A Resonant Capacitive Test Structure for Biomolecule Sensing

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A RESONANT CAPACITIVE TEST STRUCTURE FOR BIOMOLECULE SENSING

Thesis

Submitted to

The School of Engineering of the

UNIVERSITY OF DAYTON

In Partial Fulﬁllment of the Requirements for

The Degree of

Master of Science in Electrical Engineering

Danielle Nichole Bane

UNIVERSITY OF DAYTON

Dayton, Ohio

August, 2015 A RESONANT CAPACITIVE TEST STRUCTURE FOR BIOMOLECULE SENSING

Name: Bane, Danielle Nichole

APPROVED BY:

Guru Subramanyam, Ph.D. Karolyn M. Hansen, Ph.D. Advisor Committee Chairman Advisor Committee Member Professor and Chair, Department of Professor, Department of Biology Electrical and Computer Engineering

Partha Banerjee, Ph.D Committee Member Professor, Department of Electrical and Computer Engineering

John G. Weber, Ph.D. Eddy M. Rojas, Ph.D., M.A., P.E. Associate Dean Dean School of Engineering School of Engineering

ii c Copyright by

Danielle Nichole Bane

2015 ABSTRACT

A RESONANT CAPACITIVE TEST STRUCTURE FOR BIOMOLECULE SENSING

Name: Bane, Danielle Nichole University of Dayton

Advisors: Dr. Guru Subramanyam and Dr. Karolyn M. Hansen

Detection of biomolecules in aqueous or vapor phase is a valuable metric in the assessment of health and human performance. For this purpose, resonant capacitive sensors are designed and fabricated. The sensor platform used is a resonant test structure (RTS) with a molecular recognition element (MRE) functionalized guanine dielectric layer used as the sensing layer. The sensors are designed such that the selective binding of the biomarkers of interest with the MREs is expected to cause a shift in the test structure’s resonant frequency, amplitude, and phase thereby indicating the biomarker’s presence. This thesis covers several aspects of the design and development of these biosensors. Guanine biopolymer is characterized using capacitive test structures (CTSs) and RTSs with guanine dielectric layers. From this characterization, the dielectric constant and loss tangent of guanine are found to be 5.345 ± 0.294 and 0.015 ± 0.001 respectively. The resonance of the

RTS with guanine dielectric layer is 3.148 ± 0.079 GHz with a notch depth of 7.472 ± 0.330 dB.

To further characterize guanine, contact angle measurements with water were performed to determine the hydrophobic/hydrophilic properties. The contact angle is 62.07◦ ± 3.029◦ indicating the

guanine thin ﬁlms are slightly hydrophobic in comparison to glass (contact angle is 41.4◦ ± 2.72◦).

Additionally, a chemical functionalization method for guanine is developed. In this method, a cross- linker is simultaneously and covalently bound to the surface of the guanine and to the biomolecule iii thereby creating a covalent tether. Tests employing a biotin-streptavidin model indicate the chemical functionalization method is viable.

In addition to the resonant capacitive sensor, two radio frequency (RF) test structures are developed: an RF bridge and a half-wavelength resonator. Both of these test structures have gaps in the transmission lines that will be bridged with MRE functionalized carbon nanotubes (CNTs). Any binding event between an analyte of interest and MRE is expected to cause a shift in the resonance of the test structures. The test structures are designed for a resonance at 8 GHz and are simulated in

Applied Wave Research Design Environment (AWR) software. For the RF bridge, the simulation’s resonance is at 9.2 GHz with a notch depth of 20.31 dB. The simulation of the half-wavelength resonator shows a resonance at 8 GHz with a pass band peak of 0.3395 dB. Additionally, the RF bridge test structure is fabricated. Measurements pre-CNT integration reveal a resonance at 13.23

GHz with a notch depth of 29.42 dB. The resonance of the fabricated test structures post-CNT integration have varying resonances (16.43 GHz with a notch depth of 31.2 dB for one test structure and 13.58 GHz with a notch depth of 30.1 dB for the other). The variation is due to non-uniform

CNT bridges across the transmission line.

iv Dedicated to my parents, James and Beth Bane

v ACKNOWLEDGMENTS

I would like to express my gratitude to all the people whose support and encouragement has helped me complete my master’s thesis. First, I would like to thank my advisors Dr. Guru Sub- ramanyam and Dr. Karolyn M. Hansen. Throughout my time as an undergraduate and graduate student, Dr. Subramanyam has been advising me and has taught me much about microelectronics.

When I met Dr. Hansen at the beginning of my time as a graduate student, she took me under her wing and taught me a great deal of biology. I am very grateful to both my advisors for their patience, expertise, and guidance. Additionally, I want to thank Dr. Banerjee for agreeing to be a part of my thesis committee.

I am appreciative of the Dayton Area Graduate Studies Institute (DAGSI) Fellowship and the

Graduate Student Summer Fellowship programs for providing funding for my thesis. I am very grateful to Dr. Elizabeth Downie, the director of DAGSI, and Dr. Joshua Hagen, my DAGSI sponsor, for giving me this opportunity.

I would like to give a special thanks to Dr. Fahima Ouchen and Dr. James Grote at the Air Force

Research Lab’s Material’s Directorate. Dr. Ouchen graciously opened her lab to me, helped in the fabrication of the test structures, and taught me much about biopolymers. I would like to thank Dr.

Eunsung Shin for doing the metal deposition for the structures. I would like to thank Hailing Yue for teaching me how to perform measurements with the Vector Network Analyzer, teaching me how to use the Applied Wave Research Design Environment software, and answering any questions I had. I would like to thank Kuan-Chang Pan for his help. I would like to thank Kaushik Annam for guiding me in the carbon nanotube alignment. I would like to thank Dr. Jayne Robinson and Caitlin

vi Bojanowski for granting me access to and helping me use the Epi-Fluorescent Microscope. Also, I would like to thank Dr. Charles Browning and Dr. Li Cao for allowing me to use the Kruss Drop

Shape Analysis System 100.

I would also like to thank my dogs Bella, Lotus, and Summer for our sessions of dog therapy.

Lastly, I would like to thank my parents, James and Beth Bane, for all their love and support.

Their constant encouragement and occasional prodding have driven me to complete my thesis.

vii TABLE OF CONTENTS

ABSTRACT...... iii

DEDICATION...... v

ACKNOWLEDGMENTS...... vi

LIST OF FIGURES...... xi

LIST OF TABLES...... xiii

LIST OF ABBREVIATIONS AND NOTATIONS...... xiv

I. INTRODUCTION...... 1

1.1 Objectives...... 2 1.2 Signiﬁcance...... 3 1.3 Outline...... 4

II. BACKGROUND...... 5

2.1 Capacitive Sensors...... 5 2.2 Resonant Sensors...... 6 2.3 Carbon Nanotubes (CNTs)...... 7 2.4 Biopolymers...... 7 2.4.1 Silk...... 8 2.4.2 Deoxyribonucleic Acid-Cetyltrimethylammonium Chloride (DNA-CTMA)8 2.4.3 Guanine...... 9 2.5 Proteins...... 10 2.5.1 Structure...... 10 2.5.2 Protein-Based Sensors...... 12 2.6 Biomolecule Immobilization...... 13 2.7 Model Analytes...... 14 2.7.1 Orexin A...... 14 2.7.2 Trimethylamine...... 15 2.8 Summary...... 16 viii III. TEST STRUCTURES...... 18

3.1 Capacitive Test Structure (CTS)...... 18 3.2 Resonant Test Structure (RTS)...... 19 3.3 Electrical Model...... 21 3.4 RF Bridge...... 23 3.5 Half-Wavelength Resonator...... 24 3.6 Summary...... 25

IV. EXPERIMENTAL PROCEDURE...... 26

4.1 Test Structure Analysis...... 26 4.2 Contact Angle Measurements...... 27 4.3 Biotin-Streptavidin Model...... 28 4.4 CNT Bridge Formation...... 31 4.5 Summary...... 32

V. RESULTS AND DISCUSSION...... 33

5.1 CTS Results...... 33 5.2 RTS Results...... 35 5.3 Contact Angle Measurement Results...... 37 5.4 Biotin-Streptavidin Model Results...... 37 5.5 RF Bridge Simulation...... 40 5.6 Fabricated RF Bridge Test Structure Results...... 42 5.7 Half-Wavelength Resonator Simulation...... 45 5.8 Summary...... 46

VI. CONCLUSIONS AND FUTURE WORK...... 47

6.1 Resonant Capacitive Sensor...... 47 6.2 RF Bridge and Half-Wavelength Resonator...... 50

BIBLIOGRAPHY...... 51

Appendices:

A. CTS MEASUREMENTS...... 57

B. RTS MEASUREMENTS...... 58

ix C. CONTACT ANGLE MEASUREMENTS...... 59

D. BUFFER AND CHEMICAL PROTOCOLS...... 60

x LIST OF FIGURES

2.1 Structure of guanine...... 9

2.2 (a) General structure of an amino acid, the R represents the side-chain, (b) Structure of the amino acid alanine...... 11

2.3 Two amino acids forming a peptide bond...... 11

3.1 (a) 2D view of the CTS, the top left is the bottom electrode (Metal 1), the top right is the top electrode (Metal 2), the bottom is an overlay of Metal 1 and Metal 2 (b) 3D view of the CTS...... 19

3.2 (a) 2D view of the metal electrode for the RTS (b) 3D view of the RTS...... 20

3.3 Electrical model of the CTS and RTS...... 21

3.4 2D image of the RF bridge...... 24

3.5 2D image of the half-wavelength resonator...... 25

4.1 The test structure analysis setup with VNA and probe station...... 26

4.2 The DSA used for measuring contact angles...... 28

4.3 The Epi-Fluorescent Microscope used for taking ﬂuorescent images...... 31

5.1 Frequency response of the CTS with guanine dielectric layers of 500 nm and 1 µm 34

5.2 Frequency response of the RTS with guanine dielectric layers of 250 nm, 500 nm, 750 nm, and 1 µm...... 35

5.3 Relative phase of the RTS with guanine dielectric layers of 250 nm, 500 nm, 750 nm, and 1 µm...... 36

5.4 Fluorescent images of the amino silanized glass group...... 38 xi 5.5 Fluorescent images of the guanine thin ﬁlm group...... 40

5.6 The simulation frequency response of the RF Bridge...... 41

5.7 The simulation phase of the RF Bridge...... 41

5.8 Frequency response of an RF bridge test structure without CNTs...... 42

5.9 Relative phase of an RF bridge test structure without CNTs...... 43

5.10 Frequency response of two RF bridge test structures with CNTs...... 44

5.11 Relative phase of two RF bridge test structures with CNTs...... 44

5.12 The simulation frequency response of the half-wavelength resonator...... 45

5.13 The simulation phase of the half-wavelength resonator...... 46

6.1 Chemical functionalization through covalent cross-linking...... 49

xii LIST OF TABLES

2.1 Amino acids...... 10

5.1 Dielectric constant and loss tangent of guanine...... 34

5.2 Resonant frequency and notch depth of guanine at different thicknesses...... 36

6.1 Biomarker and corresponding peptide binding sequence...... 48

A.1 The results from the frequency response measurements of the CTS...... 57

B.1 The results from the frequency response measurements of the RTS...... 58

C.1 Contact angle measurements for different surfaces...... 59

xiii LIST OF ABBREVIATIONS AND NOTATIONS

AWR Applied Wave Research Design Environment

CNT Carbon nanotube

CPW Coplanar waveguide

CSF Cerebral spinal ﬂuid

CTS Capacitive test structure

DC Direct current

DNA Deoxyribonucleic acid

DNA-CTMA Deoxyribonucleic acid-cetyltrimethylammonium chloride

DSA Drop Shape Analysis System

DSP Dithiobis(succinimidyl propionate)

EBL Electron blocking layer

ELISA Enzyme-linked immunosorbent assay

FET Field-effect transistor

GC-MS Gas chromatography-mass spectrometry

HPLC High performance liquid chromatography xiv MRE Molecular recognition element

NHS N-hydroxysuccinimide

NHS-PEG4-Biotin N-hydroxysuccinimide-polyethylene glycol-biotin

NHS-PEG4-CH3 N-hydroxysuccinimide-polyethylene glycol-methyl

NMR Nuclear magnetic resonance

OLED Organic light emitting diode

PBS Phosphate buffer saline

PTSD Post-traumatic stress disorder

PVD Physical vapor deposition

QCM Quartz crystal microbalance

RF Radio frequency

RIA Radioimmunoassay

RNA Ribonucleic acid

RTS Resonant test structure

S11 Input reﬂection coefﬁcient

S21 Forward transmission coefﬁcient

SOLT Short-open-load-thru

S-parameters Scattering-parameters

VNA Vector Network Analyzer xv VOC Volatile organic compound

xvi CHAPTER I

INTRODUCTION

Biological sample analysis methods can often be restrictive. Common methods such as chromatography, electrophoresis, and enzyme-linked immunosorbent assay (ELISA) require pretreatments, laboratory equipment, and technical expertise [1]. These factors increase feedback time, increase expense, and prohibit in situ detection, and as a result, these methods slow progress in many ﬁelds. In some cases, the slow response time can lead to devastating consequences when it pertains to diagnostics or biohazard monitoring. In contrast to laboratory analyses, many inexpensive biosensors are capable of rapid, in situ detection without pretreatments.

Sensors consist of a recognition element and transducer. The interaction between an analyte of interest and recognition element generates a stimulus that the transducer converts into a signal

[2]. This signal can then be used for a simple yes/no categorical analysis and in some instances even quantitative analysis. In the case of biosensors, the recognition element is a biological or biomimetic material [3]. Some commonly used recognition elements for biosensors are ions, afﬁnity interactions, nucleic acids, enzymes, biological cells and tissues, and vapor sorption [3]. Typical sensors are thermometric, mechanical effects, resistive, capacitive, electrochemical, and optical [3].

From determining Escherichia coli in food in 30 minutes [4] for quality control purposes to

determining organophosphates in real-time [5] and other volatile organic compounds (VOCs) [6]

for security purposes, biosensors have a wide array of potential applications for in situ, real-time 1 measurements. Perhaps the ﬁelds in which biosensors potentially have the largest impact are diagnostics and medical research. For example, contactless dielectric microsensors with amperometric electrodes have the potential to help understand disease causing microbial bioﬁlms on medical devices [7]. For pharmaceuticals, quartz crystal microbalance (QCM) sensors can help analyze drug effects on cardiomyocyte cluster beating [8]. For genomic research, microcantilevers can detect a single nucleotide polymorphism in less than half an hour [9] which is much quicker than the conventional method of denaturing gel electrophoresis [1, 10]. Many biosensor technologies for medical applications have already been transferred into the marketplace. Common, commercial biosensors for health assessment are pregnancy strip tests capable of categorical results within half an hour and glucose monitors often used by diabetics for real-time analysis of blood glucose levels.

1.1 Objectives

The main objective of this thesis is the development of biosensors for the detection of biomolecules in aqueous and vapor phase. The sensor platforms include a resonant test structure (RTS) with a molecular recognition element (MRE) functionalized guanine dielectric layer and carbon nanotube

(CNT)-integrated radio frequency (RF) test structures. For the RTS, interaction of an analyte with the MRE is expected to cause a change in permittivity which will result in a shift in the resonant frequency, amplitude, and phase of the RTS thereby signaling the presence of the analyte. For the

CNT-integrated RF test structures, the CNTs are functionalized with MREs and interaction with an analyte is expected to cause a shift in the resonance.

Objective 1. Characterization of the guanine biopolymer. Guanine is employed as a dielectric layer and as an interface for tethering of MREs for eventual analyte capture. Electrical characterization of the guanine biopolymer is achieved using the capacitive test structure (CTS) and RTS. Char- acterization includes determination of the dielectric constant and loss tangent over a wide frequency

2 range. The resonance of the RTS with guanine dielectric layer is determined. The guanine layer is further characterized using contact angle measurements to determine the hydrophobic/hydrophilic properties of the biopolymer.

Objective 2. Chemical functionalization of the guanine biopolymer. The guanine biopolymer has pendant primary amines that may be available as tether points for covalent immobilization of MREs. A biotin-streptavidin model system is used to determine the viability of chemically functionalizing guanine with MREs.

Objective 3. Characterization of an RF bridge test structure and a half-wavelength resonator test structure for biosensing applications. Two CNT-integrated RF test structures are designed and developed for biosensing: an RF bridge and half-wavelength resonator. Design simulations, performed in Applied Wave Research Design Environment (AWR) software, are employed to model the frequency response of the CNT-integrated RF test structures. The RF bridge is fabricated and the frequency response is measured pre- and post-CNT integration.

1.2 Signiﬁcance

Guanine is both abundant and inexpensive [11], yet the potential of guanine for electronics has barely been investigated. This thesis covers new ground for the application of guanine as an alternative to inorganic polymers commonly used in electronics. For the ﬁrst time, guanine is used as the dielectric layer for the CTS and RTS. Additionally, the chemical functionalization method described herein is the ﬁrst demonstration of covalently binding biomolecules to guanine.

RF devices are increasingly being incorporated into biosensors for signal transduction [12]. As will be discussed in Section 2.2, many RF biosensors have the potential for label free, rapid detection on a miniaturized platform. The test structures described in this thesis have potential applications

3 for biomolecule recognition, and as with many RF biosensors, the detection would be label free and rapid.

1.3 Outline

The ordering of the following chapters is Background, Test Structures, Experimental Proce- dure, Results and Discussion, and Conclusions and Future Work. The Background covers different aspects of the sensor design: capacitive sensors, resonant sensors, CNTs, biopolymers, proteins, biomolecule immobilization, and model analytes. The Test Structures chapter describes the fabrication and design of the CTS, RTS, RF bridge, and half-wavelength resonator. The Experimental

Procedure discusses the process for measuring the test structures’ frequency response, determining the hydrophobic/hydrophilic properties of guanine through contact angles, testing the chemical functionalization method through a biotin-streptavidin model, and bridging electrodes with CNTs.

The Results and Discussion chapter presents the outcomes of the guanine characterization, chemical functionalization, and CNT-integrated RF test structure analysis. The Conclusions and Future Work section discusses the implications of the results and future work to be done.

4 CHAPTER II

BACKGROUND

The purpose of this chapter is to provide background over the various aspects of sensor design.

Sensors, biosensors in particular, encompass many different ﬁelds. In order to provide an overview, this chapter covers capacitive sensors, resonant sensors, CNTs, biopolymers, proteins, biomolecule immobilization, and model analytes for future detection.

2.1 Capacitive Sensors

A capacitor is formed by two parallel plates separated by a dielectric material. When the plates are oppositely charged, the dielectric becomes polarized [2]. Polarization weakens the electric ﬁeld across the plates thereby reducing the voltage across the capacitor and increasing the voltage stored

[2]. The capacitance is a function of the area of the capacitor’s plates, the distance between the plates, and the relative permittivity of the dielectric also known as the dielectric constant. These variables can be used to tune the capacitance.

One method for developing capacitive sensors is through a change in the distance between the capacitor plates. This method has been used to develop a glucose capacitive sensor wherein the binding of glucose to the molecularly imprinted polymer surface increases the distance [13]. Another method for developing capacitive sensors is through changes in the permittivity of the dielectric.

For example, capacitors with selective dielectric materials are capable of detecting VOCs through 5 a permittivity induced capacitance change when these VOCs are absorbed by the dielectric [14].

Changes in the dielectric properties have been a demonstrated method for detecting deoxyribonucleic acid (DNA) hybridization with single nucleotide polymorphism sensitivity [15] and measuring water concentration [16]. Also, capacitive sensors have been used to detect cardiovascular risk biomarkers through changes in the dielectric properties caused by antibody-antigen binding [17].

2.2 Resonant Sensors

The series resonance of a device happens when the reactance is zero which occurs when the capacitance and inductance are equal in magnitude and opposite in sign [18]. At series resonance, the device acts as a short circuit [18]. In the case of parallel resonance, the net susceptance is zero causing the device to act as an open circuit [18]. This behavior is beneﬁcial in developing resonant sensors. The resonant frequency can be tuned by altering the surface of the metal electrodes or by altering the dielectric material of the device. Biological materials such as proteins have electrical properties [19]. Therefore, through protein functionalization, the properties of a device can be altered [19]. Analyte binding to this protein further changes the properties of the device indicating the presence of the analyte.

Resonant structures have been implemented for biosensing. For a microwave cavity sensor, resonant frequency has been a sensing parameter for determining glucose levels in pig blood [20].

Resonant frequency is also the parameter of interest for asymmetric split-ring resonator sensors implemented for the detection of free cortisol and α-amylase [21]. An RF biosensing system has been demonstrated to detect streptavidin and complementary DNA by functionalizing biotin and single-strand DNA on a resonator [22]. Another interesting RF sensor uses shifts in the resonance to determine glucose concentrations in human serum [23]. The dielectric constant of the serum is

6 dependent on the glucose concentration therefore different glucose concentrations result in different resonant frequencies [23].

2.3 Carbon Nanotubes (CNTs)

CNTs are hollow cylinders of graphene typically with the end caps removed [3]. Through π-π

stacking, noncovalent binding can occur between biological materials and the side walls [3]. Due

to this binding, CNTs can be functionalized with single-stranded DNA and proteins for biosensing.

Additionally, CNTs enhance the sensitivity of enzymatic sensors by promoting electron transfer

between the enzyme and electrode [3].

An interesting trait is CNTs ability to form conductive chains. These chains are formed by

dropping a suspension of CNTs between two electrodes and then applying dielectrophoresis to the

electrodes [24, 25, 26]. Dielectrophoresis causes the CNTs to align along the electric ﬁeld and form

chains [25]. After dielectrophoresis, annealing is applied to remove the surfactants from the CNT

suspension [26]. CNTs have been shown to reduce resistivity [25] and have been used to bridge CNT

contact electrodes [27]. Furthermore, biosensors have been developed in which the CNTs acted as a

bridging material. Chemiresistor sensors for the detection of cortisol [24] and Escherichia coli [26]

have been developed by bridging functionalized CNTs across electrodes.

2.4 Biopolymers

Research into biopolymers for electronics is an expanding ﬁeld. This section discusses the prop-

erties and various applications of a few biopolymers (silk, deoxyribonucleic acid-cetyltrimethylammonium

chloride (DNA-CTMA), and guanine) in order to provide insight as to why biopolymers are becom-

ing increasingly applied to electronics.

7 2.4.1 Silk

Silk is naturally produced by Bombyx mori silkworms [28, 29, 30]. Silk ﬁlms are mechanically strong [28] and can be deposited through conventional techniques such as layer-by-layer assembly

[28] and casting [29, 30]. Additionally, silk can be easily absorbed by the body without triggering an immune response [29]. In addition to biodegradability, the rate of degradation for silk can be controlled [30]. Due to these qualities, silk is an ideal candidate for in vivo devices. For example, silk has been used as a dissolvable thin ﬁlm to facilitate in the conformal attachment of an electrode array to a brain [29]. Dissolvable, antibiotic-loaded silk has been used in implantable, wirelessly controlled, therapeutic devices to prevent post-surgery infections [30]. Also, silk has been used as a dielectric material for the RTS [31].

2.4.2 Deoxyribonucleic Acid-Cetyltrimethylammonium Chloride (DNA-CTMA)

DNA-CTMA is formed by an ion exchange reaction between puriﬁed DNA and CTMA [32].

DNA has a negative charge along its sugar phosphate backbone; the CTMA provides a cation to form a neutral structure. DNA-CTMA is more mechanically stable than DNA [33], and unlike

DNA, DNA-CTMA is water insoluble and can be dissolved in solvents for spin-coating [32, 33].

Due to the abundance of the naturally occurring materials involved, DNA-CTMA is inexpensive

◦ [33]. DNA-CTMA is also tunable at 25 C with applied voltage (r = 6.2855 and tanδ at 15 GHz =

0.2721 at 0 V, r = 5.4606 and tanδ at 15 GHz = 0.3282 at 20 V) [34]. DNA-CTMA has been used

in various applications. This biopolymer has been demonstrated as a potential cladding material

for electro-optic devices [32]. DNA-CTMA has been used as an electron blocking layer (EBL)

in organic light emitting diodes (OLEDs) with improved luminance and luminous efﬁciency over

OLEDs without a DNA-CTMA EBL [35]. DNA-CTMA has been used as a dielectric material in the

8 CTS [34] and RTS [31]. Also, DNA-CTMA has been integrated into ﬁeld-effect transistors (FETs) as the gate dielectric [33, 36] and as a semiconducting material [34].

2.4.3 Guanine

Guanine is a ribonucleic acid (RNA) and DNA nucleobase, as shown in Figure 2.1.

Figure 2.1: Structure of guanine

Because this material is found in many organisms, it is an abundant, inexpensive, replenishable resource [11]. Additionally, guanine is slightly hydrophobic (contact angle of water is 62.07◦, see

Section 5.3) when compared to other materials such as glass. Therefore, guanine should be pene- trable by the biomarkers yet less prone to humidity based swelling than more hydrophilic polymers which is advantageous in preventing unreliable measurements. Additionally, guanine has a high dielectric breakdown field of 3.5 MV cm-1 [11] and is thermally stable up to 465 ◦C in air [37]. Fur- thermore, guanine films, along with the other DNA nucleobases, can be formed by physical vapor deposition (PVD) [37] and can withstand electron-beam evaporation of gold due to the high thermal stability. While the other DNA nucleobases can also withstand the high temperatures needed for deposition, guanine forms a smoother film when vacuum processed [11].

Though limited, guanine has been utilized in electronic applications. For a sensor platform, glassy carbon electrodes have been modiﬁed with guanine for insulin detection [38]. Guanine has been incorporated into the gate dielectric of graphene transistors [37], organic FETs [11, 39], and

9 inorganic FETs [39]. Also, guanosine, a modiﬁed form of guanine, has been incorporated into FETs

[40]. In some cases, guanine has been shown to improve the properties of devices. Embedding guanine into the dielectric of an organic FET and inorganic FET improved the stability of the gate

[39]. Also, guanine has been shown to increase the electric ﬁeld breakdown of a DNA-CTMA-sol- gel capacitor [37].

2.5 Proteins

Proteins are often integrated into biosensors as MREs. This section gives a brief overview of the structure of proteins and applications in sensors.

2.5.1 Structure

Proteins are constructed of amino acids. Table 2.1 lists the 20 common amino acids with their three letter and one letter code.

Table 2.1: Amino acids

Amino Acid 3 Letter Code 1 Letter Code Amino Acid 3 Letter Code 1 Letter Code Alanine Ala A Leucine Leu L Arginine Arg R Lysine Lys K Asparagine Asn N Methionine Met M Asparatic Acid Asp D Phenylalanine Phe F Cysteine Cys C Proline Pro P Glutamine Gln Q Serine Ser S Glutamic Acid Glu E Threonine Thr T Glycine Gly G Tryptophan Trp W Histidine His H Tyrosine Tyr Y Isoleucine Ile I Valine Val V

The structure of every amino acid is a carbon atom with an amino group, carboxylic acid group, hydrogen group, and side-chain bound to it [1], as shown in Figure 2.2.

10 (a) (b)

Figure 2.2: (a) General structure of an amino acid, the R represents the side-chain, (b) Structure of the amino acid alanine

The primary structure of a protein is the sequence of amino acids bound linearly by peptide bonds [1]. Peptide bonds are the binding of the carboxylic acid and amino group of sequential amino acids [1], as shown in Figure 2.3.

Figure 2.3: Two amino acids forming a peptide bond

The secondary structure consists of α-helixes, β-sheets, β turns, and Ω loops formed by hydrogen bonding between the amino group and carboxylic acid group of non-sequential amino acids [1].

The tertiary structure refers to the overall structure of the protein, i.e. the determination of which side-chains are located towards the center and which are outward [1]. The quaternary structure is the interactions between several amino acid chains [1].

Through folding, proteins such as enzymes and antibodies develop their own unique structure that determines function. The structure is what allows proteins’ recognition site to bind to an analyte in order to perform its function [1]. A change in the structure may result in the loss of functionality.

11 For applications as MREs in biosensors, the recognition site of proteins must not be structurally altered in order to maintain functionality and therefore selectivity.

2.5.2 Protein-Based Sensors

Enzymes often serve as the recognition element for sensors. Enzymes selectively catalyze a substrate [1]. When immobilized on a sensor surface, the reaction of an enzyme to a substrate results in a change in the surface potential of the device. This change in surface potential can be measured by electrical sensor platforms. For example, enzymes have been used in the development of FETs for the detection of urea [41], glucose [41], acetylcholine [41, 42], N-acetyl-L-tyrosine ethyl ester [41], and dopamine [43]. The enzymes are immobilized onto the gate electrode of the

FET, so the change in the surface potential caused by enzymes results in a current change for the

FET. In addition to electrical platforms, enzymes are used for colorimetric platforms. Enzyme catalysis of a substrate can result in a colored product [1]. An example of a colorimetric based method is paper-based sensors [44]. The degree of color change indicates the concentration of the substrate.

Another common protein utilized in sensor development is antibodies. Antibodies serve as an animals immunological defense mechanism against antigens, i.e. foreign proteins, polysaccharides, and nucleic acids [1]. Once produced, antibodies recognize and bind to a particular site on the antigen and therefore have a high affinity and selectivity to that particular antigen [1]. Due to this high affinity, antibodies are used as a recognition element to enhance selectivity. Antibodies have been used as the recognition element for QCM sensors [45], thin films [46], and cantilevers [47].

For these platforms, the antibody-antigen binding is measured through the change in the mass or resonance of the devices. Antibodies can also be immobilized onto electrodes [48], and the binding event is detected through electrochemical methods such as square wave voltammetry [48].

12 Peptides can also be incorporated into sensors as the recognition element. Peptides are small proteins (less than 50 amino acids) [1]. The smaller size of peptides allows for denser immobilization on sensing layers thereby increasing sensitivity [48]. Additionally, the tertiary structure of peptides are insigniﬁcant which reduces the complexity of incorporation into sensors in comparison to proteins [49]. Peptides are often immobilized onto electrodes [48, 50]. For these platforms, the binding of a peptide to a speciﬁc analyte changes the surface potential of the electrodes surface.

This binding induced change can be measured using square wave voltammetry [48], electrochemical impedance spectra [50], cyclic voltammogram [50], and differential pulse voltammogram [50].

In addition to recognition, peptides can be used to increase sensor signals. For example, for an amperometric sensor, peptide nanotubes were used to enhance direct electron transfer between an enzyme and the electrode thereby increasing the signal strength [5].

2.6 Biomolecule Immobilization

Biomolecule immobilization refers to the way in which biomolecules are attached to a sensor surface. This functionalization can occur through many ways. As mentioned in Section 2.3, MREs can be immobilized on the side walls of CNTs through π-π stacking. Essentially, polyaromatic molecules can be used to cross-link biological materials such as nucleic acids and proteins to the surface of CNTs [3]. Lim et al. found adding three phenylalanines, an amino acid containing an aromatic moiety, to the carboxyl terminus of a speciﬁc peptide binding sequence for trimethylamine allowed for noncovalent immobilization to the surface of CNTs through π-π stacking [49]. In this

case, the phenylalanine acts as a cross-linker. Cross-linkers can also be used to covalently func-

tionalize surfaces. For instance, adding anchoring peptides to the carboxyl terminus of a speciﬁc

binding peptide sequence for orexin A is a demonstrated method for covalent binding to the sur-

face of ZnO [51]. Peptides have also been cross-linked to a surface using dithiobis(succinimidyl

13 propionate) (DSP). For example, Huan et al. bound a selective peptide for Anthrax to a gold electrode using DSP [48]. The DSP forms a thiol bond to a gold electrode through the reduction of the disulﬁde group [48], and then the amino terminus of a peptide can bind covalently to the available

N-hydroxysuccinimide (NHS) terminus on the unbound end of the DSP [48]. Additionally, immobilization can occur through adsorption. This method is simply incubating a sensor surface with the

MRE. A drawback of adsorption is MREs can easily be removed which lowers the stability of the sensor [3]. Covalent binding creates a much stronger linkage than adsorption and therefore has a higher stability [3].

2.7 Model Analytes

As stated in the Objectives (Section 1.1), the goal is the development of biosensors for the detection of biomolecules in aqueous and vapor phase. Two biomarkers of interest are orexin A and trimethylamine. The meaningful detection of orexin A and trimethylamine takes place in aqueous and vapor phase respectively, so both are model analytes for future work. This section gives an overview of the functions and current detection methods for these two biomarkers.

2.7.1 Orexin A

Orexin A, also referred to as hypocretin-1, is a biomarker found in bodily ﬂuids such as blood plasma [52, 53], cerebral spinal ﬂuid (CSF) [52, 53], and saliva. Studies have shown orexin A is a factor in metabolic rates [52], reward pathways [54], and the prevention of neuron degeneration

[55]. Patients suffering from post-traumatic stress disorder (PTSD) had lower levels of orexin A in plasma and CSF than patients without PTSD [53]. Additionally, orexin A has been proven to be involved in sleep patterns. Nishino et al. reported low orexin A levels in the CSF of people suffering from narcolepsy-cataplexy [52]. Also, the application of orexin receptor antagonist to rats resulted in increased sleep thus indicating orexin’s involvement in sleep patterns [54]. Furthermore,

14 the adverse effects on sleep patterns caused by abnormal levels of orexin A can result in diminished cognitive function.

Orexin A levels are a proven metric for determining health and human performance. For example, orexin A levels can be used to diagnose narcolepsy [52]. The existing methods to determine orexin A levels include ELISA [53] and radioimmunoassay (RIA) [52]. Both of these methods are based on antibody-antigen binding to detect the presence of the antigen of interest [1]. These methods are effective; less than a nanogram can be detected by ELISA [1] and RIA [52]. Also, RIA can even be performed in vivo [1]. Although these methods are effective, ELISA and RIA require pre-treatments that must be performed in a laboratory [1]. Furthermore, in order to perform an

RIA, radioactive materials are needed [51]. Another difﬁculty presented in ELISA and RIA is the parameters determining antigen concentration are colorimetric and degree of radiation respectively.

Because of the nature of these parameters, the resolution is low.

A couple of sensors have been developed to detect orexin A. One such sensor is an FET that detects orexin A through selective binding peptides in the dielectric of the device and relays the presence of orexin A through a current change [51]. Another orexin A sensor platform involves graphene-modiﬁed electrodes that signal analyte’s presence through current change [56].

2.7.2 Trimethylamine

Trimethylamine is a biomarker found in exhaled breath [57] and urine [58]. High trimethylamine levels are associated with the disease trimethylaminuria. In people with trimethylaminuria, the enzyme responsible for oxidizing trimethylamine into trimethylamine-N-oxide is defective resulting in excess trimethylamine in bodily excretions [59]. Levels of trimethylamine are also abnor- mally high in patients suffering from kidney failure [57]. Bain et al. found people with end-stage renal disease had signiﬁcantly higher levels of trimethylamine in their plasma pre-dialysis than did

15 healthy people [60]. In addition to being a metric of health, trimethylamine levels are relevant in determining ﬁsh quality [61].

Common methods for determining trimethylamine levels are gas chromatography-mass spectrometry (GC-MS) [60], high-performance liquid chromatography (HPLC) [61], and nuclear magnetic resonance (NMR) spectroscopy [58]. GC-MS uses the time of flight of a compound’s ions through an electric field to identify the compound [1]. HPLC is a technique to filter compounds in a liquid based on either size, charge, or affinity [1]. In NMR, both a magnetic field and electromagnetic radiation are applied to acquire the resonance spectrum [1]. These analysis techniques are time consuming and have to be performed by trained technicians in a laboratory.

In addition to biochemical analyses, many sensors have been designed to determine trimethylamine levels. Pandeeswari and Jeyaprakash developed a sensor by using alpha-MoO3 thin ﬁlms that selectively absorb trimethylamine [62]. Bourigua et al. fabricated a trimethylamine amperometric sensor through the use of covalently immobilized enzymes on an electrode [63]. Trimethylamine can also be detected using peptide immobilization. Lim et al. developed a trimethylamine sensor by immobilizing selective peptides onto single walled-CNT FETs [49]. Also, Wu et al. fabricated a trimethylamine sensor by immobilizing selective peptides onto a piezoelectric crystal [57].

2.8 Summary

This chapter covers the various aspects of the sensors designed in this thesis: capacitive sensors, resonant sensors, CNTs, biopolymers, proteins, biomolecule immobilization, and model analytes.

Capacitive and resonant sensors are promising platforms for RF biosensors. A potential drawback is the test area for both platforms is small. A small test area lessens the amount of MREs at the site causing the detection range to be narrow. In order to overcome this drawback, MRE functionalization needs to be as dense as possible. CNTs are a promising material for biosensing. Biological

16 materials such as single-stranded DNA and proteins can be immobilized on CNTs, and CNTs have been shown to enhance sensitivity. In summary of biopolymers, guanine has promising properties for electronic applications such as high dielectric breakdown ﬁeld and high thermal stability.

In review of proteins, the structure and folding determines the function of proteins, therefore, the functionalization of proteins needs to avoid structural changes to the recognition site. Additionally, functionalization needs to keep the recognition site unbound from the sensor’s surface. An interesting method for MRE immobilization involves using a cross-linker to covalently tether biomolcules to a surface. As for analytes of interest, orexin A (a biomarker of cognition) and trimethylamine

(a biomarker of the disease trimethylaminuria) are model analytes for aqueous and vapor detection respectively and are clinically relevant for the assessment of health and human performance.

17 CHAPTER III

TEST STRUCTURES

For this thesis, the CTS is used for the characterization of guanine, and the RTS with guanine dielectric layer is used as the sensor platform. This chapter describes both of these test structures and presents the electrical model of the test structures. The electrical model is essential in calculating the dielectric constant and loss tangent of the CTS and the resonant frequency of the RTS. Addition- ally, the RF bridge and half-wavelength resonator are described herein. These two platforms have potential for biological sensing.

3.1 Capacitive Test Structure (CTS)

Guanine is integrated into the CTS for characterization purposes, as shown in Figure 3.1. The

CTS consists of four layers (the adhesion layer, Metal 1, dielectric layer, and Metal 2) deposited on a glass substrate. The adhesion layer is ﬁrst formed by sputtering 100 A˚ of chromium onto a glass substrate. Next, the bottom electrode, referred to as Metal 1, is fabricated by electron-beam evaporation of 7500 A˚ of gold through a shadow mask. This layer consists of two ground pads shunted together. Following Metal 1, the dielectric layer, in this case guanine, is deposited through

PVD. As mentioned above in Section 2.4.3, Ouchen et al. demonstrated guanine deposited by

PVD forms a defect free layer [37]. After the guanine deposition, the top electrode, Metal 2, is formed through electron-beam evaporation of 3500 A˚ of gold through a shadow mask. This layer

18 is a ground-signal-ground coplanar waveguide (CPW). The location of the ground lines in Metal 2 directly over Metal 1 creates a large ground capacitor. The test capacitor is formed by the overlap of the transmission line in Metal 2 and the shunted line in Metal 1. The ground and test capacitor are in series; therefore, the test capacitance can be taken to be the effective capacitance [64].

(a)

(b)

Figure 3.1: (a) 2D view of the CTS, the top left is the bottom electrode (Metal 1), the top right is the top electrode (Metal 2), the bottom is an overlay of Metal 1 and Metal 2 (b) 3D view of the CTS

3.2 Resonant Test Structure (RTS)

The RTS, as shown in Figure 3.2, is the sensor platform used in this thesis. This test structure, fabricated on glass substrates, consists of two layers: the dielectric layer and the metal layer.

19 (a)

(b)

Figure 3.2: (a) 2D view of the metal electrode for the RTS (b) 3D view of the RTS

For this application, the dielectric layer is guanine deposited by PVD. The second layer is an electrode deposited by the electron-beam evaporation of 3500 A˚ of gold through a shadow mask.

This metal layer consists of a ground-signal-ground CPW with the ground pads pushed outward

towards the center. A line is shunted across the ground pads with an air gap in the middle. By

pushing the ground pads outward, the shunt line can now be longer. The increased length and shape

of the shunt line increases the inductance thus lowering the resonant frequency into the desired

frequency range (0-20 GHz). Capacitive coupling occurs from the transmission line to the shunt

line due to the air gap. Functionalization of the dielectric with MREs is expected to change the

dielectric properties resulting in a shift in the resonance. Binding of the biomolecules to the MREs

further shifts the resonance. As a result of the coplanar design, the dielectric in the coupled capacitor

20 is exposed at the surface thereby eliminating any difﬁculties concerning the diffusion of the MREs and biomolecules through the dielectric to the coupled capacitor.

3.3 Electrical Model

Both the CTS and RTS can be represented using the following electrical model, as shown in

Figure 3.3.

Figure 3.3: Electrical model of the CTS and RTS

In this model, C, Rp,Rs, and L represent the effective capacitance, shunt resistance, parasitic series resistance, and parasitic series inductance respectively. These values can be determined for the test structures by tuning the electrical model to match the test structures’ measured frequency

21 response. The impedance for the shunted resistor and capacitor is:

1 −1 ZC = ( + jwC) (3.1) Rp

The impedance for the parasitic resistor and parasitic inductor in series is:

ZL = Rs + jwL (3.2)

Overall, these two impedances in series represents the input impedance for the transmission line:

1 −1 Zin = ZC + ZL = ( + jwC) + Rs + jwL (3.3) Rp

The scattering-parameters (S-parameters) used for matching are the forward transmission coefﬁcient

(S21) and input reflection coefficient (S11). The forward transmission coefficient is represented as:

vout S21 = (3.4) vin

For a lossless transmission line, the forward transmission coefﬁcient is:

( 1 + jwC)−1 + R + jwL Rp s S21 = (3.5) 50 + 2[( 1 + jwC)−1 + R + jwL] Rp s

The input reﬂection coefﬁcient is:

Zin − Z0 S11 = (3.6) Zin + Z0

The characteristic impedance is taken to be 50 Ω. For a lossless transmission line, the input reﬂec- tion coefﬁcient is: ( 1 + jwC)−1 + R + jwL − 50 Rp s S11 = (3.7) ( 1 + jwC)−1 + R + jwL + 50 Rp s The CTS is used to determine guanine’s dielectric constant and loss tangent. As mentioned in

the previous paragraph, by matching the electrical model’s frequency response (S-parameters) to a

measured frequency response, the capacitance is determined. The following equation describes the

relationship between the capacitance and the dielectric constant:

A C = r 0 (3.8) d 22 For this equation, r, 0, A, and d are the dielectric constant, permittivity of free space, test capacitor area, and the distance between the parallel plates of the test capacitor respectively. For the CTS, the distance between the parallel plates is determined by the thickness of the biopolymer (in this case guanine). Therefore, the capacitance can be increased by decreasing the biopolymer’s thickness and vice versa. Additionally, Equation 3.8 can be used to calculate the dielectric constant. Another value of interest is the loss tangent. The loss tangent is a measure of the electromagnetic energy dissipated as heat by the dielectric. The loss tangent can be calculated using the following equation:

1 tan δ = (3.9) 2πfCRp

For the RTS, the resonance is the parameter of interest. Through matching the electrical model to the measured frequency response, the capacitance and inductance are found. The following equation can then be used to determine the resonant frequency:

1 f0 = √ (3.10) 2π LC

3.4 RF Bridge

The RF bridge, as shown in Figure 3.4, is a modiﬁed form of a simple ground-signal-ground

CPW. The CPW is designed to resonate at 8 GHz. The length for the CPW is calculated using the following equation: c λ = √ 0 (3.11) f 0 effective

In Equation 3.11, λ, c0, f0, and effective represent the wavelength, speed of light, resonant frequency, and effective dielectric of the substrate respectively. For a resonant frequency of 8 GHz and effective dielectric of 7, the wavelength is calculated to be about 14200 µm. The length of the transmission line is set to the wavelength. At both ends of the transmission line, large (500 µm by 500 µm) direct

current (DC) pads are added. In this design, a 5 µm gap is located in the center of the transmission

23 line. Additionally, the center of the ground lines are pushed outward, to provide space for a well that can be etched into the dielectric where the transmission line gap is located. CNTs are deposited into the well, and then, dielectrophoresis is applied to the DC pads in order to align the CNTs and form chains. Then, the sample is annealed at 200◦C to improve the conductivity of the CNT bridge.

These CNTs bridge the gap in the transmission line. The CNTs can be functionalized with proteins or DNA for biosensing purposes. For fabrication, SiO2 is deposited onto a Si substrate. Then, through electron-beam deposition of gold through a shadow mask, the electrode layer is formed.

Figure 3.4: 2D image of the RF bridge

3.5 Half-Wavelength Resonator

The design of a half-wavelength resonator consists of a ground-signal-ground CPW with a cavity toward each end of the transmission line, as shown in Figure 3.5. This test structure is designed to resonate at 8 GHz. For a resonant frequency of 8 GHz and effective dielectric of 7, the wavelength is calculated to be 14200 µm using Equation 3.11. The central line of the resonator is half the wavelength so 7100 µm. Each periphery line (feed line) is calculated to be an eighth of the

wavelength; the periphery lines are each 1800 µm. The cavities separating the periphery lines from the center line are each 25 µm. At the center of the test structure, the transmission line has a 5

µm gap, the ground pads are pushed outward, and a well is etched into the dielectric. For this test structure, the goal is to eventually bridge the gap in the transmission line with CNT chains similar to 24 the RF bridge. Biological materials can be immobilized onto the CNTs for biosensing purposes. As with the RF bridge, the half-wavelength resonator is formed by ﬁrst depositing a SiO2 layer onto a

Si substrate. Then, through electron-beam deposition of gold through a shadow mask, the electrode layer is formed.

Figure 3.5: 2D image of the half-wavelength resonator

3.6 Summary

This chapter gives an overview of the CTS and RTS. For the CTS, the overlap of the shunt and transmission lines with sandwiched dielectric layer forms a test capacitor that can be used to characterize guanine. The RTS is the sensor platform with the test capacitor being the sensing site.

By measuring the frequency response, the resonance of the RTS with guanine dielectric layer can be determined. The electrical model can be used to describe both the CTS and RTS and is a valuable tool in calculating the dielectric constant, loss tangent, and resonant frequency. Additionally, this chapter describes the design of the RF bridge and half-wavelength resonator. Both the RF bridge and half-wavelength resonator have gaps that sever the transmission line. This gap is intentional so that MRE functionalized CNTs can be used to bridge the transmission line.

25 CHAPTER IV

EXPERIMENTAL PROCEDURE

This chapter gives an overview of the testing protocols. This overview includes a description of the test structure (CTS, RTS, and RF bridge) analysis, the setup for performing contact angle measurements, and the biotin-streptavidin model. Additionally, the process for forming CNTs chains across the transmission lines of the fabricated RF bridge test structures is described.

4.1 Test Structure Analysis

The S-parameters for all the fabricated test structures (CTS, RTS, and RF bridge) are measured using a Hewlett-Packard 8720 Vector Network Analyzer (VNA) and probe station, as shown in

Figure 4.1.

Figure 4.1: The test structure analysis setup with VNA and probe station

26 The VNA is calibrated over 0-20 GHz using the short-open-load-thru (SOLT) calibration method

[18]. Once calibrated, the VNA is used to measure the frequency response of the test structures. All the test structures are measured at room temperature. The VNA is interfaced with a computer, so after measuring the test structures, the measurements can be recorded and imported into AWR software.

Speciﬁcally for the CTS and RTS, the component values (capacitance, shunt resistance, parasitic series resistance, and parasitic series inductance) of the electrical model can be tuned in AWR software so that the model’s frequency response matches the frequency response of the test structures.

After matching the electrical model to the CTS results, the dielectric constant and loss tangent can be calculated using Equation 3.8 and 3.9. For the RTS, the resonance can be determined by viewing the frequency response. The frequency at which a sharp drop in the power and shift in phase are observed is the resonance. Additionally, by matching the electrical model, the resonant frequency can be calculated using Equation 3.10.

4.2 Contact Angle Measurements

Contact angle measurements are taken of guanine thin ﬁlms. The thin ﬁlms were deposited on glass slides through PVD. The glass slides had been previously cleaned through sonication in acetone, methanol, and isopropanol (15 minutes each) and then drying with nitrogen. The measurements are performed using a Kruss Drop Shape Analysis System 100 (DSA), as shown in Figure

4.2, interfaced with a computer. For these measurements, a drop of water 5.0 - 5.2 µL is sessile dropped onto the guanine thin ﬁlm. An image is taken of the guanine/water interface and stored on the computer. By using DSA software, the contact angle is determined. The contact angle of water is used to determine the hydrophobic/hydrophilic properties of the guanine thin ﬁlms. A contact angle over 90◦ is considered hydrophobic while under 90◦ is considered hydrophilic [65].

27 Figure 4.2: The DSA used for measuring contact angles

4.3 Biotin-Streptavidin Model

The chemical functionalization method is based on the ability of a linker to covalently bind to guanine’s pendant primary amine, Figure 2.1. In order to do so, the pendant primary amines need to be oriented upwards. To test this method, amino silanized glass slides (positive control) and guanine thin ﬁlms on glass slides are incubated with N-hydroxysuccinimide-polyethylene glycol-methyl

(NHS-PEG4-CH3) and N-hydroxysuccinimide-polyethylene glycol-biotin (NHS-PEG4-Biotin). To determine whether binding has taken place, Oregon Green labeled streptavidin is used for ﬂuores- cence testing. The following are all the exposures:

A. Glass-Silane

B. Glass-Silane-Streptavidin-Oregon Green

C. Glass-Silane-NHS-PEG4-CH3

D. Glass-Silane-NHS-PEG4-CH3-Streptavidin-Oregon Green

E. Glass-Silane-NHS-PEG4-Biotin 28 F. Glass-Silane-NHS-PEG4-Biotin-Streptavidin-Oregon Green

G. Guanine thin ﬁlms

H. Guanine-Streptavidin-Oregon Green

I. Guanine-NHS-PEG4-CH3

J. Guanine-NHS-PEG4-CH3-Streptavidin-Oregon Green

K. Guanine-NHS-PEG4-Biotin

L. Guanine-NHS-PEG4-Biotin-Streptavidin-Oregon Green

If a surface has an available amine, the NHS is expected to covalently bind to the surface. For the amino silanized glass, an amine group is available for binding. NHS is able to covalently bind to an amino silanized surface, and therefore, the glass group acts as an example of positive NHS binding to a surface that the guanine group can be compared to. The streptavidin has a high afﬁnity for biotin but not for CH3. No binding should occur between the Oregon Green labeled streptavidin and the NHS-PEG4-CH3, so ﬂuorescence is not expected. For the NHS-PEG4-Biotin, if the NHS

is bound, the biotin is present and available for binding, and therefore, the Oregon Green labeled

streptavidin will bind making the surface ﬂuorescent.

For the amino silanized glass group (A-F), glass slides are used. All the glass slides are cleaned

using the following steps:

1. 5-10 minutes in acetone

2. 5-10 minutes in ethanol

3. Rinse in piranha solution (7 parts sulfuric acid:3 parts hydrogen peroxide)

4. Rinse three times in water 29 5. Rinse in ethanol

These slides are incubated in a solution of 2% amino silane in ethanol for 30 minutes and then rinsed three times with ethanol. The slides are dried on a heating plate at 110◦C for 10 minutes to drive the silane binding, and then the slides are hydrated for at least 10 minutes in phosphate buffer saline (PBS, 100 mM NaPO4, 150 mM NaCl, pH of 7.2, used for all PBS washes unless otherwise stated). Then, C and D are incubated in a 1 mM NHS-PEG-CH3 solution for 30 minutes followed

by three rinses in PBS. Meanwhile, E and F are incubated in a 1 mM NHS-PEG4-Biotin solution

for 30 minutes and then rinsed three times with PBS. Then, B, D, and F are incubated with the

Oregon Green labeled streptavidin for 30 minutes and then rinsed three times with PBS. The slides

are stored in PBS.

All of the guanine slides (G-L) are created by depositing 200 nm thin ﬁlms of guanine on glass

slides through PVD. Prior to deposition, the glass slides are cleaned through sonication in acetone,

methanol, and isopropanol (15 minutes each) and then dried with nitrogen. The guanine coated

slides are hydrated for at least 10 minutes in PBS. Afterwards, I and J are incubated in a 1 mM

NHS-PEG4-CH3 solution for 30 minutes and then rinsed three times with PBS. Meanwhile, K and

L are incubated in a 1 mM NHS-PEG4-Biotin solution for 30 minutes and then rinsed three times

with PBS. Subsequently, H, J, and L are incubated in Oregon Green labeled streptavidin for 30

minutes followed by three rinses in PBS. The slides are placed in PBS to keep hydrated.

For both groups, ﬂuorescence images are taken using an Olympus BX51 Epi-Fluorescent Mi-

croscope, as shown in Figure 4.3.

30 Figure 4.3: The Epi-Fluorescent Microscope used for taking ﬂuorescent images.

The excitation max and emission max for the Oregon Green labeled streptavidin are 496 and 524 respectively, so a Fluorescein Isothiocyanate (excitation wavelength of 490, emission wavelength of

520) ﬁlter cube is used. The protocols for making the PBS, amino-silane solution, NHS-PEG4-CH3

solution, NHS-PEG4-Biotin solution, and Oregon Green labeled streptavidin solution are described

in AppendixD.

4.4 CNT Bridge Formation

In this thesis, the RF bridge test structure is fabricated. The transmission line is bridged with

CNT chains. These conductive CNT chains are formed by ﬁrst depositing a suspension of CNTs

in water and sodium dodecyl sulfate into the dielectric well in the area of the transmission line gap

using a microcapillary pipette. Dielectrophoresis is then applied to the DC pads. For this purpose, a

function generator is set to 3 Vpp at 1 kHz. The output of the function generator is connected to DC

probes that are applied to the DC pads of the RF bridge test structure. The signal is applied for 1

minute. Afterwards, the surface is rinsed with deionized water. The test structures are then annealed

at 200◦C to remove the surfactants from the suspension.

31 4.5 Summary

This chapter describes all the test procedures performed in this thesis. For the test structures, a

VNA is used to measure the S-parameters. For the CTS and RTS, an electrical model can then be tuned to match the measured results and thereby analyze the test structures. Also, the method for determining the hydrophobic/hydrophilic properties of guanine through contact angle measurements is described. To test the validity of chemical functionalization, a biotin-streptavidin model is used.

Lastly, the method for bridging the transmission line of the fabricated RF bridge test structures with

CNTs is described.

32 CHAPTER V

RESULTS AND DISCUSSION

This chapter discusses the results from testing. These results cover guanine characterization, chemical functionalization, and CNT-integrated RF test structure analysis. For the guanine characterization, the dielectric constant and loss tangent are calculated. The resonance is determined for the RTS with guanine dielectric layer. The hydrophobic/hydrophilic properties for guanine are also determined. For the chemical functionalization, the results from the biotin-streptavidin model are presented. For the CNT-integrated RF test structures, the simulation results are shown, and the measurements from the fabricated RF bridge test structures (pre- and post-CNT integration) are presented.

5.1 CTS Results

The S-parameters are measured from 0 - 20 GHz for the CTS with a 500 nm and 1 µm thick guanine dielectric layer, as shown in Figure 5.1.

33 Figure 5.1: Frequency response of the CTS with guanine dielectric layers of 500 nm and 1 µm

The results from tuning the electrical model indicate the dielectric constant of guanine was uniform in this range. The dielectric constant and loss tangent are calculated for several CTSs

(Table A.1). These results are averaged across thickness and overall, as shown in Table 5.1.

Table 5.1: Dielectric constant and loss tangent of guanine

Thickness (nm) Dielectric Constant, r Loss Tangent, tanδ, 10 GHz 500 5.479 ± 0.204 0.014 ± 0.001 1000 5.211 ± 0.329 0.016 ± 0.001 Average 5.345 ± 0.294 0.015 ± 0.001

The dielectric constant and loss tangent at different thicknesses are in good agreement with each other. The dielectric constant and loss tangent of guanine are determined to be 5.345 ± 0.294 and

0.015 ± 0.001 respectively. The dielectric constant has been previously measured to be 4.35 at

1 kHz [11] and 5.02 [39]. The discrepancy among previous measurements and the measurements done here could be due to the frequencies at which the dielectric constant was determined. Also, the loss tangent has been previously determined to be 7 × 10-3 at 100 mHz [11]. This previously

34 determined loss tangent was found at a much lower frequency than the calculation done for this thesis thereby explaining the difference.

5.2 RTS Results

Guanine is incorporated into the RTS as the dielectric layer. The S-parameters of the RTSs at different thicknesses (250 nm, 500 nm, 750 nm, and 1 µm) are measured from 0 - 20 GHz, as shown in Figure 5.2. The relative phase is shown in Figure 5.3.

Figure 5.2: Frequency response of the RTS with guanine dielectric layers of 250 nm, 500 nm, 750 nm, and 1 µm

35 Figure 5.3: Relative phase of the RTS with guanine dielectric layers of 250 nm, 500 nm, 750 nm, and 1 µm

The frequency responses are nearly identical from 250 nm - 1 µm. The resonance is the same for these structures at different thicknesses due to the coplanar design. Resonance is a function of capacitance which is a function of the distance between the parallel plates of the capacitor. For the coplanar RTS design, the distance is not controlled by the dielectric thickness. Therefore, the capacitance, and by extension the resonance, is not dependent on the thickness of the dielectric.

Several RTSs are measured (Table B.1), and the resonance is averaged at each guanine thickness, as shown in Table 5.2. Given the guanine thickness does not affect the resonance, the average for all measured structures is also calculated and shown in Table 5.2.

Table 5.2: Resonant frequency and notch depth of guanine at different thicknesses

Thickness (nm) Frequency (GHz) Notch Depth (dB) 250 3.096 ± 0.027 7.422 ± 0.421 500 3.240 ± 0.118 7.474 ± 0.341 750 3.147 ± 0.038 7.686 ± 0.197 1000 3.126 ± 0.045 7.349 ± 0.317 Average 3.148 ± 0.079 7.472 ± 0.330

36 The average resonance of the RTS with guanine dielectric layer is 3.148 ± 0.079 GHz with a

notch depth of 7.472 ± 0.330 dB.

5.3 Contact Angle Measurement Results

Contact angle measurements were performed to determine the hydrophobic/hydrophilic prop-

erties of guanine. The contact angle of water on guanine thin ﬁlms is determined to be 62.07◦ ±

3.029◦. For comparison, the contact angle measurement of water on glass and amino silanized glass is 41.4◦ ± 2.72◦ and 45.3 ◦ ± 4.38 respectively. In comparison to glass and amino silanized glass,

guanine thin ﬁlms are slightly hydrophobic. All measurements are shown in Table C.1.

5.4 Biotin-Streptavidin Model Results

The chemical functionalization method for guanine has not been previously demonstrated, so

testing this method must be done. To test the validity of chemical functionalization, amino silanized

glass and guanine thin ﬁlm groups are used. Figure 5.4 are the ﬂuorescent images from the amino

silanized glass group. The images of the amino silanized glass indicate no adsorption occurred when

exposed to streptavidin. When a NHS-PEG4-CH3 functionalized glass slide is exposed to a labeled

streptavidin, no ﬂuorescence is observed. The CH3 is known to have no afﬁnity to streptavidin. This

result corresponds to the NHS being bound to the amino silanized glass. This conclusion is further

supported by the images of the NHS-PEG4-Biotin functionalized slides. Streptavidin has a high

afﬁnity to biotin, so if the biotin is present on the surface, the streptavidin is expected to bind. These

images clearly show the labeled streptavidin speciﬁcally bound to the surface thereby indicating the

NHS-PEG4-Biotin is bound to the surface of the amino silanized glass slides. Figure 5.4c shows

the difference in ﬂuorescence between the NHS-PEG4-CH3 and NHS-PEG4-Biotin functionalized

slides. These images indicate the NHS is able to bind to the amino silanized surface. The results

37 match the expected, known reactions between amino silanized glass, NHS linkers, and labeled streptavidin. The amino silanized glass will act as a positive comparison for the guanine slides.

(a) i) Silane, ii) Silane-NHS-PEG4-CH3, iii) Silane- (b) i) Silane, ii) Silane-NHS-PEG4-Biotin, iii) Streptavidin-Oregon Green, iv) Silane-NHS-PEG4-CH3- Silane-Streptavidin-Oregon Green, iv) Silane-NHS- Streptavidin-Oregon Green PEG4-Biotin-Streptavidin-Oregon Green

(c) i) Silane-NHS-PEG4-CH3, ii) Silane-NHS-PEG4- Biotin, iii) Silane-NHS-PEG4-CH3-Streptavidin-Oregon Green iv) Silane-NHS-PEG4-Biotin-Streptavidin-Oregon Green

Figure 5.4: Fluorescent images of the amino silanized glass group

The ﬂuorescent images from the guanine thin ﬁlm testing are shown in Figure 5.5. The images

of the guanine thin ﬁlms show a high non-speciﬁc adsorption of streptavidin. This adsorption in-

dicates guanine’s surface is charged. For the guanine slide functionalized with NHS-PEG4-CH3, streptavidin is unable to bind to the surface. This result illustrates the NHS of this linker is able to

38 bind to guanine’s pendant primary amine, and the PEG4-CH3 is completely blocking the surface.

Furthermore, streptavidin is speciﬁcally binding to the NHS-PEG4-Biotin functionalized guanine

indicating the NHS is bound to the surface. Figure 5.5c shows a side by side of the NHS-PEG4-

CH3 and NHS-PEG4-Biotin functionalized guanine slides which illustrates the clear difference in

fluorescence between the two functionalized surfaces (the surfaces had been slightly photobleached by the time this image was taken). These results indicate the NHS linkers are able to covalently bind to guanine’s pendant primary amine. In comparing the guanine group to the amino silanized glass group, the NHS functionalized guanine thin films are nearly identical to their respective amino silanized glass slides when exposed to labeled streptavidin. Altogether, the results prove guanine’s pendant amino group is available on the surface of the thin films for binding thus chemical functionalization is a valid method for guanine.

39 (a) i) Guanine, ii) Guanine-NHS-PEG4-CH3, iii) (b) i) Guanine, ii) Guanine-NHS-PEG4-Biotin, iii) Guanine-Streptavidin-Oregon Green, iv) Guanine-NHS- Guanine-Streptavidin-Oregon Green, iv) Guanine-NHS- PEG4-CH3-Streptavidin-Oregon Green PEG4-Biotin-Streptavidin-Oregon Green

(c) i) Guanine-NHS-PEG4-CH3, ii) Guanine-NHS-PEG4- Biotin, iii) Guanine-NHS-PEG4-CH3-Streptavidin- Oregon Green, iv) Guanine-NHS-PEG4-Biotin- Streptavidin-Oregon Green

Figure 5.5: Fluorescent images of the guanine thin ﬁlm group

5.5 RF Bridge Simulation

The RF Bridge is designed in AWR software. It is not possible to simulate a CNT bridge in this

software; therefore, a thin conducting line is used to bridge the transmission line gap for simulation

purposes. The frequency response is shown in Figure 5.6. The simulation’s resonance is at 9.2 GHz

with a notch depth of 20.31 dB. The relative phase of the RF Bridge is shown in Figure 5.7. The

phase shift clearly occurs at 9.2 GHz.

40 Figure 5.6: The simulation frequency response of the RF Bridge

Figure 5.7: The simulation phase of the RF Bridge

41 5.6 Fabricated RF Bridge Test Structure Results

The RF bridge test structures are fabricated. The S-parameters are measured from 0-20 GHz before CNTs are used to bridge the transmission line, as shown in Figures 5.8 and 5.9. The resonant frequency is 13.23 GHz with a notch depth of 29.42 dB.

Figure 5.8: Frequency response of an RF bridge test structure without CNTs

42 Figure 5.9: Relative phase of an RF bridge test structure without CNTs

Two RF bridge test structures are bridged with CNTs. The S-parameters are shown in Figures

5.10 and 5.11. For one test structure, the resonance is 16.43 GHz with a notch depth of 31.2 dB. This structure also has a second harmonic at 19.34 GHz with a notch depth of 35.5 dB. The resonance of the other test structure is at 13.58 GHz with a notch depth of 30.1 dB. The second harmonic for this structure is at 16.11 GHz with a notch depth of 34.4 dB. The differences in resonance may be due to the non-uniform CNTs bridges across the transmission lines.

43 Figure 5.10: Frequency response of two RF bridge test structures with CNTs

Figure 5.11: Relative phase of two RF bridge test structures with CNTs

44 5.7 Half-Wavelength Resonator Simulation

The design of the half-wavelength resonator is done in AWR software. As mentioned above in Section 5.5, CNTs cannot be simulated in this software. For the simulation, a thin conducting line is used to bridge the gap in the transmission line. Figure 5.12 is the frequency response of the simulation. The simulation has a resonant frequency at 8 GHz with a pass band peak of 0.3395 dB. A phase shift occurs at 8 GHz, as shown in Figure 5.13. This resonance occurs at the chosen resonance for the design of the half-wavelength resonator.

Figure 5.12: The simulation frequency response of the half-wavelength resonator

45 Figure 5.13: The simulation phase of the half-wavelength resonator

5.8 Summary

In this chapter, all the results are presented. The S-parameters are measured for the CTS with guanine dielectric layer. Using the electrical model, the dielectric constant and loss tangent are calculated. The S-parameters are also measured for the RTS with guanine dielectric layer. The resonance of this structure is determined from these measurements. Contact angle measurements are taken of water on the surface of guanine thin ﬁlms to determine the hydrophobic/hydrophilic properties of guanine. The ﬂuorescent images from the biotin-streptavidin model are presented and evaluated to determine the validity of the chemical functionalization method. Finally, the S- parameters are presented for the simulated CNT-integrated RF test structures and for the fabricated

RF bridge test structures.

46 CHAPTER VI

CONCLUSIONS AND FUTURE WORK

This chapter summarizes the work done in this thesis (CTS analysis, RTS analysis, contact angle measurements, biotin-streptavidin modeling, CNT-integrated RF test structure analysis) and relates the work back to the objectives set out in Section 1.1. Additionally, future work for the resonant capacitive sensor and CNT-integrated RF test structures are discussed.

6.1 Resonant Capacitive Sensor

For the development of the resonant capacitive test structure for biological sensing, the ﬁrst objective of characterizing guanine has been completed. Using the CTS, the dielectric constant and loss tangent of guanine are determined to be 5.345 ± 0.294 and 0.015 ± 0.001 respectively. Guanine has been deposited as the guanine dielectric layer in the coplanar RTS. The frequency response of the RTS with guanine thicknesses of 250 nm, 500 nm, 750 nm, and 1 µm are nearly identical. The average resonance of the RTS with guanine dielectric layer is 3.148 ± 0.079 GHz with a notch depth of 7.472 ± 0.330 dB. Contact angles measurements with water are used to determine the hydrophobic/hydrophilic properties of guanine. The contact angle is 62.07◦ ± 3.029◦ indicating

guanine is slightly hydrophobic with respect to glass and amino silanized glass.

The second objective of determining the validity of chemical functionalization for guanine has

also been completed. Results from the biotin-streptavidin model indicate NHS is able to covalently 47 bind to guanine’s pendant primary amine. These results mean covalent cross-linking biomolecules to guanine is a valid method for functionalization.

Future work to be done in the development of the sensor is to chemically functionalize the surface of the guanine dielectric layer with the MREs for a biomolecule of interest. Two clinically relevant biomarkers for aqueous and vapor testing are orexin A (a biomarker of cognition) and trimethylamine (a biomarker of the disease trimethylaminuria) respectively. The peptide sequences shown in Table 6.1 have an afﬁnity to the biomarkers of interest.

Table 6.1: Biomarker and corresponding peptide binding sequence

Biomarker Peptide Binding Sequence Orexin A DQSNKIISLQRL [51] Trimethylamine LFLSNLSFSDLC [57]

For chemical functionalization, the linker DSP will be used to tether the peptides to the guanine.

DSP is a chain terminated at both ends with an NHS group capable of covalently binding with a primary amine. The goal is for DSP to covalently bind to both the guanine and peptides simultaneously thus forming a covalent tether, as shown in Figure 6.1.

48 Figure 6.1: Chemical functionalization through covalent cross-linking

Covalently tethering the peptides to the guanine dielectric layer is expected to increase the sensitivity, stability, and reusability of the sensors. Once the guanine has been chemically functionalized with peptides, the next step is to measure the frequency response to determine the baseline resonance. Afterwards, the sensors will be tested with both positive (orexin A in aqueous phase and trimethylamine in vapor phase) and negative controls. Binding between a positive control and the selective peptides is expected to change the permittivity of the dielectric layer resulting in a shift in the resonance. A negative control should not bind to the peptides, and therefore, no shift in resonance is expected.

49 6.2 RF Bridge and Half-Wavelength Resonator

The RF bridge and half-wavelength resonator are designed and simulated in AWR software.

Both test structures are designed for a resonance at 8 GHz. The simulated resonance of the RF bridge is 9.2 GHz with a notch depth of 20.31 dB. The RF bridge test structures have been fabricated and bridged with CNTs. The measured results show a higher resonance than simulated, and the measured resonances also vary across CNT-integrated test structures. This variation is due to the

CNT bridges being non-uniform. For the half-wavelength resonator, the simulation has a resonance at 8 GHz with a pass band peak of 0.3395 dB.

For the half-wavelength resonator, the next step is to add DC pads to both ends of the transmission line. The DC pads are used for dielectrophoresis which is needed in order to align the

CNTs and form chains. Afterwards, the half-wavelength resonator can be fabricated and the CNTs integrated into the transmission line. Additionally, future work will include developing a more uniform process for bridging the transmission line. One way of increasing uniformity is to deposit a standard amount of the CNT suspension. After developing a more uniform method for CNT bridge formation, the next step is to use MRE functionalized CNTs for biosensing purposes.

50 BIBLIOGRAPHY

[1] J. M. Berg, J. L. Tymoczko, and L. Stryer, Biochemistry. New York: W. H. Freeman and Company, 2002.

[2] J. Fraden, Handbook of Modern Sensors: Physics, Designs, and Applications. New York, USA: Springer, 2010.

[3] F.-G. Banica, Chemical Sensors and Biosensors: Fundamentals and Applications. United Kingdom: John Wiley & Sons, Ltd, 2012.

[4] P. Cheng, Z.-G. Huang, Y. Zhuang, L.-C. Fang, H. Huang, J. Deng, L.-L. Jiang, K.-K. Yu, Y. Li, and J.-S. Zheng, “A novel regeneration-free E. coli O157:H7 amperometric immunosen- sor based on functionalised four-layer magnetic nanoparticles,” Sensors and Actuators B, vol. 204, pp. 561–567, 2014.

[5] P. A. Baker, M. N. Goltz, A. M. Schrand, D. Y. Yoon, and D.-S. Kim, “Organophosphate vapor detection on gold electrodes using peptide nanotubes,” Biosensors and Bioelectronics, vol. 61, pp. 119–123, 2014.

[6] J. White, K. Truesdell, L. B. Williams, M. S. AtKisson, and J. S. Kauer, “Solid-State, Dye- Labeled DNA Detects Volatile Compounds in the Vapor Phase,” PLoS Biology, vol. 6, no. 1, pp. 30–36, 2008.

[7] J. Gottschamel, L. Richter, A. Mak, C. Jungreuthmayer, G. Birnbaumer, M. Milnera, H. Bruckl,¨ and P. Ertl, “Development of a Disposable Microﬂuidic Biochip for Multiparameter Cell Population Measurements,” Analytical Chemistry, vol. 81, no. 20, pp. 8503–8512, 2009.

[8] N. Tymchenko, A. Kunze, K. Dahlenborg, S. Svedhem, and D. Steel, “Acoustical sensing of cardiomyocyte cluster beating,” Biochemical and Biophysical Research Communications, vol. 435, pp. 520–525, 2013.

[9] K. M. Hansen, H.-F. Ji, G. Wu, R. Datar, R. Cote, A. Majumdar, and T. Thundat, “Cantilever- Based Optical Deﬂection Assay for Discrimination of DNA Single-Nucleotide Mismatches,” Analytical Chemistry, vol. 73, no. 7, pp. 1567–1571, 2001.

[10] S. G. Fischer and L. S. Lerman, “DNA fragments differing by single base-pair substitutions are separated in denaturing gradient gels: Correspondence with melting theory,” Proceedings of the National Academy of Sciences, vol. 80, no. 6, pp. 1579–1583, 1983.

51 [11] M. Irimia-Vladu, P. A. Troshin, M. Reisinger, G. Schwabegger, M. Ullah, R. Schwoediauer, A. Mumyatov, M. Bodea, J. W. Fergus, V. F. Razumov, H. Sitter, S. Bauer, and N. S. Sariciftci, “Environmentally sustainable organic ﬁeld effect transistors,” Orgnanic Electronics, vol. 11, no. 12, pp. 1974–1990, 2010.

[12] H.-J. Lee and J.-G. Yook, “Recent research trends of radio-frequency biosensors for biomolec- ular detection,” Biosensors and Bioelectronics, vol. 61, pp. 448–459, 2014.

[13] Z. Cheng, E. Wang, and X. Yang, “Capacitive detection of glucose using molecularly imprinted polymers,” Biosensors and Bioelectronics, vol. 16, no. 3, pp. 179–185, 2001.

[14] S. V. Patel, T. E. Mlsna, B. Fruhberger, E. Klaassen, S. Cemalovic, and D. R. Baselt, “Chem- icapacitive microsensors for volatile organic compound detection,” Sensors and Actuators B: Chemical, vol. 96, no. 3, pp. 541–553, 2003.

[15] B. Kang, U. Yeo, and K.-H. Yoo, “Anodized aluminum oxide-based capacitance sensors for the direct detection of DNA hybridization,” Biosensors and Bioelectronics, vol. 25, no. 7, pp. 1592–1596, 2010.

[16] M. A. Nurge and S. A. Perusich, “In-line capacitance sensor for real-time water absorption measurements,” Sensors and Actuators B: Chemical, vol. 150.

[17] A. Qureshi, J. H. Niazi, S. Kallempudi, and Y. Gurbuz, “Label-free capacitive biosensor for sensitive detection of multiple biomarkers using gold interdigitated capacitor arrays,” Biosen- sors and Bioelectronics, vol. 25, no. 10, pp. 2318–2323, 2010.

[18] L. Besser and R. Gilmore, Practical RF Circuit Design for Modern Wireless Systems. Boston, MA: Artech House, 2003, vol. 1.

[19] Y. I. Kim, T. S. Park, J. H. Kang, M. C. Lee, J. T. Kim, J. H. Park, and H. K. Baik, “Biosen- sors for label free detection based on RF and MEMS technology,” Sensors and Actuators B: Chemical, vol. 119, no. 2, pp. 592–599, 2006.

[20] S. Kim, H. Melikyan, J. Kim, A. Babajanyan, J.-H. Lee, L. Enkhtur, B. Friedman, and K. Lee, “Noninvasive in vitro measurement of pig-blood D-glucose by using a microwave cavity sensor,” Diabetes Research and Clinical Practice, vol. 96, no. 3, pp. 379–384, 2012.

[21] H.-J. Lee, J.-H. Lee, S. Choi, I.-S. Jang, J.-S. Choi, and H.-I. Jung, “Asymmetric split-ring resonator-based biosensor for detection of label-free stress biomarkers,” Applied Physics Let- ters, vol. 103, no. 5, pp. 53 702–1–53 702–5, 2013.

[22] S.-G. Kim, H.-J. Lee, J.-H. Lee, H.-I. Jung, and J.-G. Yook, “A highly sensitive and label free biosensing platform for wireless sensor node system,” Biosensors and Bioelectronics, vol. 50, pp. 362–367, 2013.

[23] N. Y. Kim, R. Dhakal, K. K. Adhikari, E. S. Kim, and C. Wang, “A reusable robust radio frequency biosensor using microwave resonator by integrated passive device technology for

52 quantitative detection of glucose level,” Biosensors and Bioelectronics, vol. 67, pp. 687–693, 2015.

[24] C. Tlili, N. V. Myung, V. Shetty, and A. Mulchandani, “Label-free, chemiresistor immunosen- sor for stress biomarker cortisol in saliva,” Biosensors and Bioelectronics, vol. 26, no. 11, pp. 4382–4386, 2011.

[25] K. Yu, Z. Zhang, Y. Liu, and J. Leng, “Carbon nanotube chains in a shape memory polymer/carbon black composite: To signiﬁcantly reduce the electrical resistivity,” Applied Physics Letters, vol. 98, no. 7, pp. 074 102–1–074 102–3, 2011.

[26] C. Garc´ıa-Aljaro, L. N. Cella, D. J. Shirale, M. Park, F. J. Munoz,˜ M. V. Yates, and A. Mulchandani, “Carbon nanotubes-based chemiresistive biosensors for detection of microor- ganisms,” Biosensors and Bioelectronics, vol. 26, no. 4, pp. 1437–1441, 2010.

[27] Y.-H. Lee, J.-H. Lee, S. J. Chung, S. Lee, and B. K. Ju, “Carrier carrying capacity of one- step grown suspended carbon nanotube bridge with carbon nanotube contact electrodes: For practical one-dimensional electronics,” Applied Physics Letters, vol. 89, no. 7, pp. 073 109–1– 073 109–3, 2006.

[28] C. Jiang, X. Wang, R. Gunawidjaja, Y.-H. Lin, M. K. Gupta, D. L. Kaplan, R. R. Naik, and V. V. Tsukruk, “Mechanical Properties of Robust Ultrathin Silk Fibroin Films,” Advanced Functional Materials, vol. 17, no. 13, pp. 2229–2237, 2007.

[29] D.-H. Kim, J. Viventi, J. J. Amsden, J. Xiao, L. Vigeland, Y.-S.Kim, J. A. Blanco, B. Panilaitis, E. S. Frechette, D. Contreras, D. L. Kaplan, F. G. Omenetto, Y. Huang, K.-C. Hwang, M. R. Zakin, B. Litt, and J. A. Rogers, “Dissolvable ﬁlms of silk ﬁbroin for ultrathin conformal bio-integrated electronics,” Nature Materials, vol. 9, pp. 511–517, 2010.

[30] H. Tao, S.-W. Hwang, B. Marelli, B. An, J. E. Moreau, M. Yang, M. A. Brenckle, S. Kim, D. L. Kaplan, J. A. Rogers, and F. G. Omenetto, “Silk-based resorbable electronic devices for remotely controlled therapy and in vivo infection abatement,” Proceedings of the National Academy of Sciences of the United States of America, vol. 111, no. 49, pp. 17 385–17 389, 2014.

[31] E. N. Jones, “Development of Biopolymer Based Resonant Sensors,” Master’s thesis, Univer- sity of Dayton, 2010.

[32] J. G. Grote, N. Ogata, D. E. Diggs, and F. K. Hopkins, “Deoxyribonucleic acid (DNA) cladding layers for nonlinear optic polymer based electro-optic devices,” in Proc. SPIE, vol. 4991, 2003, pp. 621–625.

[33] B. Singh, N. S. Sariciftci, J. G. Grote, and F. K. Hopkins, “Bio-organic-semiconductor-ﬁeld- effect-transistor based on deoxyribonucliec acid gate dielectric,” Journal of Applied Physics, vol. 100, no. 2, 2006.

53 [34] C. M. Bartsch, “Development of a Field Effect Transistor Using a DNA-Biopolymer as the Semiconducting Layer,” Ph.D. dissertation, University of Dayton, 2007.

[35] J. A. Hagen, W. Li, A. J. Steckl, and J. G. Grote, “Enhanced emission efﬁciency in organic light-emitting diodes using deoxyribonucleic acid complex as an electron blocking layer,” Ap- plied Physics Letters, vol. 88, no. 17, 2006.

[36] P. Stadler, K. Oppelt, T. B. Singh, J. G. Grote, R. Schwodiauer,¨ S. Bauer, H. Piglmayer- Brezina, D. Bauerle,¨ and N. S. Sariciftci, “Organic ﬁeld-effect transistors and memory elements using deoxyribonucleic acid (DNA) gate dielectric,” Organic Electronics, vol. 8, no. 6, pp. 648–654, 2007.

[37] F. Ouchen, E. Gomez, D. Joyce, P. Yaney, S. Kim, A. Williams, A. Steckl, N. Venkat, and J. Grote, “Investigation of DNA nucleobases-thin ﬁlms for potential applications in electronics and photonics,” in Proc. SPIE, vol. 8817, 2013.

[38] A. Salimi, A. Noorbakhash, E. Shariﬁ, and A. Semnani, “Highly sensitive sensor for pico- molar detection of insulin at physiological pH, using GC electrode modiﬁed with guanine and electrodeposited nickel oxide nanoparticles,” Biosensors and Bioelectronics, vol. 24, no. 4, pp. 792–798, 2008.

[39] J. Lee, J. H. Park, Y. T. Lee, P. J. Jeon, H. S. Lee, S. H. Nam, Y. Yi, Y. Lee, and S. Im, “DNA- Base Guanine as Hydrogen Getter and Charge-Trapping Layer Embedded in Oxide Dielectrics for Inorganic and Organic Field-Effect Transistors,” Applied Materials and Interfaces, vol. 6, no. 7, pp. 4965–4973, 2014.

[40] G. Maruccio, P. Visconti, A. Valentina, S. D’Amico, A. Biasco, E. D’Amone, R. Cingolani, R. Rinaldi, S. Masiero, T. Giorgi, and G. Gottarelli, “Field Effect Transistor Based on a Mod- iﬁed DNA Base,” Nano Letters, vol. 3, no. 4, pp. 479–483, 2003.

[41] A. B. Kharitonov, M. Zayats, A. Lichtenstein, E. Katz, and I. Willner, “Enzyme monolayer- functionalized ﬁeld-effect transistors for biosensor applications,” Sensors and Actuators B, vol. 70, no. 1-3, pp. 222–231, 2000.

[42] A. Hai, D. Ben-Haim, N. Korbakov, A. Cohen, J. Shappir, R. Oren, M. E. Spira, and S. Yitzchaik, “Acetylcholinesterase-ISFET based system for the detection of acetylcholine and acetylcholinesterase inhibitors,” Biosensors and Bioelectronics, vol. 22, no. 5, pp. 605–612, 2006.

[43] A. Kisner, R. Stockmann, M. Jansen, U. Yegin, A. Offenhausser,¨ L. T. Kubota, and Y. Mourz- ina, “Sensing small neurotransmitter-enzyme interaction with nanoporous gated ion-sensitive ﬁeld effect transistors,” Biosensors and Bioelectronics, vol. 31, no. 1, pp. 157–163, 2012.

[44] A. W. Martinez, S. T. Phillips, E. Carrilho, S. W. Thomas III, H. Sindi, and G. M. Whitesides, “Simple Telemedicine for Developing Regions: Camera Phones and Paper-Based Microﬂuidic Devices for Real-Time, Off-Site Diagnosis,” Analytical Chemistry, vol. 80, no. 10, pp. 3699– 3707, 2008. 54 [45] H. Sharma and R. Mutharasan, “Half Antibody Fragments Improve Biosensor Sensitivity without Loss of Selectivity,” Analytical Chemistry, vol. 85, no. 4, pp. 2472–2477, 2013.

[46] A. Murata, T. Ooya, and T. Takeuchi, “Simple immobilization of antibody in organic/inorganic hybrid thin ﬁlms for immunosensing,” Biosensors and Bioelectronics, vol. 43, pp. 45–49, 2013.

[47] U. Sungkanak, A. Sappat, A. Wisitsoraat, C. Promptmas, and A. Tuantranont, “Ultrasensitive detection of Vibrio cholerae O1 using microcantilever-based biosensor with dynamic force microscopy,” Biosensors and Bioelectronics, vol. 26, no. 2, pp. 784–789, 2010.

[48] T. N. Huan, V. T. T. Ha, L. Q. Hung, M.-Y. Yoon, S.-H. Han, and H. Chung, “Square wave voltammetric detection of Anthrax utilizing a peptide for selective recognition of a protein biomarker,” Biosensors and Bioelectronics, vol. 25, no. 2, pp. 469–474, 2009.

[49] J. H. Lim, J. Park, J. H. Ahn, H. J. Jin, S. Hong, and T. H. Park, “A peptide receptor-based bioelectronic nose for the real-time determination of seafood quality,” Biosensors and Bioelec- tronics, vol. 39, no. 1, pp. 244–249, 2013.

[50] P. Miao, J. Yin, L. Ning, and X. Li, “Peptide-based electrochemical approach for apoptosis evaluation,” Biosensors and Bioelectronics, vol. 62, pp. 97–101, 2014.

[51] J. Hagen, W. Lyon, Y. Chushak, M. Tomczak, R. Naik, M. Stone, and N. Kelley-Loughnane, “Detection of Orexin A Neuropeptide in Biological Fluids Using a Zinc Oxide Field Effect Transistor,” ACS Chemical Neuroscience, vol. 4, no. 3, pp. 444–453, 2013.

[52] S. Nishino, B. Ripley, S. Overeem, S. Nevsimalova, G. J. Lammers, J. Vankova, M. Okun, W. Rogers, S. Brooks, and E. Mignot, “Low Cerebrospinal Fluid Hypocretin (Orexin) and Altered Energy Homeostasis in Human Narcolepsy,” Annals of Neurology, vol. 50, no. 3, pp. 381–388, 2001.

[53] J. R. Strawn, G. J. Pyne-Geithman, N. N. Ekhator, P. S. Horn, T. W. Uhde, L. A. Shutter, D. G. Baker, and T. D. Geracioti Jr., “Low cerebrospinal ﬂuid and plasma orexin-a (hypocretin-1) concentrations in combat-related posttraumatic stress disorder,” Psychoneuroendocrinology, vol. 35, no. 7, pp. 1001–1007, 2010.

[54] C. J. Winrow, K. Q. Tanis, D. R. Reiss, A. M. Rigby, J. M. Uslaner, V. N. Uebele, S. M. Doran, S. V. Fox, S. L. Garson, A. L. Gotter, D. M. Levine, A. J. Roecker, P. J. Coleman, K. S. Koblan, and J. J. Renger, “Orexin receptor antagonism prevents transcriptional and behavioral plasticity resulting from stimulant exposure,” Neuropharmacology, vol. 58, no. 1, pp. 185– 194, 2010.

[55] Y. Feng, T. Liu, X.-Q. Li, Y. Liu, X.-Y. Zhu, J. Jankovic, T.-H. Pan, and Y.-C. Wu, “Neu- roprotection by Orexin-A via HIF-1α induction in a cellular model of Parkinson’s disease,” Neuroscience Letters, vol. 579, pp. 35–40, 2014.

55 [56] B. J. Sanghavi, W. Varhue, J. L. Chavez,´ C.-F. Chou, and N. S. Swami, “Electrokinetic Pre- concentration and Detection of Neuropeptides at Patterned Graphene-Modiﬁed Electrodes in a Nanochannel,” Analytical Chemistry, vol. 86, no. 9, pp. 4120–4125, 2014.

[57] T.-Z. Wu, Y.-R. Lo, and E.-C. Chan, “Exploring the recognized bio-mimicry materials for gas sensing,” Biosensors and Bioelectronics, vol. 16, no. 9-12, pp. 945–953, 2001.

[58] N. G. G. M. Abeling, A. H. Van Gennip, H. D. Bakker, A. Heerschap, U. Engelke, and R. A. Wevers, “Diagnosis of a new case of trimethylaminuria using direct proton NMR spectroscopy of urine,” Journal of Inherited Metabolic Disease, vol. 18, no. 2, pp. 182–184, 1995.

[59] T. Higgens, S. Chaykin, K. B. Hammond, and J. R. Humbert, “Trimethylamine N-Oxide Syn- thesis: A Human Variant,” Biochemical Medicine, vol. 6, no. 4, pp. 392–396, 1972.

[60] M. A. Bain, R. Faull, G. Fornasini, R. W. Milne, and A. M. Evans, “Accumulation of trimethylamine and trimethylamine-N-oxide in end-stage renal disease patients undergoing haemodial- ysis,” Nephrology Dialysis Transplantation, vol. 21, no. 5, pp. 1300–1304, 2006.

[61] P. Dalgaard, H. L. Madsen, N. Samieian, and J. Emborg, “Biogenic amine formation and microbial spoilage in chilled garﬁsh (Belone belone belone) - effect of modiﬁed atmosphere packaging and previous frozen storage,” Journal of Applied Microbiology, vol. 101, no. 1, pp. 80–95, 2006.

[62] R. Pandeeswari and B. Jeyaprakash, “Nanostructured α-MoO3 thin ﬁlm as a highly selective TMA sensor,” Biosensors and Bioelectronics, vol. 53, pp. 182–186, 2014.

[63] S. Bourigua, S. El Ichi, H. Korri-Youssoufi, A. Maaref, S. Dzyadevych, and N. J. Renault, “Electrochemical sensing of trimethylamine based on polypyrrole-flavin-containing monooxy- genase (FMO3) and ferrocene as redox probe for evaluation of fish freshness,” Biosensors and Bioelectronics, vol. 28, no. 1, pp. 105–111, 2011.

[64] C. M. Bartsch, G. Subramanyam, J. Grote, F. K. Hopkins, L. L. Brott, and R. R. Naik, “A New Capacitive Test Structure for Microwave Characterization of Biopolymers,” Microwave and Optical Technology Letters, vol. 49, no. 6, pp. 1261–1265, 2007.

[65] G. Bracco and B. Holst, Surface science techniques. Springer Science & Business Media, 2013.

56 APPENDIX A

CTS MEASUREMENTS

Table A.1 are the calculated capacitance, dielectric constant, and loss tangent of all the CTSs measured.

Table A.1: The results from the frequency response measurements of the CTS

Thickness (nm) Capacitance (pF) Dielectric Constant, r Loss Tangent, tanδ, 10 GHz 500 2.12 5.694 0.014 500 2.02 5.426 0.015 500 2.07 5.560 0.015 500 2.07 5.560 0.015 500 1.92 5.157 0.013 1000 0.97 5.211 0.016 1000 0.87 4.674 0.018 1000 1.02 5.479 0.016 1000 0.97 5.211 0.016 1000 1.02 5.479 0.016

57 APPENDIX B

RTS MEASUREMENTS

Table B.1 are the resonances of all the RTSs measured.

Table B.1: The results from the frequency response measurements of the RTS

Thickness (nm) Frequency (GHz) Notch Depth (dB) 250 3.091 7.502 250 3.066 6.963 250 3.091 8.059 250 3.091 7.453 250 3.141 7.133 500 3.166 7.535 500 3.414 7.564 500 3.166 7.802 500 3.215 6.994 750 3.166 7.420 750 3.091 7.666 750 3.166 7.875 750 3.166 7.783 1000 3.141 7.266 1000 3.166 7.554 1000 3.091 7.622 1000 3.166 7.467 1000 3.066 6.834

58 APPENDIX C

CONTACT ANGLE MEASUREMENTS

Table C.1 are the contact angle measurements for different surfaces.

Table C.1: Contact angle measurements for different surfaces

Surface Contact Angle (◦) Surface Contact Angle (◦) Glass 44.6 Glass 39.4 Glass 42.7 Glass 38.8 Amino Silane 51.3 Amino Silane 41.5 Amino Silane 42.6 Amino Silane 45.9 Guanine 59.7 Guanine 69.5 Guanine 61.2 Guanine 60 Guanine 61.3 Guanine 61.6 Guanine 61.4 Guanine 63.4 Guanine 63.8 Guanine 58.8

59 APPENDIX D

BUFFER AND CHEMICAL PROTOCOLS

PBS (100 mM NaPO4, 150 mM NaCl, pH of 7.2) Preparation

• The amount of NaPO4 (MW = 119.98 g/mol) needed to make 500 mL of PBS is calculated to

be 5.999 g.

• The amount of NaCl (MW = 58.44 g/mol) needed to make 500 mL of PBS is calculated to be

4.383 g.

• Weigh out the correct amount of NaPO4 and NaCl and place in a beaker capable of holding

at least 500 mL.

• Measure 500 mL of deionized water in a graduated cylinder.

• Pour deionized water into the beaker containing the NaPO4 and NaCl.

• Use stirring plate to thoroughly mix the solution in the beaker.

• Use a calibrated pH meter to determine the pH.

• Use HCl or NaOH to adjust the pH to 7.2.

• Pour PBS into a sealable bottle for storage.

60 2% Amino Silane Preparation

• Aliquot 32 mL ethanol into a centrifuge tube.

• Amino silane (Purchased from Gelest, Product Number SIA0590.5) has to be opened in a

glove bag under nitrogen.

• While in nitrogen ﬁlled glove bag, aliquot 640 µL of amino silane into a centrifuge tube.

• Remove centrifuge tube of amino silane from the glove bag.

• Aliquot the amino silane out of its centrifuge tube and into the centrifuge tube of ethanol.

• Mix solution using a Vortex Mixer.

NHS-PEG4-CH3 Preparation

• NHS-PEG4-CH3 (Purchased from Thermo Fisher Scientiﬁc, Product Number PI-22341) has

to be opened in a glove bag under nitrogen.

• While in a nitrogen ﬁlled glove bag, resuspend the NHS-PEG4-CH3 in 1.1 mL of dimethyl

sulfoxide (250 mM stock).

• In order to make a 10 mL solution of 1 mM NHS-PEG4-CH3 (MW = 333.33 g), aliquot 40

µL of the resuspension into a centrifuge tube.

• Remove the centrifuge tube from the glove bag.

• Aliquot 9.96 mL of PBS into the centrifuge tube containing the NHS-PEG4-CH3 resuspen-

sion.

• Mix solution using a Vortex Mixer.

61 NHS-PEG4-Biotin Preparation

• The amount of NHS-PEG4-Biotin (Purchased from Thermo Fisher Scientiﬁc, Product Num-

ber PI-21330, MW = 588.67 g) needed to make a 10 mL of a 1 mM solution is calculated to

be 5.89 mg.

• NHS-PEG4-Biotin has to be opened in a glove bag under nitrogen.

• While in a nitrogen ﬁlled glove bag, deposit 5.89 mg of the NHS-PEG4-Biotin into a cen-

trifuge tube.

• Remove the centrifuge tube from the glove bag.

• Aliquot PBS into the centrifuge tube until the tube contains 10 mL.

• Mix solution using a Vortex Mixer.

Streptavidin-Oregon Green Preparation

• Aliquot 15 µL of Streptavidin-Oregon Green (Purchased from Thermo Fisher Scientiﬁc,

Product Number PI-21832) into a centrifuge tube.

• Aliquot 15 mL of PBS into the centrifuge tube.

• Mix solution using a Vortex Mixer.

• Store solution in the dark to avoid photobleaching the Oregon Green.