicted that Young’s modulus decreased over the stretching process. These results could be potentially due to the lamella structure of the elastomer is hard to yield a wellorientated structure without an elaborate-designed heat setting process.28,29 And slow quenching is also required to obtain higher crystallinity of the polymer physical structure.30,31 In our study, we did not use the slow quenching process after stretching. Lastly, we used uniaxial heat flow by a heat gun to melt biphasic Ga/PEBAX core/shell structure that generates thermal imbalance of PEBAX shell during melting and stretching process, resulting in partial break of the polymer chain segments. Crystallization of the PEBAX with different stretching ratios can be evaluated by 30 DSC measurements (Figure 2.2 C). The enthalpy of melting around 160-170oC induced from the hard polyamide region supports estimating a mechanical property of deformed PEBAX.32–34 The DSC scan and enthalpy of melting (Figure 2.2 D) demonstrated Young`s modulus attenuation with the stretching over 300%. The lack of uniform heating enabled the reduction of Young`s modulus of the PEBAX tubing after elongation (500%, ~1 MPa) with a 50-60 µm scale. To evaluate the tissue penetration ability of the core-shell structure, we conducted penetration tests on agarose gel and ex vivo bat brain. We observed the minimum thickness for brain penetration by in vitro agarose gel (0.6 wt%) (Figure 2.3 A) and ex vivo rat`s brain model is approximately 60 µm in diameter when Ga is solidified (Figure 2.3 B). Although it is easier to penetrate the brain tissue with larger diameter wires, we should minimize Ga wires thickness to reduce brain tissue damage and obtain a high signal-to-noise ratio during physiological signal recording. Ga/PEBAX wires with different core/shell ratios were prepared to evaluate the effect of core/shell ratios on Young`s modulus (Figure 2.3 C). While the ratio of 1:1 exhibited Young`s modulus change of three orders (0.6 GPa to 0.7 MPa), the higher ratio (4:1) (cross-section image can be seen in Figure 2.3 D) produced a mechanical switch of four orders (4 GPa to 0.3 MPa). The larger Gallium core allowed the extremely low Young`s modulus (0.3 MPa) at the liquid state, which is similar or less than hydrogel. The liquid-like softness of the Ga/PEBAX structure enabled it to comply with external shape (Figure 2.3 E). 31 2.4 Conclusion In summary, we prepared the Ga/PEBAX core/shell structure to develop the thermal responsive platform. The core/shell structure exhibits biocompatibility and softness by shell material selection. The mechanical property of PEBAX can be attenuated by stretching deformation due to the lack of crystallization that demonstrated DSC scans. The deformation of 500% stretching reduced Young`s modulus of PEBAX to ~1 MPa from 20 MPa. The dominant portion of the liquid core even allowed the lower mechanical property to Young`s modulus of 0.3 MPa that is similar to or less than hydrogel. Liquefied Ga/PEBAX core-shell structure opens up numerous opportunities to obtain brain tissue-like electrodes. (Young`s modulus of 1-10 kPa). 32 2.5 References (1) Velliste, M.; Perel, S.; Spalding, M. C.; Whitford, A. S.; Schwartz, A. B. 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Acta 2002, 391 (1–2), 271–277. https://doi.org/10.1016/S0040-6031(02)00189-2. 36 Figure 2.1. Preparation of Ga-based liquid metal wire and its core/shell structured electrode by heating, stretching, and cutting processes. 37 Figure 2.2. Mechanical property of PEBAX by stretching ratio. A) Stress-strain curve of PEBAX stretching. B) Chemical structure of PEBAX we used. C) 1st cooling curves of stretched PEBAX DSC scan for crystallization behavior. D) 2nd heating curves of stretched PEBAX DSC scan for melting behavior. Enthalpy of melting is shown in the table. 38 Figure 2.3. Penetration and thermal responsive property of Ga/PEBAX structure. Penetration properties through A) agarose 0.6 wt% gel and B) ex vivo rat brain. C) Stressstrain curve of Ga/PEBAX core/shell structure with the ratio. D) Cross-section image of Ga/PEBAX structure (4:1 with 60 µm thickness). E) Image showing mechanical compliance of the Ga/PEBAX wire. CHAPTER 3 SURFACE CHEMISTRY OF GALLIUM WITH INCUBATION IN PHYSIOLOGIC CONDITION 3.1 Introduction Gallium (Ga) and its alloys, called liquid metals, are recognized as promising materials for a range of devices from stretchable electronics to transformable machines. Gallium has several interesting materials such as low melting point, unique surface chemistry, negligible vapor pressure, and excellent electrical conductor.1,2 the fluidic and viscous properties of Ga-based materials provide intrinsic active stretchability without plastic deformation and self-healing after cutting, which is attracted interest to implement resilient electrical circuits.3–5 These material properties can offer the possibility to fabricate tissue-compliant bioelectronic devices attracting considerable attention in developing implantable electronic devices.6–11 Bioelectronic devices often require soft, stretchable, and reconfigurable material properties to interact with the soft tissues in the biological system and curvilinear skin in the dynamic system. Although there are significant interests in Ga-based bioelectronic devices, our current knowledge on the interaction of Ga and the physiological system is limited. Ga is a non-noble and transient metal. A thin oxide layer can be formed when Ga is exposed to an atmosphere above 1 ppm of oxygen concentration.12 Since oxygen 40 solubility in water is approximately 7 ppm at the body temperature, oxide layer formation on the Ga surface is an inevitable phenomenon in physiologic conditions.13 Several studies have investigated the oxidation behavior of Ga. Studies revealed the mechanical and viscous properties of the Ga oxide layer on the Ga surface have solid mechanical characteristics. The thickness of the oxide layer is critical to control the mechanical properties of Ga and provide the rheological behavior of Ga.14–16 Also, it is known that the oxide layer growth on the Ga surface is in good agreement with logarithmic kinetic limiting with time, which is a common transport phenomenon.17,18 In addition to Ga oxidation, studies have demonstrated Ga has minimal cytotoxic effect with various cell cultures, such as HeLa, stem cells, and fibroblasts, under the mM concentration of Ga ions.19–21 Other researchers have studied the electrical properties of Ga alloys to apply soft and stretchable electronics through the thickness control of the oxide layer on EGaIn and GaInSn droplets.22,23 Furthermore, surface treatments of Ga alloys were conducted to improve their reactivity with redox electrochemical and UV/ozone treatments.14,24 Other studies have dealt with the oxidation and reduction behavior of Ga to characterize dielectric functions.25,26 Although previous studies have elucidated many aspects of Ga material chemistry, there are still knowledge gaps for understanding the interaction between Ga and the physiological environment. First, the physiologic environment is an aqueous mixture of cells, biomolecules, soluble gases, and ions, etc. It is crucial to understand the interaction between Ga and biological molecules. Second, previous Ga ion measurements in cell culture have indicated the release of Ga ion into an aqueous environment.21 However, the mechanism of this chemical process is not yet understood. Next, previous 41 literature recognized that the Ga oxidation is self-limited in air and forms a nano-scale oxide layer on the Ga surface.12,14 However, the self-limitation mechanism in the physiologic condition is not yet investigated. Lastly, the long-term chemical dynamics between Ga and physiologic buffer have not yet been reported in the literature. This study demonstrates a comprehensive approach by combining different spectroscopic analyses to understand the interfacial chemical changes between the Ga surface and the physiologic buffer over 45 days. The study was conducted in vitro by Dulbecco`s modified eagle medium (DMEM), a common cell culture medium consisting of various amino acids, carbohydrates, and ions. Ga was incubated in DMEM buffer at body temperature (37oC) over 45 days. We first measured the pH and the Ga ion in the buffer solution over time to understand the kinetic of the Ga ion release process, then observed the surface change by images. To understand the mechanism of Ga ion release, Fourier-transform infrared spectroscopy (FT-IR) and X-ray photoelectron spectroscopy (XPS) were performed that help to investigate Ga surface composition and materials deposited on the Ga surface in the DMEM. The elemental composition was measured over time to monitor the time-dependent chemical changes. Lastly, we confirm electrochemical performance with incubation that provides critical information to monitor physiologic signals from muscle, cardiogram, and neural activity. Electrochemical impedance spectroscopy (EIS) and cyclic voltammetry (CV) measurements confirmed electrochemical performance degradation with incubation. These fundamental studies would be a cornerstone to understand the complex interaction between the Ga surface and the biological environment. 42 3.2 Materials and Methods 3.2.1 Materials Ga (99.99999% purity) was purchased from Sigma-Aldrich, and biocompatible polyether block amide (PEBAX) tubing (55D) was provided from Zeus Industrial Products (SC, USA) as a shell of Ga wire; Ga was injected into the PEBAX tubing for fabricating Ga wires that enable to evaluate electrochemical properties and control surface area. Phosphate buffered saline (PBS, pH 7.4, GibcoTM, 1x) and DMEM (GibcoTM, 1x) were purchased from Thermo Fisher Scientific. 3.2.2 Methods 3.2.2.1 SEM images All the measurements were carried out after gentle rinsing with DI water to remove unattached impurities on the Ga oxide surface. Ga surface change was characterized by the Scanning electron microscopy (SEM, TM3030, Hitachi, Japan) images captured by backscattered electrons (BSE) and Energy dispersion X-ray spectroscopy (EDAX, from Hitachi S-4800, Japan) detector. All the samples were kept in DMEM for the time-dependent analysis and purged by liquid N2 to prevent further oxidation just before measurement. The samples were exposed to an electron beam of an accelerating voltage of 20kV, which can penetrate up to 1 μm from the surface.27 The Ga surface was incubated in DMEM at 37oC with the Ga surface side up, preventing the liquefied Ga from leaking (Image to analyze liquefied Ga was shown in the Appendix). 43 3.2.2.2 Elemental analysis of dissolved Ga ions Both 60 and 150 μm thickness of Ga wires and 45 μL of Ga droplet were placed in 50 mL glass vials filled with 40 mL of DMEM at 37oC. Exposed areas of each sample are 1,589.6, 11,304, and 78,000,000 μm2, respectively. The samples of the aqueous solution with the released Ga ions were taken with time. The Ga ion concentrations of the samples were measured using an inductively coupled plasma-mass spectrometer (ICPMS, 8900 Triple Quadrupole, Agilent, USA). pH changes of water and DMEM were detected until having plateaued using a benchtop pH meter (Oakton pH 700, USA). 0.2 mL of Ga droplets were applied to generate pH changes of water and DMEM. The physiologic buffers were maintained at 37oC. 3.2.2.3 FT-IR analysis Surface change of Ga droplet was monitored by Fourier-transform infrared spectroscopy (FT-IR, Nicolet iS10, Thermo Fisher Scientific, USA) with attenuated total reflection (ATR, DuraScopeTM, SensIR technologies, USA) accessories. DMEM solution was used as background to monitor Ga change only. The surface of the Ga droplet was measured in real-time in DMEM. Only the initial spectrum (for pristine Ga) was obtained in the air with HCl vapor to remove the oxide layer.28 Also, a Ga droplet was incubated in 1M H2O2 solution to accelerate GaGaOOH reaction, which is a typical method.29 The sample incubated in H2O2 a day was used as a reference to compare the bands from GaOOH. 44 3.2.2.4 XPS elemental analysis X-ray photoelectron spectroscopy (XPS, Kratos Axis Ultra DLD, UK) was performed to verify the chemical composition of the Ga oxide skin and the contamination deposited on the skin. The analysis was carried out with a monochromated Al (1486.6 eV) source operating 15 kV at ~1 x 10-9 torr. XPS spectra were analyzed using the CasaXPS software. All spectra were calibrated with hydrocarbon (C 1s, C-C) photoemission set to 284.6 eV binding energy. The high-resolution analysis was used to acquire core spectra of the C 1s, K 2p, Ga 2p, Cl 2p, and P 2p levels. 3.2.2.5 Characterization of electrochemical property The electrochemical analysis was performed in a three-electrode electrochemical cell using an SP-150 (Bio-Logic, France) with EC-Lab V11.10 software. Electrochemical impedance spectroscopy (EIS) analysis was carried out between 10 Hz and 0.5 MHz with 20 mV of sinus amplitude to obtain Bode impedance. Cyclic voltammetry (CV) measurement was conducted in the range of -0.5 to 0.5 V with 0.1 V/s. To obtain charge storage capacity (CSC), we conducted different CV measurements in the range of -0.6 to 0.8 V with 0.5 V/s. Then the CSC of all the Ga wires was calculated from the full area integration under the CV curve. Electrochemical properties of the Ga wires were evaluated in DMEM (GibcoTM, 1X) and 1xPBS (0.01M PBS, pH 7.4, GibcoTM, Thermo Fischer Scientific). The Nyquist impedance parameters were obtained from ZView 3.5h software (Scribner Associates Inc.) with instant and equivalent circuit fitting. 45 3.3 Results and Discussions 3.3.1 Theoretical background to understand surface chemistry of Gallium in physiologic condition Based on previous studies, several chemical reactions can influence the composition of the Gallium surface in a physiological buffer. The physiological buffer consists of water, buffered chemicals, biomolecules, and dissolved gases. Gallium would likely react with dissolved oxygen, water, and hydroxyl ions. Ga can be oxidized by dissolved oxygen, growing various forms of Ga oxide (GaxOy), including Ga2O3, Ga2O, and Ga(OH)3, which are generally defined as “Ga oxide group.”30–33 The solid-like oxide form on the liquid Ga surface has a high viscosity that indicates the surface is active and unstable; hence the surface can interact with different ions.12,34 In an aqueous condition, unstable Ga oxides (GaxOy) could react with hydroxyl ions and then convert to GaOOH directly.22,35–37 A recent study confirmed that crystallite GaOOH was formed on the surface of Ga-based liquid metal in water after a day at room temperature, verified by their X-ray diffraction measurement.38 Lastly, Ga3+ can potentially be released in an aqueous system from a possible GaOOH and water reaction (GaOOH + H2O ↔ Ga3+ + 3OH-).21,39 The reaction generates reactive hydroxyl free radicals on the Ga surface that can react with other ions in physiological buffer,40–42 resulting in biofouling that often occurs on the metal surface in a physiologic buffer.43–46 Therefore, the Ga surface under a physiological buffer would develop different surface chemical composition profiles, oxide layer thickness, and the degree of Ga3+ release. This study uses a series of complementary analytical tools to a more comprehensive understanding of the surface chemistry of Ga in physiological buffer. 46 3.3.2 Elemental change in physiologic condition Although a previous study was able to determine Ga ion in physiologic buffer, it is unclear the kinetics and the reaction pathways of Ga ion release in a physiologic buffer. We first measured Ga ion concentration over time in the physiological buffer to understand the kinetics of the Ga ion release. Ga metals were incubated in DI water (Figure 3.1 A) and a cell culture medium (DMEM) (Figure 3.1 B), which is composed of common biomolecules in physiologic conditions, respectively. And the Ga ion measurements in the physiologic buffer were conducted by ICP-MS over time at 37oC (to mimic body temperature), which would determine the Ga ion release rate from the Ga surface. The four-week measurement results showed the concentration of Ga ion in the solution follows a root square growth, which indicates the Ga ions in the solution reach a saturation over time. The saturation behavior of Ga ions in the physiological buffer suggested the surface corrosion has similar kinetics as diffusion-limited Ga oxide growth in the presence of oxygen. In addition to Ga ion measurement, we also measured the pH values of the incubation media (water and DMEM). The results indicate the pH values of the physiological buffers increase over time (Figure 3.1 C). The pH values of DMEM and DI water without Ga droplet were stabilized to 7.9 (initial DMEM: 7.4) and 5.7 (initial DI water: 6.4), respectively, after 37oC incubation. These changes were generally caused by CO2 consumption and dissolution. From the baseline of pH in each solution, the pH increases of DMEM and DI water with Ga droplet were confirmed to 0.5 and 2.0, respectively. Both Ga ion concentration and pH measurement indicated the possible corrosion pathway from the Ga oxide group (GaxOy) to GaOOH, then Ga3+ and OHrelease (GaOOH + H2O ↔ Ga3+ + 3OH). The saturation behavior of Ga ions and pH in 47 the physiological buffer suggested the Ga surface is passivated over time and the surface reaction eventually stopped. To further understand the chemical change of Ga surfaces in physiologic buffer, we first imaged the Ga surface over time. The images from optical (Figure 3.2 A) and backscattered electron (BSE, Figure 3.2 B) microscopy of the Ga surface over time in DMEM suggested the Ga surface reactions might be affected by a passivation layer. Unlike pristine Ga surface oxidation in the air (Figure 3.2 C), a passivation layer was deposited on the Ga oxide surface from 2-day incubation, unveiling an agglomeration of Cl/Ca ions and carbon-based organic materials by EDAX elemental mapping (Figure 3.2 D). The passivation layer gradually started to cover the Ga oxide surface and eventually wrapped up the whole surface after 45-day. The thickness of the contamination layer was predicted over a micron scale, considering the principle of the BSE observation that can monitor up to 1-3 µm.27 The image observation indicates that Ga in physiologic conditions showed different behavior compared to the oxidation in air. Further spectroscopic analysis is needed to understand the dynamic composition changes on the Ga surfaces. 3.3.3 Spectroscopic monitoring on the Gallium surface in physiologic condition To understand the change in surface composition and passivation layer growth of the Ga surface in physiologic buffer, we first used the attenuated total reflection (ATR) technique from FT-IR spectroscopy to continuously investigate the dynamic of the chemical change on the Ga surface in physiologic buffer. To obtain the FT-IR results 48 from the Ga surface, the FT-IR spectrum of DMEM was used as the background correction. The results from background-corrected FT-IR spectroscopy are shown in Figure 3.3 A. We categorized the change of the Ga surface monitored by FT-IR into four stages. We also prepared GaOOH-rich Ga surface by incubating the Ga surface in H2O2. The FT-IR spectrum of GaOOH-rich Ga surface is shown in Figure 3.3 B as a reference for discussion. In the first stage (Figure 3.3 C, oxidation initiation), the pristine HCl-etched Ga droplet has no significant band.28 However, various bands were generated as soon as the Ga droplet was immersed in DMEM. In the first stage, a strong CO2 peak around 1400 cm-1 and oxygen-related bands around 3500 cm-1 and at 550-600 cm-1, indicate oxidation on the Ga surface was initiated. After an hour, in the second stage (Figure 3.3 D), unique O-H bands at 2900 cm-1 were measured, which is clear evidence of the GaOOH existence and propagation.37,47 The band was also observed in GaOOH dominated Ga oxide surface induced by H2O2 (Figure 3.3 B). And the GaOOH propagation continued for 6 hours (Figure 3.3 D). For 6 hours to 2 days, NH stretching bend kept increasing (third stage, Figure 3.3 E), and stabilization after three days (fourth stage, Figure 3.3 F). During the third stage, Ga-O bands generated from unstable Ga-O and volatile Ga2O around 500-600 cm-1 and at 950 cm-1 were increased after 6 hours.37,48,49 Another important feature from the third stage is the increase of the typical N-H bending at 3200-3300 cm-1, which can overlap the O-H band generated from GaOOH. The typical N-H stretching was generated from various organic materials in physiologic buffer, which is evidence of organic passivation layer deposition.50–52 After two days (fourth stage), although the intensity of the spectra was increased, the bands of the spectra from N-H stretching (3300 cm-1), 49 alkene stretching and N-H bending (around 1600 cm-1), and C-H peaks (around 600-800 cm-1) have no significant change after a week. The FT-IR results provide more detailed chemical dynamics on oxidation propagation and organic passivation growth of Ga surface. The FT-IR results correspond to the results of the Ga ion, pH, and microscopic measurements. Lastly, the chemical composition of the Ga surface was determined using XPS measurements. The XPS measurements were conducted after removing the Ga droplet from the physiological buffer. Since the penetration depth of XPS measurement is approximately 10 nm, it is ideal for probing the chemical composition of the outermost surface.53 Ga 2p peaks are useful to demonstrate the oxide layer formation naturally grown on the bare Ga element with incubation. The core-level XPS of the Ga 2p doublets are shown in Figure 3.4 A for the first 5 days of incubation in the physiologic buffer. The binding energy (BE) values of these doublets, Ga 2p3/2 and Ga 2p1/2 signals are 1141.7 and 1114.9 eV, respectively. Both doublets consist of native oxide (red, positively shifted, approximately 2 eV) and element (blue). We can determine the oxidation state of Ga by the ratio between both native oxide and base element signals.54 Figure 3.4 B summarized the ratios of both 2p 1/2 and 3/2 after baseline subtraction and deconvolution. The increase of the ratio represents the positive shift from Ga element to native oxide peak. Since the positive shift of the peaks in the Ga XPS survey was induced by oxidation, the results determined oxide layer was grown on the Ga surface over the five days. We confirmed that XPS spectra from the Ga 3d (Figure 3.4 C), 3s (Figure 3.4 D), and O 1s (Figure 3.4 E) were also positively shifted, which indicates an oxide layer was formed on the Ga surface for five days.33,38,55–58 50 From the wide range, XPS survey in the binding energy (BE) range of 0-1200 eV over a 45-day incubation period in DMEM, the intensities of Ga related peaks after five days were attenuated and disappeared, as shown in Figure 3.5 A. Although bare Ga has no C and O that was confirmed by the EDAX study (Table 3.1), the XPS survey observed the presence of C and O on the bare Ga surface. This is the typical result due to CO2 in the test chamber and the contamination during measurement, we can therefore ignore the signals of bare Ga.59–61 Instead of all the Ga signals, including Ga 2p, LMM, 3s, 3p, and 3d, other elemental signals such as Na 1s (1071 eV), Cl 2p (199 eV), P 2p (133 eV), and C-NH2 (400 eV) appeared after five days (Figure 3.5 B), which were also confirmed by EDAX (Table 3.1) analysis. The change suggests two possible scenarios of Ga incubation in a physiologic buffer. First, abundant OH- and Ga3+ on the Ga oxide surface verified by FT-IR study react with many ions in the physiologic buffer. And the thickness of the passivation layer should be thicker than the 20 nm scale, which is the limitation thickness detected by XPS measurement.53,62,63 The growth of the contamination layer blocked the X-ray analysis of the inside Ga core; hence no prominent Ga-related peaks were detected after five days. The XPS study detected various organic compounds on the Ga oxide surface in DMEM after five days. The C 1s core-level spectra (Figure 3.5 C) showed an increase in the C-O and COO after five days (the ratio between COO and C-O with C-C peaks were tabulated in Table 3.2). The peaks are different from the default C 1s core-level spectrum from bare Ga. Hence the change offers a strong clue to prove the absorption of organic molecules on the Ga oxide surface. Other compounds were also specified the incubation period between 5 and 45 days, such as Na 1s (Figure 3.5 B), K 2p (Figure 3.5 C), Cl 2p 51 (Figure 3.5 D), and P 2p (Figure 3.5 E), which are already mentioned above. However, there is no trend of contamination layer formation with incubation periods. Since the layer formed naturally and at random without manipulation steps, the formation process is far from pursuing uniformity. Also, organic compounds displayed different deconvoluted peaks from the XPS survey from tiny molecule structure differences.64,65 Overall, the XPS results can prove the oxidation on the Ga surface with incubation and the contamination layer formation on the Ga oxide surface after five days. And the diminishing Ga oxide signal in XPS also suggests the passivation layer is over 20 nm thickness. Overall, the SEM (Figure 3.2) and XPS (Figure 3.5) measurements can confirm that the passivation layer was formed on the Ga oxide surface after the 4-day incubation. Plentiful OH- and Ga3+ ions were generated from the theoretical reaction first; GaOOH + H2O → Ga3+ + 3OH- (Figure 3.1). The reaction rate was slowed down with incubation time and ultimately reached a saturated OH- and Ga3+ ion concentration on the surface. This saturated behavior is likely due to the Ga oxide surface being passivated by various ions and organic materials in physiologic buffer, as shown in Figure 3.2. 3.3.4 Electrochemical characterization of Ga surface in physiologic condition The Ga-based liquid metal experienced oxidation and then was contaminated in physiologic conditions as stated in above Chapter 3.3. Cyclic voltammetry (CV) and electrochemical impedance spectroscopy (EIS) were used to investigate the effect of surface oxidation and further corrosion following immersion in DMEM cell culture media. When we compared the anodic limit of CV curves over time, the Ga wires had an 52 initial anodic onset potential of 0.18 V; the onset points were increased to 0.33 V over five days (Figure 3.6 A). The increase of the onset potential suggested that the bandgap of the conductive materials was increased, which resulted in lower electrical conductivity and limited charge transfer at the electrode/electrolyte interface.24 The anodic onsets showed little change after five-day passivation in physiologic buffer (Figure 3.6 B). In addition to the CV analysis, EIS measurements were also performed over 90 days. The Nyquist plots from EIS data illustrated the real and imaginary components of the electrochemical impedance. They are often used to understand the electrochemical behavior of the outermost surface called the “skin effect66.” The Warburg impedance (tails of the Nyquist plots) of the liquid metal electrode can be observed in Figure 3.6 C. The slope of the Warburg impedance decreased over time with oxidation. The shape of the plot was gradually changed from linear to a curve of semi-circle after 90 days. The appearance of a semi-circle in the Nyquist plot is often an indication of a corrosion process at the electrode surface, which generates a different circuit-equivalent model compared to the standard Randles circuit model (Figure 3.6 D). The Nyquist impedance results demonstrated that the passivated oxide layer also participates in the electrochemical reaction in an aqueous solution, which leads to resistance increase (Figure 3.7).67,68 The charge transfer resistance (Rct) obtained from Nyquist impedance plots were tabulated in Table 3.3. The resistance parameter fitting results indicated that the corrosion circuit model suggested in Figure 3.6 D provided more accurate fitting results after 15-day incubation than the typical Randles circuit model. The equivalent circuit model selection confirmed that the Rct of Ga was increased from 102 to 106 Ω cm2 with physiologic incubation. Both CV and EIS studies demonstrated the instability and 53 limited electrochemical properties of the liquid metal surface after long-term exposure to an aqueous, physiological environment. 3.4 Conclusion The goal of this study is to elucidate the complex interactions between Gallium and biological molecules. Gallium was incubated in DMEM at body temperature over 45 days and monitored the interactions from various methods. The dynamic study between Gallium and physiologic buffer confirmed that release of Ga ions and pH change showed logarithmic growth with incubation time, which reactions support typical GaOOH +H2O → Ga3+ + 3OH- pathway. The optical study demonstrated the surface change of Gallium in the physiologic buffer is different from that in air. The incubation process generated further oxidation beyond the self-limited oxidation in air and contamination on the Gallium oxide surface. Three different spectroscopic studies were used to determine the chemical composition on the Gallium surface. FT-IR study verified the GaOOH-rich formation on the surface within six-hour incubation then ionized to Ga3+ and OH-. The chemical composition change was stabilized after four days, which trend is similar to kinetic study results. XPS analysis indicated that surface oxidation progressed, and then a contamination layer was formed on the Gallium oxide surface. The contamination layer may be introduced by the interaction between ionized GaOOH and various ions in DMEM. Lastly, the RAMAN study demonstrated GaOOH dominated the Gallium surface within two days, then the contamination layer grew on the Gallium oxide surface. Considering the penetration limit of the x-ray, laser, and electron beam, the thickness of 54 the contamination layer could be expected over 1-2 µm. 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B 2017, 121 (16), 4220–4225. https://doi.org/10.1021/acs.jpcb.7b02174. (66) Burton, C. E.; David, R. M.; Portnoy, W. M.; Akers, L. A. The Application of Bode Analysis to Skin Impedance. Psychophysiology 1974, 11 (4), 517–525. https://doi.org/10.1111/j.1469-8986.1974.tb00581.x. (67) Fazel, M.; Salimijazi, H. R.; Shamanian, M. Improvement of Corrosion and Tribocorrosion Behavior of Pure Titanium by Subzero Anodic Spark Oxidation. ACS Appl. Mater. Interfaces 2018, 10 (17), 15281–15287. https://doi.org/10.1021/acsami.8b02331. (68) Deyab, M. A.; Mele, G. Stainless Steel Bipolar Plate Coated with Polyaniline/ZnPorphyrin Composites Coatings for Proton Exchange Membrane Fuel Cell. Sci. Rep. 2020, 10 (1), 3277. https://doi.org/10.1038/s41598-020-60288-9. 62 Figure 3.1. Ga ion mass transfer in water and physiologic buffer at 37oC. Ga ion release in A) DI water and B) DMEM with incubation. C) pH change in water (black) and DMEM (red) with incubation. 63 Figure 3.2. Image validation of Ga with incubation. A) Confocal (scale bar: 70 µm) and B) SEM (BSE, scale bar: 10 µm) images in DMEM with incubation for 45 days. C) Ga surface change in air for 45 days (scale bar: 10 µm). D) EDAX analysis to confirm the elements bonded to the Ga surface with physiologic incubation (scale bar in EDAX images are arbitrary from the device. We can ignore the scale). 64 Figure 3.3. FT-IR spectra of Ga droplet in DMEM with incubation. A) For a week and B) GaOOH-rich phase promoted by H2O2 etching (1 day). The spectra were grouped under four difference stages with incubation period; C) Oxidation initiated, D) GaOOH propagation, E) OH bending increased, and F) stabilization. 65 Figure 3.4. XPS survey scans of Ga surface in DMEM for 5 days. A) Ga 2p core level spectra. B) Summary of the ratio between native oxide and element peak from Ga 2p core level spectra. XPS survey scans of C) Ga 3d, D) Ga 3s, and E) O 1s for 5 days. 66 Figure 3.5. XPS survey scans of Ga surface for 45 days. A) All spectra, B) magnified after 5-day, C) C 1s core level spectra, D) Cl 2p core level spectra, and E) P 2p core level spectra. 67 Figure 3.6. Electrochemical performance degradation in DMEM with incubation. CV curves of Ga wires A) for 5 days and B) after 5 days. C) Nyquist plots and D) circuit models to show the change of the equivalent circuit model from normal Randles to corrosion. 68 Figure 3.7. Nyquist plot change in DMEM for 3 months. A) Erratic for 5 days, B) stabilized for 5-25 days, and C) after 45 days with resistance increasing. 69 Table 3.1. EDAX data of the Ga surface for 45 days * All the results are average obtain from five times measurements. a Preparation was explained in Section 3.2.2.5. 70 Table 3.2. Summary of the XPS parameters from C 1s core level spectra 71 Table 3.3. Nyquist plot parameters of Ga with physiologic incubation in DMEM (from Figure 3.7) CHAPTER 4 GALLIUM SURFACE MODIFICATION BY ELECTROCHEMICAL DEPOSITION 4.1 Introduction Liquid metals from Ga and its alloys have emerged as a promising material for soft bioelectronics due to their excellent electrical conductivity and no mechanical stiffness at body temperature.1,2 However, the liquid metals are prone to oxidization at a ppm level oxygen concentration that deteriorates the electrical properties of liquid metals.3–5 The liquid metal oxidation progressed further in an aqueous solution, causing Ga ion release and ultimate surface degradation.6,7 The surface degradation also negatively impacts the electrochemical properties of liquid metals, inhibiting ion reactions and resulting in low reactivity of the liquid metal electrode as demonstrated in Chapter 3. Also, excessive Ga ions near tissue and organs cause cell degeneration.8,9 Therefore, liquid-based bioelectronics requires hermetic encapsulation to eliminate direct contact with biological fluids. This encapsulation strategy can effectively prevent oxidation from biological fluids, improve biocompatibility, and increase the biostability of liquid metals under physiological conditions. Electrochemical deposition of the materials on the Ga surface can be an effective strategy to improve the biostability and electrochemical performance of liquid metal in 73 physiologic conditions. There are two main challenges for the electrochemical deposition process on Ga. First, electrodeposition from an aqueous electrolyte is preferred for biological applications due to the availability of various dopant agents in aqueous solution and the potential cytotoxicity of organic solvents.10–12 However, Ga alloys undergo hydrolytic oxidation and degradation in an aqueous solution over a positive current.13 Second, electrodeposited materials can often create heterogeneous and porous structures. It is difficult to deposit onto the entire Ga surface uniformly.14 In this study, we have considered suitable materials selection and their suitable electrochemical deposition strategies to improve electrochemical performance and chronic biostability of Ga surface. Gold was selected as the first representative encapsulating metal due to its high electrochemical property and biocompatibility. Gold nanoparticles (Au NPs) can be deposited on the Ga surface from hydrogen tetrachloroaurate hydrate (HAuCl4) by electrochemical deposition.15 Since the deposition of Au NPs requires negative potential that generates a reduction process, Au NPs could be deposited on the Ga surface without hydrolytic oxidation. The reduction process even removes the presented oxide layer on the Ga surface, the method can therefore be harnessed to change surface chemistry and composition of Ga with deposition time.16,17 The next strategy is a carbon nanotube (CNT)/polymer composite layer deposition. Carbon nanotubes (CNT) involve a high surface area and electrochemical properties. Many studies have utilized CNT to modify the metal surface for high current density and low impedance applications.18,19 In addition, polymeric material wrapping on the CNT can provide ionized atoms such as N+ and O- on the CNT surface and robust adhesion with the metal surface due to its viscous property. CNT/polymer composite with 74 nonporous structures have excellent flexibility and electrochemical stability to improve the electrochemical performance and biostability of Ga alloys.20,21 The other deposition strategy has been designed by a tailorable conducting polymer that improves not only the biostability and electrochemical performance of the Ga surface but also softness due to the viscoelastic polymer characteristics. A conducting polymer such as poly(3,4-ethylene dioxythiophene) (PEDOT) showed low impedance, high capacitance, and biocompatible; thus, the PEDOT was used in several electroanalytical applications.22–29 However, previous studies have indicated that it is difficult to polymerize a high-performance PEDOT coating on a non-noble metal surface and avoid oxidation during electrodeposition in an aqueous environment.30–32 To address these issues, we prepared liquid metal electrode surfaces by electrochemical deposition of PEDOT: BF4 (tetrafluoroborate) on the Ga surface in an organic solvent electrolyte. The organic solvent approach enabled successful electrochemical deposition without Ga surface oxidation or degradation. Lastly, we have designed a multilayer deposition strategy to obtain the higher electrochemical performance and biostability of the Ga surface. For example, metal nanoparticle decoration on the CNT sidewalls is an effective strategy to enhance electrochemical properties further for the various sensor applications.33–35 Particularly, the gold nanoparticles (Au NPs) have been used for high-performance electrochemical biosensors due to their biocompatibility and stability.36–38 After the electrochemical deposition process, we performed dopamine (DA) sensing experiments in vitro as a representative neurotransmitter to compare the performance with other electrochemical biosensors. Also, we confirm a novel 75 electrodeposition strategy to address stretchable bioelectronics, which has great attraction recently. Suggested different material selection and deposition strategies on the Ga surface are helpful to improve biostability, biocompatibility, the electrochemical performance of the Ga-based soft bioelectronics. 4.2 Materials and Methods 4.2.1 Materials For the gold deposition, HAuCl4 (99.999%) and potassium chloride (KCl. 99.5%) were purchased from Acros organics and Sigma-Aldrich, respectively. For the CNT deposition, single-wall carbon nanotubes (SWCNT, P2-SWNT) was purchased from Carbon solutions, Inc. Poly(diallydimethylammonium chloride) (PDDA, MW 200,000350,000, 20 wt% in H2O) and poly(styrenesulfonate sodium salt) (PSS-Na+, MW 1,000,000) were purchased from Sigma-Aldrich for the polymer wrapping on the CNT surface. For the PEDOT deposition, 3,4-ethylenedioxythiophene (EDOT, 97%), tetraethylammonium tetrafluoroborate (TEABF4, 99%), and propylene carbonate (PC, anhydrous, 99.7%) were purchased from Sigma-Aldrich. For the control of Ga-based liquid metal wires, 125 µm perfluoroalkoxy coated Pt wire (75 µm, 99.99 purity%) was purchased from A-M systems. 4.2.2 Methods 4.2.2.1 Gold nanoparticles deposition The aqueous gold electrolyte consisted of 0.2 mM HAuCl4 and 0.1 M KCl to deposit Au NPs on the Ga surface. The electrochemical deposition was accomplished on 76 the Ga surface from -2.0 V for 10-30 min, corresponding to 80-700 nm thickness. 4.2.2.2 CNT/polymer deposition PDDA was dissolved in DI water with 24h stirring then CNT was introduced in the polymeric solution. We confirmed a 15 min sonication is optimum to provide excellent CNT dispersion with minimum defects of CNT characteristics. The electrochemical deposition was accomplished from -2.7 V for 30-50 min to the hermetic coating on the Ga surface. As a control group, CNT/PSS electrolyte was also prepared by the same experimental setup. The CNT/PSS was deposited on the Ga surface from +2.5V for 40-60 min to encapsulate the entire Ga surface. 4.2.2.3 PEDOT deposition Unlike Au NPs and CNT deposition processes, the PEDOT deposition was performed from the organic solvent-based electrolyte. 0.12 M of TEABF4 was dissolved in PC, and then 0.01 M of EDOT was dropped in the PC/TEABF4 solution with 24h stirring. A nanoporous PEDOT:BF4 was deposited on the Ga surface from +1.3 V for 25 sec. As a control group, PEDOT:PSS was also deposited on the Ga surface, representing the PEDOT structure. 0.1 M PSS-Na+ was dissolved in DI water with 24h stirring, then 0.01M EDOT was dropped in an aqueous PSS polymeric solution. Constant +1.7 V of potential was applied for electrochemical deposition for 10-15 min. 77 4.2.2.4 Characterization of electrochemical properties The electrochemical analysis was performed in a three-electrode electrochemical cell using an SP-150 (Bio-Logic, France) with EC-Lab V11.10 software. The surfacemodified Ga wires were used as a working electrode. Pt and Ag/AgCl were used as a counter and a reference electrode, respectively. Electrochemical impedance spectroscopy (EIS) analysis was carried out between 10 Hz and 0.5 MHz with 20 mV of sinus amplitude to obtain bode impedance. Cyclic voltammetry (CV) measurement was conducted in the range of -0.6 to 0.6 V with 0.1 V/s. To obtain charge storage capacity (CSC), we conducted different CV measurement in the range of -0.6 to 0.8 V with 0.5 V/s. Then the CSC of all the Ga-based microelectrodes was calculated from the full area integration under the CV curve. Standard electrochemical properties were evaluated in phosphate buffered saline electrolyte (0.01M 1xPBS, pH 7.4, GibcoTM, Thermo Fischer Scientific). 4.2.2.5 Characterization of morphologies The cross-section of the Ga wires after electrochemical deposition was characterized by scanning electron microscopy (SEM, Hitachi S-4800, Japan) at 20 kV accelerating voltage. Backscattered electrons (BSE) images and energy dispersion X-ray spectroscopy (EDAX) were obtained by SEM (TM3030, Hitachi, Japan) at 15kV accelerating voltage to examine Ga-based liquid metal surface degradation and electrochemical deposition. Nanostructured PEDOT was characterized by a field emission SEM (Helios Nanolab 650, FEI, Thermo Fisher Scientific, USA). 78 4.2.2.6 Characterization of stability The deposited materials stability following Ga-based liquid metal wires exposure to the physiologic buffer was confirmed by SEM (Hitachi S-4800, Japan). After aging, the Ga ion concentration was quantified using an inductively coupled plasma-mass spectrometer (ICP-MS, 8900 Triple Quadrupole, Agilent, USA). The bare and surface modified Ga wires were placed in 50 mL glass vials filled with 40 mL of PBS at 37oC for a month to examine the release of Ga ions from the electrodes immersed in a physiological buffer at body temperature. In addition, we evaluated the electrochemical stability after heating and bending tests at 1 kHz of impedance, which is the general standard frequency to record mammalian neural signals generated from the action potential.39 The heating test was designed to mimic the physiologic condition, including ion conductive PBS/Agarose gel at 37oC for three months. 500 times of 120o bending test was carried out using a tensile tester (MARK-10, NY, USA). Image to analyze bending stability of Ga wires was shown in Appendix A.2. The electrochemical stability of all the microelectrodes was confirmed by multicycle CV curves using an SP-150 with EC-Lab V11.10 software. CV measurements were conducted in the range of -0.5 to 0.5 V with a scan rate of 1.0 V/s in PBS electrolyte in a three-electrode electrochemical cell. 4.2.2.7 Dopamine (DA) detection for sensitivity and selectivity Various concentrations of dopamine hydrochloride (DA, Sigma-Aldrich) were prepared in 0.1M of PBS (10x PBS, pH 7.4) to confirm the sensitivity of the liquid metalbased DA sensors. The measuring window of the CV is within -0.6 to 0.6 V with 50 mV/s of scan rate in a three-electrode electrochemical cell. L-ascorbic acid (AA, 99%) 79 and uric acid (UA, >99%) were purchased from Sigma-Aldrich to mimic human fluid culture. The 2.0 mM AA and 1.0 mM UA were introduced in the PBS solution with the 0.01 mM DA to confirm the selectivity of the liquid metal-based DA sensors, which is the typical ratio in human body fluids.40 4.2.2.8 Fast scan cyclic voltammetry (FSCV) determination FSCV was conducted by CV advanced monitoring in potentiostat (SP-150, Biologic, France) with Ultra-Low-Current modules, EC-Lab V11.10 software, and a twoelectrode system, including the liquid metal-based DA sensors as working electrodes and chloridized Ag/AgCl wire as a reference. A triangular waveform cycle from -0.4 to 1.3 V then back to -0.4 V at 400 V/s was designed to perform the FSCV. The scan used a holding potential of -0.4 V between scans to set a cycle frequency of 10 Hz, considering the accumulation and adsorption process of the DA on the sensor surface. The different amounts of DA dissolved in artificial cerebrospinal fluid (ACSF) were injected into the test tubing through the flow at 1ml/min. The ACSF was prepared by the rule of Cold Spring Harbor Protocols, doi:10.1101/pdb.rec065730. The fifth cycle of six successive scans was used to obtain a stabilized result. The test setup follows now acknowledged to be the FSCV rule by recent studies.41,42 4.3 Results and Discussions 4.3.1 Gold deposition on the Ga surface Based on previous studies of Gallium chemistry, enhancement of both electrochemical and biostable properties of Ga should be required by a conductive 80 nanomaterial encapsulation. Since Au and other noble metals are common materials for enhancing the electrochemical properties of the bioelectronics, the Au was deposited on the liquid metal first from the HAuCl4 solution by the negative current (Figure 4.1 A), and a simplified reaction equation is below 𝐻 𝐴𝑢𝐶𝑙 3𝑒 → 𝐴𝑢 4𝐶𝑙 𝐻 𝑅𝑒𝑑𝑢𝑐𝑡𝑖𝑜𝑛 𝑜𝑛 𝑤𝑜𝑟𝑘𝑖𝑛𝑔 𝑒𝑙𝑒𝑐𝑡𝑟𝑜𝑑𝑒 (1) Although the Au electrochemical deposition provides improved electrochemical properties initially, such as high charge storage capacity (Figure 4.1 B) and lower impedance (Figure 4.1 C) to Ga surface, the Au NPs layer displayed poor stability when the Ga wire was incubated in physiologic buffer at 37 ºC. We observed a significant leakage of liquid metal through the electrodeposited Au layer (Figure 4.1 A, right). Therefore, viscous material support is needed to improve the biostability of the Ga-based bioelectrodes, preventing liquid metal leaking as well as Ga degradation. 4.3.2 CNT/PSS deposition by positive current CNT/polymer composite could be a solution due to the high electrochemical stability of the CNT and viscous characteristics of polymers.43–45 To test the effect of deposition current, we choose two polymers including opposite charges in an aqueous solution. Poly(4-styrene sulfonate) (PSS) is a typical polymer equipped with a watersoluble property that enables water dispersion of conductive polymers or carbon materials. And it has been used for different carbon-based nanocomposites.46,47 Also, the PSS is used as a dopant for conductive polymers or provides sulfur catalyst on CNT surface to improve electrical properties.48,49 Hence, we performed CNT/PSS electrochemical deposition as a conductive and biostable layer on Ga and typical 81 platinum (Pt) wires. The CNT/PSS composite was deposited by positive current (+2.5 V) because PSS is negatively charged in an aqueous solution. We observed CNT/PSS layer was successfully deposited on both Pt (Figure 4.1 D) and Ga surfaces (Figure 4.1 E). The CNT composite showed an order of impedance decrease with the thickness of the CNT layer on the Pt surface (Figure 4.1 F). In addition, charge storage capacitance (CSC) of Pt wire was increased from 0.10 to 1.23 mC/cm2 by 40 min CNT/PSS deposition (Figure 4.1 H). However, the nanocomposite on Ga leads to poor electrochemical property (Figure 4.1 G and I). These results demonstrate the significant impact of the oxidation process on the Ga surface during electrodeposition. Unlike noble metal, Ga-based liquid metal can therefore undergo hydrolytic oxidation during anodic electrochemical deposition. Therefore, an alternative deposition strategy is needed to overcome the oxidation of the Ga surface during the deposition process. 4.3.3 Optimal CNT/PDDA deposition Water-soluble PDDA wrapping on the CNT surface is another effective method to prepare an aqueous CNT solution electrolyte. Moreover, PDDA is negatively charged in an aqueous solution, enabling electrochemical deposition by negative current (see chemical structure in Figure 4.2 A). We performed spectroscopic measurements to determine the optimal ratio between CNT and PDDA. Fourier-transform infrared spectroscopy (FT-IR) measurement confirmed the successful blend of CNT/PDDA with the ratios (Figure 4.2 B). The PDDA spectrum displayed 1470 and 1630 cm-1 representing bands attributed by the C=C bond and a unique 2130 cm-1 band.50 82 RAMAN analysis represented the difference of CNT characteristics with the ratios, resulting in the change of electrochemical performance of the CNT composites (Figure 4.2 C). Moreover, G band analysis that is the combination of G+ at 1625 cm-1 and G- at 1600 cm-1 provided the diameter change and metallicity of the CNT51. We observed CNT/PDDA of the 1:3 mixture displayed the downshift of the G+ band and the relatively low intensity of the G- band that indicates a diameter increase of CNT and a loss of metallicity of the composite, respectively (Figure 4.2 D).50,51 In contrast, the CNT/PDDA (1:10) represented a smaller diameter increase of CNT and higher metallicity than the CNT/PDDA (1:3) due to incomplete blending and PDDA agglomeration.51,52 The agglomerated excessive PDDA as an insulator can cause the increase of the impedance and the low charge storage capacitance (CSC) (Figure 4.3 A and B). The spectroscopic and electrochemical measurements demonstrated that the CNT/PDDA (1:3) is an optimal ratio used as a CNT/PDDA solution in this study. We also determined an optimal CNT/PSS (1:2) solution in the same method for the preliminary study (see Figure 4.3 C and D). From the aqueous CNT/PDDA electrolyte, we succeeded in depositing the composite on the Ga surface can be seen in Figure 4.4 A. The whole CNT/PDDA structure on the Ga surface displays the nanocomposite leads to high encapsulation, and the magnified CNT/PDDA structure indicates the average diameter of SWCNT/PDDAs was measured approximately 25 nm, which is five times higher than that of pristine SWCNTs (4-5 nm) (Figure 4.4 B). The electrochemical properties of CNT/PDDA microelectrode were characterized by Bode impedance and CV. Impedance (Z) measurements showed a progressive decrease in Z with deposition time (Figure 4.4 C). 83 CV measurements also exhibited increased capacitance with increasing CNT/PDDA deposition (Figure 4.4 D), indicating that more conductive CNTs were mounted on the Ga surface. The CSC values of the CNT/PDDA microelectrodes are 0.14 for 20 min and 1.18 mC/cm2 for 60 min, respectively. The electrochemical property trend of the CNT/PDDA system was totally different from that of CNT/PSS, as stated in Chapter 4.3.2. Hence, we conclude that PDDA is a suitable supporter of depositing CNT on the Ga surface with high biostability and electrochemical performance. Surface analysis of Ga after CNT deposition provided evidence that oxidation during CNT/PSS deposition leads to poor electrochemical performance (Figure 4.5 A). CNT layers generated by PSS and PDDA were removed to observe the interspace between the liquid metal surface and removed CNT layers. Tetrahydrofuran (THF, Sigma-Aldrich) was used to remove the PEBAX shell and CNT layer. Figure 4.5 B shows bulk oxide layer was formed on the liquid metal surface after CNT/PSS layer deposition while rarely oxidized after CNT/PDDA layer deposition. Positive potential causes the oxide layer growth, leading to electrochemical performance degradation.30,53 Whereas the negative potential was typically used to remove the nature-grown oxide layer on the liquid metals.16,17 Nyquist plot displayed electrochemical evidence. Figure 4.6 A showed that both polymeric materials, PDDA and PSS, yield to reduce real impedance (Re Z) after electrochemical deposition on the Pt surface, which indicates the decrease of charge transfer resistance (Rc) between electrode and electrolyte. CNT/PDDA deposition on the Ga surface also offered the decrease of Re Z with deposition time (Figure 4.6 B). In contrast, CNT/PSS electrochemical deposition by anodic current formed 84 another resistance interface, typically generated by oxidation and degradation (Figure 4.6 C).54–56 We calculated all the parameters such as solution resistance (Rs), oxide layer resistance and capacitance (Ro and Co), including Rc using another valid equivalent circuit model as shown in Table 4.1 (corrosion circuit). We compared the Rc after CNT/PSS deposition exhibited a four-order magnitude greater than that after CNT/PDDA deposition. Hence, our strategy effectively enhances the electrochemical property of the Ga-based liquid metal, restricting the oxide layer growth. 4.3.4 Biostable and high-performance multilayer nanocomposites deposition Although we have determined the electrochemical deposition strategy and the optimal CNT/PDDA solution electrode deposition, the electrochemical impedance improvement is still limited compared to Au NPs performance as directed in Chapter 4.3.1. Here, we suggested a CNT/Au multilayer nanocomposite on the Ga surface to ensure biostability by CNT/PDDA conductive encapsulation and high electrochemical property by Au NPs decoration. Au NPs were deposited by the same method suggested in Chapter 4.2.2.1 on the CNT/PDDA structure. SEM results confirmed that CNT/PDDA and Au NPs were deposited on the Ga surface in sequence (Figure 4.7 A) and zoomed in SEM images of the CNT and Au layers (Figure 4.7 B). We determined the potential of 2.0 V is suitable to deposit Au NPs onto the sidewall of CNT, and at least 20 min was required to obtain uniform Au NPs structure on the CNT/PDDA layer (Figure 4.7 C). Electrochemical impedance spectroscopy (EIS) and CV measurements were performed to characterize the electrochemical performance of the CNT/Au NPs on the Ga surface. Two-magnitude of impedance was drastically decreased after Au deposition and 85 continuously decreased with an increase in deposition time (Figure 4.8 A). The CV measurements also indicated increased CSC after CNT and Au NPs deposition (Figure 4.8 B). The CSCs after CNT for 40 min and Au NPs deposition for 20 min were calculated to 0.11 and 1.73 mC/cm2, respectively. Overall, CNT/Au nanocomposites installation improved the electrochemical properties of the Ga surface, which could enhance electrochemical sensing performance. The stability of the structure is an important factor for chronic bioelectronic applications. We conducted various stability tests to demonstrate the surface stability and electrochemical property under in vitro physiologic buffer solution. Figure 4.8 C displays the CNT/Au NPs structure was maintained in PBS/agarose gel at 37oC for a month. The Ga ions in the PBS/agarose gel were measured to demonstrate that the nanocomposites can prevent the release of the liquid metal ions (Figure 4.8 D). Since the measurements showed no detectable amounts of Ga ion after CNT structure deposition and after Au decoration for 4-week incubation, we concluded the CNT/PDDA layer could protect the exposure of the liquefied Ga to the outside physiologic condition. Electrochemical stabilities were also evaluated under the long-term physiological and multiple mechanical stresses. The magnitude of the impedance at a frequency of 1 kHz was plotted over time and bending cycles, respectively. Figure 4.8 E shows both CNT and CNT/Au nanostructures kept in PBS/agarose gel at 37oC generated an impedance change of less than a magnitude for three months. The Ga wire bending (positive and negative 60o) was conducted in the water bath at body temperature to simulate the micromotion of soft tissues. The EIS measurements were performed over the cycles, and the results indicated that the impedance of both CNT and CNT/Au 86 nanocomposites was maintained for 500-cycle bending stresses (Figure 4.8 F). Lastly, multiple cycles of CV measurements were performed to evaluate the repeated redox stability of the CNT/Au NPs structure on the liquid metal. We determined little change in the shape of the CV cycle within 600 successive cycles (Figure 4.8 G). In summary, structural and electrochemical stabilities of nanostructured liquid metal biosensors were evaluated by a series of heating, mechanical, and electrochemical reaction tests under in vitro physiologic conditions. All the stability tests demonstrated the CNT/Au nanocomposites could provide excellent biostability and electrochemical properties for liquid metal surfaces. 4.3.5 High-performance PEDOT deposition Conducting polymer electrochemical deposition may be a potential strategy to protect the liquid metal surface, improve its electrochemical properties, and ultimately prepare ideal soft bioelectronics consisting of all soft materials. PEDOT doped with poly(styrene sulphonate) (PEDOT:PSS) is typically used for bioelectronic applications due to its low impedance, high capacitance, and demonstrated biocompatibility.22–29 Furthermore, the softer mechanical properties of the PEDOT layer (Young`s modulus of 10-200 MPa)57–59 that can easily comply with outside shape may decrease the mechanical mismatch between the liquefied Ga and the PEDOT layer to minimize delamination. First, a conventional electrochemical polymerization method of PEDOT: PSS was applied from EDOT and PSS dispersion in water based on a previously published method.60 PEDOT:PSS deposition led to high electrical conductivity (typically 300 S/cm, prepared by electrochemical polymerization), high electrochemical performance (under 87 105 ohm of impedance), and biocompatibility; thus, the PEDOT:PSS is widely used for biosensor studies at neural and tissue interface. PEDOT can be formed by oxidative polymerization under negatively charged PSS- solution. Both PEDOT and PSS have a strong electrostatic association, resulting in p-doped conductive PEDOT formation.57,61 Although PEDOT:PSS successfully deposited on the liquid metal surface, the deposition failed to improve the electrochemical properties of the liquid metal electrodes (Figure 4.9 A-D). Previous studies of PEDOT:PSS deposition on non-noble metals such as Ag, Cu, and stainless steel also demonstrated a decrease in the electrochemical properties.30,62 Unlike noble metals, such as Pt, active metals showed higher free corrosion potential after PEDOT:PSS coating due to surface oxidation by acidic PEDOT:PSS property. We confirmed that Ga-based liquid metal underwent oxidation on the surface during PEDOT:PSS electrochemical deposition on Ga surface (2x2 mm), as shown in the elemental maps of Gallium surface after deposition (Figure 4.9 E). We observed elemental oxygen in the bare Gallium area. Since PEDOT:PSS polymerization requires positive anodic potential in an aqueous environment, it can facilitate hydrolytic oxidation.32 In an attempt to avoid surface oxidation during PEDOT deposition and polymerization, tetrafluoroborate (BF4) was selected as an alternative dopant to polymerize PEDOT in an organic solvent;32,63 This method may improve the stability of the Ga-based liquid metal by avoiding aqueous degradation while preventing hydrolytic oxidation during deposition. Propylene carbonate (PC) was selected as an organic solvent for its low cytotoxicity than typical acetonitrile (ACN) and dimethyl sulfoxide (DMSO).12 Figure 4.10 A shows that one can achieve uniform deposition of nanoporous 88 PEDOT:BF4 on the Ga surface with complete coverage. EIS and CV measurements were conducted to characterize the electrochemical properties of the PEDOT functionalized liquid metal electrodes. The Bode plots of the EIS measurements showed a decrease in the magnitude of impedance over deposition time (Figure 4.10 B). There was a three-fold magnitude decrease in impedance over a 50s deposition time. The phase diagram revealed decreasing phase lagging with increasing PEDOT deposition (Figure 4.10 C). The CV measurements also exhibited increased charge storage capacity (CSC) with PEDOT deposition time (Figure 4.10 D). The CSC of the PEDOT functionalized liquid metal electrode also was significantly improved (126.83 mC/cm2 for a 50s deposition) compared to the bare liquid metal electrode (0.03 mC/cm2). Overall, PEDOT: BF4 deposition significantly improved the electrochemical performance of the liquid metal electrode. The stability of the PEDOT functionalized liquid metal electrode is also important for chronic bioelectronic applications. And the stability of the PEDOT layer is highly dependent on the substrate characteristics.25,64 Therefore, various in vitro stability tests were performed to evaluate the morphology and electrochemical properties of the PEDOT functionalized liquid metal electrode. First, Ga-based electrodes were kept in PBS/Agarose gel at 37oC for a month to confirm surface morphological stability. Figure 4.11 A shows the PEDOT deposited under optimum conditions (+1.3V for 25 sec) maintained its structure without any physical change or delamination. The Ga ion concentration also was measured in the PBS buffer by ICP-MS (Figure 4.11 B). The ICPMS results showed no detectable amount of Ga ion in the PBS buffer after 1-month incubation of the PEDOT: BF4 deposited Ga electrodes, in contrast to liquid metal 89 electrodes without PEDOT deposition (Figure 4.11 B black line). These results indicated that PEDOT deposition could successfully protect the tips of liquid metal electrodes from corrosion under physiological conditions. Long-term electrochemical measurements were conducted to evaluate the electrochemical and mechanical stability of PEDOT: BF4 functionalized liquid metal electrodes. The electrodes were kept in PBS-agarose gel at body temperature to mimic physiological conditions for three months. EIS measurements were conducted over time. The magnitude of the impedance at a frequency of 1 kHz was plotted over time (Figure 4.11 C). We observed the magnitude of impedance for the PEDOT functionalized liquid metal electrodes initially increase from 1x104 Ω to 3x104 Ω over a week, then maintained this constant value at approximately 3x104 Ω over three months. CV cycle tests also were performed to evaluate the repeated charging stability of the PEDOT: BF4 functionalized liquid metal electrodes. We found little change in the form of the CV cycle over 1000 successive scans from new PEDOT functionalized liquid metal electrodes (Figure 4.11 D); Based on similar stability studies by other groups, the CV stability tests support the idea that such electrodes can provide stability needed for chronic biosensing application.65,66 Lastly, mechanical stress on the encapsulated Ga electrode may lead to delamination of the deposited PEDOT layer. To examine this, the mechanical stability of the PEDOT deposited on the liquid metal electrode was tested by a repeated bending cyclic test. The PEDOT functionalized liquid metal electrodes were subjected to 500 cycles of a positive and negative 60o bending (the image can be seen in Appendix A.2). EIS measurements were conducted over time in PBS buffer. The results (Figure 4.11 E) 90 indicated that impedance of PEDOT functionalized liquid metal electrodes was maintained during such repeated 120o multiple bending tests. Overall, we have conducted a series of analyses on in vitro stability to demonstrate our PEDOT deposition approach on liquid metal can maintain excellent stability over time and protect the liquid metal from oxidation and corrosion under simulated physiological conditions. 4.3.6 Electrochemical detection of dopamine (DA) in vitro Dopamine (DA) is a well-known neurotransmitter with key neurological significance. The dysregulation of the DA may cause neurologic disorders, such as Parkinson`s disease and ADHD. Hence the monitoring of DA activities is important for the diagnose and fundamental understanding of many neurological conditions. Since the monitoring of the DA neurotransmission requires fast response, high sensitivity, and chemical selectivity, DA sensing is an excellent model to evaluate the electrochemical performance of the biosensors. Au and PEDOT facilitate the DA oxidation to odopaminoquinone (DAQ) and reduction vice versa; thus, the nanostructured Au and PEDOT are widely used to improve DA sensing capabilities.67–73 Here, the sensitivity of the DA detection is measured to demonstrate the feasibility of the CNT/Au NPs and PEDOT structure on the liquid metal for biosensing applications. The liquid metal-based DA biosensor displayed distinct redox electrochemical reactions with 1 mM of DA (under 0.1M PBS) around 0.22 V (anodic) and 0.15 V (cathodic), corresponding to the DA oxidation and reduction, respectively (Figure 4.12 A and B). The anodic peaks of CV curves with the DA concentrations were plotted from 25 nM to 5 mM to show a continuous trend for CNT, CNT/Au NPs, and 91 PEDOT:BF4 structure (Figure 4.12 C). Although all nanostructures showed a linear trend with concentration up to 1 mM, CNT/Au NPs and PEDOT:BF4 nanostructures showed the non-linear response at larger concentrations due to diffusion limitation by excessive adsorption of the DA on CNT/Au NPs and PEDOT:BF4.74 Linearity was confirmed by linear fit over 0.99 of R-squared under the DA concentration of 1µM, which is often used in the previous studies.67,69,72,75,76 Figure 4.12 D highlighted the linear fit of all the nanocomposites to determine the LOD of the DA monitoring. Au NPs and PEDOT:BF4 functionalization improved linearity from 75 nM and 100 nM, comparing to the CNT layer itself (450 nM). The liquid metal-based DA sensors exhibited a sensitivity of 0.236 ± 0.013 µm/µM and a limit of detection (LOD) of 23.2 nM for CNT/Au NPs and 0.116 ± 0.022 µm/µM and a LOD of 79.6 nM for PEDOT:BF4. The LOD was determined by the S/N=3 criteria, where S is the standard deviation of the signal and N is the slope of the corresponding noise calibration curve.72,77 Both values are still compatible with the previous studies after calibration of the platform scale (from 25 to 10 µm), which is the scale of the general carbon fibers.69,72,75 Most literature determined that cerebral fluids consisted of 1-10 µM concentrations of DA and a plethora of ascorbic acid (AA) and uric acid (UA), which are byproducts of metabolism. The concentrations of AA and UA are 100-200 times higher than that of DA;40,67,78 the AA and UA generated anodic peaks on the CNT/Au NPs and PEDOT surface at 0.15 (under 1 mM for CNT/Au NPs) and 0.40 V, respectively (Figure 4.13 A and B). Since all the acids have different anodic voltages, the liquid metal-based microelectrodes would have selectivity under human fluids. Figure 4.13 A and B present the current responses from 10 µM DA are selected under a mixture including 200 times 92 AA (2 mM) and 100 times UA (1 mM). We confirmed the anodic peaks of the DA under a large concentration of AA and UA also have a linear trend in the range of 2 to 50 µM (Figure 4.13 C and D); the linear range is similar to the typical range of granular Au NPs and PEDOT shown in the previous studies.31,65,67,69 Both sensitivity and selectivity results suggest the CNT/Au NPs and PEDOT structure on the soft liquid metal-based platform could be used as an excellent DA biosensor. 4.3.7 Dopamine (DA) monitoring by fast scan cyclic voltammetry (FSCV) FSCV biosensing application should be considered because it can detect subseconds changes of biochemicals both in vitro or in vivo. Venton`s and other research groups established in vitro FSCV biosensing parameters for DA based on the carbon fiber platform; the setup has been widely used for in vivo DA sensors.41,42,68,79 However, there has been a need for a more tissue-compliant DA biosensing platform to reduce tissue damage.80,81 This study compared the FSCV DA detection performance in vitro of the CNT/Au NPs and PEDOT functionalized liquid metal electrodes to the conventional carbon fiber electrodes. Unfortunately, PEDOT functionalized liquid metal structure failed to obtain DA signals under FSCV conditions even the structure provided extremely low impedance, high capacitance, and low DA detection sensitivity from a general scan rate of 50 mV/s. The porous nanostructures deposited on top of the liquid metal surface hamper the transport of the electroactive ions to the liquid metal. The porous PEDOT:BF4 structure induces a tortuous ion diffusion pathway, resulting in the limitation of the fast electrochemical reaction that is generally required for real-time biosensing applications.14,82 Although we confirmed the PEDOT functionalized Ga wire can monitor 93 DA peak with 100 V/s, which is a quarter of the general FSCV scan rate, the structural inadequacy limits the DA sensing application of the PEDOT:BF4 structure (Figure 4.14 A). While the CNT/Au NPs nanocomposite provides nonporous adhesion between the CNT/PDDA interlayer and liquid metal surface that favors structure to transfer electroactive ions. Thus, we selected CNT/Au NPs nanocomposite as a suitable deposition strategy on the Ga-based liquid metal to use DA sensing application. FSCV scan results with the amounts of DA can be seen in Figure 4.14 B and C, which represent background-included (under ACSF) and background-subtracted CV scan, respectively. The current response of DA after background-subtracted was measured at 42 ± 2 nA for 1 µM DA (0.76 V) and 261 ± 13 nA for 10 µM DA (0.85 V). The sensitivity by the FSCV test was also obtained from the linear fit of the anodic peaks corresponding to the DA concentrations (Figure 4.14 D). The peaks showed a linear trend in the range between 25 and 200 nM, which was determined over R2=0.99. The linear fit of the DA peaks provided a sensitivity of 44.9 ± 0.51 nA/µM that is approximately 2-3 times higher than the sensitivities obtained from recent studies such as PEDOT:Nafion, nanodiamond, and Au NPs.42,79,83 The experiments were also repeated under 0.1M PBS (see Figure 4.14 E). Overall, the FSCV measurements demonstrated that the CNT/Au NPs functionalized liquid metal platform has great potential as a versatile biosensing platform. 4.4 Conclusion Although Ga-based liquid metal platform is an effective material strategy to fabricate soft bioelectronics, the biodegradation and limited electrochemical properties 94 have significantly hindered the biosensing applications of liquid metal. Various electrochemical deposition trials were performed to deposit gold, CNT, and PEDOT on the liquid metal surface with high electrochemical performance, biostability and biocompatibility. Gold deposition provided conductively and biocompatibility on the Gabased liquid metal, but poor biostability in physiologic conditions restricted the use for bioelectronics. CNT/PDDA nanocomposite enabled conductive and nonporous layer on the liquid metal surface to improve biostability and electrochemical performance. Positive-charged PDDA allowed the CNT deposition by a cathodic reaction to prevent hydrolytic oxidation during electrochemical deposition. PEDOT, a representative conducting polymer, was also studied to improve biostability, electrochemical property, and softness due to polymeric characteristics. To prevent hydrolytic oxidation during PEDOT polymerization, we used biocompatible organic solvent as an electrolyte and then deposited PEDOT:BF4 on the Ga surface that included high stability and electrochemical performance. Multilayer composite fabrication is also a suitable strategy to improve biostability and electrochemical performance. CNT/Au nanocomposite on the liquid metal from suggested electrochemical deposition methods was prepared for the soft DA biosensors. The study demonstrated the DA sensor showed a sensitivity of 0.236 ± 0.013 µm/µM and a LOD of 23.2 nM. The biosensor can also detect DA under 200 times AA and 100 times UA concentration in a physiological buffer, the typical concentration ratio in a physiologic environment. Finally, the FSCV test was performed under the scan rate of 400 V/s and the frequency of 10 Hz to demonstrate the fast detection capacity of this platform. The test provided the anodic peaks obtained from FSCV scans yield a linear 95 trend in the range between 25 and 200 nM with a sensitivity of 44.9 ± 0.51 nA/µM. All the results indicated the CNT/Au NPs functionalized liquid metal could have great potential as the next generation tissue compliant electrochemical biosensing platform. Gold nanoparticles can be deposited on the liquid metal to provide high conductivity, biocompatibility, and uniform electrical field on the Ga surface. 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Bode impedance of CNT/PSS functionalized H) Pt and I) liquid metal wires. 106 Figure 4.2. PDDA wrapping on the CNT surface. A) Chemical structure of PDDA on SWCNT. B) FT-IR and C) RAMAN spectra of SWNT and SWNT/PDDA composites with ratios. D) High resolution G band of RAMAN spectra SWNT/PDDA with the ratios. 107 Figure 4.3. Optimization of the CNT and polymer ratios. A) Bode impedance and B) CV curves of CNT/PDDA composites with ratios. C) RAMAN spectra of SWNT and SWNT/PSS composites with ratios. D) G band of RAMAN spectra of SWNT/PSS with ratios. 108 Figure 4.4. CNT/PDDA composite preparation. A) SEM images and B) magnified CNT/PDDA composite structure. C) Bode impedance and D) CV curves of CNT/PDDA composites with deposition time. 109 Figure 4.5. Optical confirmation of Ga surface oxidation after CNT/polymer composite deposition. A) BSE images of Ga wires after removing shell (PEBAX, by THF); after CNT/PSS (left) and CNT/PDDA (right) electrochemical deposition on the Ga surface. B) EDAX analysis of magnified Ga surface (side view for CNT/PSS and top-view for CNT/PDDA electrochemical deposition). White scale bar is 5 µm. Oxygen and Gallium displayed red and yellow, respectively. 110 Figure 4.6. Nyquist impedance after CNT/polymer electrochemical deposition. A) CNT/PSS and CNT/PDDA deposition on Pt surface (for 40 min). B) CNT/PDDA (Inlet is magnified impedance form B and C) CNT/PSS on Ga surface for 40 min. 111 Figure 4.7. CNT/Au NPs multilayer nanocomposite preparation. A) Illustration showing liquid metal wire and nanocomposites layer. B) SEM images of (a) bare Ga wire, (b) CNT/PDDA and then (c) Au NPs layer (Scale bar: 20 µm). (d) Magnified the surface of CNT/Au NPs, (c). C) SEM images of Au NPs deposition (a) for 5 min (magnified to b) and (c) for 10 min. 112 Figure 4.8. Electrochemical properties of nanostructure functionalized Ga electrodes. A) Bode impedance and B) CV curves with CNT and Au deposition time (inlet: magnified to observe CNT performance). C) SEM images of the CNT/Au nanostructure kept in PBS/Agarose gel at 37oC (scale bar: 20 µm) and D) the amounts of released Ga ions in the PBS gel for four weeks from Ga wire after CNT and CNT/Au functionalization. Heat and mechanical stabilities confirmation from 1 kHz impedance changes during E) heating and F) bending test for three months and for 500 cycles at 37oC in PBS (see Appendix A.2). G) electrochemical stability for 600 cycles with 1.0 V/s scan rate. 113 Figure 4.9. PEDOT:PSS on the Ga surface. A) SEM image after PEDOT:PSS electrochemical deposition on the Ga surface and B) magnified. C) Bode impedance and D) CV curves after PEDOT:PSS deposition on the Ga surface. E) SEM and EDAX analysis after PEDOT:PSS deposition on the Ga surface (white scale bar: 100 µm). Gallium, Sulfur, and Oxygen displayed yellow, blue, and red, respectively. 114 Figure 4.10. PEDOT:BF4 nanostructure preparation. A) SEM images of PEDOT functionalized Ga-based microelectrodes and magnified (right). B) Bode impedance, C) phase angle, and D) CV curves of PEDOT:BF4 on the Ga surface with electrochemical deposition time. 115 Figure 4.11. Physiologic stabilities of PEDOT:BF4 structure on the Ga surface. A) SEM images of PEDOT nanostructure kept in PBS at 37oC and B) the amounts of released Ga ions in the PBS for four weeks. Heat stability confirmation from 1kHz impedance changes during C) heating for three months at 37oC in PBS/Agarose gel. D) Electrochemical stability for 1000 cycles with 1.0 V/s scan rate. E) Mechanical stability confirmation from 1kHz impedance changes during 120o bending test with 500 cycles (see Appendix). 116 Figure 4.12. Dopamine (DA) monitoring by surface functionalized Ga wires. CV curves (scan rate 50 mV/s) of the A) CNT/Au and B) PEDOT functionalized Ga sensors with 1mM DA. C) Trend of anodic peaks as a function of DA concentration measured by CNT (black), CNT/Au NPs (red), and PEDOT (blue) nanocomposites. D) Linear fit result over R2=0.98 in the range of 25 nM to 1 µM DA concentration for calculating the sensitivity. 117 Figure 4.13. Dopamine (DA) monitoring for selectivity under physiologic hindrance. CV curves with 0.01 mM DA under 100-200 times amounts AA and UA of A) CNT/Au NPs and B) PEDOT:BF4 nanostructures on the Ga surface. Their linear fit results over R2=0.98 in the range of 50 nM to 50 µM DA concentration on the C) CNT/Au NPs and D) PEDOT:BF4 nanostructures to confirm DA selectivity. 118 Figure 4.14. Fast scan CV (FSCV) test results to confirm in vivo test availability. A) DA detection of PEDOT:BF4 structure with scan rate. Under 100 V/s scan rate can detect 5 µM DA. B) FSCV test results with the amounts of DA on the CNT/Au NPs B) with background and C) after background subtraction. D) A linear fit of anodic peaks with the amounts DA under 200 nM range to calculate the sensitivity from FSCV. E) FSCV test results under 0.1M PBS after background subtraction. 119 Table 4.1. Fitting results from Randles and Corrosion equivalent circuits in Nyquist plot (Figure 4.6) CNT/PDDA by Randles circuit CNT/PSS by Corrosion circuit Circuit model Compo. /Unit Rs 100 Rc 101 Cc 10-6 * Ω • cm2 Ω • cm2 F • cm-2 Ga wire 337.6 94.8 0.04 CNT 20 min 180.0 29.7 CNT 40 min 113.0 13.7 ** Σ𝑋𝑋 2 Rs 100 Ro 101 Rc 104 Ω • cm2 Co 10-6 * Σ𝑋𝑋 2 Ω • cm2 Ω • cm2 5.82 (364.7) - (109.4) - (0.04) 6.92 8.33 2.42 84.1 133.2 378.2 0.52 4.21 2.28 9.12 3.87 109.8 102.6 420.7 0.75 8.62 2.03 ×10-3 F • cm-2 Cc 10-8 F • cm-2 ×102 * : Sum of squares errors for all the parameters (calculated by ZView 3.5h software) ** : Parameters of Ga wires were obtained by only Randles circuit. CHAPTER 5 GALLIUM-BASED BIOMEDICAL APPLICATIONS AND THEIR PERFORMANCE IN VIVO 5.1 Introduction Gallium (Ga)-based liquid metals enable soft implantable bioelectronics to minimize the mechanical mismatch between electronics and brain tissues. Also, suggested various electrochemical deposition methods on the liquid metals provide higher electrochemical performance, biocompatibility, and biostability that are required for the chronic implantable bioelectronics.1–6 Our preliminary studies validated the electrochemical property and biostability of the electrochemically functionalized liquid metal microelectrodes by structural and impedance stabilization after thermal incubation, bending stability, and redox reaction stability.7,8 Especially, poly(3,4-ethylene dioxythiophene) (PEDOT) encapsulation is the most preferred method to fabricate entire soft bioelectronics in physiologic condition, which consist of flexible PEBAX shell, liquefied Ga, and soft polymer-based conductive material. In this study, we selected a biostable PEDOT layer from organic solvent electrolyte (propylene carbonate, PC) and tetraethylammonium tetrafluoroborate (TEABF4) dopant to perform in vivo physiologic signals recording.9,10 For the implantable bioelectronics, we performed action potential recording in a 121 nonhuman primate and invertebrate model to assess the feasibility of PEDOT functionalized liquid metal electrodes for high-performance bioelectronic applications. To our knowledge, this is the first demonstration of single-unit neural recording from liquid metal bioelectronic devices to date. First, acute neural recording experiments were conducted from a rhesus monkey (macaque mulatta). The recording results demonstrated PEDOT functionalized liquid metal electrodes can record single-unit action potential in nonhuman primates. Next, single-unit action potential spikes were recorded in an invertebrate model (Lumbricus Terrestris), then repeated for four weeks to evaluate the recording stability of the electrodes. The in vivo recording experiments confirmed PEDOT:BF4 deposition can improve the functionality of liquid metal electrodes for use in living tissue and allow stable recording performance in a complex physiological environment. Moreover, this study demonstrated the electrochemical deposition of PEDOT: BF4 can improve the material properties of liquid metal under physiological conditions, enable future development of liquid metal-based biosensors, and open numerous design opportunities for next-generation liquid metal-based bioelectronic applications. 5.2 Materials and Methods 5.2.1 Materials Liquefied Ga (99.995% purity, Sigma-Aldrich) or GaIn (≥99.99% trace metals basis, Sigma-Aldrich) was injected into the polyether block amide (PEBAX, Zeus Industrial Products, SC, USA) tubing. The liquid metal included PEBAX tubing was stretched by thermal drawing to fabricate microscale electrodes. 125 µm thickness 122 perfluoroalkoxy coated Pt wire (99.99% purity, A-M systems, WA, USA) and 200 µm tin-plated Cu wire were used as reference wires to compare recording performance. 5.2.2 Methods 5.2.2.1 Acute action potential recording from nonhuman primate An adult male rhesus monkey (macaque mulatta, 10 kg) was used in this study. All experimental procedures and protocols were compliant with guidelines from the University of Utah`s Institutional Animal Care and Use Committee (IACUC). Single unit action potential (AP) measurements were performed from the prefrontal cortex using PEDOT liquid metal-based electrodes (LMEs) and commercial tungsten electrode (200 µm diameter, tapered, 300Ω@1kHz, FHC, USA) was used to determine a neuron position first. The electrodes were positioned to neurons in the brain by a cannulaassisted setup and the fabrication method was introduced in Appendix A.3 (fabrication steps for the reference tungsten electrode setup were introduced in the reference.11 The physiologic signals were obtained from a microelectrode AC amplifier (Model 1800, A-M Systems, USA). Then the obtained signals were filtered by a neural signal processor (NSP, CerebusTM, Blackrock Microsystem, USA). The recording was performed for 8 min after signal stabilization. All the signals were analyzed by custom code written in Matlab (MathWorks). The signals were filtered from 0.1 (low-pass) to 2.5 kHz (high-pass) for the neural data acquisition and resampled at 1 kHz. Signal to noise ratio (SNR) was obtained from 8 min filtered data. Averaged signal and noise amplitudes are used to calculate SNR. Offline Spike Sorter (BOSS, v1.03.00, Blackrock Microsystems, USA) was utilized for principal component analysis (PCA) to sort 123 distinctive single-unit action potentials. 5.2.2.2 Craniotomy and surgical procedure The liquid metal wires were temporarily solidified in -25oC chiller before surgery to penetrate brain tissues. The PEDOT-based liquid metal electrode of 1.5 mm was exposed from the cannula to record single-unit APs by mechanical screwing work (see Appendix A.4). The penetration is generally conducted within 1 min, which is enough to ensure a solid-state of liquid metal. An animal was surgically implanted with a stainless steel headpost (Gray Matter Research, Bozeman, USA), attached to the skull using orthopedic titanium screws and dental acrylic. All surgical procedures were performed under anesthesia using Isoflurane and strict aseptic conditions. Stainless steel recording chambers (interior 19 mm diameter) were also mounted on the skull with screws and dental acrylic fixation. After the surgery, recording electrodes were lowed into the prefrontal cortex using a hydraulic Microdrive (Narishige, Japan). The neuron position was determined by conventional Tungsten electrode (FHC, USA) first then the PEDOT LME was positioned on the same site. The method was also specified in the reference.12 5.2.2.3 Long-term electrophysiological signals determination from invertebrate model The earthworms (Lumbricus Terrestris) were purchased from a local vendor (provided by DMF Bait Co., USA). The earthworms were stored at 4oC, and they moved to a room temperature laboratory an hour before the experiments. The earthworms were anesthetized in a 10% ethanol solution for 4 min then pinned on a corkboard for the 124 electrode implantation. The PEDOT LMEs were positioned to nerve cords of the worms using a needle "shuttle" to penetrate the skin and muscular wall of the worms. The appropriate position was determined by the shape of the signal, which displayed both positive and negative peaks within 1 ms.13 Electrical stimulation was selected to generate electrophysiologic response that is the method to apply and obtain the identical scale stimuli and action potentials, respectively. We previously performed the electrical stimulation to the earthworms by invasive and non-invasive (to the earthworms` skin). However, the multiple direct electrical stimulations to the earthworms caused deterioration of the physical condition, ultimately resulting in death. Hence, we designed indirect electrical stimulation through DI water bath and a single stimulus is composed of 100 mV for 50 ms, which can be seen in Appendix A.5. Both bare and PEDOT functionalized liquid metal electrodes were inserted into the earthworms weekly for action potential recording during a 4-week aging test. The electrodes were incubated in PBS/agarose gel at 37oC, removed from a heated PBS/agarose gel each week and reinserted into new and undamaged earthworms to guarantee identical physical performance and viability. We also confirmed the multiple insertions of the needle “shuttle” deteriorated the physical performance of the earthworms, like the multiple electrical stimulations; the weakened earthworms could generate attenuated action potential signals. Stimulus artifacts and action potentials were recorded by a toolkit for neural signal recording (The Neuron SpikerBox Pro, Backyard brains, MI, USA). Recording tests were carried out in a Faraday cage with a band-pass filter from 0.1 to 2.5 kHz and a ground electrode connected to the worm's tail with 50 Hz notch filter to minimize noise. 125 We calibrated the amplitude of the electrophysiological signals obtained from the recording toolkit to voltage difference using a two-electrode electrochemical system of SP-150 (Bio-Logic, France) with EC-Lab V11.10 software. Distinct action potential signal was obtained after 0.1 s from stimulus artifacts by electric stimulation. The experiments were repeated three times with independent earthworms. At least 100 electrical stimuli elicited the action potentials to the soft electrodes then each signal was aligned and averaged by Matlab (MathWorks, USA) and version 2.4.2 of Audacity (R) recording and editing software. Signal to noise ratio (SNR) was obtained from the equation below 𝑆𝑁𝑅 20 log 𝐴 ⁄𝐴 (1) where Asignal and Anoise are average amplitude from representative 50 signals and average noise, respectively. We obtained both information from the BYB Spike Recorder software (Backyard brains). In addition, the number of high-quality signals (SNR > 6) as a function of the electrode aging time was counted from 100 electrical stimuli to figure out time-dependent electrophysiology studies. 5.3 Results and Discussion 5.3.1 Acute single-unit action potential recording from nonhuman primate Once we demonstrated that the PEDOT:BF4 deposition could enhance electrochemical properties and stability of liquid metal electrodes in vitro, we conducted in vivo functionality tests of the liquid metal electrodes in the prefrontal cortex of one nonhuman primate. The ability to acutely record action potentials in awake behaving nonhuman primates is challenging in general. The setup usually requires the placement of 126 a chamber on the skull and removal of the bone inside the chamber. Thus, to record action potentials, one either needs a stiff electrode capable of penetrating dura and the granulation tissue overgrowth on top of it or an electrode small enough to be protected inside a cannula to serve as a shuttle piercing the dura. We chose the second option: to take advantage of our ability to manufacture long yet consistently narrow electrodes inside a cannula, similar to the idea previously used in the nonhuman primates literature.14,15 The customized arrangement enabled us to use the PEDOT functionalized liquid metal electrode to record the brain activity in an acute awake behaving nonhuman primate setup. We recorded the spiking activities from the PEDOT enable liquid metal electrode to confirm the acquisition ability of single-unit action potentials. The specific operation can be seen in Appendix A.4. The filtered (100-2,500 Hz) neural signals demonstrated the PEDOT functionalized liquid metal-based electrodes can record physiologic signals successfully, which are displayed in Figure 5.1 A. Piled waveforms from an 8-min acute recording (Figure 5.1 B) indicated that the recorded neural signals were generated from action potentials, which consist of 1kHz.16 The averaged waveform from the 8-min recording (Figure 5.1 C) represented the signal and noise levels to obtain the signal-to-noise ratio (SNR) of 21.8. Principal component analysis (PCA) showed the PEDOT functionalization on the liquid metal platform can detect various single-unit action potentials, which can be sorted by different time course and shape (Figure 5.1 D). The sorted waveforms (Figure 5.1 E) demonstrated that at least one unit displayed a distinctive single unit signal different from the noise signal (the blue one). Although the implanted electrode showed slightly degraded electrochemical properties after implantation (Figure 5.2 A-D), the acute recording test result demonstrated the PEDOT 127 functionalized liquid metal electrode can monitor single-unit action potentials with highquality SNR over 6. 5.3.2 Long-term single-unit action potential recording from invertebrate model The earthworm (Lumbricus Terrestris) has been used as an invertebrate model to record single-unit action potentials and investigate motor functionality with the liquid metal electrodes incubation time due to its large size easy-accessible giant nerves.17 Such a simple invertebrate neural model can significantly reduce the surgical complications and movement variables associated with vertebrate models. It allowed us to obtain consistent APs spikes in different experimental trials to examine recording performance in a simple model. The liquid metal electrodes were positioned adjacent to the giant nerve cords of the worms using a needle shuttle to penetrate the hard skin and muscular wall of the worm, which is an effective and well-developed implantation method for soft and flexible neural electrodes.18,19 The needle shuttle was removed from the worm's body after guiding the electrodes to the targeted giant nerve cord region (Figure 5.3 A, also can be seen in Appendix A.4). Electrical stimulation was used to apply a 100mV voltage pulse over 50ms to initiate consistent AP spikes from the giant nerve fibers (Figure 5.3 B). The signal artifacts from electrical stimulation and the corresponding single unit APs were recorded (Figure 5.3 C and can be seen in Appendix A.5). The entire recording data was sorted solely to identify the single unit action potential waveform (Figure 5.3 D). Both bare and PEDOT functionalized liquid metal electrodes were incubated in 37oC PBS/Agarose gel for four weeks. The electrodes were removed from a heated 128 PBS/agarose gel each week and inserted into the earthworms for action potential recording. Action potential recording for each type of electrode was repeated three times with undamaged worms. Average spike signals were obtained from all the trials. Corresponding AP spikes (left) were aligned and averaged (right) as shown in Figure 5.4 to illustrate recording consistency and compare spike amplitude, respectively. Both liquid metal electrodes with and without PEDOT functionalization could record AP spikes after initial implantation. However, the recording performance of the electrodes showed significant differences with incubation time. The bare liquid metal electrodes generated high noise after a week, then failed to record signals three weeks after implantation. In addition to analyzing spike amplitude, we also calculated the signal-to-noise ratio (SNR) of the recorded data and identified high-quality recorded spikes (SNR over 6) over 100 stimulation trials (Figure 5.5 A). We compared the numbers of high-quality recording spikes among bare liquid metal, PEDOT functional liquid metal, and platinum (Pt) electrodes. The PEDOT functionalized liquid metal electrodes showed a 94% highquality recording performance after four weeks, while bare liquid electrodes could not assess high-quality recording after the initial implantation. Moreover, the recording performance of the PEDOT functionalized liquid metal also was higher than Pt electrodes (see Figure 5.5 B). Overall, the long-term recording study from the invertebrate model confirmed PEDOT:BF4 deposition can allow stable recording performance in a complex physiological environment over time. 129 5.3.3 Discussion for the action potential recording Acute action potential recording measurements in awake behaving nonhuman primates showed stable and high-quality single-unit neural recording, which can be seen in Figure 5.1. This is the first demonstration of single-unit neural recording in awake behaving nonhuman primates using Gallium-based liquid metal bioelectronic devices to date. PCA study isolated each single-unit action potential and demonstrated at least one action potential was recorded simultaneously. The waveform of each neuron that can be seen in Figure 5.1 E is the typical waveform of APs verified by previous studies.5,20–23 The high SNR (21.8) suggests the liquid metal electrodes with a biostable conductive polymer interface can enable high-performance implantable bioelectronics. Repeated recording experiments from aged liquid metal electrodes were used to demonstrate the stability of electrodes, as shown in Figure 5.4. Although liquid metal surface itself could record action potentials first, the surface lost the recording function after two-week incubation due to surface oxidation and corrosion. Surface images and EDAX analysis showed severe oxidation and impurities on the bare liquid metal surface after implantation (Figure 5.6). Copper-based electrode, which is from a representative non-noble metal, showed similar performance degradation under the same test condition (Figure 5.5 B). We conclude non-noble metals are unsuitable to use bioelectronics without surface treatment (Figure 5.5 A and B). Post-implantation analysis showed biofouling on the electrode surface (Figures 5.2 and 5.6). Although the conductive polymer functionalization could maintain high-quality recording performance after four-week incubation with an SNR of 12.2±1.2 (Initial: 22.7±0.4), the recording performance was decreased to 65% from initial SNR within 130 initial 2-week incubation, which may be seen in Figure 5.4. We confirmed PEDOT crystallinity was decreased with incubation due to the removal of residual solvent (Figure 5.7). This also affected the performance decrease as well as biofouling on the PEDOT surface. In summary, this study demonstrated that the electrochemical deposition of PEDOT: BF4 improves the material chemistry of liquid metal under physiological conditions, an approach that can enable future development of liquid metal-based biosensors and open up numerous design opportunities for next-generation liquid metalbased bioelectronics. 5.4 Conclusion Acute and four-week single-unit action potential recording results indicate the feasibility of the chronic action potential recording from a liquid metal platform. This is the first demonstration of single-unit neural recording from liquid metal bioelectronic devices to date. The biostable PEDOT:BF4 functionalization led to an SNR of 21.8 from nonhuman primates, which is higher than thick and stiff conventional Tungsten electrodes (SNR=14.3). Although we confirmed the SNR from the invertebrate model was decreased after four weeks (repeated insertion), the SNR was stabilized to 12.2, which is high-quality recording performance (over 6). The combination of liquid metal platform and soft PEDOT encapsulation helps minimize the mechanical mismatch between the electrodes and fragile brain tissue, hence the PEDOT liquid metal electrodes have great potential for chronic bioelectronics. 131 5.5 References (1) Kozai, T. D. Y.; Langhals, N. B.; Patel, P. R.; Deng, X.; Zhang, H.; Smith, K. L.; Lahann, J.; Kotov, N. A.; Kipke, D. R. Ultrasmall Implantable Composite Microelectrodes with Bioactive Surfaces for Chronic Neural Interfaces. Nat. Mater. 2012, 11 (12), 1065–1073. https://doi.org/10.1038/nmat3468. (2) Deng, J.; Yuk, H.; Wu, J.; Varela, C. E.; Chen, X.; Roche, E. T.; Guo, C. F.; Zhao, X. Electrical Bioadhesive Interface for Bioelectronics. Nat. Mater. 2021, 20 (2), 229–236. https://doi.org/10.1038/s41563-020-00814-2. 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Magnetic Actuation of Flexible Microelectrode Arrays for Neural Activity Recordings. Nano Lett. 2019, 19 (11), 8032–8039. https://doi.org/10.1021/acs.nanolett.9b03232. (7) Lim, T.; Kim, H. J.; Zhang, H.; Lee, S. Screen-Printed Conductive Pattern on Spandex for Stretchable Electronic Textiles. Smart Mater. Struct. 2021, 30 (7), 075006. https://doi.org/10.1088/1361-665X/abfb7f. (8) Lim, T. H.; Yeo, S. Y.; Lee, S. H. Multidirectional Evaluations of a Carbon Air Filter to Verify Their Lifespan and Various Performances. J. Aerosol Sci. 2018, 126 (September), 205–216. https://doi.org/10.1016/j.jaerosci.2018.09.009. (9) Malmström, J.; Nieuwoudt, M. K.; Strover, L. T.; Hackett, A.; Laita, O.; Brimble, M. A.; Williams, D. E.; Travas-Sejdic, J. Grafting from Poly(3,4Ethylenedioxythiophene): A Simple Route to Versatile Electrically Addressable Surfaces. Macromolecules 2013, 46 (12), 4955–4965. https://doi.org/10.1021/ma400803j. 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Science (80-. ). 1971, 173 (3997), 652–654. https://doi.org/10.1126/science.173.3997.652. (22) Fan, D.; Rossi, M. A.; Yin, H. H. Mechanisms of Action Selection and Timing in Substantia Nigra Neurons. J. Neurosci. 2012, 32 (16), 5534–5548. https://doi.org/10.1523/JNEUROSCI.5924-11.2012. (23) Lu, C.; Park, S.; Richner, T. J.; Derry, A.; Brown, I.; Hou, C.; Rao, S.; Kang, J.; Mortiz, C. T.; Fink, Y.; Anikeeva, P. Flexible and Stretchable Nanowire-Coated Fibers for Optoelectronic Probing of Spinal Cord Circuits. Sci. Adv. 2017, 3 (3), e1600955. https://doi.org/10.1126/sciadv.1600955. 134 Figure 5.1. Single-unit action potential recording from nonhuman primates. A) Highpass filtered (100-2,500 Hz) action potentials (APs) signals with schematic illustrations showing the concept to obtain APs from PEDOT LMEs. B) Piled and C) averaged singleunit APs from 8 min successive recordings. D) results from principal component analysis (PCA) showing distinct clusters from E) different shapes of waveforms (blue: noise). 135 Figure 5.2. Electrochemical properties of PEDOT LMEs after implantation (red). A) Bode impedance, B) phase angle, and C) CV curves. D) SEM images after implantation (after 10 min recording). 136 Figure 5.3. Schematic illustrations showing the method to obtain action potentials (APs). A) Images showing the experimental setup. B) Raw data of signals generated from electrical stimulation (100 mV for 50 ms). C) Magnified each signal (red square in B). D) magnified APs (red square in C). 137 Figure 5.4. Repeated action potentials (APs) recordings from invertebrate model. Filtered AP signals excluding electrical stimulation (left), and aligned and averaged spikes from 50 single units with aging period (right). 138 Figure 5.5. Action potential recording stability from PEDOT LMEs. A) The numbers of high-quality recording spikes (SNR > 6) of bare liquid metal wires, Pt wires, and PEDOT LMEs for four weeks. B) Action potentials signals excluding electrostimulation of typical metal wires, tin plated Cu wires (left) and Pt wires (right) with aging periods. 139 Figure 5.6. Post-implantation surface images. SEM and EDAX analysis of the bare liquid metal wires and PEDOT LMEs after aging in agarose (at body temperature) and implanted in live earthworms. 140 Figure 5.7. Validation of the initial recording performance attenuation. FT-IR studies in the range of A) 900-1300, and B) 1500-1900 cm-1 to confirm propylene carbonate residue inside PEDOT structure. C) XRD study to compare crystallinity of PEDOT:BF4 with incubation and annealing. CHAPTER 6 CONCLUSIONS AND FUTURE DIRECTIONS 6.1 Conclusions Failure of implantable electronic devices imposes tremendous challenges for the future development of chronically sustainable bioelectronic/tissue interfaces. The design of the electronics for implantable applications needs a transition from the generally rigid and static architectures to softer but sturdy adhesion among the assembly layers. Implementation of these features needs novel materials selection and fabrication strategies. Ga is introduced in this dissertation to address soft and tissue-compliant microelectrodes. The metal`s unique melting point between room and body temperature enables thermal responsive characteristics after implantation. The liquefied Ga can provide no mechanical stiffness that minimizes tissue inflammation and scar formation in the vicinity of the bioelectronics. Although extensive studies have been performed recently to use Ga-based liquid metals for wearable bioelectronics, intensive studies are required for the implantable electrodes such as surface chemistry, material stability, biocompatibility, and electrochemical performance in physiologic conditions. Therefore, we prepared Ga-based microelectrodes then investigated specific mechanical, chemical, physical, and biological aspects step by step. In Chapter 2, we present Ga/PEBAX core/shell structure preparation expressing 142 thermal responsive features. Optimal shell material selection enabled to have biocompatibility, biostability and softness in physiologic conditions. Through mechanical analysis, the stretched and formulated Ga/PEBAX core/shell structure exhibited extremely low Young`s modulus (0.3 MPa), which is similar to or less than conventional hydrogel structure. The softness was demonstrated through thermal analysis by DSC scan. Decreased enthalpy of melting with stretching confirmed lack of crystallinity that indicates hard segments of PEBAX were attenuated by incomplete stretching process. The thickness and ratio of core/shell structure were optimized to meet requirements for use as an implantable electrode, minimizing disruption and penetrating brain surface at the same time. We obtained the ratio (core/shell) of 4 that is the maximum to date and then selected 60-80 µm thickness through in vitro and ex vivo penetration simulation. In Chapter 3, we present Ga surface chemistry in physiologic buffer with incubation at body temperature. Kinetic and elemental studies elucidated that Ga ions (Ga3+) and OH- were released, leading to a pH increase. Although the chemical reaction triggered further oxidation than in air, the reaction between Ga and physiologic buffer was also limited after 5 days; hence, we investigated the reason by various spectroscopic measurements expressing surface chemistry change. FT-IR study verified the GaOOHrich formation on the surface within six-hour incubation then ionized to Ga3+ and OH-. The composition change was stabilized after four days, which trend is similar to kinetic study results. XPS analysis suggested surface contamination on the Ga oxide layer by a plethora of ions and organic materials in the physiologic buffer. Lastly, the RAMAN study demonstrated GaOOH dominated the Ga surface within two days then the contamination layer grew on the Ga oxide surface over micron scale. The oxidation and 143 contamination attenuated the electrochemical performance of Ga that was confirmed by impedance and CV measurements. Nyquist plot analysis indicated the typical electrochemical circuit changed to corrosion circuit with incubation. Also, the anodic onset of CV curves was increased with incubation that demonstrates a reactivity decrease. Overall, the oxidation and contamination pathway of the Ga surface in physiologic conditions limited electrochemical performance, thereby surface modification should be considered to use Ga for implantable bioelectronics. In Chapter 4, we present various surface modification strategies to overcome the poor biostability of Ga. A main focus of the surface modification is to improve electrochemical performance and sturdy encapsulation on the Ga surface. The electrochemical deposition was used to modify the Ga surface, and three different materials, gold, CNT, and PEDOT, were selected for high-performance and biostable features. Although gold provided high electrochemical performance, poor biostability restricted the use solely for implantable bioelectronics. CNT was deposited with a viscous polymeric material, PDDA, to improve the biostability of the encapsulation. PDDA was ionized in an aqueous electrolyte, showing a positive charge that enabled to deposit on the Ga surface by negative potential without hydrolytic oxidation. Despite the modified biostability after CNT/PDDA encapsulation, unsatisfactory electrochemical property improvement may limit the use of high-performance electrochemical biosensors. PEDOT was also considered to exhibit high biostability, electrochemical properties and ultimately maintain softness due to polymeric characteristics. To prevent hydrolytic oxidation during PEDOT polymerization as well as deposition, we selected biocompatible organic solvent and then deposited PEDOT:BF4 on the Ga surface that 144 included high stability and electrochemical performance. Electrochemical dopamine (DA) sensing was performed by CNT/Au multilayer nanocomposite for biomedical device application. CNT/PDDA was first encapsulated on the Ga-based liquid metal surface then Au nanoparticles were decorated on the CNT layer. The in vitro sensing study demonstrated the CNT/Au multilayer nanocomposite displayed compatible sensitivity (0.236 ± 0.013 µm/µM), LOD (23.2 nM), and selectivity under excessive hindrance materials, mimicking physiologic fluids. Finally, the nonporous CNT/PDDA interlayer and high-performance Au NPs enabled DA detection by FSCV test performed under the scan rate of 400 V/s and the frequency of 10 Hz. All the results indicated the CNT/Au NPs functionalized liquid metal could have great potential as the next generation tissue compliant electrochemical biosensing platform. Gold nanoparticles deposition by reduction process can also be removed presented oxide layer at the same time. The function therefore enabled to fabricate stretchable encapsulation on the liquid metal surface. An instantaneously formed solid Ga oxide layer can lead to a wrinkled liquid metal surface suitable for stretching deformation. The oxide layer can be replaced with Au NPs, maintaining a wrinkled structure. Poor biostability of gold nanolayer can be overcome by PEDOT:BF4 deposition on the Au NPs, hence multilayer nanocomposite deposition strategy is useful for wearable and stretchable bioelectronics. Last Chapter 5 presents various in vivo physiologic signals recording, showing feasibility for chronically implantable bioelectronics. Acute and four-week single-unit action potential recording results verified the feasibility of the chronic action potential recording from a liquid metal platform. This is the first validation of single-unit neural 145 recording from liquid metal bioelectronic devices to date. The biostable PEDOT:BF4 functionalization obtained an SNR of 21.8 from nonhuman primates, which is higher than the conventional tungsten electrode (SNR=14.3). Although the SNR obtained from the invertebrate model was decreased to 65% of initial SNR after 4-week incubation, the stabilized SNR (12.2) showed high-quality recording performance (over 6). The liquid metal platform and soft PEDOT encapsulation help to minimize the mechanical mismatch between the electrodes and fragile brain tissue, the PEDOT liquid metal electrodes therefore have great potential for chronic bioelectronics. 6.2 Future Directions This dissertation illustrated our materials-based approaches to manufacture Gabased bioelectronics, expression thermal responsive mechanical behavior, and biocompatible characteristics. To use implantable bioelectronics, we have validated the characteristics of Ga in the physiologic conditions through elemental, kinetic, optical, spectroscopical, and electrochemical measurements. The poor biostability of Ga in physiologic conditions overcame by various encapsulation methods from electrochemical deposition. Although we confirmed the recording performance of surface-modified Ga electrodes in vivo and in vitro, further studies will be required to demonstrate electrochemical performance stability, biocompatibility, and less inflammation to prepare chronically implantable bioelectronics. Detailed steps to achieve these goals will be described here. In Chapter 2 of this dissertation, Ga/PEBAX core/shell structure was introduced to address the thermal responsive concept, including stiffness to penetrate brain surface 146 and softness to minimize the mechanical mismatch between the electrode and fragile brain tissue. The Young`s modulus was minimized to 0.3 MPa, which is the lowest value to date from a core/shell ratio of 4 with 60 µm thickness. Despite the effect to minimize Young`s modulus, which is similar to hydrogel and less than other materials, it still displays a gap to address the softness of brain tissue (Figure 6.1). Liquefied Ga has theoretically no mechanical stiffness, we can therefore reduce existing Ga-based electrodes by 1) higher ratio between core Ga and shell material, 2) new shell material involving higher thermal drawing property, and 3) higher softness than rubber materials such as hydrogel-like polymers. The shell material optimization process would accompany the drawing property by DSC study and crystallinity conformation by XRD scan for mechanical property determination. In Chapter 5 of this dissertation, now in vivo test by 18-month-old rats is performing to demonstrate histological advantage due to softening mechanical property as well as recording performance stability. The thermal responsive implantable microelectrodes should exhibit the excellent potential to minimize mechanical discrepancy, resulting in less inflammation and scar formation around the electrodes. A microelectrodes array was prepared by multiple steps, which can be seen in Figure 6.2. Each microelectrode has a slanted structure to improve penetration performance with the miniaturized thickness (less than 60 µm of Ga portion) and six-microelectrodes (4: PEDOT encapsulated Ga wire, 2: Pt bare wire as reference) were stagger-aligned on the PCB array to enhance the possibility of neural detection. Also, the prepared microelectrodes array exhibited excellent penetration ability under 1.0% agarose in vitro test (Figure 6.3 A). Each electrode has an impedance of 103 ohm and outstanding 147 capacitance due to nanoporous structure (Figure 6.3 B); those are enough to record single-unit action potentials from in vivo trial, which is confirmed by nonhuman primates test. The arrays were sterilized by cold 70% ethanol to prevent mechanical degradation by typical ethylene oxide (ETO) treatment under high temperatures. The implanted microelectrode array is fixed by medical cement with a skull to ensure implantation stability. The test will be proceeded for three-month to obtain histologic change with implanted periods, supported by Dr. Tresco`s previous strategies1–4. Also, long-term single-unit action potential recording is being accompanied by the same implantation period. This may be achieved within this year and provide insight into the benefit of thermal responsive microelectrode for chronical use. This will be the first demonstration of long-term single-unit neural recording and histologic study from liquid metal-based bioelectronic devices. 148 6.3 References (1) Meng, F.; Hlady, V.; Tresco, P. A. Inducing Alignment in Astrocyte Tissue Constructs by Surface Ligands Patterned on Biomaterials. Biomaterials 2012, 33 (5), 1323–1335. https://doi.org/10.1016/j.biomaterials.2011.10.034. (2) Polikov, V. S.; Tresco, P. A.; Reichert, W. M. Response of Brain Tissue to Chronically Implanted Neural Electrodes. J. Neurosci. Methods 2005, 148 (1), 1– 18. https://doi.org/10.1016/j.jneumeth.2005.08.015. (3) Biran, R.; Martin, D. C.; Tresco, P. A. Neuronal Cell Loss Accompanies the Brain Tissue Response to Chronically Implanted Silicon Microelectrode Arrays. Exp. Neurol. 2005, 195 (1), 115–126. https://doi.org/10.1016/j.expneurol.2005.04.020. (4) Nolta, N. F.; Christensen, M. B.; Crane, P. D.; Skousen, J. L.; Tresco, P. A. BBB Leakage, Astrogliosis, and Tissue Loss Correlate with Silicon Microelectrode Array Recording Performance. Biomaterials 2015, 53, 753–762. https://doi.org/10.1016/j.biomaterials.2015.02.081. 149 Figure 6.1. Schematic illustration showing Young`s modulus of bioelectrode materials. 150 Figure 6.2. PEDOT functionalized microelectrodes array fabrication step. 151 Figure 6.3. PEDOT functionalized microelectrodes array performance. A) Penetration (a) before and (b) after implantation. Electrochemical performance: B) bode impedance, C) Phase, D) CV, and E) CV with PEDOT deposition time. APPENDIX A.1 Acknowledgement Thanks to Dr. Diego Pernandez (Geology and Geophysics, University of Utah) for ICP-MS analysis, Dr. Paulo Perez (Nanofab, University of Utah) for XPS analysis, Kimberley Watts (Material Characterization Lab, Material Science Engineering, University of Utah) for SEM, FT-IR, and XRD use, Dr. Thang Tran (Manager of chemical engineering senior Lab, University of Utah) for FT-IR, DSC, critical dryer use, Eddie Polanco (Chemical Engineering, University of Utah) to assist CNC drill use, Dixon Laser institute, Nanofab, and Cruz center to figure out confocal RAMAN and SEM use, Minju Kim, Kanghun Choi and Jungkyu Kim (Mechanical Engineering, University of Utah) to assist action potential recording form invertebrate model and PDMS based EMG sensor fabrication, Dr. Behrad Noudoost and Amir Akbarian (Department of Ophthalmology and Visual Sciences, University of Utah) to record action potentials from nonhuman primates, Hunter Strathman (Biomedical Engineering, University of Utah) to help PCB designing and fabrication. Thanks to new laboratory members, Seoyeon Won and Qian for assisting my future work to fabricate physiologic signal monitoring and cytotoxicity test. 153 A.2 Electrochemical Properties Measurement for Liquefied Sample Liquefied Ga can be leaked from the microscale (over 200 µm) PEBAX tubing by squeezing of stretched PEBAX tubing and gravity of Ga metal. We prepared the electrochemical analysis setup to prevent the leakage through the Ga surface to upside can be seen in Figure A.1. A.3 Bending Stability Confirmation of The Liquified Ga Wires A 500-time of 120o bending test was carried out using a tensile tester (MARK-10, NY, USA). The Ga wires were liquefied in DI water at 37oC. Then the tests were performed showing in Figure A.2. A.4 PEDOT Liquid Metal-based Electrodes (PEDOT LMEs) Preparation for Acute Action Potential Recording from Nonhuman Primate To penetrate duraskin on the brain surface of the rhesus monkey, we prepared cannula-assisted PEDOT liquid metal based electrodes (PEDOT LMEs) in this study. Liquid metal wire placed in the Parylene-insulated cannula (Vitaneedle, stainless steel, customized Parylene coating) to prevent “short” with PEDOT LMEs. Fabrication steps for conventional tungsten wire based electrodes were specified in J Vis Exp 2019 153 e60365. The brief fabrication steps of PEDOT LMEs were shown in Figure A.3. The PEDOT LMEs can be protruded by ferrule rotating after penetrating stiff duraskin attached on the nonhuman primates` brain surface. 154 A.5 PEDOT liquid metal based electrodes (PEDOT LMEs) preparation for long-term recordings from invertebrate model The PEDOT LMEs were positioned to nerve cords of the worms using a needle "shuttle" to penetrate the skin and muscular wall of the worms, which can be seen in Figure A.4. Also, we selected an electrical stimulation to generate electrophysiologic response that is the method to apply and obtain the identical scale stimuli and action potentials, respectively. We previously performed the electrical stimulation to the earthworms by invasive and non-invasive (to the earthworms` skin). However, the multiple direct electrical stimulations to the earthworms caused deterioration of the physical condition, ultimately resulting in the death. Hence, we designed indirect electrical stimulation through DI water bath and a single stimulus is composed of 100 mV for 50 ms, and we may see the action potential corresponded electrical stimulations in Figure A.5. Both bare and PEDOT functionalized liquid metal electrodes were inserted into the earthworms weekly for action potential recording during a 4-week aging test. The electrodes were incubated in PBS/agarose gel at 37oC, removed from a heated PBS/agarose gel each week and reinserted into new and undamaged earthworms to guarantee identical physical performance and viability. We also confirmed the multiple insertions of the needle “shuttle” deteriorated the physical performance of the earthworms, like the multiple electrical stimulations; the weakened earthworms could generate attenuated action potential signals. 155 Figure A.1. Electrochemical analysis setup in DMEM buffer. 156 Figure A.2. 120o bending deformation test setup. 157 Figure A.3. Cannula assisted PEDOT LMEs setup. 158 Figure A.4. Action potential recording setup from invertebrate model. 159 Figure A.5. Physiologic electrical signals set including electrical stimulation artifacts (red) and corresponded action potentials (blue).