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Engineering solutions for biomaterials: self-assembly and surface-modification of polymers for clinical applications

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Engineering solutions for biomaterials: self-assembly and surface-modification of polymers for clinical applications

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Content ENGINEERING SOLUTIONS FOR BIOMATERIALS – SELF-ASSEMBLY AND SURFACE-
MODIFICATION OF POLYMERS FOR CLINICAL APPLICATIONS

By
Victoria Sun

A Dissertation Presented to the
FACULTY OF THE USC GRADUATE SCHOOL
UNIVERSITY OF SOUTHERN CALIFORNIA
In Partial Fulfillment of the Requirements for the Degree
DOCTOR OF PHILOSOPHY
(Chemical Engineering)

December 2018

Copyright 2018 Victoria Sun

For my mother,
who taught me that unconditional love overcomes all adversity.

ii

Acknowledgements
The completion of a PhD is never accomplished alone, or without the guidance and
generosity of those around us. I have been the fortunate recipient of such support from professors,
advisors, family, and friends, to whom I owe immeasurable gratitude.
I would like to thank my advisor, Professor Andrea M. Armani, for her continued guidance
and relentless support throughout the past seven years. Always on board with crazy ideas in the
lab, allowing me to completely shift my thesis during my fourth year of graduate school, essentially
beginning a new dissertation topic, Professor Armani has taught me the value of exploring new
fields and pursuing your dreams. Andrea is a beloved advisor who believed in me long before I
could do so on my own, and she has been my greatest champion. She has modeled to me what it
takes to be truly dedicated to her students and to her I owe an immense amount of gratitude.
I would like to thank my qualifying and dissertation committee members: Professors Wei
Wu, Nick Graham, Jay Ravichandran, Steve Nutt, Pin Wang, and Michael Kassner. Thank you to
Dr. Mitchell Gross for allowing me into your clinic to meet your incredible patients. Observing
and working alongside the clinical oncology team at Keck, a door of possibility was opened for
me to imagine how I could use my training in engineering to study the most complex system of all
– the human body. Additionally, I would like to thank Prof. Muhammad Sahimi for all that he
taught me of fluid dynamics and Professor Anna Wenzel and Professor Scott Banta, for allowing
me to work in their labs as an undergraduate.
The Armani Research Group is truly a family – knowledge and support are shared freely
and generously without expectation. Thank you to Dr. Xiaoqin Shen, René Zeto, Dongyu Chen,
Andre Kovach, Hyungwoo Choi, and Jinghan He. Thank you to past members of the group: Dr.
Ashley Maker, Dr. Simin Mehrabani, Dr. James Thompson, Dr. Jason Gamba, Dr. Mark Harrison,
iii

Dr. Soheil Soltani, Dr. Vinh Diep, Dr. Alexa Hudnut, Dr. Eda Gungor, Dr. Michele Lee, Dr. Erick
Moen, and Kelvin Kuo.
To my collaborators, Dr. Ploy Natnaree Siriwon, and Professors Cecilia Lopez and Pin
Wang, thank you for everything you have taught me about in vivo studies and assays. To the
students I have had the joy of working with directly: Alejandra Rios, Emma Meinke, Tara Assi,
Leah Tsui, Lea Fang, Joelle Burkhardt, Sydney Agus, Josh Neutel, Nic Murillo, and Sarah Damico,
thank you for your endless patience and enthusiasm in the lab. Special thanks to the following
technicians and their expertise offered through USC’s shared lab facilities: Ryan Park, Ivetta
Vorobyova, John Curulli, Dr Donghai Zhu, and Dr. Shuxing Li.
To the Breakfast Club, you have walked with me through seasons of joy as well as hardship,
always encouraging me in my moments of self-doubt. Thank you to Yi-Ki Cheng, Lorine Chang,
and Jennifer Lui. Special thanks to the following women I have had the privilege to call my dear
friends: Athena Last, Nieves Maldonado, Lydia Ngai, Dr. Adebola Adeniran, Samantha Conway,
Sabrina Liao, Pamela Dong, Lydia Wada, and Anni Wong. Thank you for your guidance and
continued support.
I have been blessed with loving parents and two incredible sisters. Thank you to my
parents, who have provided me with a home and life full of love and most importantly, have taught
me where to place my value and worth. Lastly, to my sisters, thank you for being my first friends,
for dreaming with me, and for being my voice of reason. Every joyful memory and every turning
point in life has occurred with you both by my side, providing me with the security and confidence
to step into the unknown. Without you, I would not have made it to the end of the PhD, and I
certainly would not have had the courage to pursue the path to medicine. DJJ and M, I love you.

iv

Table of Contents
Acknowledgements ......................................................................................................................... ii
Table of Contents ........................................................................................................................... iv
Table of Figures ............................................................................................................................ vii
List of Tables ................................................................................................................................ xv
Abstract ........................................................................................................................................... 1
Chapter 1. Overview ....................................................................................................................... 3
1-1 Introduction ........................................................................................................................... 3
1-2 Chapter Overview ................................................................................................................. 3
1-3 Chapter 1 References ............................................................................................................ 6
Chapter 2. Background ................................................................................................................... 7
2-1 Active vs. Passive Transport Across Cell Membranes ......................................................... 7
2-2 Label-free Sensing of Cell Membrane Behavior ................................................................ 10
2-3 Testing Setup ...................................................................................................................... 17
2-4 Chapter 2 References .......................................................................................................... 19
Chapter 3. Lipid Bilayer Optical Biosensors ................................................................................ 21
3-1 Background and Significance ............................................................................................. 21
3-2 Materials and Methods ........................................................................................................ 23
3-3 Results and Discussion ....................................................................................................... 26
3-4 Integrating the bilayer platform for heavy metal sensing ................................................... 34
3-5 Chapter 3 References .......................................................................................................... 46
Chapter 4. Transmembrane Protein Study .................................................................................... 50
4-1 Background and Significance ............................................................................................. 50
4-2 Engineering of the Protein Construct .................................................................................. 51
4-3 Development of the Detection System ............................................................................... 53
v

4-4 Chapter 4 References .......................................................................................................... 59
Chapter 5. Enzyme-mediated Hydrogels ...................................................................................... 60
5-1 Background and Significance ............................................................................................. 60
5-1.1 Motivation .................................................................................................................... 65
5-2 Proteins of Interest .............................................................................................................. 68
5-3 Materials and Methods ........................................................................................................ 69
5-4 Results and Discussion ....................................................................................................... 74
5-5 Chapter 5 References .......................................................................................................... 78
Chapter 6. Enhanced CAR-T cell proliferation using an enzyme-mediated crosslinking hydrogel
scaffold loaded with IL-15/IL-15Rα ............................................................................................. 81
6-1 Background and Significance ............................................................................................. 81
6-2 Protein of Interest ................................................................................................................ 82
6-3 Materials and Methods ........................................................................................................ 84
6-4 Results and Discussion ....................................................................................................... 88
6-5 Chapter 6 References .......................................................................................................... 96
Chapter 7. Live Imaging using Novel Fluorescent Hydrogels ..................................................... 98
7-1 Background and Significance ............................................................................................. 98
7-2 Dyes of Interest ................................................................................................................. 100
7-3 Materials and Methods ...................................................................................................... 104
7-4 Results and Discussion ..................................................................................................... 106
7-4 Organ Distribution Study .................................................................................................. 113
7-5 Future Work ...................................................................................................................... 113
7-6 Chapter 7 References ........................................................................................................ 115
Appendix A. Detection of Post Translational Modifications with Optical Biosensors .............. 116
Appendix A-1. Introduction .................................................................................................... 116
vi

Appendix A-2. Materials and Methods ................................................................................... 118
Appendix A-3. Results and Discussion ................................................................................... 118
Appendix A-4. Future Work .................................................................................................. 123
Appendix A-5. Appendix A References ................................................................................. 125
Appendix B. Porous Silica Sol-gel Microresonators .................................................................. 126
Appendix B-1. Introduction .................................................................................................... 126
Appendix B-2. Materials and Methods ................................................................................... 127
Appendix B-3. Results and Discussion ................................................................................... 129
Appendix B-4. Conclusions/Future Work .............................................................................. 135
Appendix B-5. Appendix B References .................................................................................. 137

vii

List of Figures
Figure 2-1. The fluid mosaic model ................................................................................................ 8
Figure 2-2. Schematic of varying modes of transport across the plasma membrane ..................... 9
Figure 2-3. Schematic of a WGM cavity where the green arrows (left) depict constructive
interference with the light entering the device (“on resonance”) and the red arrows (right) depict
light out of phase, resulting in non-continuous circulation (“off resonance”) .............................. 12
Figure 2-4. Microresonator fabrication: (a) an optical microscope image of a silica microsphere,
and (b) a scanning electron micrograph of a silica microtoroid ................................................... 16
Figure 2-5. A schematic of the optical resonator and characterization setup. .............................. 18
Figure 3-1. Chemical structures of the detergents used: (a) sodium dodecyl sulfate (SDS), and (b)
Triton X-100 (TX-100). ................................................................................................................ 22
Figure 3-2. Optical resonant cavity excitation .............................................................................. 24
Figure 3-3. Schematic of micelle saturation on a bare silica microsphere ................................... 27
Figure 3-4. Wavelength shifts of bare silica microcavity saturated in: (a) 0.2 mM TX-100, and (b)
15 mM SDS................................................................................................................................... 27
Figure 3-5. The resonant frequency shift due to liposome fusion and bilayer formation of 0.6 mM
DOPC is indicated by the red curve, and subsequent solubilization via 0.5 mM TX-100 (green)
and 15 mM SDS (purple) are shown. ............................................................................................ 30
Figure 3-6. Effects of material absorption on the quality factor of the sensor. The measurements
of the photon lifetime or Q factor was performed using the two detergents: (a) SDS and (b) TX-
100................................................................................................................................................. 31
Figure 3-7. DLS histograms of lipid-detergent micelle mixtures from the experimental wash-off
of (a) SDS, and (b) TX-100 rinses. ............................................................................................... 32
viii

Figure 3-8. UV-Vis absorption spectra of control samples: (a) 15 mM SDS, (b) 0.5 mM TX-100,
and (c) 0.6 mM SUVs. .................................................................................................................. 33
Figure 3-9. A cartoon schematic detailing the proposed mechanism of lipid bilayer solubilization
using the two detergents: SDS (below) and TX-100 (above). ...................................................... 33
Figure 3-10. A cartoon schematic of the proposed optical detection platform. ............................ 36
Figure 3-11. A cartoon schematic of the proposed mechanism: (1) self-assembly of a phopholipid
bilayer with NTA-modified headgroups, (2) a nickel ion is bound to its interaction pair, (3) a third
chemical, EDTA, displaces the nickel from the surface. .............................................................. 37
Figure 3-12. The coordinate complexes of (a) single and (b) double EDTA ligands chelated to a
nickel(II) ion. ................................................................................................................................ 38
Figure 3-13. The proposed wavelength shifts correlating to the varying concentrations of analytes
and their interaction pairs. ............................................................................................................. 39
Figure 3-14. Overall resonant frequency shifts due to the different solutions injected into the
chamber. The δλ's of interest include: small unilamellar vesicles (red), varying concentration of Ni
ions (blue, turquoise, pink, green), and EDTA (maroon). ............................................................ 40
Figure 3-15. The resonant frequency shift due to liposome fusion and bilayer formation of 1 mg/ml
1:1 DOPC:DGS-NTA is indicated by the red curve, and subsequent injection of 1X PBS is shown
in blue. The buffer-initiated solubilization suggests bilayer instability. ....................................... 41
Figure 3-16. Control experiments on a bare silica device without the presence of a lipid bilayer
containing NTA in a basic environment (pH 10). A resonant frequency shift is not observed when
the platform is exposed to buffer (black) or EDTA (red). ............................................................ 42
ix

Figure 3-17. The resonant frequency shift due to liposome fusion in multiple scans (indicated by
the red, blue, and purple curves). The signals from buffer (black and green curves) suggest that the
flowrate plays an important role in the detection signal. .............................................................. 43
Figure 3-18. The resonant frequency shift due to liposome fusion and bilayer formation of 1 mg/ml
1:1 DOPC:DGS-NTA is indicated by the black curve. ................................................................ 43
Figure 3-19. Confocal microscopy image of the small unilamellar vesicles containing NTA-
modified head groups. ................................................................................................................... 44
Figure 3-20. UV-Vis absorption spectra taken of the liposomes in buffer solution. .................... 44
Figure 3-21. UV-Vis absorption spectra of nickel(II) ions in solution (black), 25 mM EDTA (red),
and the EDTA-Ni complex (blue). ................................................................................................ 45
Figure 4-1. A schematic of the proposed assembly pathway of α-hemolysin. The nature and
kinetics of diffusing inter-monomer contacts is not yet understood (permission to use this figure
from his dissertation granted by Dr. James B. Thompson). .......................................................... 51
Figure 4-2. A hemolytic assay demonstrating the activity of the αHL-WT- R(GS)2S2G2S2H6
against intact rabbit erythrocytes. The graph represents a light-scatter vs. time plot, therefore
monitoring the hemolysis process. ................................................................................................ 52
Figure 4-3. A Coomassie-stained 12% BIS-TRIS SDS-PAGE gel following electrophoresis
showing crude IVTT products in empty pDEST14 vector with T7 promoter (Lane 1) and encoded
αHL gene (Lane 2), the product of a crude Ni2+ -NTA IMAC purification of empty vector (Lane
3) and αHL-WT + -R(GS)2S2G2S2H6 (Lane 4), and the subsequent wash off empty vector (Lane
6), and αHL-WT + -R(GS)2S2G2S2H6 (Lane 7). Lane 5: blank. ................................................... 53
Figure 4-4. A cartoon schematic depicting: (a) a solid supported lipid bilayer, and (b) the proposed
tethered lipid bilayer, allowing sufficient space for protein insertion into the membrane. .......... 55
x

Figure 4-5. Overall reaction scheme for the N-hydroxysuccinimide (NHS) functionalization
chemistry on the solid silica support. ............................................................................................ 57
Figure 4-6. Fluorescence microscopy images of samples with 2.5 µM streptavidin across the same
wafer. In (a), a silica pad is uniformly coated with a bilayer (red box), and in (b) uniform small
unilamellar vesicles (SUVs) are observed using the extrusion technique (red circles). ............... 58
Figure 5-1. Composition of the proposed hydrogel: (a) gelatin-PEG-tyramine, and (b) Heparin-
Pluronic®. The material may be loaded with therapeutics such as growth factors or cytokines prior
to mixing of the two polymer conjugates (PEG: polyethylene glycol; TA: tyramine). ................ 68
Figure 5-2.
1
H NMR spectra of the modified end group of Pluronic in dimethylsulfoxide (DMSO).
....................................................................................................................................................... 69
Figure 5-3.
1
H NMR spectra of the gelatin-PEG-tyramine conjugate in dimethylsulfoxide
(DMSO). ....................................................................................................................................... 71
Figure 5-4. UV-Vis spectra of the polymer conjugates and the gel matrix. ................................. 72
Figure 5-5. Swelling ratios of the hydrogel matrix in buffer ........................................................ 73
Figure 5-6. Release profile of β-NGF demonstrates that the hydrogel is capable of sustained release
of growth factor over the course of 3.5 weeks ex-vivo. Control: no growth factor loaded in the gel
matrix. Dual GF: β-NGF and bFGF. ............................................................................................. 75
Figure 5-7. Numbered subcutaneous injection sites in female Sprague Dawley rat. Rat was
euthanized one day post injection and the excised samples were imaged. ................................... 76
Figure 5-8. LEFT: Injection of hydrogel into bladder wall, RIGHT: Injection of hydrogel into
urethral lining. Red arrows indicate presence of hydrogel. .......................................................... 76
Figure 5-9. Comparing the material properties and synthetic route for two different injectable
therapies: a heparin-modified hydrogel vs a heparin-modified microsphere. .............................. 77
xi

Figure 6-1. ELISA results of cumulative release where (c) depicts the amount released at 18 hours,
and (d) at 192 hours. ..................................................................................................................... 89
Figure 6-2. Comparison of release from heparin-coated microspheres vs. heparin-coated
hydrogels. ...................................................................................................................................... 90
Figure 6-3. Scanning electron microscope images of freeze-dried samples of the hydrogels. ..... 90
Figure 6-4. Cell density plots of peripheral blood samples collected on day 4, from mice treated:
(a) without gel or cytokine; - gel / - IL15, (b) unloaded hydrogel; + gel / - IL15, (c) loaded hydrogel;
+ gel / + IL15, and (d) with cytokine in solution; - gel / + IL15 .................................................. 91
Figure 6-5. T cell population found in the peripheral blood (PB) of the treated mice at: (a) 4, (b) 6,
(c) 8, (d) 12, (e) 15, and (f) 18 days. ............................................................................................. 92
Figure 6-6. (a) schematic diagram of the in vivo experiment, (b) peripheral blood is collected and
assayed for T cells, (c) T cell populations of the hydrogels loaded with cytokine (red), without
cytokine (black), and cytokine without hydrogel (blue), (d) cell populations in splenocytes and
femoral bone marrow for the three groups. .................................................................................. 92
Figure 6-7. Cell density plots of isolated splenocytes from mice treated with: (a) hydrogel without
cytokine; + gel / - IL15, and (b) cytokine-loaded hydrogel; + gel / +IL15. ................................. 94
Figure 7-1. The novel dye can be incorporated into a biocompatible polymer hydrogel and injected
into the animal to evaluate the release of material over time. .................................................... 100
Figure 7-2. Fluorescent dyes demonstrate capacity for sustained emission (over 34 days) while
encapsulated in hydrogels. Gel 1: TPPy (blue); Gel 2: TPVPy (green). .................................... 101
Figure 7-3. DLS results showing the hydrodynamic radius (Rh) of the dye-encapsulated hydrogels.
(HP = heparin-Pluronic hydrogel; HP + Dye 1 = hydrogel with encapsulated TPPy dye; HP + Dye
2 = hydrogel with encapsulated TPVPy dye). ............................................................................ 102
xii

Figure 7-4. Mobility graphs indicate slightly negative surface charges on the samples. (HP =
heparin-Pluronic hydrogel; HP + Dye 1 = hydrogel with encapsulated TPPy dye; HP + Dye 2 =
hydrogel with encapsulated TPVPy dye). ................................................................................... 102
Figure 7-5. Subcutaneous injections in a chicken wing demonstrate that the fluorescent hydrogel
is visible under a UV lamp with minimal absorption. ................................................................ 103
Figure 7-6. Photograph of the animals treated with hydrogels, 1 hour post-administration. ..... 107
Figure 7-7. Lifetime measurements of the fluorescence intensity in the first 12 hours post-
administration, with increasing concentrations (0.05, 0.10. and 0.15 mg/ml) of (a) TPVPy, and (b)
TPPy dye ..................................................................................................................................... 108
Figure 7-8. Lifetime measurements of the fluorescence intensity in the first 24 hours post-
administration, with increasing concentrations (0.05, 0.10. and 0.15 mg/ml) of (a) TPVPy, and (b)
TPPy dye ..................................................................................................................................... 109
Figure 7-9. Excitation: 430 nm; Emission: 600nm; undoctored raw image files; (a) 0h, (b) 1h, (c)
2h, (d) 4h, (e) 6h, (f) 12h, (g) 24 h, (h) 48 h, (i) 72 h, (j) 120 h, (k) 168h .................................. 110
Figure 7-10. Excitation: 430 nm; Emission: 600nm; A reference region was selected to account
for background fluorescence fluctuations due to feedings; (a) 0h, (b) 1h, (c) 2h, (d) 4h, (e) 6h, (f)
12h, (g) 24 h, (h) 48 h, (i) 72 h, (j) 120 h, (k) 168h .................................................................... 111
Figure 7-11. Normalized total radiant efficiency of the fluorescent hydrogels over the course of
one week post-administration, with increasing concentrations of dye (0.05, 0.10. and 0.15 mg/ml)
..................................................................................................................................................... 112
Figure 7-12. Normalized total radiant efficiency of the fluorescent hydrogels in the first 24 h post-
administration, with increasing concentrations of dye (0.05, 0.10. and 0.15 mg/ml) ................. 112
xiii

Figure 7-13. Organ distribution study of the following samples from top left: lungs, fatty tissue,
liver, kidneys, small and large intestines, urine, and blood. ....................................................... 113
Figure A-1. Arginine methylation by protein arginine methlytransferases (PRMTs). ............... 117
Figure A-2. Optical microscope image of the detection platform, a silica whispering gallery mode
resonator. ..................................................................................................................................... 118
Figure A-3. The real-time detection of arginine at varying concentrations to determine the limit of
detection, where (a) shows raw resonant wavelength shifts, (b) shows the line angle diagram of L-
arginine, and (c) summarizes the overall shift at each concentration. ........................................ 119
Figure A-4. Real-time detection of ADMA at varying concentrations where (a) shows the raw
resonant wavelength shifts in real-time, and (b) summarizes the overall shift at each concentration.
..................................................................................................................................................... 120
Figure A-5. Real-time detection of arginine in PBS, where: (a) depicts raw resonant wavelength
shifts, and (b) summarizes the overall shift at each concentration. ............................................ 121
Figure A-6. Real-time detection of ADMA in PBS across a range of concentrations. .............. 121
Figure A-7. Desorption measurement of arginine, where (a) overall resonant wavelength shifts,
and (b)indicates reproducibilty of the detection system ............................................................. 122
Figure A-8. Desorption measurements of arginine with surface-modified detection platform, where
(a) illustrates the mechanism by which a silane group conjugates to amine groups, and (b) shows
the real-time detection of 0.55mM arginine solutions. ............................................................... 123
Figure B-1. Two surfactants of varying hydrocarbon tail lengths were investigated. ................ 128
Figure B-2. Optical micrograph images of the silica pads after spin-coating. ........................... 131
Figure B-3. Scanning electron microscope images of the spun-coated sol-gels on silica pads on
silicon. (b) interface of silica pads on the silcon wafer is indicated by the white arrow. ........... 131
xiv

Figure B-4. Scanning electron microscopy images of the spun-coated silica disks elevated on a
silicon wafer. ............................................................................................................................... 132
Figure B-5. Scanning electron microscope images indicate cracking of the toroid device and wafer,
due to the annealing procedure. Images taken at varying objectives: (a) X 75, (b) X 250, (c) X 650,
and (d) X 1,200. .......................................................................................................................... 133
Figure B-6. Scanning electron microscopy images taken of the porous sol-gel thin films at varying
objectives: (a) X 5,000, (b) X 10,000, (c) X 20,000, and (d) X 30,000 ...................................... 134
Figure B-7. Scanning electron microscopy image of porous sol-gels on a microtoroid via spin-
coating technology. Pores are observed only on the inner edge of the toroid, as well as on the
silicon pillar and substrate. ......................................................................................................... 135
Figure B-8. Scanning electron microscopy image of a microsphere on a silicon wafer at: (a) X
1,600, and (b) X 7,000. ............................................................................................................... 136

xv

List of Tables
Table 2-1. Examples of various types of membrane transport proteins found in human cells ..... 10
Table 3-1. Effects of a supported lipid bilayer on Q and refractive index. ................................... 28
Table 3-2. Refractive Index Measurements .................................................................................. 29
Table 3-3. List of drinking water contaminants and Maximum Contamination Level (MCLs) ... 35
Table 5-1. Biomaterials find applications across a wide range of healthcare products ................ 61
Table 5-2. Summary of the parameters used in hydrogel design, structure, and characterization 61
Table 5-3. Hydrogels can be composed of natural, synthetic, and biohydrid hydrophilic polymers
....................................................................................................................................................... 62
Table 5-4. Addressing challenges and material property requirements (HRP: horse radish
peroxidase; H2O2: hydrogen peroxide). ........................................................................................ 67
Table 6-1. Summary of the four treatment groups employed for sc-injections (n=3). Notation is
indicated as a superscript .............................................................................................................. 87
Table 7-1. DLS results summary of the hydrogel samples ......................................................... 103
Table 7-2. Caliper measurements of the hydrogel injections ..................................................... 107
Table A-1. Varying pH values of L-Arginine and ADMA solutions in both MilliQ water and PBS.
..................................................................................................................................................... 119
Table B-1. Classification of Porous Materials Based on Pore Diameters .................................. 126
Table B-2. Sol-gel solutions at varying parameters, and their characterized pore sizes. ........... 130

1

Abstract
Biomaterials are substances engineered to interact with biological systems for a medical
purpose. The study of biomaterials and biomimetic systems spans many fields including medicine,
biology, chemistry, tissue engineering, and materials science, and they find a range of applications
like optical diagnostics, imaging, and therapeutics. These applications rely heavily on self-
assembly, biocompatibility, and surface modification of polymers. During the course of this PhD,
technical and scientific contributions were made to all three areas. Specifically, novel methods for
modifying the surface of optical sensors to understand cell membrane behavior were developed,
and biodegradable materials which permit the prolonged release of a therapeutic over an extended
period of time were demonstrated.
Over the course of my graduate study, this body of work has carried me through a variety
of seemingly disparate experiments, from controlled environments outside of living organisms, or
in vitro, to experimentation using whole living organisms, or in vivo. In whole however, it has
yielded insight into different methods of performing experiments in both artificial environments
as well as in cell cultures and animals, transitioning from fundamental basic science research to
applied science. As such, these projects share a unifying goal of engineering solutions to clinical
challenges such as the mechanisms of drug delivery and finding treatments for human disease and
disorders such as cancer.
In the first part of the dissertation, optical microresonators with lipid membrane coatings
are designed to observe cell membrane behavior on a nanoscale, in vitro platform. The self-
assembly of a lipid bilayer from a solution of micelles is observed in real-time on an optical
platform, demonstrating its capacity to operate as a biosensor. Observing interactions on the
nanoscale is best performed in parallel with studies in living systems, or in vivo. This is explored
2

in the second part of the dissertation with animal studies, where the administration of a biomaterial,
a hydrogel, was utilized to deliver protein therapeutics. This was complemented by a live imaging
study, observing the trajectory of the implantable material in vivo using dye-encapsulated
hydrogels, demonstrating biocompatibility as well as its potential as a candidate in future clinical
studies.

3

Chapter 1. Overview
1-1 Introduction
Over the past 50 years, biomaterials have gained recognition for their clinical relevance.
They include a variety of materials, which can be either natural or synthetic, alive or lifeless, and
are commonly made up of multiple components. While biomaterials are used in a wide spectrum
of human diseases, the projects pursued as part of this thesis cover only some of the applications
found in medicine today. In this dissertation, the types of biomaterials range from coatings for
label-free sensing to implantable devices to bioimaging probes for live animal studies.
Within the study of biomaterials, the rapidly evolving interdisciplinary field of
nanomedicine has led to significant advances in the treatment and understanding of human disease.
Nanoparticles, which include liposomes, micelles, and polymeric systems, have revolutionized
drug and gene delivery to specific disease sites in the body
1-7
. Due to their small size, ease of
production and administration, increased stability, biocompatibility, and capacity for
functionalization, these materials make ideal “carriers of choice” for drug and gene delivery. In
the following chapters, the use of various commercially available materials, including
phospholipids and polymers, are explored and briefly discussed below. From developing optical
sensing platforms to modifying polymeric gels to afford longer circulation time and targeting
potential to imaging real-time clearance of nanoparticles from the circulation, this dissertation
explores how nanoparticles can be tailored to address specific challenges in medicine.
1-2 Chapter Overview
Chapter 2 introduces the fluid mosaic model and provides background on the lipid bilayer
and active vs. passive transport across the cell membrane. It details several forms of label-free
sensing to observe these behaviors in biological systems, highlighting the use of optical
4

microresonators, which is the sensing platform on which both experimental and theoretical
measurements are performed. Optical microresonators, their basic features, the fabrication process,
and the operation of such devices, are discussed especially as they relate to biological systems of
interest developed later in the dissertation.
Chapter 3 explores the use of an optical biosensor platform to observe the self-assembly of
a lipid bilayer and its subsequent solubilization, setting the groundwork for label-free sensing of
biological systems. By establishing the capacity to observe in real-time lipid bilayer dynamics,
without the need for fluorescence detection, further investigations into the ability to detect heavy
metals on a lipid bilayer coated optical resonator were explored. This work included investigations
into surface-modified phospholipids and tuning the bilayer properties to enhance detection without
compromising the integrity of the lipid membrane biomaterial.
Chapter 4 further explores the optical biosensor platform as a means for tracking the self-
assembly of a membrane protein, alpha hemolysin. This transmembrane protein is commonly used
in studies of controlled delivery of ions and small organic compounds across plasma membranes.
Efforts here were focused on engineering the protein construct.
In Chapter 5, efforts were shifted to address a specific healthcare challenge as I began the
transition from basic to applied science research. In a collaboration with the Department of
Urology at USC, the goal of this project was to develop an implantable device for the treatment of
stress urinary incontinence (SUI). A hydrogel composed of biocompatible polymers that have been
surface modified with functional groups to enhance protein affinity was synthesized and evaluated
as a treatment model for rat models of SUI. This chapter details the synthesis and preliminary
evaluations of biocompatibility in a living organism, setting the groundwork for using the hydrogel
5

scaffold as a vehicle in the delivery of protein therapeutics, enabling the expansion to other
applications and clinical disease models.
Chapter 6 details the use of the previously synthesized hydrogel to deliver the cytokine,
IL-15, known for its potential as a cancer therapeutic. In this project, cytokine-loaded hydrogels
were subcutaneously injected into immunodeficient mice treated with CAR-T cell therapy.
Enhanced T cell proliferation in the peripheral blood of the animals was measured, demonstrating
its capacity as an adjuvant cancer therapy in in vivo studies.
In Chapter 7, we report the use of two fluorescent hydrogel-based probes for long-term
tracking and imaging applications. As the hydrogel degrades in vivo, the images reveal that the
hydrogel-based fluorescent probes preferentially accumulate into the brown fat regions of the fatty
tissue in the back, and are excreted through the digestive tract in approximately one week. The
development of fluorescence emissive materials makes the hydrogel a promising candidate for
bioimaging applications, especially as an alternative to more traditional fluorophores (e.g.
rhodamine B, FITC, and quantum dots), which are susceptible to photobleaching and suffer from
limited tissue penetration.
The following Appendices detail projects that were undertaken by students that I mentored.
Strategies for completing the research are discussed. In Appendix A, we investigated post-
translational modifications using silica whispering gallery mode resonators. Results from the real-
time detection of small molecules in solution are discussed. Appendix B explores the challenges
and methods employed in coating a whispering gallery mode resonator with porous sol-gels. These
investigations in tuning the sol-gel thin film properties were conducted by modifying synthesis
and device fabrication procedures.

6

1-3 Chapter 1 References
1. Kommareddy, S.; Amiji, M., Biodistribution and pharmacokinetic analysis of long‐
circulating thiolated gelatin nanoparticles following systemic administration in breast cancer‐
bearing mice. Journal of pharmaceutical sciences 2007, 96 (2), 397-407.
2. Torchilin, V., Targeted polymeric micelles for delivery of poorly soluble drugs. Cellular
and Molecular Life Sciences CMLS 2004, 61 (19-20), 2549-2559.
3. Torchilin, V. P., Recent advances with liposomes as pharmaceutical carriers. Nature
Reviews Drug Discovery 2005, 4, 145.
4. Shenoy, D.; Little, S.; Langer, R.; Amiji, M., Poly (ethylene oxide)-modified poly (β-
amino ester) nanoparticles as a pH-sensitive system for tumor-targeted delivery of hydrophobic
drugs: part 2. In vivo distribution and tumor localization studies. Pharmaceutical research 2005,
22 (12), 2107-2114.
5. Shenoy, D. B.; Amiji, M. M., Poly (ethylene oxide)-modified poly (ɛ-caprolactone)
nanoparticles for targeted delivery of tamoxifen in breast cancer. International journal of
pharmaceutics 2005, 293 (1-2), 261-270.
6. Kommareddy, S.; Amiji, M., Preparation and evaluation of thiol-modified gelatin
nanoparticles for intracellular DNA delivery in response to glutathione. Bioconjugate chemistry
2005, 16 (6), 1423-1432.
7. Kaul, G.; Amiji, M., Tumor-targeted gene delivery using poly (ethylene glycol)-modified
gelatin nanoparticles: in vitro and in vivo studies. Pharmaceutical research 2005, 22 (6), 951-961.
7

Chapter 2. Background
2-1 Active vs. Passive Transport Across Cell Membranes
Biological membranes, made up of a complex mixture of phospholipids, proteins, sterols,
and other components, spatially define the cell and govern numerous fundamental biological
processes. As the gatekeepers of signal recognition, transduction and amplification, biological
membranes are responsible for maintaining a charge and concentration gradient, and serve as the
primary source of communication between what stays “inside” versus “outside” (e.g. intake of
food and excretion of waste). Furthermore, dynamic protein complexes are embedded throughout
the membrane, giving rise to numerous interdependent physical phenomena. For example,
substances such as neurotransmitters interact directly with specific protein receptor sites, while
anesthetics freely diffuse across the hydrophobic regions of the membrane. The delivery of
therapeutic agents across the plasma membrane is fundamental to medicine and therefore,
developing a comprehensive understanding of its structure and function is critical.
The cytosol of nearly all prokaryotes and eukaryotes is surrounded by a phospholipid
bilayer called the plasma membrane, with its inner and outer layers referred to as leaflets. Since
the forces holding the entire cellular membrane together are intermolecular, the membrane is
considered fluid. Its parts, containing phospholipids, cholesterol and proteins, can move laterally
but do not separate due to hydrophobic and hydrophilic interactions between individual lipids.
This is known as the fluid mosaic model of a membrane (Figure 2-1). Embedded within the plasma
membrane are proteins, which act as transporters, receptors, attachment sites, identifiers, adhesive
proteins, and enzymes. Membrane proteins are distributed asymmetrically throughout the
membrane and between the leaflets. As transporters, these proteins select which solutes can enter
or leave the cell. They also act as receptors that are responsible for receiving chemical signals from
8

the cellular environment. Other membrane proteins act as identifiers, which other cells recognize,
and even other membrane proteins aid in adhesion of one cell to another.

Figure 2-1. The fluid mosaic model

Natural membranes are semipermeable to most compounds to some degree. The two
aspects of a compound affecting its semipermeability include size and polarity. The larger the
molecule, or the greater its polarity, the less permeable it becomes. A natural membrane is
generally impermeable to polar molecules with a molecular weight greater than 100 without some
type of assistance. For example, very large lipid soluble (nonpolar) molecules such as steroids can
easily pass through the membrane, while sodium ions, which possess a complete charge, have
more difficulty moving through the membrane, even though they are smaller than water molecules.
Most diffusion of polar or charged molecules across a natural membrane takes place
through leakage channels, or incidental holes, created by the irregular shapes of integral proteins
(Figure 2-2). This is called passive diffusion
1
. In order to transport larger or more charged
molecules however, transport or carrier proteins are designed to facilitate the diffusion of these
necessary molecules across the membrane. There are several mechanisms used by transport
proteins, but in facilitated diffusion, diffusion occurs down the electro-chemical gradient. Most
human cells rely on facilitated diffusion for their glucose supply.
9

= Na
+
= K
+
facilitated
diffusion
passive
diffusion
active transport
Na
+
/K
+
pump
= glucose
= steroid
ATP

Figure 2-2. Schematic of varying modes of transport across the plasma membrane

A living organism must also be able to maintain a concentration gradient of some nutrients
against its electrochemical gradients. In contrast to the previously discussed facilitated diffusion,
active transport requires the expenditure of energy to transport a compound against its
electrochemical gradient. Active transport can be accomplished by the direct expenditure of ATP.
In further cases, it can also be accomplished indirectly by using ATP to create an electrochemical
gradient, and then use energy of the gradient to acquire or expel a molecule. This is called
secondary active transport. Below, Table 2-1 includes examples of the different types of membrane
transport proteins found in the human cell.

10

Table 2-1. Examples of various types of membrane transport proteins found in human cells
Transport Mech. Type Examples Significance
Passive transport Simple Drug delivery Many small molecule
pharmaceutical drugs
including steroids are
thought to transport across
biological membranes via
passive diffusion at a rate
related to their lipophilicity.
Channel Voltage-gated
ion channels
Voltage-gated calcium
channels play an important
role in neurotransmitter
release in pre-synaptic nerve
endings.
Facilitated Glucose In red blood cells (RBCs),
glucose concentration is
carefully regulated so that it
is normally higher than
intracellular concentrations,
and is transported across the
membranes by facilitated
diffusion
Active transport Primary Na
+
/K
+
pump The Na
+
/K
+
-ATPase helps
maintain resting potential,
effect transport, and regulate
cell volume.
Secondary G-proteins When neurotransmitters are
released into the synapse, the
second messenger system is
preferred for prolonged
changes involved in memory.

2-2 Label-free Sensing of Cell Membrane Behavior
Over the past several decades, various techniques have been explored to study
biomolecular interactions within the membrane and within lipid bilayers, the most popular of
which involves fluorescent or radioisotope labeling. However, experimental results based solely
on fluorescent techniques can be confounded due to interactions between the probe molecules and
the lipids. In addition, the disadvantages of isotopic labeling methods include: costliness, time
constraints due to their short half-lives, health and environmental hazards, potential occurrences
11

of alterations in proteins due to the chemistry involved, and limitation to the number of proteins
and variables that can be examined at one time. Therefore, label-free methods are desirable.
Two label-free techniques which have successfully detected lipid bilayer dynamics are
mechanical cantilevers
2
and surface plasmon resonance (SPR) sensors
3-4
. SPR offers a sensitivity
advantage over mechanical cantilevers due to its capacity to maintain performance in aqueous
environments. However, SPR signals are known to be dependent on probe density, surface
uniformity, and temperature and refractive index changes in the bulk solution. At present, there
are two approaches to circumvent such limitations: (1) change the detection mechanism (e.g., use
imaging methods)
5-7
, or (2) reduce the evanescent field length (e.g., use alternative evanescent
field methods). In the presented work, efforts have been focused on the latter, with the aim of
reducing bulk effects.
The specific sensor platform studied in the present thesis is the whispering gallery mode
(WGM) optical cavity
8
. Resonant cavities confine specific wavelengths of light, also known as
the resonant wavelengths, which are defined by the device’s optical and geometrical properties.
Because the circulating optical field partially interacts with the environment, these devices are
ideally suited for detection applications
9-10
. In previous work, resonant cavity sensors have
demonstrated ultra-sensitive detection of polymer film morphology changes and protein binding
kinetics by monitoring changes in the photon lifetime of the cavity and the resonance wavelength
11-
13
.
2-2.1 Whispering Gallery Mode Resonators
In a whispering gallery mode cavity, the refractive index contrast between the cavity device
and its surrounding environment dictates what wavelengths remain “on resonance”. These
wavelengths of light are known as resonance wavelengths, as shown in Figure 2-3. When the
12

device is “on resonance”, the total round trip is equal to an integral number of wavelengths, in
which constructive interference with the light entering the device is observed.

Figure 2-3. Schematic of a WGM cavity where the green arrows (left) depict constructive
interference with the light entering the device (“on resonance”) and the red arrows (right) depict
light out of phase, resulting in non-continuous circulation (“off resonance”)

Whispering gallery mode resonators are most frequently fabricated into the following
geometries: microrings, microdisks, microspheres, and microtoroids, with microspheres and
microtoroids being the most efficient due to their high photon lifetimes. They can be fabricated
from a variety of materials including silicon, silica, silicon nitride, etc. The projects discussed in
the present report use silica microspheres and microtoroids as the devices of interest. Their
increased photon storage time makes them more desirable in biological and chemical detection
applications
14-16
.
Silica microspheres are fabricated by a facile method of melting optical fibers under a CO2
laser. Advantages include an efficient, high-throughput fabrication protocol, and high quality,
atomically smooth devices optimal for biodetection. Silica microtoroids are fabricated on silicon
wafers using photolithography. While the fabrication may be more costly and time-intensive,
microtoroids have the advantage of being easily integrated with other electronic/optical
components.
13

2-2.2 Quality Factor
A device’s efficiency is calculated by the unit-less variable Q, also known as the quality
factor. It is used to characterize the photon storage lifetime. The higher the quality factor, the
longer the circulation of photons within the periphery of the cavity. For example, silica
microspheres and microtoroids have a quality factor of Q > 10
8
-10
10
. The quality factor is governed
by many possible loss mechanisms, comprised of both intrinsic and extrinsic losses
8
. These loss
mechanisms affect the circulation time of photons in a WGM optical cavity and can be described
by the following equation:
Qtot
-1
= Qloaded
-1
= Qmeasured
-1
= Qint
-1
+ Qext
-1
= Qmat
-1
+ Qss
-1
+ Qrad
-1
+ Qcont
-1
+ Qcoup
-1
where,
Qmat
-1
= material loss
Qss
-1
= surface scattering loss
Qrad
-1
= radiation loss (curvature or bending loss)
Qcont
-1
= contamination loss
Qcoup
-1
= coupling loss

In the above equation, the material (Qmat
-1
), surface scattering (Qss
-1
), radiation (Qrad
-1
), and
contamination (Qcont
-1
) losses contribute to the intrinsic quality factor (Qint
-1
), while the extrinsic
quality factor (Qext
-1
) is due to coupling losses (Qcoup
-1
). High efficiency waveguides, such as
tapered optical fibers, reduce extrinsic losses. As a result, the dominant intrinsic loss mechanism
is the material loss of the cavity, allowing the approximation Qload~Qmat to be made
17-18
. Qmat can
be analytically calculated according to the expression: Qmat = 2πneff /λαeff, where αeff is the effective
material absorption and neff is the effective refractive index. neff can be calculated by adding up the
product of refractive index and the portion of the optical field present in each part of the system
(the cavity and its surrounding environment). For example, for a silica cavity in air, the effective
refractive index is calculated by the following:
neff = βnsilica + δnair
14

β + δ = 1
where nsilica and nair are the refractive indices of silica and air, respectively, β is the portion of the
optical field in the silica cavity, and δ is the portion in the surrounding air. αeff can be calculated
by adding up the product of material attenuation and the portion of the optical field present in each
part of the system (cavity and surrounding environment). For the same cavity as described above,
in air the effective material loss can be calculated by the following:
αeff = βαsilica + δαair
β + δ = 1
Qss, the surface scattering loss, is expressed by a combination of two scattering losses, surface
scattering (Qss) and internal scattering (Qis):

𝑄𝑄 𝑠𝑠𝑠𝑠 =
𝐾𝐾 𝐾𝐾 + 1
3 𝜆𝜆 3
𝑅𝑅 8 𝑛𝑛 𝑒𝑒𝑒𝑒𝑒𝑒
𝜋𝜋 2
𝐵𝐵 2
𝜎𝜎 2

𝑄𝑄 𝑖𝑖 𝑠𝑠 =
3 𝜆𝜆 3
𝐾𝐾 4 𝜋𝜋 2
𝑝𝑝 2
𝜅𝜅 Т 𝛽𝛽 𝑇𝑇 𝑛𝑛 𝑒𝑒𝑒𝑒𝑒𝑒
7

where K defines the internal reflection condition, λ is the wavelength of light, R is the radius of
the cavity, neff is the effective refractive index, p is the Pocklet coefficient, κ is the Boltzmann
constant, T is the melting temperature, βT is the isothermic compressibility, σ and B are the root
mean square (rms) size and the correlation length of surface inhomogeneities, respectively
19
. Since
the devices fabricated in the proposed projects are all reflowed under a CO2 laser resulting in
atomically smooth surfaces, the loss due to surface scattering can be neglected.
Qrad is the radiation loss, which can be eliminated by an appropriate selection of device
size. It is known that Qrad
-1
vanishes exponentially with increasing microcavity size and that once
15

the device diameter approaches a minimum size threshold, total internal reflection no longer
occurs, resulting in radiation loss.
Qcont
-1
can be eliminated by avoiding contact with external contaminates including moisture
and dust. Contamination on the surface of the cavity can be avoided by proper storage of the
device, as well as treatment with oxygen plasma prior to experiments.
2-2.3 Device Fabrication
The silica microspheres are fabricated from optical fiber purchased from Newport (Irvine,
CA), using a conventional CO2 laser reflow method
20
. Briefly, an optical fiber is stripped of the
polymer coating, cleaved, and reflowed under a CO2 laser, forming a spherical, atomically smooth
silica device. The resulting spheres range between 150 and 200 microns in diameter (Figure 2- 4).
Silica microtoroids are fabricated using a well-established photolithography process in the
cleanroom
21
. First, a silicon wafer with 2 µm thick thermally grown silica is cleaned with a series
of solvents in the following order: acetone, methanol, and isopropanol. The wafer is dried with an
air gun and placed on a 120˚C hotplate for 2 minutes to further dehydrate the surface. In the next
step, a 2 minute vapor deposition of hexamethyldisilazane (HMDS) using a small volume of liquid
HMDS under a glass chamber enhances the adhesion between the silica wafer and photoresist.
Positive photoresist S1813 (Shipley) is spun coated on the surface of the wafer for 45 seconds at
3000 rpm, after which the sample undergoes a soft bake at 95˚C for 2 minutes. The wafer is placed
on a Karl Zeiss MJB 3 photomask aligner underneath a chrome light-field mask containing arrays
of circular pads. Upon exposure to UV (80mJ/cm
2
), the sample is placed in MF-321 developer
(Shipley). Since S1813 is a positive photoresist, all sections of the wafer exposed to UV light are
washed off with the developer, with the photoresist covered circular pads remaining. The sample
is further rinsed with deionized water (DIW) and dried with an air gun. A hard bake at 110˚C for
16

2 minutes reflows the pad edges of remaining photoresist and further hardens the sample. The
sample is immersed into a solution of buffered oxide etchant (BOE) (Transene), during which all
the silica exposed by the developer is removed, leaving elevated pads on silicon. Once the wafer
has become hydrophobic, indicating that all the uncovered silica has been etched away, the sample
is removed from the BOE and rinsed with DIW. The samples are rinsed with acetone, methanol,
and isopropanol to wash off the photoresist, dried with an airgun, and cut into small wafers each
containing approximately 5-8 pads using a diamond scribe. Next, the samples are placed into a
xenon difluoride (XeF2) etching chamber, which isotropically etches silicon, resulting in elevated
silica microdisks on silicon pillars. The chemical reaction of the etching process is shown below:
2XeF2 (s) + Si (s)  2Xe (g) + SiF4 (g)
Xenon difluoride, a white crystalline solid at room temperature, will sublimate at pressures lower
than its vapor pressure at room temperature (~3800 mTorr), and are introduced to the samples
through a cyclic (pulsed) process. The number of pulses is determined based on the pad size and
desired undercut. In the last step, the microdisks are reflowed under CO2 equipped with a Synrad
10.6µm wavelength, 70 W laser with a beam intensity profile of an approximate Gaussian
distribution. The disk is centered under the CO2 laser and irradiated, inducing surface tension in
the silica and resulting in an atomically smooth toroidal geometry (Figure 2- 4).

Figure 2- 4. Microresonator fabrication: (a) an optical microscope image of a silica microsphere,
and (b) a scanning electron micrograph of a silica microtoroid
17

2-3 Testing Setup
High efficiency waveguides, such as tapered optical fibers, reduce extrinsic losses. As a
result, the dominant intrinsic loss mechanism is the material loss of the cavity, allowing the
approximation Qload~Qmat to be made
17-18
. A tapered optical fiber waveguide is coupled to the
microcavity from a narrow-linewidth (300 kHz) tunable laser by Newport Velocity Series (Irvine,
CA).
Fabrication of the waveguide involves heating a stripped optical fiber over a hydrogen
torch while simultaneously stretching the fiber using a motorized two-axis stage controller by
Sigma Koki (Tokyo, Japan), until the fiber is tapered with a ~700 nm diameter waist region, or a
value close to the wavelength of the selected laser.
To align the waveguide with the cavity, top and side view camera systems are used in
conjunction with nanopositioning stages. The output signal from the waveguide is collected by a
photodetector, and sent to a high-speed digitizer/oscilloscope to view the transmission spectra.
This evanescent field approach allows for minimal coupling loss, maintaining the high intrinsic Q
of the cavity. As the laser is continuously scanned, a resonant wavelength is identified and tracked
on a high-speed oscilloscope and integrated digital PCI card. The Q is measured in the under-
coupled regime to minimize any linewidth distortion due to thermal non-linear effects. A
schematic of the optical setup can be found in Figure 2-5. The signal is recorded using a LabView
program and the data is further analyzed in Origin.
By fitting the resonant peaks to a Lorentzian curve, the full-width-half-max (FWHM) is
used to determine the loaded quality factor (Q) of the device according to Qload = λ/δλ, where λ =
resonant wavelength, and δλ = full-width-half-max (FWHM) of the spectra corresponding to the
linewidth of the microcavity.
18

Figure 2-5. A schematic of the optical resonator and characterization setup.

19

2-4 Chapter 2 References
1. Baldwin, S. A., Mammalian passive glucose transporters: members of an ubiquitous family
of active and passive transport proteins. Biochimica et Biophysica Acta (BBA) - Reviews on
Biomembranes 1993, 1154 (1), 17-49.
2. Pera, I.; Fritz, J., Sensing Lipid Bilayer Formation and Expansion with a Microfabricated
Cantilever Array. Langmuir 2007, 23 (3), 1543-1547.
3. Glasmästar, K.; Larsson, C.; Höök, F.; Kasemo, B., Protein Adsorption on Supported
Phospholipid Bilayers. Journal of Colloid and Interface Science 2002, 246 (1), 40-47.
4. Keller, C. A.; Glasmästar, K.; Zhdanov, V. P.; Kasemo, B., Formation of Supported
Membranes from Vesicles. Phys Rev Lett 2000, 84 (23), 5443-5446.
5. Nikitin, P. I.; Gorshkov, B. G.; Nikitin, E. P.; Ksenevich, T. I., Picoscope, a new label-free
biosensor. Sensors and Actuators B: Chemical 2005, 111–112 (0), 500-504.
6. Özkumur, E.; Needham, J. W.; Bergstein, D. A.; Gonzalez, R.; Cabodi, M.; Gershoni, J.
M.; Goldberg, B. B.; Ünlü, M. S., Label-free and dynamic detection of biomolecular interactions
for high-throughput microarray applications. Proceedings of the National Academy of Sciences
2008, 105 (23), 7988-7992.
7. Daaboul, G. G.; Vedula, R. S.; Ahn, S.; Lopez, C. A.; Reddington, A.; Ozkumur, E.; Ünlü,
M. S., LED-based Interferometric Reflectance Imaging Sensor for quantitative dynamic
monitoring of biomolecular interactions. Biosensors and Bioelectronics 2011, 26 (5), 2221-2227.
8. Matsko, A. B.; Ilchenko, V. S., Optical Resonators with Whispering-Gallery Modes-Part
I: Basics. IEEE Journal of Selected Topics in Quantum Electronics 2006, 12 (1), 3-14.
9. Hunt, H. K.; Armani, A. M., Label-Free Biological and Chemical Sensors. Nanoscale
2010, 2 (9), 1544-1559.
10. Luchansky, M. S.; Bailey, R. C., High-Q Optical Sensors for Chemical and Biological
Analysis. Analytical Chemistry 2012, 84, 793-821.
11. Choi, H. S.; Ismail, S.; Armani, A. M., Studying polymer thin films with hybrid optical
microcavities. Optics Letters 2011, 36 (12), 2152-2154.
12. Soteropulos, C.; Hunt, H.; Armani, A. M., Determination of binding kinetics using
whispering gallery mode microcavities. Applied Physics Letters 2011, 99 (10), 103703.
13. Cheema, M. I.; Mehrabani, S.; Peter, Y.-A.; Armani, A. M.; Kirk, A. G., Simultaneous
measurement of quality factor and wavelength shift by phase shift microcavity ring down
spectroscopy. Optics Express 2012, 20 (8), 9090-9098.
14. Armani, A. M., Label-free, single-molecule detection with optical microcavities (August,
pg 783, 2007). Science 2011, 334 (6062), 1496-1496.
15. Armani, A. M.; Vahala, K. J., Heavy water detection using ultra-high-Q microcavities.
Optics Letters 2006, 31 (12), 1896-1898.
16. Vollmer, F.; Arnold, S., Whispering-gallery-mode biosensing: label-free detection down
to single molecules. Nature methods 2008, 5 (7), 591-596.
17. Gorodetsky, M. L.; Savchenkov, A. A.; Ilchenko, V. S., Ultimate Q of optical microsphere
resonators. Optics Letters 1996, 21 (7), 453-455.
18. Armani, A. M.; Armani, D. K.; Min, B.; Vahala, K. J.; Spillane, S. M., Ultra-high-Q
microcavity operation in H2O and D2O. Applied Physics Letters 2005, 87 (15), 151118.
19. Zhang, X.; Choi, H. S.; Armani, A. M., Ultimate quality factor of silica microtoroid
resonant cavities. Applied physics letters 2010, 96 (15), 153304-153304-3.
20

20. Vernooy, D. W.; Ilchenko, V. S.; Mabuchi, H.; Streed, E. W.; Kimble, H. J., High-Q
measurements of fused-silica microspheres in the near infrared. Optics Letters 1998, 23 (4), 247-
249.
21. Armani, D.; Kippenberg, T.; Spillane, S.; Vahala, K., Ultra-high-Q toroid microcavity on
a chip. Nature 2003, 421 (6926), 925-928.

21

Chapter 3. Lipid Bilayer Optical Biosensors
3-1 Background and Significance
Understanding transport across and motion within the cell membrane can enable numerous
advances in drug design and improve our fundamental understanding of biological processes.
However, because of the inherent complexity and cell-to-cell variation, the experimental results
are inconsistent and convoluted. One approach which is commonly used to address this challenge
is constructing artificial membranes or lipid bilayers using self-assembly techniques
1-5
.
The most straightforward method for creating a lipid bilayer is to use commercially
available phospholipids and rely on the hydrophobic/hydrophilic interactions between the fatty
acyl chains of phospholipid and/or glycolipid molecules with the surrounding environments
2
. Van
der Waals interactions among the hydrocarbon chains favor close packing of these hydrophobic
moieties, while hydrogen bonding and electrostatic interactions between the polar head groups and
water molecules also stabilize the lipid bilayer. Depending on the relative concentration and tail
length, different suspended structures, such as vesicles and micelles, can self-assemble. Creating
a solid-supported lipid bilayer in an aqueous environment is slightly more complex. First,
electrostatic forces must be overcome so that the vesicles can be brought in close proximity to the
substrate, and second, the boundary between the hydrophilic and hydrophobic portion of the
bilayer is destabilized once the lipids of the proximal leaflets are able to interact
6
. There are many
approaches for bringing the lipid structures close to the solid substrate, for example relying on
diffusion or using Langmuir-Blodgett methods
7-12
. One advantage of the vesicle fusion method or
spontaneous self-assembly based on diffusion of vesicles to the surface is that planar bilayers are
readily formed on hydrophilic substrates.
22

The numerous methods of creating lipid bilayers have inspired a wide range of
experimental techniques to be developed. The majority of the detection approaches are focused on
understanding the behavior and evolution of the membranes. The two main strategies rely on either
detecting the electrical properties of the membrane in real-time or on imaging the motion of
fluorescent labels within the bilayers
8, 13-26
. In comparing these, label-free methods are attractive,
because they avoid the unwanted side effect of phospholipid modification. Previous studies have
shown that some potentiometric styryl dyes can increase membrane dipole potential, changing the
lipid dynamics
27
. However, reliance on electrical signals places restrictions on experimental
conditions; for example, swings in pH can cause signals.
In the present chapter, resonant cavity sensors are used to study the self-assembly and
detergent-assisted solubilization of lipid bilayers in real-time. The change in lipid bilayer
conformation is confirmed by monitoring both the wavelength shift and the photon lifetime change
in parallel. Two different detergents are used to solubilize the bilayer (SDS and TX-100). Due to
their structural differences as indicated in Figure 3-1, non-ionic SDS and anionic TX-100 are
understood to behave uniquely with the lipid bilayer
28-29
.

Figure 3-1. Chemical structures of the detergents used: (a) sodium dodecyl sulfate (SDS), and (b)
Triton X-100 (TX-100).

23

3-2 Materials and Methods
Chemicals
The phospholipid DOPC (1,2-dioleoyl-sn-glycero-3-phosphocholine) is purchased from
Avanti Polar Lipids (Alabaster, AL) and used without further purification. According to the
manufacturer, the concentration of the phospholipids is 10mg/ml in anhydrous chloroform. PBS
(phosphate buffer solution), SDS (sodium dodecyl sulfate) and TX-100 (Triton X-100) are
purchased from Sigma Aldrich (St. Louis, MO). Solutions of the following detergents are prepared
in 1X PBS: 15 mM SDS and 0.5 mM TX-100, which are well above the critical micellar
concentrations (CMCs)
29-31
.
Preparation of SUVs
The small unilamellar vesicles (SUVs) are prepared using the extrusion and vesicle fusion
method discussed
6, 8, 32-35
. Aliquots of DOPC phospholipids are dried under nitrogen, forming a
thin film. In order to completely remove the chloroform, the lipid film is further evaporated under
vacuum overnight, then rehydrated in 1X PBS and resuspended by vortexing. From the vortexed
solution, a volume of 0.6 mM of DOPC in PBS is extruded through a 100 nm pore diameter
membrane. The solution is passed through the polycarbonate membrane several times until the
volume of liposomes changes from an opaque to a transparent solution. Vesicle size is measured
and verified via dynamic light scattering instrumentation by Wyatt Technology (Santa Barbara,
CA).
Resonator set-up
The silica resonant cavities are fabricated using the method discussed in Chapter 2. The
resulting spheres range between 150 and 200 microns in diameter. The fabricated device is further
treated under an oxygen plasma unit from Anatech (Union City, CA). Plasma treatment of the
24

device serves two purposes: (1) removal of potential contaminants from the atmosphere, and (2)
optimization of surface adhesion. Oxygen plasma treatment yields a uniform monolayer of
hydroxyl functional groups, whose hydrophilicity promotes vesicle fusion and rupture, forming
the lipid bilayer across the surface of the microcavity. The devices are characterized using a tunable
765nm laser coupled to a tapered optical fiber waveguide.
Detection of lipid bilayer dynamics
The clean microresonator is placed into a ~200μl volume aqueous chamber filled with 1X
PBS, as shown in Figure 6a. The SUVs are introduced into the chamber at a concentration of 0.6
mM in 1X PBS with a continuous volumetric flowrate of 40 μl/min, using a motorized syringe
pump. As the vesicles approach the surface of the microresonator, they rupture and fuse, resulting
in a self-assembled lipid bilayer
10-11, 23
. Subsequent detergent injections are introduced to the
chamber under identical conditions.

Figure 3-2. Optical resonant cavity excitation
Two independent detection mechanisms are used to confirm the formation and detergent-
assisted solubilization of the lipid bilayers: 1) a refractive index change, and 2) a material loss or
material absorption change. Both methods rely on the optical mode directly interacting with the
25

lipid bilayer. As can be seen in the finite element method simulation in Figure 3-2, the evanescent
tail of the optical field strongly interacts with the lipid bilayer.
In the first approach, as the bilayer self-assembles on the cavity surface, the effective
refractive index that the optical mode experiences increases because lipids have a higher index
than water. Because the resonant wavelength (λ) depends on the refractive index experienced by
the mode (neff) according to the expression 2πRneff=mλ, where R is the device radius and m is the
optical mode number, when the index changes, the wavelength changes.
In the second method, because the optical absorption of lipids is higher than water, as the
bilayer forms and displaces the water, the effective optical loss increases, resulting in a decrease
in Qmat. Therefore, by monitoring the Q degradation and recovery and the shift in resonant
frequency in real-time, the bilayer formation and solubilization are confirmed.
To perform resonant frequency shift measurements, the previously described testing set-
up is used with a slight modification. Specifically, the resonant wavelength and transmission are
continuously tracked and recorded using a custom LabView program which only records the
minimum point. For the present measurements, the data acquisition rate is 600 samples/min.
However, to measure the Q, it is necessary to download the entire spectrum, which is very time-
consuming. Therefore, the Q spectra are taken at key transition points in the experiment. The wash-
off solution is also collected for further analysis.
In addition, several control experiments are performed to determine the effect of changing
the solutions on the detected signal. If the solvents have significantly different indices, the change
in the solvent could yield a false positive signal. First, the refractive index of each solution (PBS,
TX 100 in PBS, SDS in PBS) is measured using a handheld refractometer. Next, the microcavity
is sequentially exposed to each solution, and the resonant wavelength and Q are tracked. Because
26

a lipid bilayer is not present, the TX-100 and SDS solutions should have minimal impact on the Q
and resonant wavelength.
Dynamic Light Scattering (DLS) and UV-Vis analysis
DLS is used to characterize the size and size distribution of the micelles and particles in
the initial SUV, SDS and TX-100 solutions as well as in the experimental wash-off. DLS is ideally
suited for these measurements because it is able to characterize multiple sub-populations
simultaneously.
UV-Vis spectroscopy is also used to characterize the concentration of the SUVs in the
solution. In highly conjugated organic compounds such as DOPC, the wavelengths of absorption
peaks can be correlated with the types of bonds in a given molecule, making it a valuable tool for
confirming known functional groups within the molecule.
3-3 Results and Discussion
3-3.1 Interaction of optical field with WGM device
Upon introduction of SUVs into the aqueous chamber, a resonant frequency shift is
detected as the vesicles fuse to the microcavity and form a stable planar lipid bilayer. It is important
to note that the red shift observed is specific to vesicle rupture and bilayer adhesion, and not due
to flow effects or other similar experimental artifacts. Experiments introducing phosphate buffer
solution at the same volumetric flow rate (40µl/min) reveal that there is no red shift detected. All
results are summarized in Table 3-1. Control experiments containing only detergent solutions are
expected to reveal a resonant frequency shift that is distinct from that due to bilayer adhesion, an
indicated by the schematic in Figure 3-3. Additionally, the control experiments indicated that the
resonant frequency shifts eventually plateaus after the chamber is completely saturated. The
background noise due to the detergent was minimal (∆λ ~10-15pm) in comparison to the shift due
27

to bilayer formation/detergent solubilization (∆λ ~250-300pm), and likely is due to instability from
the injection pump (Figure 3-4).
introduce detergent micelles
(0.22 mM TX-100)

Figure 3-3. Schematic of micelle saturation on a bare silica microsphere

Figure 3-4. Wavelength shifts of bare silica microcavity saturated in: (a) 0.2 mM TX-100, and (b) 15
mM SDS

28

Table 3- 1. Effects of a supported lipid bilayer on Q and refractive index.
Combination (+ PBS) Q ∆λ (pm)
Control Experiments
PBS 7.95E6 5.61E-3
SDS 4.88E6 8.93
TX-100 8.93E6 15.0
SDS Experiments
PBS 3.62E7 0.201
bilayer 1.85E5 + 245.6
bilayer + SDS 2.40E7 - 245.6
TX-100 Experiments
PBS 4.68E7 0.478
bilayer 1.13E5 +300.3
bilayer + TX-100 1.21E7 - 302.1

It is important to note that once the bilayer has formed, upon introduction of pure buffer to
wash out the SUV-saturated chamber, a slight blue shift is observed. Initially, two possible
explanations were considered: (1) a blue shift is induced by a slight change in environment due to
either temperature or the refractive index of PBS, and (2) the PBS may be washing off
multilamellar vesicles/layers adsorbed to the bilayer, which are not a part of the laterally diffusing
main bilayer membrane. However, the temperatures of all solutions are maintained at a constant
37˚C, and the refractive index measurements (Table 3-2) indicated a negligible difference between
29

the solutions used in the experiments. Therefore, most likely, the observed blue shift is due to the
removal of residual liposomic micelles that may have adhered to the device.
Table 3-2. Refractive Index Measurements
Combination (+ PBS) Refractive index
PBS 1.3349
0.6 mM SUVs 1.3349
0.5 mM TX-100 1.3342
15 mM SDS 1.3345

3-3.2 Solubilization of bilayers
Membrane solubilization by detergents is a widely used procedure for solubilizing
biological membranes or extracting and purifying membrane proteins. It is also used to
demonstrate lipid fluidity and confirmation of a biomimetic membrane. Micelle formation and
dissociation occur at fast rates, with rate constants higher than 10 s
-1
, and solubilization occurs
quite rapidly as well
36
. It is expected that introduction of a detergent at concentrations above the
critical micelle concentration will solubilize the membrane, washing off all the phospholipids from
the substrate, as indicated in Figure 3 - 5. When the detergents are added for both experiments, the
resonant wavelength rapidly returns to its initial position, indicating that the entire lipid bilayer
was solubilized. In the control experiments, the shift due to just the detergents are over an order
of magnitude smaller.
30

Figure 3 - 5. The resonant frequency shift due to liposome fusion and bilayer formation of 0.6 mM
DOPC is indicated by the red curve, and subsequent solubilization via 0.5 mM TX-100 (green) and
15 mM SDS (purple) are shown.
Additionally, unlike in previous protein dissociation experiments which showed uniform
exponential decay indicative of a single kinetic coefficient, the lipid bilayer dissociation or
solubilization wavelength shift contained intermediate steps. These intermediate plateaus are
indicative of multiple pathways, and perhaps, semi-stable but short-lived intermediate states.
However, while the system can easily detect the 5 nm thick bilayer, the precise nature of
the detergent-induced shape changes could not be quantified with the set-up, due to signal blurring
from the rapid and abundant motion of the lipids. Because the lateral diffusion of phospholipids
and detergent micelle insertions occur on a scale within the same order of magnitude, the
membrane is deforming in multiple areas simultaneously, making the distinction between the
different mechanisms difficult to achieve.
3-3.3 Secondary confirmation via Q measurements
The second method employed to detect bilayer adhesion and detergent solubilization
involves the measurement of the evolution of the device’s quality factor (Q) throughout the
experiments (Figure 3 - 6). When the lipid bilayer forms on the cavity surface, the Q decreases
from above 10
7
to approximately 10
5
. This decrease of nearly two orders of magnitude is primarily
31

due to the material absorption of the bilayer. However, upon detergent initiated solubilization and
removal of the bilayer, the Q is recovered and returns to its initial value. Additionally, the control
experiments verify that the addition of the detergent, without the bilayer present, has minimal
impact on the Q. Therefore, this Q change is due solely to the formation and solubilization of a
lipid bilayer.

Figure 3 - 6. Effects of material absorption on the quality factor of the sensor. The measurements of
the photon lifetime or Q factor was performed using the two detergents: (a) SDS and (b) TX-100.

3-3.4 Characterization of lipid-detergent mixtures
The DLS results from the experimental wash-off experiments, as well as the controls, are
summarized in Figure 3 - 7 and Figure 3 - 8, respectively. By comparing the control and the
experimental results, it is clear that the SDS and the TX-100 interact with the lipid bilayer in
different ways.
Exposure to anionic SDS results in a DLS peak centered at 10nm which is distinct from
the control measurements, and therefore, can be uniquely attributed to the SDS-lipid mixture.
These results were expected, as the bilayer is disrupted upon SDS intercalation.
32

In contrast, the DLS measurements from TX-100 indicate lipid-detergent structures on the
order of ~100nm and larger only (Figure 3 - 7). Due to the chemical structure of TX-100, and its
non-ionic characteristic, it can be assumed that its integration into the bilayer may be distinct from
other detergents like SDS. Additionally, one may speculate that stereochemistry plays a role in
how the molecules insert into the bilayer. For instance, the aromatic hydrocarbon lipophilic group
may integrate directly into the hydrophobic inner leaflet of the bilayer, and rather than breaking
up the bilayer into smaller micelles, it could form larger geometries such as vesicles and sheets.
There may be competing interactions between the hydrophilic polyethylene oxide chain and the
hydrophobic 4-phenyl group that is not observed with the anionic SDS.
The solution of SUVs was further characterized using UV-Vis spectroscopy. The π-π* and
n-π* bonding interactions of DOPC can be used to confirm that there are no additional
contaminants introduced to the chamber through the syringe pump. Two distinct absorption peaks
observed at 190 nm and 250 nm, are in moderate agreement with the type of bonds specific to
DOPC
37
(Figure 3 - 8). The absorption peak at 190 nm can be attributed to the π-π* type of
transition from C=C groups and the absorption peak at 250 nm is attributed to the n-π* type of
transition due to C=O and N=O bonds
38-40
.

Figure 3 - 7. DLS histograms of lipid-detergent micelle mixtures from the experimental wash-off of
(a) SDS, and (b) TX-100 rinses.
33

Figure 3 - 8. UV-Vis absorption spectra of control samples: (a) 15 mM SDS, (b) 0.5 mM TX-100,
and (c) 0.6 mM SUVs.
In summary, given the combination of DLS and resonant cavity detection results, we
propose the pair of solubilization mechanisms shown in Figure 3 - 9.

Figure 3 - 9. A cartoon schematic detailing the proposed mechanism of lipid bilayer solubilization
using the two detergents: SDS (below) and TX-100 (above).

34

By monitoring both the resonant frequency and the quality factor, the formation and
solubilization of lipid bilayers are detected in real-time using resonant cavities. Both non-ionic
(TX-100) and anionic (SDS) detergent are used to initiate the solubilization, and distinctly
different profiles are observed. After the lipid bilayer is completely removed from the device
surface, it returns to its initial state. This recovery demonstrates the recyclability and repeatability
of biosensing experiments with the proposed device.
Through modifications of the device surface or utilization of different types of
microcavities, such as liquid droplets, these methods could be extended to membranes with liquid-
liquid and solid-liquid phase coexistence. Therefore, the present work sets the foundation for
further investigation into membrane dynamics in real-time. Additionally, the present study lays
the groundwork for future transmembrane protein studies using resonant cavity sensors. Protein
behaviors, such as folding and active vs. passive transport mechanisms
25, 41-44
, are expected to
occur at a distinct range from the lateral diffusion of phospholipids and are expected to be
detectable by the optical cavity biosensor.
3-4 Integrating the bilayer platform for heavy metal sensing
According to the Environmental Protection Agency (EPA), toxic metals including the
“heavy metals,” which include copper (Cu), zinc (Zn), lead (Pb), mercury (Hg), nickel (Ni), cobalt
(Co), and chromium (Cr) are individual metals and metal compounds that can negatively affect
human health. While very small amounts are necessary to support human life, in larger amounts,
such metals may build up in biological systems, becoming a significant health hazard. Heavy metal
pollution is an obvious concern for global sustainability, with the EPA listing many of them as
highly toxic and carcinogenic even at trace levels. For example, the consumption of drinking water
containing cadmium levels >0.005 (mg/L)
2
can lead to severe kidney damage. Table 3-3 lists the
35

different types of heavy metals listed by the EPA and their corresponding effects on the organs
and surrounding environment, most of which are catastrophic to human health. By adapting our
current system of a lipid membrane mounted on the microsphere, a novel biomimetic method to
detect small concentrations of such heavy metals can have a significant impact in the water safety
industry.
Table 3-3. List of drinking water contaminants and Maximum Contamination Level (MCLs)
Contaminant MCL
(mg/L)
2

Potential Health Effects from
Ingestion of Water
Sources of Contaminant in
Drinking Water
Cadmium 0.005 Kidney damage Corrosion of galvanized pipes;
erosion of natural deposits;
discharge from metal refineries;
runoff from waste batteries and
paints
Chromium 0.1 Allergic dermatitis Discharge from steel and pulp
mills; erosion of natural deposits
Copper 1.3 Short term exposure:
gastrointestinal distress
Long term exposure: liver or
kidney damage
Corrosion of household
plumbing systems; erosion of
natural deposits
Lead 0.015 Infants and children: delays in
physical or mental
development; children could
show slight deficits in attention
span and learning abilities
Adults: kidney problems; high
blood pressure
Corrosion of household
plumbing systems; erosion of
natural deposits
Mercury 0.002 Kidney damage Erosion of natural deposits;
discharge from refineries and
factories; runoff from landfills
and croplands

Recent advances in optical, electrochemical, and field-effect transistor sensors for heavy
metal detection have demonstrated great potential. For example, optofluidic devices integrate
optical components with microfluidic chips to develop portable sensors
45-46
. In addition, small
molecules, DNA, proteins, and bacteria have been integrated with inorganic materials to
36

selectively bind heavy metals as the molecular recognition probes. In a recent publication, a
microcantilever sensor based on the Ni-NTA-EDTA sample system demonstrated a nickel ion
detection limit of 0.5 µM
47
. As an initial proof-of-concept experiment, a similar sample system
was adapted onto optical microspheres in an attempt to detect the same or lower concentrations of
nickel ions. By demonstrating reasonable sensitivity in these initial experiments, the device can be
optimized to detect other heavy metals as well.
In this project, a head-group modified lipid bilayer is employed as a molecular recognition
probe (Figure 3 - 10). This technique, which is biomimetic, may even enable us to distinguish how
different metals exhibit distinct toxicity toward the cell membrane. Small unilamellar vesicles are
prepared using the extrusion and vesicle fusion method discussed in Chapter 2-2 with only one
modification in liposomic solution. Aliquots of DOPC and the headgroup-modified 1,2-dioleoyl-
sn-glycero-3-[(N-(5-amino-1-carboxypentyl)iminodiacetic acid)succinyl] (DGS-NTA) are dried
out in a 1:1 molar ratio.

Figure 3 - 10. A cartoon schematic of the proposed optical detection platform.

Despite the rapid development of label-free physical, chemical, and biological sensors in
recent decades, the principles of detection used remain limited to the adsorption of the targeted
analyte, usually based on specific molecular recognition pairs such as antibody-antigen and
37

receptor-ligand/antagonist
47
. The principle of detection is quite simple: chemical B is immobilized
on the surface of the detection device and a detection signal is generated when analyte A binds,
forming an interaction pair. In contrast, in the displacement method of detection, a third chemical,
chemical C, is also employed to detect analyte A. Because chemical C has a stronger affinity to A
compared to B, it can displace A from the surface forming A-C, resulting in a signal distinct from
that of the A-B interaction pair (Figure 3 - 11). It is expected that the presence of the nitrilotriacetic
acid (NTA) chelating group on the device will enhance the signal of nickel ions, and EDTA, a
ligand capable of donating a pair of electrons making it a good Lewis Base, has a greater affinity
to form a coordinate covalent bond with a metal ion.

Figure 3 - 11. A cartoon schematic of the proposed mechanism: (1) self-assembly of a phopholipid
bilayer with NTA-modified headgroups, (2) a nickel ion is bound to its interaction pair, (3) a third
chemical, EDTA, displaces the nickel from the surface.
Ethylenediaminetetraacetic acid (EDTA) is a tetradentate ligand composed of a “claw-like”
molecular structure, capable of chelating heavy metals and other toxins. Chelation involves the 2
nitrogen atoms and 2 oxygen atoms in separate (-COO
-
) groups, forming a coordinate complex
depicted by the reaction below:
38

EDTA + M  EDTA-M
where M = metal cations (e.g. Hg
2+
, Mg
2+
, Ca
2+
, Al
3+
, Fe
2+
, Fe
3+
, Cu
2+
, Ni
2+
)
In the proposed structure of the Ni-EDTA complex, a single molecule of EDTA can form two
bonds to Ni
2+
. The nickel(II) ion can form six such bonds, so a maximum of 3 EDTA molecules
can be attached to one Ni
2+
ion (Figure 3 - 12).
H
2
C
H
2
C
H
N
N
H
Ni
CH
2
OH
2
OH
2
OH
2
H
2
C
H
2
C
H
N
N
H
Ni
OH
2
OH
2
H
N
N
H
CH
2
CH
2
2+
2+
chelate with 1 EDTA ligand
chelate with 2 EDTA ligands

Figure 3 - 12. The coordinate complexes of (a) single and (b) double EDTA ligands chelated to a
nickel(II) ion.

Using the testing setup described in Chapter 2-2, the three steps discussed above will be
performed and an optical detection measured. In Figure 3 - 13, we propose the following
mechanism will occur:
39

Figure 3 - 13. The proposed wavelength shifts correlating to the varying concentrations of analytes
and their interaction pairs.
The project aim was to demonstrate the feasibility of using the previously described solid-
support system in order to detect heavy metal ions, such as Ni
2+
. The composition of the uniform
bilayer membrane can be modified with a chelating agent like nitrilotriacetic acid (NTA), which
is capable of cleaving heavy metal ions, including Ni
48-50
. The lipid bilayer is assembled with the
commercially available head-group modified NTA phospholipids, which have been demonstrated
to interact strongly with divalent transition metal ions including Ni, Cu and Co. As a last step,
ethylenediaminetetraacetic acid (EDTA) is introduced to the detection cell in order to release the
adsorbed Ni ions from the surface of the device. Operating as a secondary chelating agent, EDTA
competes with NTA, cleaving Ni from NTA. The Ni-NTA-EDTA sample system is used as a
preliminary proof-of-concept tool with plans to measure the LOD for high toxicity metals ions
commonly found in industrial water supplies
47
.
The phospholipids 1,2-dioleoyl-sn-glycero-3-phosphocholine (DOPC) and 1,2-dioleoyl-
sn-glycero-3-[(N-(5-amino-1-carboxypentyl)iminodiacetic acid)succinyl (DGS-NTA) are
purchased from Avanti Polar Lipids and used without further purification. Liposomic solutions of
NTA
NTA
NTA
NTA
NTA
EDTA
Time
∆ λ
40

DOPC:DGS-NTA in a 1:1 molar ratio are prepared in an overall mass concentration of 1 mg/ml.
The liposomes are prepared using the extrusion technique, which is widely studied and has been
well documented
6, 51-52
. In preliminary experiments, liposome fusion and bilayer formation
resulted in a ~20 pm signal, indicating the formation of a uniform and stable bilayer, while overall
pH of the aqueous environment was not taken into consideration. Upon introduction of EDTA, a
detection signal greater than the bilayer signal is generated, which is an unexpected result. In
Figure 3 - 14, the injection of EDTA induces an overall red shift of ~40 pm, which is approximately
twice that of the SUVs. Additionally, a blue shift in resonant frequency and subsequent return to
baseline due to buffer is observed, as shown by the pink curve in Figure 3 - 14.
0 2000 4000 6000
0.00
0.01
0.02
0.03
0.04
0.05
0.06
0.07

1X PBS
SUVs
1X PBS
0.1 uM NiSO4
0.25 uM NiSO4
0.5 uM NiSO4
1X PBS
42.7 mM EDTA
1X PBS
δλ (nm)
Time (seconds)

Figure 3 - 14. Overall resonant frequency shifts due to the different solutions injected into the
chamber. The δλ's of interest include: small unilamellar vesicles (red), varying concentration of Ni
ions (blue, turquoise, pink, green), and EDTA (maroon).

A control experiment was conducted, introducing EDTA to a bare silica device without the
presence of a bilayer, demonstrating a similar signal change. Further investigation into the physical
properties of this ligand revealed that, due to its four carboxyl groups, the pH of the solution must
be maintained in basic conditions (pH~10) in order to deprotonate the acids. At high pH conditions,
41

protons are stripped from the carboxylic acid groups, allowing EDTA to bond to the cation. The
real-time bilayer formation experiments were repeated in basic conditions. While the signal
consistent with vesicle fusion was observed, the signal did not plateau, as shown in Figure 3 - 15.
Several additional control experiments were required. The same experimental conditions
(concentration of chelating agent, solution pH, etc.) were performed on the bare silica device,
without the presence of a supported lipid bilayer. As indicated by the results from Figure 3 - 16, it
is important to establish that no resonant shift is detected in the control experiments. Further
experiments will involve varying the molar concentration of headgroup-modified lipids and
allowing more time for liposome fusion to occur.

Figure 3 - 15. The resonant frequency shift due to liposome fusion and bilayer formation of 1 mg/ml
1:1 DOPC:DGS-NTA is indicated by the red curve, and subsequent injection of 1X PBS is shown in
blue. The buffer-initiated solubilization suggests bilayer instability.
42

Figure 3 - 16. Control experiments on a bare silica device without the presence of a lipid bilayer
containing NTA in a basic environment (pH 10). A resonant frequency shift is not observed when
the platform is exposed to buffer (black) or EDTA (red).

Additional experiments with efforts focused on optimizing working conditions (controlled
pH environment, solution flowrate, concentration, etc.) were performed. In Figure 3 - 17, initial
experiments indicated that the presence of vesicles resulted in a red shift, but the signal was not
very stable, due to a high flowrate. Figure 3 - 18 indicates that an NTA-modified lipid bilayer can
be stably supported on the microsphere. Confocal microscopy images were also taken of the
extruded liposomes prior to incubation with the detection platform. As can be seen in Figure 3 -
19, the liposomes are uniform in size (~100 nm in diameter) and do not form clusters. Further
characterization was conducted on the liposomes via UV-Vis Absorption spectra, as shown in
Figure 3 - 20.
43

Figure 3 - 17. The resonant frequency shift due to liposome fusion in multiple scans (indicated by
the red, blue, and purple curves). The signals from buffer (black and green curves) suggest that the
flowrate plays an important role in the detection signal.

Figure 3 - 18. The resonant frequency shift due to liposome fusion and bilayer formation of 1 mg/ml
1:1 DOPC:DGS-NTA is indicated by the black curve.
44

Figure 3 - 19. Confocal microscopy image of the small unilamellar vesicles containing NTA-
modified head groups.

Figure 3 - 20. UV-Vis absorption spectra taken of the liposomes in buffer solution.

In Figure 3 - 21, the solutions of analyte (Ni
2+
), its interaction pair (NTA), and the method
of displacement (EDTA), as well as their coordinate complexes, were further characterized using
UV-Vis spectroscopy to ensure that there were no absorption peaks in the 633 nm or 765 nm range.
45

Figure 3 - 21. UV-Vis absorption spectra of nickel(II) ions in solution (black), 25 mM EDTA (red),
and the EDTA-Ni complex (blue).

This project established the feasibility of depositing a stable lipid bilayer on an optical
resonator for heavy metal sensing purposes. By introducing Ni and EDTA to the detection platform
in varying concentrations, a limit of detection may be achieved. This opens the door to future
studies of additional metal ions of interest, including heavy metals such as cadmium, manganese,
lead, and mercury.

46

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50

Chapter 4. Transmembrane Protein Study
4-1 Background and Significance
In the following project, efforts were focused on tracking the assembly of multimeric
membrane protein pores as they are introduced to the microsphere-supported bilayer. The self-
assembling bacterial toxin, α-hemolysin from staphylococcus aureus, is an ideal candidate due to
its stability over a wide range of pH and temperature values. Its transmembrane pore remains open
at normal conditions and previous studies have shown that it can bind to various biological or
synthetic lipid bilayers spontaneously, without the requirement of specific ionic conditions. Alpha
hemolysin’s transmembrane pore is used in studies of controlled delivery of ions and small organic
compounds across plasma membranes and through walls of synthetic lipid vesicles. In the present
study, we expect that alpha hemolysin will self-incorporate into the microresonator-supported lipid
bilayer upon exposure.
Recently, research groups have observed the rapid assembly of α-hemolysin using single-
molecule fluorescent imaging (Figure 4 - 1)
1-2
. However, the mechanisms by which the heptameric
ring-shaped intermediate, termed as a prepore, spontaneously inserts a membrane-spanning β-
barrel domain into the lipid bilayer, is not well understood
3
. By monitoring the resonant frequency
of the optical cavity, label-free detection of pore formation will be performed. The experiment will
demonstrate the ability of our device to track the trajectory of α-hemolysin monomers as they
permeate and form pores in a lipid bilayer.
51

Figure 4 - 1. A schematic of the proposed assembly pathway of α-hemolysin. The nature and
kinetics of diffusing inter-monomer contacts is not yet understood (permission to use this figure
from his dissertation granted by Dr. James B. Thompson).
Potential modifications can be made to the cavity-supported bilayer in the future study of
additional proteins. To accommodate larger domain transmembrane proteins, a soft support layer
can be inserted between the bilayer and the cavity surface
4
. This method involves the fusion of
small unilamellar vesicles (SUVs) to a pre-deposited monolayer of either protein or synthetic
polymer. When designing polymer cushions, potential materials of interest include: dextran,
cellulose, chitosan, polyelectrolytes, lipopolymer tethers, and protein-ligand linkers.
4-2 Engineering of the Protein Construct
The gene-encoding pDEST 14 vector containing αHL-wild type with-R(GS)2S2G2S2H6
was designed and engineered by phosphoramidate chemistry. A hexahistidine amino acid affinity
tag could be engineered by PCR after the C-terminal end of the α-hemolysin wild-type gene,
allowing for high-stringency purification. Following PCR and in vivo recombination of the
plasmids, the products were mini-prepped and sequenced by dideoxy sequencing to verify the
results. Coupled in vitro transcription and translation (IVTT) to produce nanomolar concentrations
of the α-hemolysin protomer in small volumes is possible using a pT7 vector-encoded gene using
52

E. coli extracts. IVTT allows for the protein to be produced in small volumes suitable for assaying
the activity. A hemolysis assay was performed, and by monitoring the optical density over time,
the progress of hemolysis was followed, demonstrating that the IVTT-produced α-hemolysin
protein is active.
Due to its high hemolytic properties against intact rabbit erythrocytes
5
, α-hemolysin is
straightforward to assay for its activity. Upon incubation of rabbit erythrocytes with α-hemolysin
monomers, the erythrocytes will be spontaneously lysed at a concentration-dependent rate due to
pore formation and hence induced osmotic stress. The cell lysis results in light scattering, which
can be monitored in a UV/VIS spectrophotometer (Figure 4 - 2).

Figure 4 - 2. A hemolytic assay demonstrating the activity of the αHL-WT- R(GS)2S2G2S2H6
against intact rabbit erythrocytes. The graph represents a light-scatter vs. time plot, therefore
monitoring the hemolysis process.
The α-hemolysin was prepared by overexpression in transformed E.coli and purified by
immobilized metal-affinity chromatography using a Ni
2+
:NTA column. Overexpression trials were
performed using E. coli strains, which were transformed with the plasmids generated by PCR,
encoding the αHL vector constructs. The bacterial strain BL21 was chosen due to its faster growth
in comparison to other strains, such as that of the JM109(DE3) strain. BL21(DE3)pLysS bacteria
was initially selected due to the pLysS plasmid encoding T7 lysozyme. Lysozyme controls the
53

expression more tightly, down-regulating the expression of the target-gene under the control of the
T7 promoter without interfering with its expression post-induction with IPTG. Its presence in the
cytosol of the bacteria also facilitates a more straightforward recovery of bacterial cytosolic
components, easing cell lysis after overexpression. However, in the preparation of a bulk glycerol
stock, the previously selected bacterial strain was unsuitable for protein transformation. The same
competent cells without the pLysS plasmid were used instead, allowing for successful batch
expression of the desired protein.

Figure 4 - 3. A Coomassie-stained 12% BIS-TRIS SDS-PAGE gel following electrophoresis showing
crude IVTT products in empty pDEST14 vector with T7 promoter (Lane 1) and encoded αHL gene
(Lane 2), the product of a crude Ni2+ -NTA IMAC purification of empty vector (Lane 3) and
αHL-WT + -R(GS)
2
S
2
G
2
S
2
H
6
(Lane 4), and the subsequent wash off empty vector (Lane 6), and
αHL-WT + -R(GS)
2
S
2
G
2
S
2
H
6
(Lane 7). Lane 5: blank.
4-3 Development of the Detection System
A tunable laser is coupled to the microsphere cavity using a tapered optical fiber, which
will excite a whispering gallery mode within the device. As α-hemolysin monomers are introduced
to the device and move toward the surface, shifts of the resonance wavelength can be monitored
in real time.
54

By combining fluorescence and optical resonance methods, it will be possible to study the
motion of small molecules within the lipid bilayer in four dimensions, x-y-z plus time. The
whispering gallery mode of the cavity creates an evanescent field which decays exponentially from
the surface of the cavity. Any fluorescent probe residing in the evanescent tail will therefore be
excited, allowing for x-y motion to be detected. Z-axis motion of small molecules within the
bilayer can be detected via the resonant frequency, or the specific wavelength of light which is
confined within the optical cavity. As molecules move through the bilayer, the refractive index
changes, which in effect, changes the resonant frequency.
The resonant frequency shift can be detected and used to study the position of molecules
as they pass through the bilayer. It is important to verify that there is a strong interaction of the
optical field with the lipid bilayer. The quality factor of a lipid bilayer coated resonant cavity will
be determined in both the visible (633 nm) and the near-IR (980 nm) using a pair of single-mode,
tunable external cavity continuous wavelength lasers. The lasers are coupled to the resonant
cavities using tapered optical fiber waveguides. Top and side view camera systems monitor the
adjustable distance between the tapered waveguide and the microcavity. The Q factor is
determined by scanning the wavelength of the single-mode laser and measuring both the resonant
power transmission and the loaded linewidth in the under-coupled regime.
Preliminary data demonstrating the in-situ assembly of a bilayer onto a microsphere has
been acquired (Chapter 3). It has been previously shown that the bilayer, upon introduction of a
detergent (Triton X-100), is washed off the resonator, regenerating a bare silica sphere. The
concept of tethering a lipid bilayer membrane to a solid support via a polymer cushion, a peptide,
protein-, or oligosaccharide-coupling layer has been widely explored in recent years
6-8
. Unlike a
supported lipid bilayer, which is physisorbed on a solid support, a tethered lipid bilayer consists
55

of a planar bilayer immobilized on a solid support or soft cushion
9
. Such a system offers potential
for practical applications including biosensing purposes capable of accommodating membrane-
integral receptors. Unlike the transmembrane protein α-hemolysin, many integral proteins of
interest require adequate spacing (greater than the ~10-20 Å spacing provided by a non-tethered
bilayer) between the membrane and solid support, in order to prevent protein degradation (Figure
4 - 4). Another advantage of tBLMs is their long-term stability on the order of days, weeks or even
months
9
.

Figure 4 - 4. A cartoon schematic depicting: (a) a solid supported lipid bilayer, and (b) the proposed
tethered lipid bilayer, allowing sufficient space for protein insertion into the membrane.
Due to its demonstrated biocompatibility, a protein-ligand bond such as biotin-streptavidin
is expected to provide the necessary spacer between the substrate and a lipid bilayer. Unlike
covalent bonds, protein-ligand bonds are neither strong nor permanent, but rather held together by
intermolecular forces such as ionic bonds, hydrogen bonds, and van der Waals forces. However,
it serves as a convenient preliminary experiment, and bilayer fluidity will be verified via Förster
resonance energy transfer (FRET) and fluorescence recovery after photobleaching (FRAP)
measurements. Fluorescence microscopy is useful for determining the homogeneity of dye-doped
planar bilayer systems, and FRAP enables quantitative measurement of the lateral diffusion of
phospholipids within such membranes. Many challenges are presented with the proposed project,
the greatest of which involves diffusional constraints to the lipid molecules of the tBLM. By
56

varying the concentration of biotin linkers, the fluidity and lateral diffusivity of lipids can be
controlled.
Circular silica pads on silicon are used as the solid support for initial fluorescence
microscopy analysis. The SiO2 pads are fabricated on silicon wafers via the photolithography
procedure described in Chapter 2. The wafers are cleaned through a series of three organic solvent
rinses (acetone, methanol, isopropyl alcohol, respectively), dried with an air gun and treated with
a simple 3-step process. First, oxygen plasma treatment deposits a monolayer of hydroxyl groups
to the surface, enhancing the hydrophilicity of the substrate. Second, the hydroxylated wafers are
functionalized with primary amine groups via chemical vapor deposition of the reagent 3-
aminopropyltrimethoxysilane (APTMS) under a vacuum dessicator for 1 hour. Lastly, the wafers
are incubated in a solution of N-hydroxysuccinimidobiotin (NHS-Biotin) in dimethylsulfoxide
(Figure 30), resulting in a uniform layer of stable amide bonds on the substrate surface, with
varying concentrations of biotin. After a PBS rinse, the samples are then suspended in a solution
of streptavidin for 30 minutes (2.5 µM) at room temperature. After a second rinse of PBS the
samples are incubated in solution at 4°C overnight. Small unilamellar vesicles composed of L-α-
phosphatidylethanolamine, transphosphatidylated (Egg Trans PE), 1% w/v 1,2-dipalmitoyl-sn-
glycero-3-phosphoethanolamine-N-(biotinyl) (Biotinyl PE) and 1% w/v L-α-
Phosphatidylethanolamine-N-(lissamine rhodamine B sulfonyl) (Egg Liss Rhod PE) are prepared
in 1mg/ml concentrations using the standard extrusion method discussed in Chapter 3.

57

Figure 4 - 5. Overall reaction scheme for the N-hydroxysuccinimide (NHS) functionalization
chemistry on the solid silica support.
Initial proof-of-concept experiments demonstrated the feasibility of this approach for
tethering a lipid bilayer to the microresonator, as shown in Figure 4 - 6, where uniform SUVs
(small unilamellar vesicles) are observed in solution.
58

Figure 4 - 6. Fluorescence microscopy images of samples with 2.5 µM streptavidin across the same
wafer. In (a), a silica pad is uniformly coated with a bilayer (red box), and in (b) uniform small
unilamellar vesicles (SUVs) are observed using the extrusion technique (red circles).

(a) (b)
(b)
59

4-4 Chapter 4 References
1. Thompson, J. R. Imaging the Assembly of the Staphylococcal Pore-Forming Toxin α-
Hemolysin. Dissertation, Oxford University, 2009.
2. Thompson, James R.; Cronin, B.; Bayley, H.; Wallace, Mark I., Rapid Assembly of a
Multimeric Membrane Protein Pore. Biophysical Journal 2011, 101 (11), 2679-2683.
3. Schuster, B.; Pum, D.; Braha, O.; Bayley, H.; Sleytr, U. B., Self-assembled α-hemolysin
pores in an S-layer-supported lipid bilayer. Biochimica et Biophysica Acta (BBA) - Biomembranes
1998, 1370 (2), 280-288.
4. Kibrom, A.; Roskamp, R. F.; Jonas, U.; Menges, B.; Knoll, W.; Paulsen, H.; Naumann, R.
L. C., Hydrogel-supported protein-tethered bilayer lipid membranes: a new approach toward
polymer-supported lipid membranes. Soft Matter 2011, 7 (1), 237-246.
5. Hildebrand, A.; Pohl, M.; Bhakdi, S., Staphylococcus aureus alpha-toxin. Dual mechanism
of binding to target cells. Journal of Biological Chemistry 1991, 266 (26), 17195-17200.
6. Sinner, E. K.; Knoll, W., Functional tethered membranes. Current Opinion in Chemical
Biology 2001, 5 (6), 705-711.
7. Achalkumar, A. S.; Bushby, R. J.; Evans, S. D., Cholesterol-based anchors and tethers for
phospholipid bilayers and for model biological membranes. Soft Matter 2010, 6 (24), 6036-6051.
8. Atanasov, V.; Knorr, N.; Duran, R. S.; Ingebrandt, S.; Offenhäusser, A.; Knoll, W.; Köper,
I., Membrane on a Chip: A Functional Tethered Lipid Bilayer Membrane on Silicon Oxide
Surfaces. Biophysical Journal 2005, 89 (3), 1780-1788.
9. Jackman, J.; Knoll, W.; Cho, N.-J., Biotechnology Applications of Tethered Lipid Bilayer
Membranes. Materials 2012, 5 (12), 2637-2657.

60

Chapter 5. Enzyme-mediated Hydrogels
5-1 Background and Significance
In the field of medical diagnostics and therapeutics, the need to improve patient care is
always present. Thus, there is continued effort in the academic research community to enhance
methods, materials, and devices. Hydrogels, the crosslinked form of hydrophilic polymers, are a
class of biomaterials that have demonstrated great potential for biological and medical
applications. In the field of nanotechnology, hydrophilic polymers frequently lie at the center of
research emphasis due to their perceived stimuli-responsive “intelligence”, finding a wide range
of biomedical/biological applications including: thin films, therapeutic scaffolds, and nanoparticle
delivery.
One example of such “intelligent materials” is the commercially available P-DERM®
Hydrogel Adhesive. These silicone hydrogels are designed to offer a soothing and gentle solution
for burns, device fixation, blisters, and other skin contact applications by holding moisture at the
surface. Further, the manufacturers, Polymer Science, Inc., have demonstrated the capacity to load
cosmeceutical, pharmaceutical or antimicrobials into their hydrogels. In recent years, the
development of biomaterials and their application to medical problems has dramatically improved
the outcome in treating numerous diseases. Below, Table 5-1 summarizes some of these benefits
and challenges associated with manufacturing biomaterials.

61

Table 5-1. Biomaterials find applications across a wide range of healthcare products
Advantages Disadvantages
 High performance materials (e.g.
polymers, ceramics, metals)
 Widespread use in medicine
 > 40,000 pharmaceutical preparations
currently in use
 > 8,000 medical devices and 2,500
diagnostic products employ
biomaterials
 Many biomaterials lack desired
functional properties to interface with
biological systems
 Not engineered for optimized
performance

The suitability of hydrogels as biomedical materials and their performance in a specific
application depend to a large extent on their bulk structure. The most important parameters used
to characterize the network structure of hydrogels are: (1) the polymer volume fraction in the
swollen state, υ2,s, (2) the molecular weight of the polymer chain between two neighboring
crosslinking points, 𝑀𝑀 �
c, and (3) the corresponding mesh size, ξ.
1
The parameters, which can be
determined theoretically or through the use of a variety of experimental techniques, are
summarized below in Table 5-2.
Table 5-2. Summary of the parameters used in hydrogel design, structure, and
characterization
Parameter Definition Measurement of
υ2,s Polymer volume fraction in
swollen state
The amount of fluid
imbibed/retained by the
hydrogel
𝑀𝑀 �
c Molecular weight of polymer
chain b/w two neighboring
crosslinking points
The degree of crosslinking of
the polymer
ξ Corresponding mesh size The space available between
the macromolecular chains
(e.g. for drug diffusion)

There are two categories of hydrophilic polymers: natural and synthetic. Natural polymers,
derived from tissues or other natural sources, mimic aspects of the native microenvironment.
62

Commonly used natural polymers for biomedical applications include: hyaluronic acid, chitosan,
heparin, alginate, fibrin, collagen, chondroitin sulfate, and silk. Synthetic polymers are fabricated
using organic chemistry and molecular engineering principles. Some examples include:
polyethylene glycol (PEG), polyvinyl alcohol (PVA), poly(N-isopropylacrylamide) (PNIPAAm),
and polycaprolactone (PCL). Aside from their relative low cost and commercial availability,
advantages of synthetic polymers include their versatility for chemical modification, low batch-
to-batch variation, and ease of tuning mechanical properties. However, since they lack the inherent
biochemical cues for interaction with cells, synthetic polymers are frequently used in combination
with natural polymers for biomimetic peptides to facilitate cell adhesion, migration and protein
secretion. The two categories of hydrophilic polymers are summarized in Table 5-3 below.
Table 5-3. Hydrogels can be composed of natural, synthetic, and biohydrid hydrophilic
polymers
Natural polymers Synthetic polymers
 Derived from tissues of other natural
sources
 Mimic aspects of native
microenvironment
 Examples: hyaluronic acid, chitosan,
heparin, alginate, fibrin, collagen
 Fabricated using organic chemistry
and molecular engineering principles
 Commercially available, low batch-to-
batch variation
 Versatility for chemical modification
 Ease of tuning mechanical properties
 Examples: polyethylene glycol,
polyvinyl alcohol, polycaprolactone

Hydrogels have also been synthesized to contain functional groups for enhancing cell
adhesion. The addition of such modalities has proven to dramatically alter the properties of the
hydrogels. For example, in a study performed by Hern et al, hydrogels functionalized with amino
acid sequences derived from natural proteins (e.g. RGD) were shown to enhance cellular
adhesion
2
. Stimuli-responsive hybrid materials consisting of hydrogels and genetically engineered
proteins have also been previously demonstrated. In a 2005 study performed by Ehrick et al,
calmodulin (CaM), a calcium-binding protein, was selected as a biological recognition element in
63

a stimuli-responsive hydrogel due to its ability to undergo large conformational changes
3
. The
stimuli-responsive hydrogel exhibits gating and controlled transport of biomolecules across the
network, demonstrating its potential for microfluidics and drug delivery.
Hydrogels have been used as scaffolds for tissue engineering and as immune-isolation
barriers for microencapsulation technology
1
. In microencapsulation, allogeneic or xenogeneic
cells are protected from the host’s immune system by a semipermeable membrane with the
capacity to separate out the immune components. Lim et al demonstrated the use of calcium ion
crosslinked alginate polymers for the treatment of diabetic animal models
4
. In addition, polymeric
microcapsules containing cells can be immobilized in agarose hydrogels to enhance the
functionality of transplanted constructs.
Hydrogels have widely been applied as intelligent carriers in controlled drug-delivery
systems. In order to optimize such qualities including: (1) permeability (e.g. sustained release), (2)
enviro-responsive nature (e.g. pulsatile release), (3) surface functionality (e.g. PEG coatings for
stealth release), (4) biodegradability (e.g. bioresorbable applications), and (5) surface
biorecognition sites (e.g. target release and bioadhesion applications), much effort has been
focused on engineering the physical and chemical properties at the molecular level. In a 2005 study
conducted by Sershen et al, the collapse and reswelling of gold-colloid composite hydrogels and
gold-nanoshell composite hydrogels in response to irradiation demonstrated its capacity to be
implemented into a microdevice as a means to control flow
5
. Other methods currently being
investigated include photoactive, temperature-dependent, or electrically and chemically sensitive
polymers. Using soft lithography, the integration of PEG hydrogels within microfluidic channels
has also been shown to control the aggregation of proteins and cells within the microfluidic
channel, as demonstrated by Langer at al
6
. Here, PEG hydrogels were used to form microstructures
64

capable of capturing and localizing cells in regions of low shear stress, demonstrating the capacity
to design controlled microreactors.
Furthermore, hydrogels can also be synthesized with gradients of signaling or adhesive
molecules or with varying crosslinking density across the material for applications in spatial drug
release, inducing directed cell migration/adhesion, and in studying biological systems. In addition
to pumps and valves, hydrogels can be used as integrated sensors within microdevices, due to their
ability to be integrated using photolithography, molding, or other approaches. Recently,
microelectromechanical systems (MEMS) sensor platforms, specifically those based on
microcantilevers, have found use in a wide variety of applications due to their miniature size and
ultrahigh sensitivity. For example, environmentally responsive hydrogels have been
micropatterned onto silicon microcantilevers to develop an ultrasensitive bioMEMS sensor
platform, demonstrating the first time actuation could be controlled by an intelligent polymer
network
7
.
In addition to its wide applications in the field of bionanotechnology, the use of hydrogel-
based biomaterials for the delivery and recruitment of cells to promote tissue regeneration in the
body is of great interest in the scientific community. Such medical applications currently being
explored include bone regeneration, cartilage repair, cardiac regeneration, and spinal cord repair.
Most recently, in work conducted by bioengineers at the University of Maryland, the authors
designed heart-injectable hydrogels containing microRNA to promote cardiomyocyte proliferation
in mice after myocardial infarction
8-9
. Beyond miRNA, a variety of cargo can include therapeutic
cells, growth factors, and therapeutic peptides and proteins. Due to their hydrophilicity, the water-
swollen polymer networks formed from a variety of natural and synthetic polymeric building
blocks can be engineered to enable crosslinking by chemical reaction or through physical
65

interactions in the presence of cells and proteins. Such reactions can proceed rapidly enough for
injection and in situ hydrogel formation. Chemical crosslinking can occur via well-established
reactions such as: (1) azide-alkyne, (2) thiol-ene, (3) Diels-alder, and (4) oxime reactions. Physical
crosslinking via ionic interactions, thermoresponsive materials, and hydrophobic interactions have
also been explored.
5-1.1 Motivation
The following project, conducted in collaboration with Dr. Larissa Rodriguez of the
Urology department at Keck, addresses challenges in the clinical treatment of stress urinary
incontinence, or SUI. SUI is caused by anatomical changes in urethral support, resulting in
dysfunction of the urethra and subsequent uncontrolled bladder voiding. Presently, there are a
variety of non-surgical treatments including behavioral therapy, biofeedback and medications.
However, there lacks a minimally invasive long-term treatment and the mainstay of treatment for
significant SUI is surgery. Such treatment plans include surgical slings, or the less invasive
alternative, injectable bulking agents. Bulking agents, typically made of collagen or a
biocompatible polymer, are implanted into the urethral wall, providing immediate support to the
urinary tract
10
. Results however, may only last between 1-5 years and requires multiple injections.
Efforts in this project are focused on developing a dual-purpose injectable capable of: (1) providing
the “bulking effect” of current clinically available injectable therapies, and (2) releases proteins to
induce regeneration of nerves and smooth muscles in the urethra. In the lab, controlled release of
such proteins, called growth factors, have demonstrated the ability to promote the growth of cells.
Various properties must be addressed in selecting an adequate material for clinical applications.
Two materials were evaluated as potential scaffolds: (1) a polymeric hydrogel, and (2) a more
conventionally accepted therapeutic composed of polymeric beads.
66

As previously mentioned, there is a variety of crosslinking mechanisms for inducing
gelation that have been explored, including crosslinking by irradiation (i.e. e-beam, gamma, or x-
ray), pH, light, or heat
11-12
. Of great interest has been the in situ enzyme-mediated hydrogel system,
which induces chemical crosslinking of both natural and synthetic polymers. These have
demonstrated a promising outlook in biomedical, biopharmaceutical, and biofabrication
applications
13
. Enzymatic reactions are an attractive approach to forming injectable hydrogels due
to the physiological conditions under which they occur, and the rate and ease at which the gel
stiffness can be adjusted. A growing interest in injectable hydrogel systems formed by enzymes
such as horseradish peroxidase is due in large part to the facile tuning of the gelation rate and
crosslinking density
14
.
A particular advantage is the in situ crosslinking mechanism, by which aqueous solutions
of gel precursors and bioactive agents can be injected using a syringe, upon which either physical
or chemical crosslinking occurs in situ, or inside the body. This eliminates the need for surgical
implantation as conventional preformed hydrogels previously required. This reaction can be finely
tuned by adjusting the amount of enzyme added.
Hydrogels formed by the enzymatic activity of the plant-derived protein horseradish
peroxidase (HRP), have garnered much attention in recent literature. The catalytic mechanism of
the HRP-mediated crosslinking reaction has been discussed in detail in prior literature, dating as
early as 2001
13, 15-16
, and recent biomedical applications of HRP-crosslinked hydrogels include:
(1) growth factor and protein drug delivery
17-20
, (2) stem cell differentiation, (3) cartilage and
smooth muscle regeneration, (4) tissue engineering
21-24
, and (5) dermal wound closure
25-28
. Briefly,
the crosslinking reaction of polymer-phenol conjugates in the presence of hydrogen peroxide
(H2O2) occurs when an oxidized form of HRP reacts with phenolic hydroxyl (Ph) moieties,
67

generating two equivalents of phenoxy radicals in one catalytic cycle. The phenoxy radicals react
with each other through a radical coupling reaction, generating a C-C or C-O bond as a stable
cross-link
13
. The reaction is represented below:
2 AH2 + H2O2  2 AH* + 2 H2O
in which AH2 and AH* represent the reducing agent and its radical product, respectively. This
oxidative coupling of phenols in the presence of H2O2 has garnered much attention in the field of
developing chemically crosslinked hydrogels. A wide range of both natural and synthetic polymers
has been modified with tyramine for the formation of hydrogels by HRP
14, 22, 29-35
. Advantages to
using polymer-phenol conjugates like the one employed in the present study include long term
storage stability, facile tunability of gelation rate and crosslinking density, and minimal oxidizing
during synthesis.
Various properties must be addressed in the selection of an adequate material for clinical
applications. Table 5-4 summarizes several challenges in the initial selection process and the
subsequent justification for the material selected.
Table 5-4. Addressing challenges and material property requirements (HRP: horse radish
peroxidase; H
2
O
2
: hydrogen peroxide).
Challenge/Material Property Solution
Biocompatible and biodegradable Gelatin
Stability, naturally occurring, high affinity for
therapeutic
Heparin
Material occurs under physiological conditions Enzyme-mediated crosslinking reaction
Physicochemical properties (e.g. gelation rate,
stiffness, mechanical strength) can be tailored
Adjustable enzyme ratio [HRP:H2O2]

The injectable hydrogel matrix is composed of modified polymer conjugates, as shown
below in Figure 5-1. The hydrogel forms a colorless, viscous gel upon mixing two biocompatible
materials via an “enzyme-mediated oxidative reaction”: (1) a cross-linkable gelatin poly(ethylene
glycol)-tyramine, and (2) a heparin-modified Pluronic® precursor.
68

Figure 5-1. Composition of the proposed hydrogel: (a) gelatin-PEG-tyramine, and (b) Heparin-
Pluronic®. The material may be loaded with therapeutics such as growth factors or cytokines prior
to mixing of the two polymer conjugates (PEG: polyethylene glycol; TA: tyramine).
Enzymatic reactions are mild and readily take place under physiological conditions. In the
proposed study, the reaction is catalyzed using horseradish peroxidase (HRP) and hydrogen
peroxide (H2O2). Physicochemical properties such as gelation rate, degradation time, and
mechanical strength can be changed by altering the enzyme to solvent ratio. For example,
increasing HRP decreases the gelation time, while increasing H2O2 reduces the swelling ratio of
the hydrogel, indicating an increase in crosslinking density
30, 36-37
. Furthermore, polymer-phenol
conjugates like GPT are not easily oxidized during synthesis and are stable for long-term storage.
5-2 Proteins of Interest
In the treatment of SUI, several growth factors have been identified as excellent candidates.
These are bFGF and β-NGF which have been shown to induce regeneration of nerves and smooth
muscles in the urethra.
20
In several studies, controlled release of such proteins, called growth
factors, have demonstrated the ability to promote the growth of cells
37-38
. Basic fibroblast growth
factor, or bFGF, plays an important role in the control of smooth muscle cell migration in vivo.
Studies published as early as 1993 have implicated bFGF as a mediator of replication in the arterial
media, showing that systemic injections of the growth factor significantly stimulated the rate of
smooth muscle cell migration to injured arteries in rats
39
. Nerve growth factor-beta, or β-NGF,
contains nerve growth stimulating properties and is involved in the regulation of growth and the
69

differentiation of sympathetic and sensory neurons, which are nerve cells that transmit pain,
temperature, and touch sensations. By loading the hydrogel with single bFGF, or dual bFGF and
β-NGF, in one or both of the polymer conjugates, the capacity to release proteins over a sustained
amount of time was evaluated using an enzyme-linked immunosorbent assay (ELISA).
5-3 Materials and Methods
Synthesis of Polymer
Commercially available Pluronic® F127 (BioReagent, Sigma-Aldrich) is modified with its
end group terminated by a carboxylic acid and verified with
1
H NMR, as shown in Figure 5-2.
Pluronic® F127 is diluted in a solution of 1,4-dioxane and a 1.2 :1: 1 molar ratio of succinic
anhydride : 4-Dimethylaminopyridine (DMAP) : triethylamine (TEA) is added to the reaction
flask. The reaction is allowed to stir for 24 hours and filtered with chloroform through a 0.45 µm
pore membrane filter. The resulting filtrate is precipitated in diethyl ether and lyophilized, yielding
a solid white precipitate.

Figure 5-2.
1
H NMR spectra of the modified end group of Pluronic in dimethylsulfoxide (DMSO).

70

The carboxylated Pluronic® is further conjugated to heparin via a carbodiimide-mediated
reaction. Heparin, a highly negatively charged glucosaminoglycan or GAG, is known to have a
high binding affinity for growth factors, and its use in the effective and stable delivery of proteins
has shown very promising results
38, 40
. Briefly, the carboxylated Pluronic® is dissolved in MES
buffer (pH 5.6) until the solution turns cloudy, indicating polymer aggregation. A 4:2 molar ratio
of EDC : NHS (1-ethyl-3-(3-dimethylaminopropyl)carbodiimide : N-hydroxysuccinimide) is
added to the solution and stirred for thirty minutes. Heparin is added to the reaction flask and
stirred at room temperature for 24 hours, after which the solution is filtered and dialyzed for 4 days
to remove unwanted byproducts and unreacted reagents.
Synthesis of the second polymer conjugate, the Gelatin-PEG-tyramine graft copolymer,
involves a three-step procedure. Polyethylene glycol (PEG 4.0 K) and a 1:3 molar ratio of PEG :
DMAP as well as PEG : triethylamine is dissolved in a solution of methylene chloride. The reaction
is allowed to stir at room temperature for 15 minutes to activate the terminal hydroxyl groups of
PEG. Excess 4-Nitrophenyl chloroformate (PNC) dissolved in methylene chloride is added
dropwise to the activated PEG solution and stirred under a nitrogen atmosphere at room
temperature for 24 hours (molar ratio PEG : DMAP : TEA : PNC; 1:3:3:3). The reaction is filtered
and the solvent is removed by rotary evaporation. The concentrated solution is precipitated in cold
ether, filtered, and dried overnight, yielding a white powder. The PNC-PEG-PNC conjugate is
dissolved in DMSO at room temperature under a nitrogen atmosphere. A 1:1 molar solution of
tyramine (TA) to PNC-PEG-PNC is added dropwise to the reaction flask and stirred at room
temperature under nitrogen for 6 hours, resulting in an amine-reactive poly(ethylene glycol)-
tyramine (PNC-PEG-TA) solution. A gelatin solution suitable for cell culture (gel strength: 300 g
Bloom) is added to the PNC-PEG-TA solution and stirred at 50 ⁰C under nitrogen for 24 hours,
71

yielding a yellow color. The solution is purified by dialysis against distilled water (molecular
cutoff = 6-8 kDa) for 3 days to remove any unconjugated molecules. A white gel begins to
precipitate in the dialysis tubing, changing from yellow to cloudy white. PNC salts are removed
by filtration, and the solution is lyophilized and characterized by
1
H NMR, as shown in Figure 5-
3. In Figure 5-4, the polymer precursors and their conjugated gel matrix were further verified with
UV-Vis Spectroscopy, indicating absorption peaks at ~300 nm.

Figure 5-3.
1
H NMR spectra of the gelatin-PEG-tyramine conjugate in dimethylsulfoxide (DMSO).
72

Figure 5-4. UV-Vis spectra of the polymer conjugates and the gel matrix.
Measurement of Swelling Characteristics
A swelling test was performed by immersing a mesh strainer containing a 200 µl sample
of the hydrogel in deionized water at room temperature to a constant weight. The swollen samples
were weighed after wiping out excess water from the surface of the mesh strainer and the swelling
properties were determined through the following equation:
Swelling Ratio (%) = [(Ws – Wd)/Wd] x 100%
where Ws and Wd are the weights of the hydrogels at swelling state and dry state, respectively.
Measurements were taken daily for 8 days, indicating the stability of the gel, shown in Figure 5-5.
200 400 600 800 1000
0.0
0.5
1.0
1.5
2.0
2.5
3.0
3.5
4.0
Abs.
Wavelength (nm)
heparin-Pluronic
gelatin-PEG
hydrogel matrix
73

Figure 5-5. Swelling ratios of the hydrogel matrix in buffer
Measurement of Gel Content
Gel content was estimated by measuring the insoluble part after immersion in deionized
water at room temperature and calculated as follows:
Gel Content (%) = (Wd/Wi) x 100%
where Wd is the weight of the dried sample and Wi is the initial weight of the dried sample after
crosslinking.
Scanning electron microscopy
The swollen membrane samples were cut into 12 mm circular sections with a thickness of
2-3 mm. The swollen samples were individually placed on an aluminum stub of 15 mm diameter
and a thickness of 22 mm. The stub was carefully placed in liquid nitrogen without immersing the
sample and allowed to equilibrate for approximately 5 minutes. The sample was sputtered with
palladium and placed in the scanning electron microscope (Hitachi S-4800 SEM). The operating
conditions for the frozen samples were: (i) working voltage of 15 kV, (ii) load current of 100 µA,
and (iii) a gun bias of 1.
1 2 3 4 5 6 7 8
0
10
20
30
40
50
Swelling Ratio
Day
74

Growth factor encapsulation
Therapeutics can be loaded into the Heparin-Pluronic® precursor via thermodynamic self-
assembly. A solution of 4 wt% Heparin-Pluronic® is dissolved in 100 µl of sterile phosphate buffer
saline (PBS) containing 0.1 mg/ml horseradish peroxidase (HRP). Separately, a 4 wt% sample of
gelatin-PEG-tyramine is dissolved in 100 µl of sterile phosphate buffer saline in H2O2 solution
(0.1 wt%). A sample of 150 ng of growth factor (bFGF and β-NGF) were added to each polymeric
solution for dual-loaded growth factor hydrogels. The two solutions were mixed thoroughly,
yielding a transparent gel matrix. By optimizing the HRP/H2O2 concentration, the gelation rate
was optimized to form in 2 minutes.
In vitro release studies
The in vitro release of β-NGF from the delivery system was analyzed using a commercially
available ELISA kit (R&D Systems). To determine the release from the hydrogel system, either
single-, double-, loaded growth factors were loaded into a 200 µl solution of the polymer matrix,
and suspended in a microcentrifuge tube containing sterile phosphate buffer saline (PBS). All
samples were maintained at 37 °C for up to 35 days. At each time point, the supernatant was
collected for analysis and the hydrogel scaffolds were replenished with new PBS. Percent
cumulative release was determined by normalizing the cumulative release of growth factors at
each time point with the total cumulative release over the course of 35 days.
5-4 Results and Discussion
In vitro release of β-NGF
The in vitro release of the single-loaded β-NGF, or the dual-loaded β-NGF and b-FGF,
from the hydrogel matrix was determined using ELISA, as shown below in Figure 5-6. The
releasate was analyzed daily until the hydrogel had completely dissolved into solution. As
75

indicated by Figure 5-6, cumulative release plateaus after approximately two weeks. These results
indicate that regardless of loading strategy (one vs. two growth factors), increasing amounts of the
proteins are released from the gel matrix over time.

Figure 5-6. Release profile of β-NGF demonstrates that the hydrogel is capable of sustained release
of growth factor over the course of 3.5 weeks ex-vivo. Control: no growth factor loaded in the gel
matrix. Dual GF: β-NGF and bFGF.
In order to evaluate the material’s potential for SUI treatment, in-vivo studies using our
collaborators’ animal injury model to observe immune responses were started. In order to evaluate
the biocompatibility of the scaffold, the prepared hydrogels were injected subcutaneously into
adult female Sprague Dawley (SD) rats, as shown in Figure 5-7.
76

Figure 5-7. Numbered subcutaneous injection sites in female Sprague Dawley rat. Rat was
euthanized one day post injection and the excised samples were imaged.
The hydrogel was also injected into the bladder and urethral walls to be collected for staining, as
shown in Figure 5-8.

Figure 5-8. LEFT: Injection of hydrogel into bladder wall, RIGHT: Injection of hydrogel into
urethral lining. Red arrows indicate presence of hydrogel.
Because the injected volume is no longer visible by observation alone after one week,
indicating that the material is either dissolving or migrating, the location of the gel cannot be
tracked. This challenge is further explored in Chapter 7, where efforts have been focused on
loading the gel with a hydrophobic fluorescent tracer.
For the treatment of SUI, standard therapies may include the injection of polymeric beads
along the urethral walls, rather than a gel. This approach motivated a parallel study using poly(D,L-
77

lactic-co-glycolic acid), or PLGA, microspheres as the material of interest. Due to the high water
content of most hydrogels, it is expected that the release profile is much shorter than those achieved
with microspheres based on more hydrophobic polymers like PLGA
41
. Therefore, an interesting
parallel study evaluating the effectiveness of heparin-modified hydrogels vs. heparin-modified
polymer beads was conducted and later discussed in the following Chapter 6. The impetus for
comparing a gel with the standard porous microsphere is summarized in Figure 5-9 below.

Figure 5-9. Comparing the material properties and synthetic route for two different injectable
therapies: a heparin-modified hydrogel vs a heparin-modified microsphere.

Unfortunately, due to loss of funding and collaborator disinterest in the SUI project, further
in vivo studies with our collaborators came to an end. The use of both materials described is
discussed in the following Chapter for another clinical application.

78

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81

Chapter 6. Enhanced CAR-T cell proliferation using an enzyme-
mediated crosslinking hydrogel scaffold loaded with IL-15/IL-15Rα
6-1 Background and Significance
Chimeric antigen receptor (CAR)-engineered T cells have seen great clinical success in
treating hematological malignancies, with CAR-T therapies yielding remarkable results by
harnessing the patient’s own immune system to induce long-term remission. Great efforts have
been placed on engineering cytokine production to promote CAR functions, such as IL-2 and IL-
15. Cytokines (~5-20 kDa) are proteins produced by a broad range of immune, endothelial, and
stromal cells, that play an important role in cell signaling. As immunomodulating agents, cytokines
act through cell receptors, modulating immune responses, and regulating cell populations. More
importantly, the role of cytokines in host responses to cancer has been of great interest in the field,
as they play an important role in cell-mediated immunity.
Further, some cytokines have been developed into protein therapeutics using recombinant
DNA technology. However, there are many challenges involved in recombinant cytokine therapy,
most important of which include toxicity effects and the narrow therapeutic window of
interleukins. Thus, there are numerous ongoing clinical trials of cytokines in patients with cancer
evaluating the potential for demonstrating antitumor activity. Several recombinant cytokines have
been approved for use as supportive agents in cancer. For example, recombinant Interleukin-2 (IL-
2) (trade name: Aldesleukin/Proleukin), a protein that regulates the activities of leukocytes or
white blood cells, has been approved by the Food and Drug Administration and in several
European countries for use as a surgical adjuvant treatment of high-risk malignant melanoma and
renal cell cancer. Although various dosages across the U.S. and around the world are used, the
efficiency and side effects of different dosages are frequently a point of contention
1-2
. In the United
82

States, large intermittent doses are typically administered intravenously on an inpatient basis
3-4
.
An alternative lower dose regimen can be administered via injection subcutaneously on an
outpatient basis
5
.
6-2 Protein of Interest
Due to demonstrated systemic toxicity and low objective response rate, some clinical
studies have recently demonstrated the potential failure of IL-2 as a standard therapy. Another type
of common γ-chain cytokine is IL-15, which has gained significant clinical interest as a highly
promising immunotherapeutic agent. IL-15 is understood to be an important growth factor similar
to IL-2 in its role promoting the activation and proliferation of natural killer (NK) cells and CD8
+

T cells
6
. Similar in structure to IL-2, IL-15 is a member of the four α helix bundle family of
cytokines secreted by mononuclear phagocytes following infection. Characterized as a T cell
growth factor, IL-15 is known to induce cell proliferation of the innate immune system’s natural
killer (NK) cells, whose primary objective is to eliminate virally infected cells. IL-15 demonstrates
excellent potential as an immunotherapeutic agent for modulating immune responses
7
.
In the literature, IL-15 was initially discovered in an adult T cell leukemia cell line and
simian kidney epithelial cell line as a 14-16 kDa protein with the capacity to stimulate CTLL (a
subclone of T cells derived from a C57bI/6 mouse) and peripheral blood T cell proliferation, and
induce peripheral blood mononuclear cell effector function in vitro
8-10
. In pre-clinical models, IL-
15 has been shown to enhance the anti-tumor immunity of CD8+ T cells
11
. At the National
Institutes for Health, a Phase I clinical trial to evaluate the safety, dosing, and anti-tumor efficacy
of IL-15 in metastatic melanoma and renal cell carcinoma has begun enrolling patients. Due to its
short half-life, there are currently two varieties of IL-15 superagonist available: (1) IL-15 SA, and
(2) ALT-803. IL-15 SA, a complex of IL-15 and IL-15Rα-Fc, is currently being evaluated for
83

antiviral and anticancer activities, in addition to enhancing immunotherapy and vaccination. The
anti-tumor efficacy of the IL-15/IL-15Rα complex, which is understood to enhance IL-15 half-life
and bioavailability over IL-15 alone, has been demonstrated in a variety of pre-clinical tumor
models
7, 12-15
and is currently being investigated as a potential adjuvant/therapeutic and inducer of
homeostatic proliferation, without the need for prior immunodepletion. ALT-803, a fusion protein
of IL-15, was recently granted fast track status by the FDA in 2017 and Phase III trials in bladder
cancer are currently being prepared.
Although the immunological effects of IL-15 overlap those of IL-2, the motivation for
selecting IL-15 as the therapeutic agent over IL-2 derives from the difference in toxicities and cell
maintenance and differentiation. In comparison to IL-2, preclinical studies have shown that there
is little vascular capillary leak observed in IL-15, suggesting that IL-15 exhibits a therapeutic index
superior to IL-2 in cancer immunotherapy
11, 16
. Due to the narrow therapeutic window and
relatively short half-life of such cytokines however, dose escalation, combination regimens, and
sustained release are all methods to improve response rates. The incidence of adverse side effects
including vascular leak syndrome and hypotension has precluded the use of higher doses
17
.
Further, clinical data using combination regimens have shown only a modest increment in response
rates
18
. Therefore, the demonstrated sustained release of new T cell growth factors such as the IL-
15 complex, exhibiting therapeutic indices superior to IL-2 or IL-15 alone, serves as the primary
motivation for the present work.
Previous studies have established that subcutaneous (sc) administration of cytokines lessen
the constitutional adverse effects (AEs) by lowering peak serum levels associated with immediate
cytokine release
6
. Thus, intravenous (iv) administration of CAR-T therapy, followed by a
subcutaneous injection of the IL-15/IL-15Rα complex embedded within a hydrogel delivery
84

system, is investigated in the present study to evaluate the potential of inducing comparable
proliferation and activation of CD3
+
and CD8
+
T cells. The gradual absorbance into the blood
circulation and lower maximal blood levels in sc-injected mice, support the administration of
cytokines in patients with various malignancies and diseases found in previous literature.
6-3 Materials and Methods
Synthesis of Polymer
The synthesis of the polymer conjugated has been detailed in Chapter 5-3.
Cytokine encapsulation
Therapeutics can be loaded into the HP nanogel via thermodynamic self-assembly. A
solution of 4 wt% HP is dissolved in 100 µl of sterile phosphate buffer saline (PBS) containing 0.1
mg/ml horseradish peroxidase (HRP). Separately, a 4 wt% sample of GPT is dissolved in 100 µl
of sterile phosphate buffer saline in H2O2 solution (0.1 wt%). Soluble IL-15/IL-15Rα complexes
were prepared as described below, and added to the solubilized solution of Heparin-Pluronic. The
two solutions were mixed thoroughly, yielding a transparent gel matrix. The addition of enzyme
and hydrogen peroxide was tuned in order to achieve a gelation time of precisely 90 seconds upon
mixing the two polymer conjugates.
Preparation of the IL-15 complex
Lyophilized ampoules of Human IL-15 (R&D Systems) and Recombinant Human IL-15 R
alpha Fc Chimera Protein (R&D Systems) were reconstituted in PBS. The complex was mixed and
incubated at 37 ⁰C for 30 minutes, as previously reported in the literature. Upon mixing, the
combined IL-15/IL-15Rα was added to the heparin-Pluronic polymer conjugate prior to
hydrogelation. Each mouse received 1.25 µg IL-15 precomplexed with 7.5 µg rmIL-15Rα-Fc in
sterile PBS, unless otherwise noted.
85

In vitro release studies
The in vitro release of IL-15 from the delivery system was analyzed using a commercially
available ELISA kit (R&D Systems). To determine the release from the hydrogel system, 4 ng of
IL-15 not complexed with Receptor alpha was loaded into a 200 µl solution of the polymer matrix,
and suspended in a microcentrifuge tube containing T cell media (TCM). All samples were
maintained at 37 °C for up to 8 days. At each time point (hours 2, 4, 6, 18, and days 2, 3, 5, 6, 7,
and 8), the supernatant was collected for analysis and the hydrogel scaffolds were replenished with
new TCM. Percent cumulative release was determined by normalizing the cumulative release of
IL-15 at each time point with the total cumulative release of IL-15 over the course of 8 days. The
values reported are an average of 3 experiments, with a sample size of n = 3 for each experiment.
In vitro T cell proliferation
An in vitro cell proliferation assay using T cells was employed to assess the IL-15
biological activities of the complex. To study the role of the IL15 loaded hydrogel on T cell
proliferation, T cells were cultured in 12-well plates under different conditions: without gel (- gel
/ - IL15), unloaded gel (+ gel / - IL15), or loaded gel (+ gel / + IL15), where IL15 refers to the IL-
15/IL-15Rα complex prepared as described above. T cell proliferation was measured by taking
daily cell counts, starting at 0.5 million cells per 1 ml of TCM on day 0. The cell count was taken
with both hemocytometers and FACS (MACSQuant® Analyzer 10 Flow Cytometer) analysis.
Preparation of T cells and adoptive transfer
Human peripheral blood mononuclear cells (PBMCs) were obtained from AllCells
(Alameda, CA). Frozen PBMCs were thawed and cultured in TCM containing X-VIVO™ 15
serum-free medium (Lonza, Allendale NJ), 5% (vol/vol) GemCell human serum antibody AB
(Gemini Bio-Products, West Sacramento CA), 1% (vol/vol) Glutamax-100X (Gibco Life
86

Technologies, Grand Island, NY), 10 Mm HEPES buffer (Corning, Corning NY), 1% (vol/vol)
penicillin/streptomycin (Corning, Corning NY), and 12.25 mM N-acetyl-L-cysteine (Sigma
Aldrich, St. Louis MO). A solution of 10 nm/ml human IL2 in PBS was added to supplement the
culture (Peprotech, Rocky Hill NJ). A bead:PBMC ratio of 3:1 of Dynabeads® human T-expander
CD3/CD28 (Invitrogen, Carlsbad CA) was used to activate and expand the T cells. The culture
medium was replenished every 48 hours, maintaining the T cell density between 0.5-1E6 cells/ml
until cell proliferation ceased, which was approximately 14-16 days post-activation.
Mice
All mice were purchased from the Jackson Laboratory (Farmington, CT) and housed under
specific pathogen-free conditions. Young adult (6 to 8 weeks) female, NOD.Cg-
Prkdc
scid
IL2Rγ
tm1Wj1
/SZ (NSG) mice were used as donors of peripheral blood, spleen, and bone
marrow in all experiments. All animal studies were performed in accordance with the Animal Care
and Use Committee guidelines of the NIH and were conducted under protocols approved by the
Institutional Animal Care and Use Committee at the University of Southern California.
In vivo biodistribution study
NSG mice were randomly assigned to treatment groups with 3 mice in each group: (i)
Group 1 mice received only PBS (
-/-
), (ii) Group 2 mice were treated with hydrogels and no
cytokines (
+/-
), (iii) Group 3 mice were treated with cytokine-loaded hydrogel (
+/+
), and (iv) Group
4 mice received an injection of cytokines in PBS solution without a hydrogel (
-/+
), where the
notation for the 4 treatment groups is summarized in Table 1. All results reported, unless otherwise
noted, were normalized to Group 1 mice as defined in Table 6-1.

87

Table 6 - 1. Summary of the four treatment groups employed for sc-injections (n=3). Notation
is indicated as a superscript
Group 1 Group 2 Group 3 Group 4
Hydrogel No Yes Yes No
IL-15/IL-15Rα complex No No Yes Yes
Notation - /- +/- +/+ -/+

Each mouse received intravenously 1 x 10
7
CAR-T cells through a tail infusion, followed
by an sc-injection of the hydrogel treatment groups. The groups treated with cytokines received
1.25 µg IL-15 precomplexed with 7.5 µg rmIL-15Rα-Fc in sterile PBS, a concentration determined
from published literature. A 100 µl volume of the prepared hydrogel (with or without cytokine)
was administered subcutaneously into the flank. Blood samples were collected after T-cell infusion
over the course of 18 days for flow cytometry analysis. Red blood cells were lysed with cell lysis
buffer (155 mM NH4Cl, 12 mM NaHCO3, and 0.1 mM EDTA in MilliQ water) and stained with
anti-hCD3 (clone HIT3a), CD8 (clone SK1), and CD45 (clone HI30), and quantified with FACS
analysis. Data were collected from a two-parameter (dual color fluorescence) density plot. Cells
were gated for appropriate forward and side scatter. Data were analyzed with FlowJo.
The spleen and bone marrow were collected 18 days following adoptive T-cell transfer and
made into single-cell suspensions for ex vivo analyses. Single-cell suspensions of the spleen were
created in ex-vivo media by homogenizing resected spleens through a 70 µm nylon mesh. Single-
cell suspensions of the femoral bone marrow were obtained by flushing TCM with a 27 G ½ needle
(Becton Dickinson) and syringe through the cracked bone samples obtained from the mice. Red
blood cells were lysed in lysis buffer for 10 minutes at room temperature and washed once with
PBS. Splenocytes and femoral bone marrow cells were centrifuged at 2,000 rpm for 5 minutes at
88

4 °C and resuspended in PBS for further analysis. T cells were defined as CD3
+
and CD8
+
. Cells
were analyzed using FACS analysis.
Statistics
Gel content and swelling ratio were reported as means ± standard deviation. One-way
ANOVA with Tukey’s multiple comparison tests were performed to assess the differences among
groups in the in vivo assays, using OriginPro software. A P value less than 0.05 was considered
statistically significant. Significance of findings was defined as: ns = not significant, p>0.05; *,
p<0.05. Data are presented as mean ± SD.
6-4 Results and Discussion
In vitro release of IL-15
The in vitro release of the IL-15 complex from the hydrogel matrix was determined using
ELISA, as shown in Figure 6 - 1b. The releasate was analyzed at hours 2, 4, 6, and 18, on days 2,
3, 5, 6, 7, and 8, until the hydrogel had completely dissolved into solution. As indicated by the
concentration amounts shown in Figure 6 - 1c and Figure 6 - 1d, the release profile reveals an
initial burst release of 26.1 ± 1.01%, which was observed at hour 18, followed by a negligible
release of 4.42 ± 1.40% at day 8. After 8 days, the accumulated release of IL-15 was 54.1 ± 1.80%,
gradually arriving to a plateau. As cell proliferation typically ceases approximately 14-16 days
post-activation, the capacity of the hydrogel to release cytokines for no greater than 2 weeks was
optimal.
89

Figure 6 - 1. ELISA results of cumulative release where (c) depicts the amount released at 18 hours,
and (d) at 192 hours.
As expected, the burst release was delayed in a hydrogel scaffold, as opposed to the high
burst release typically observed within the first several hours in delivery scaffolds containing
microspheres of a similar polymer. The synthesis of heparin-conjugated PLGA porous
microspheres was previously discussed in Chapter 5-4. In PLGA microspheres conjugated with
heparin, almost all of the IL-15 loaded is released in the first several hours, as shown in Figure 6 -
2. Of key importance concerning the release profile of biologicals is the hydrogel pore size, which
in the case of enzyme-mediated hydrogels, is largely dependent on the HRP concentration.
Notably, SEM images of the hydrogel showed pore sizes appropriate for encapsulation of
therapeutic biologics (i.e., proteins) between 5-50 µm in diameter. The morphology of the
hydrogel scaffold and the presence of dense porosity throughout the polymer matrix are shown in
Figure 6 - 3.
90

Figure 6 - 2. Comparison of release from heparin-coated microspheres vs. heparin-coated
hydrogels.

Figure 6 - 3. Scanning electron microscope images of freeze-dried samples of the hydrogels.

Further, because the polymer matrix is conjugated with heparin, which is understood to
have a high affinity for growth factors and cytokines, the present scaffold can be broadly utilized
for the delivery of a wide range of cytokines as well as other therapeutic proteins.
IL-15 complex enhances CAR-T Cell Proliferation In Vivo
Flow cytometry of peripheral blood lymphocytes revealed dramatic efflux of memory
CD3
+
and CD8
+
T cells from the circulating blood 4 days post IL-15/IL-15Rα administration,
followed by influx and hyperproliferation yielding 1.97-fold expansions of CD3
+
and CD8
+
cells
91

that ultimately returned to baseline. Peripheral blood sample results indicating enhanced T cell
proliferation in the cytokine-loaded hydrogel are shown below in Figure 6 - 4, where statistical
significance was present across all groups.

Figure 6 - 4. Cell density plots of peripheral blood samples collected on day 4, from mice treated:
(a) without gel or cytokine; - gel / - IL15, (b) unloaded hydrogel; + gel / - IL15, (c) loaded hydrogel;
+ gel / + IL15, and (d) with cytokine in solution; - gel / + IL15

All peripheral blood samples taken from each time point over the course of 18 days have
been summarized in Figure 6 - 5, where Groups 2-4 have been normalized over the negative control
group mice treated with neither hydrogel nor cytokine. In comparing all groups, both CD3
+
and
CD8
+
T cells were found to be elevated in peripheral blood in significant values for the unloaded
hydrogel (Group 2
+/-
; 1.50-fold), and for the negative control (Group 1
-/-
; 6.63-fold). The results
are summarized in Figure 6 - 6.

92

Figure 6 - 5. T cell population found in the peripheral blood (PB) of the treated mice at: (a) 4, (b) 6,
(c) 8, (d) 12, (e) 15, and (f) 18 days.

Figure 6 - 6. (a) schematic diagram of the in vivo experiment, (b) peripheral blood is collected and
assayed for T cells, (c) T cell populations of the hydrogels loaded with cytokine (red), without cytokine
(black), and cytokine without hydrogel (blue), (d) cell populations in splenocytes and femoral bone
marrow for the three groups.

93

Studies in mice lacking IL-15 or their receptors IL-15Rα have shown that their presence is
vital for the maintenance of memory CD8
+
T cells, implicating IL-15’s role as a selective growth
factor, impacting their generation and survival
19-22
. The present results were found to be in
agreement with such studies, which have also demonstrated that the activity of cancer patient-
derived NK cells was increased by treatment with the IL-15 complex
23
. Cytotoxic CD3
+
and CD8
+
T cells are understood to be potent mediators of host protection against disease like infection and
cancer. Studies from mouse models and in humans have confirmed the effects of CD8
+
T cells in
protecting recipients against infections and leukemic relapses
24-25
. Further, CD3
+
T cells are known
to play a critical role in eliminating infected or malignant cells from the body
26
.
The animals were sacrificed at day 18, as the in vitro results indicated that cell proliferation
ceased approximately 14-16 days post-activation. The spleen, which consists of a variety of
immune function cell populations such as T and B lymphocytes, is frequently evaluated for
immunomodulatory and anti-inflammatory activities. Bone marrow derived dendritic cells are also
routinely employed in cell based assays, as it plays an important role in the long-term maintenance
of memory CD8
+
T cells
21
. Resection of the spleen and femoral bone marrow indicated the absence
of T cell homing, a result in agreement with previous studies for tumor-free mice.
Splenocytes revealed a mean cell count of 2.89 ± 1.09 for the Group 3
+/+
mice treated with
a cytokine-loaded hydrogel, 1.55 ± 0.11 for the Group 2
+/-
mice treated with an unloaded hydrogel
(1.89-fold), and 3.89 ± 2.22 for the Group 4
-/+
mice treated with an sc-injection of IL-15 complex
in aqueous solution (1.34-fold). One-way ANOVA with Tukey’s multiple comparison tests were
performed, revealing no statistically significant differences across all groups, as shown in Figure
6 - 6, and the cell density plots of the isolated splenocytes from the hydrogel-treated mice are
shown in Figure 6 - 7.
94

Figure 6 - 7. Cell density plots of isolated splenocytes from mice treated with: (a) hydrogel without
cytokine; + gel / - IL15, and (b) cytokine-loaded hydrogel; + gel / +IL15.
In the bone marrow, CD3
+
and CD8
+
cell populations were found to be approximately the
same across all groups with no statistical significance: 1.04 ± 0.13 the Group 2
+/-
mice, 1.16 ± 0.24
for the Group 3
+/+
mice, and 1.13 ± 0.12 for the Group 4
-/+
mice. It is important to note that the
frequencies of CD3
+
cells, and that of CD8
+
T cells, were significantly lower in bone marrow
compared to the spleen. These results are found to be in agreement with prior results in wild-type
C57BL/6J mice
21
. Further, the present results indicate that during homeostatic, steady-state
conditions, most CD8
+
T cells freely distribute between the circulation and secondary lymphoid
organs
27
. Collectively, these results suggest that the IL-15/IL-15 Rα complex-loaded hydrogel
system promotes target antigen-independent CAR-T cell proliferation in circulation in tumor-free
NSG mice.

95

Conclusion
IL-15/IL-15Rα-loaded hydrogels could be safely administered during in vivo mouse
studies, markedly altering homeostasis of lymphocyte subsets in blood, with CD3
+
and CD8
+
cells
most dramatically affected. To reduce toxicity and increase efficacy, alternative dosing strategies
in the form of an sc-administered enzyme-mediated hydrogel scaffold was initiated, eliminating
the need for repeated intravenous infusions. The results presented in this study provide a clear
rationale for testing the cytokine-loaded injectable hydrogel in disease models in combination with
CAR-T immunotherapy. True validation of the capacity of the hydrogel scaffold on lowering
tumor burden in a xenograft mouse model is on-going.

96

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98

Chapter 7. Live Imaging using Novel Fluorescent Hydrogels
7-1 Background and Significance
As discussed in Chapter 6, there exist many developments in the field of biomedical
applications using hydrogels, with a rising interest in sustained drug delivery
1-5
. However, very
few have reached the market. Little information is known about where the hydrogels migrate
within in the body as they degrade and eventually disintegrate and dissolve. Studies evaluating the
trajectory of such materials are necessary to better understand the potential localization or
excretion of the device. One approach is to monitor the location using imaging agents, like
fluorescent dyes. However, successful translation to clinical applications require the fluorescent
hydrogels demonstrate superior performance to current commercially available contrast agents, or
organic dyes. The purpose of this project was to develop and validate a method for imaging the
location of the hydrogel in living systems to enable real-time tracking.
Here, we have developed an injectable organic fluorescent hydrogel for long-term, non-
invasive, in vivo tracing. Progress in the field of in vivo optical imaging has led to the design and
development of many probes for imaging biological processes. They typically include two
components: a binding element to allow for specificity and a fluorescent or light emitting
component. Common binding agents are small peptides and bulky monoclonal antibodies, and
common emission materials are inorganic imaging dyes, fluorescent dye-doped silica
nanoparticles, quantum dots, and gold nanoparticles
6-8
.
There are several limitations to the imaging agents already explored in the literature. For example,
conventional contrast agents, such as organic dyes, are susceptible to poor photostability, low
quantum yield, and insufficient in vitro and in vivo stability
6
. In the field of inorganic imaging
agents and nanoparticles, toxicity effects become evident. The nonbiodegradable nature of
99

particles like carbon nanotubes, as well as their needle-like structure, have been associated with
tissue damage and chronic toxicity in mice
9
. Further, near-infrared (700-2,500nm) quantum dots
contain highly toxic semiconductor compounds such as PbS, PbSe, InAs, or HgTe, limiting their
use for in vivo applications.
The present work is focused on improving the imaging agent. In the efforts to initiate
practical biomedical applications of supramolecular hydrogels such as drug delivery and
regenerative medicine, the in vivo biocompatibility and stability of the material must be carefully
evaluated. Several novel fluorescent probes synthesized in the Armani lab have been demonstrated
to be viable candidates, exhibiting strong fluorescence with emission across a wide wavelength
range well into the near infrared (350 – 800 nm). Fluorescence emissions from the visible to near
IR range are of great interest in the field of in vivo optical imaging. Studies have shown that for in
vivo imaging, light-emitting probes with emission in the red to infrared (>600 nm) are preferred
due to the low absorption in tissues at these wavelengths. In living tissue, absorption can be
affected by the presence of hemoglobin, which absorbs in the blue-green, while it is relatively
transparent in the red
7-8
. Organs with high vascular content have the lowest transmission, with a
cutoff below 600 nm due to absorption of light by OxyHb and DeoxyHb.
Important properties to evaluate are the extinction coefficient (the efficiency for absorbing
light; usu. 5,000 – 200,000 mol
-1
cm
-1
), quantum yield (the ratio of the number of fluorescence
photons emitted), and photostability (the rate of destruction of molecules in the excited state due
to bond breakage). These properties are particularly dependent on the environment. In vivo
imaging with fluorescent probes follows the same idea as conventional fluorescence microscopy,
except that imaging takes place in real-time in a living subject. In the field of in vivo fluorescence
imaging, the use of green fluorescent protein (GFP) as a reporter of gene expression in living cells
100

is a well-established technique
10-12
. In such studies, gene delivery mechanisms or transgenic mice
must be used to introduce GFP into cells. Because the signal generation is endogenously
incorporated into the animal, enabling in vivo measurements of transgene expression, expression
levels can be highly variable depending on the construct. Additionally, the emission wavelength
(508 nm) is not ideal for deep tissue penetration. In the proposed study, a novel fluorescent
hydrogel eliminates the need for genetically modified animals or cells. We propose incorporating
the dyes into a hydrogel that can be injected into the animal, either subcutaneously or
intramuscularly, as illustrated in Figure 7- 1.

Figure 7- 1. The novel dye can be incorporated into a biocompatible polymer hydrogel and injected
into the animal to evaluate the release of material over time.
7-2 Dyes of Interest
The three dyes of interest investigated in this project as potential in vivo tracers
are triphenylvinylphenylpyridine (TPPy; blue), [(triphenylvinyl)phenyl vinyl]pyridine (TPVPy;
green) and [methoxyphenylphenylvinyl)phenyl vinyl]pyridine (MTPVPy; red). Because of their
solvent polarity dependence, the dyes can be encapsulated into a hydrogel composed of both
natural and synthetic polymers, forming biocompatible polymer nanoparticles. The biodegradable
gel, as previously discussed in Chapter 5-2, is primarily composed of gelatin, PEG, and heparin-
conjugated Pluronic F-127, and has a high-water content. Fluorometry measurements were
101

obtained using a Horiba steady-state fluorescence spectrofluorometer, demonstrating the ability
for the gel to fluoresce for up to 34 days, as indicated in Figure 7- 2. Preliminary results
demonstrate that two of the three dyes, the blue TPPy and green TPVPy, emit strong fluorescence
while encapsulated in the hydrogel. Despite the strong fluorescence of the red dye MTPVPy in
solid form, preliminary attempts to encapsulate the dye into the hydrogel resulted in quenching of
the dye. This is due to the solvent polarity dependence. Efforts are currently being focused on
modifying a functional group in the dye to make it compatible with the hydrogel.

Figure 7- 2. Fluorescent dyes demonstrate capacity for sustained emission (over 34 days) while
encapsulated in hydrogels. Gel 1: TPPy (blue); Gel 2: TPVPy (green).
For the two dyes, TPPy and TPVPy, further characterization was performed with a Wyatt
Dynamic Light Scattering (DLS) instrument. The hydrodynamic radius was measured, and the
blue dye TPPy had a similar particle distribution as the control hydrogel loaded without dye, as
shown from the overlapping correlation function and histogram plots in Figure 7- 3. Interestingly,
the sample loaded with the green dye TPVPy had a larger average particle size. It is possible that
this dye forms large particles or induces aggregation. Mobility graphs were acquired to further
102

evaluate the zeta potential, showing that adding blue dye TPPy causes the zeta potential to increase
by a factor of three, as indicated by Figure 7- 4. In comparison with the control gel and green dye
TPVPy, which have a slight negative surface charge, the sample with blue dye TPPy has the
steepest V-Curve, suggesting that it may be more stable than the other samples. The results have
been summarized in Table 7-1.

Figure 7- 3. DLS results showing the hydrodynamic radius (Rh) of the dye-encapsulated hydrogels.
(HP = heparin-Pluronic hydrogel; HP + Dye 1 = hydrogel with encapsulated TPPy dye; HP + Dye 2
= hydrogel with encapsulated TPVPy dye).

Figure 7- 4. Mobility graphs indicate slightly negative surface charges on the samples. (HP =
heparin-Pluronic hydrogel; HP + Dye 1 = hydrogel with encapsulated TPPy dye; HP + Dye 2 =
hydrogel with encapsulated TPVPy dye).

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Table 7- 1. DLS results summary of the hydrogel samples
Sample Radius (nm) Zeta Potential (mV)
Hydrogel (-) control 39.1 ± 0.7 -4.8 ± 0.6
Hydrogel + TPPy; blue 37.4 ± 1.5 -16 ± 2
Hydrogel + TPVPy; green 106 ± 4 -8.4 ± 0.4

Since the gels have demonstrated the capacity for sustained fluorescence in aqueous
environments ex-vivo, efforts were focused on achieving similar results in a tissue sample. As a
preliminary study, the gels were injected subcutaneously into a chicken wing obtained from the
poultry store. The goal of this experiment was to determine if the fluorescent hydrogels could be
adequate tracers and visible beneath the skin, prior to submission for an IACUC protocol to begin
live animal studies. As shown in Figure 7- 5, the hydrogel was visible under a UV lamp and
remained intact with minimal spreading and minimal absorption. Additionally, the emission
remained over the 24hr timeframe. While an interesting result for a rudimentary experiment, the
chicken wing is physiologically unnatural and in vivo studies are required to proceed. Moving
forward, understanding the kinetics and routes for tuning release in live animals will be important
to evaluate the clinical utility. An IACUC protocol to image live mice with the proposed material
was written and approved.

Figure 7- 5. Subcutaneous injections in a chicken wing demonstrate that the fluorescent hydrogel is
visible under a UV lamp with minimal absorption.
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Cytotoxicity of functionalized biomaterials such as fluorescent hydrogels can be evaluated
either via cell studies, or in vivo, with the latter option being explored in the present study. In this
contribution, we report two self-assembling hydrogels loaded with novel molecules that
demonstrate superb performance in tracking and imaging applications. The results suggest that this
hydrogel has low toxicity, which is essential for subsequent applications in disease xenograft
models. This study offers guidelines to design and yield fluorescent supramolecular hydrogels for
the purpose of in vivo imaging.
7-3 Materials and Methods
Experimental
Mice
Young adult (4 to 5 weeks) female, SKH1-Elite mice (strain code: 477) were provided by
Charles River (Wilmington, MA) as part of the Charles River Animal Model Evaluation Program.
All animals were housed under specific pathogen-free conditions and weighed daily and observed
for any behavioral abnormalities. All procedures performed in accordance with the Animal Care
and Use Committee guidelines of the NIH and were conducted under protocols approved by the
Institutional Animal Care and Use Committee at the University of Southern California.
Synthesis of polymer gel precursors and fluorescent molecule encapsulation
All commercially available starting reagents and buffers were purchased from Sigma
Aldrich. The hydrogel was comprised of two gel precursors; the syntheses of which have been
previously discussed in detail throughout Chapter 5.

105

Characterization of the hydrogel-based fluorescent probes in vitro
Fluorometry measurements of the hydrogel-based fluorescent probes were obtained using
a Horiba steady-state fluorescence spectrofluorometer, as previously discussed. Daily spectra of
the prepared hydrogels were taken in order to demonstrate sustained emission in vitro.
In Vivo biocompatibility imaging
All procedures were performed with the animals placed under anesthesia (isoflurane 1-
4%). SKH1 Elite hairless mice were treated with four subcutaneous injections into the following
locations: upper left and right shoulder, and lower left and right flank. Each mouse received
subcutaneously a 50 µl volume of the hydrogel probe into all four sites, producing a measurable
bolus of approximately 0.385 mm
2
per injection. All animals were weighed daily, and the injected
hydrogels were measured with calipers until it was no longer detectable. Real-time, noninvasive
imaging of the mice was performed to monitor and record optical activity in vivo, using the
Xenogen IVIS-200 imaging system at the USC Molecular Imaging Center. Three varying
concentrations of each dye was prepared: 0.05, 0.10. and 0.15 mg/ml.
An integrated gas manifold allowed for rapid and temporary anesthesia for simultaneous
imaging of all four animals. The animals were placed into a light-tight imaging chamber coupled
to a highly sensitive CCD camera system cooled to -95 ºC, capable of quantitating single-photon
signals originating within the tissue of living mice. The animals were excited at 430 nm, and
emission was collected at the following wavelengths, 500, 520, 540, 560, 580, 600, and 620 nm.
Imaging timepoints were collected over the course of one week at the following at hours post-
administration: 0, 1, 2, 4, 6, 12, 24, 48, 72, 120, and 168.
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24 hours post-administration, the mice were housed together a clean cage to recover with
unlimited access to food and water, and cared for inside the USC vivarium at the health science
campus. Quantitative estimation of the fluorescence intensity was performed using the Living
Image® software package.
Organ distribution study
At 168h post-administration, the mice were sacrificed and the main organs as well as fatty
tissue, urine, and blood samples were dissected out. For probing potential in vivo toxicity of the
injected materials, major organs such as liver, kidney, lungs, heart, gall bladder, and small and
large intestines were collected at the day of sacrifice, and imaged. Animals were euthanized under
anesthesia according to the Principles of AVMA (American Veterinary Medical Association)
guidelines, followed by cervical dislocation. Quantitative estimation of the fluorescence intensity
was performed using the Living Image® software package.
7-4 Results and Discussion
The injected mice were imaged at time zero, where the hydrogel bolus was visible beneath
the skin, as shown in Figure 7- 6. Caliper measurements of the hydrogel injections were taken until
they were no longer visible at hour 72. Results of each injection dimensions are summarized in
Table 7-2.
.
107

Figure 7- 6. Photograph of the animals treated with hydrogels, 1 hour post-administration.

Table 7- 2. Caliper measurements of the hydrogel injections
Day 1 – 0 h
Right shoulder Left shoulder Right flank Left flank
Mouse 1 0.50 x 0.75 0.60 x 0.90 0.60 x 0.50 0.55 x 0.75
Mouse 2 0.60 x 0.70 0.60 x 0.70 0.70 x 0.60 0.55 x 0.75
Mouse 3 0.55 x 0.75 0.60 x 0.80 0.55 x 0.75 0.55 x 0.75
Mouse 4 0.80 x 0.60 0.55 x 0.75 0.90 x 0.65 0.60 x 0.55
Day 2 – 24h No change
Day 3 – 48 h
Right shoulder Left shoulder Right flank Left flank
Mouse 1 0.30 x 0.50 0.20 x 0.40 0.20 x 0.40 0.20 x 0.30
Mouse 2 0.40 x 0.30 0.20 x 0.30 0.10 x 0.10 0.30 x 0.20
Mouse 3 0.20 x 0.30 0 x 0 0.20 x 0.20 0.40 x 0.30
Mouse 4 0.20 x 0.10 0.10 x 0.10 0.20 x 0.20 0.50 x 0.60
Day 4 – 72 h Hydrogels undetectable

The 600 nm emission filter was selected for all presented measurements. Fluorescence
emission lifetime for the first 12 hours post-administration are shown in Figure 7- 7, indicating
that while much of the fluorescence drops in the first 12 hours, the dyes are still detectable in vivo.
108

Figure 7- 7. Lifetime measurements of the fluorescence intensity in the first 12 hours post-
administration, with increasing concentrations (0.05, 0.10. and 0.15 mg/ml) of (a) TPVPy, and (b)
TPPy dye
24 hours post-administration of the fluorescent hydrogels, the mice were fed and given free
access to food and water for the duration of the study. This was indicated by a significant increase
in total radiant efficiency, as shown in Figure 7- 8. Further, upon closer examination of the optical
images of the mice, as shown in Figure 7- 9g, increased fluorescence was observed after the
animals were fed. The fluorescence of food present in the mice was confirmed by Figure 7- 9h,
where a bright spot located at the stomach was observed. This was detectable in all subsequent
time points, after which the animals were provided with access to food, and is consistent with
animals that have not been placed on a restricted diet.
109

Figure 7- 8. Lifetime measurements of the fluorescence intensity in the first 24 hours post-
administration, with increasing concentrations (0.05, 0.10. and 0.15 mg/ml) of (a) TPVPy, and (b)
TPPy dye
110

Figure 7- 9. Excitation: 430 nm; Emission: 600nm; undoctored raw image files; (a) 0h, (b) 1h, (c)
2h, (d) 4h, (e) 6h, (f) 12h, (g) 24 h, (h) 48 h, (i) 72 h, (j) 120 h, (k) 168h

A reference region was selected to account for the background fluorescence fluctuations
due to feedings. This was performed by selecting four regions of interest (ROI) around the heads
of all animals at every collection time. By taking the difference of the mean average radiant
efficiencies [p/s/cm
2
/sr]/[µW/cm
2
] and dividing by the mean average radiant efficiency of the
ROIs at 0 hours, this calculated correction factor was used to adjust the fluorescence intensity color
scale. The corrected background fluorescence is shown in Figure 7- 10.
111

Figure 7- 10. Excitation: 430 nm; Emission: 600nm; A reference region was selected to account for
background fluorescence fluctuations due to feedings; (a) 0h, (b) 1h, (c) 2h, (d) 4h, (e) 6h, (f) 12h,
(g) 24 h, (h) 48 h, (i) 72 h, (j) 120 h, (k) 168h
Replotting the fluorescence lifetime, the results are more in line with expected results,
where the intensity continues down a decreasing trend as the fluorescent hydrogels are cleared
through the excretory pathway. The lifetime measurements are summarized in Figure 7- 11 and
Figure 7- 12.

112

Figure 7- 11. Normalized total radiant efficiency of the fluorescent hydrogels over the course of one
week post-administration, with increasing concentrations of dye (0.05, 0.10. and 0.15 mg/ml)

Figure 7- 12. Normalized total radiant efficiency of the fluorescent hydrogels in the first 24 h post-
administration, with increasing concentrations of dye (0.05, 0.10. and 0.15 mg/ml)

-20 0 20 40 60 80 100 120 140 160 180
0.5
1.0
1.5
2.0
2.5
3.0
3.5
Norma. Total Rad. Efficiency
Time (h)
Control
Green 1
Blue 1
Green 2
Blue 2
Green 3
Blue 3
0 20
0.0
0.5
1.0
1.5
2.0
2.5
3.0
3.5
Norma. Total Rad. Efficiency
Time (h)
Control
Green 1
Blue 1
Green 2
Blue 2
Green 3
Blue 3
113

7-4 Organ Distribution Study
Preliminary studies indicate that the fluorescent dyes pass through the excretory system,
with most fluorescence detectable in the urine, large intestine, and fatty tissue in the back. These
results are similar to those found in literature, where researchers have investigated the
biodistribution and urinary excretion of different surface-modified silica nanoparticles (SiNPs) in
mice
13-14
.

Figure 7- 13. Organ distribution study of the following samples from top left: lungs, fatty tissue,
liver, kidneys, small and large intestines, urine, and blood.
7-5 Future Work
Further studies that fully investigate the excretion of fluorescent hydrogels, by employing
larger sample populations, will be required. While preliminary results indicate efficient excretion,
an organ distribution study performed at various time points post-administration would better
reveal the trajectory and potential temporal localization of the material in vivo. Furthermore, the
long-term fate of fluorescent hydrogels that remain in the body is still largely unknown.
Translation to clinical applications will require us to demonstrate that the fluorescent hydrogels
114

are not only biocompatible, but also can out-perform other short-wavelength probes. Regardless,
the implications are substantial and much remains to be explored.

115

7-6 Chapter 7 References
1. Sharma, P.; Brown, S.; Walter, G.; Santra, S.; Moudgil, B., Nanoparticles for bioimaging.
Advances in colloid and interface science 2006, 123, 471-485.
2. Smith, A. M.; Mancini, M. C.; Nie, S., Bioimaging: second window for in vivo imaging.
Nature nanotechnology 2009, 4 (11), 710.
3. Weissleder, R., A clearer vision for in vivo imaging. Nature Publishing Group: 2001.
4. Bhattarai, N.; Gunn, J.; Zhang, M., Chitosan-based hydrogels for controlled, localized drug
delivery. Advanced drug delivery reviews 2010, 62 (1), 83-99.
5. Hoare, T. R.; Kohane, D. S., Hydrogels in drug delivery: Progress and challenges. Polymer
2008, 49 (8), 1993-2007.
6. Masteikova, R.; Chalupova, Z.; Sklubalova, Z., Stimuli-sensitive hydrogels in controlled
and sustained drug delivery. Medicina 2003, 39 (2), 19-24.
7. He, C.; Kim, S. W.; Lee, D. S., In situ gelling stimuli-sensitive block copolymer hydrogels
for drug delivery. Journal of controlled release 2008, 127 (3), 189-207.
8. Vinogradov, S. V.; Bronich, T. K.; Kabanov, A. V., Nanosized cationic hydrogels for drug
delivery: preparation, properties and interactions with cells. Advanced drug delivery reviews 2002,
54 (1), 135-147.
9. Liu, Z.; Davis, C.; Cai, W.; He, L.; Chen, X.; Dai, H., Circulation and long-term fate of
functionalized, biocompatible single-walled carbon nanotubes in mice probed by Raman
spectroscopy. Proceedings of the National Academy of Sciences 2008, 105 (5), 1410-1415.
10. Yang, M.; Baranov, E.; Jiang, P.; Sun, F.-X.; Li, X.-M.; Li, L.; Hasegawa, S.; Bouvet, M.;
Al-Tuwaijri, M.; Chishima, T., Whole-body optical imaging of green fluorescent protein-
expressing tumors and metastases. Proceedings of the National Academy of Sciences 2000, 97 (3),
1206-1211.
11. Michalet, X.; Pinaud, F.; Bentolila, L.; Tsay, J.; Doose, S.; Li, J.; Sundaresan, G.; Wu, A.;
Gambhir, S.; Weiss, S., Quantum dots for live cells, in vivo imaging, and diagnostics. science
2005, 307 (5709), 538-544.
12. Filonov, G. S.; Piatkevich, K. D.; Ting, L.-M.; Zhang, J.; Kim, K.; Verkhusha, V. V.,
Bright and stable near-infrared fluorescent protein for in vivo imaging. Nature biotechnology
2011, 29 (8), 757.
13. He, X.; Nie, H.; Wang, K.; Tan, W.; Wu, X.; Zhang, P., In vivo study of biodistribution
and urinary excretion of surface-modified silica nanoparticles. Analytical chemistry 2008, 80 (24),
9597-9603.
14. Burns, A. A.; Vider, J.; Ow, H.; Herz, E.; Penate-Medina, O.; Baumgart, M.; Larson, S.
M.; Wiesner, U.; Bradbury, M., Fluorescent silica nanoparticles with efficient urinary excretion
for nanomedicine. Nano letters 2008, 9 (1), 442-448.

116

Appendix A. Detection of Post Translational Modifications with
Optical Biosensors
Appendix A-1. Introduction
The objective of this study is to improve the detection of protein arginine modifications
with the broader goal of developing new medical diagnostic tools. This project uses the optical
detection platform detailed in Chapters 2 and 3 to detect the amino acid, L-arginine, and its
methylated products. L-arginine, an essential amino acid crucial for building protein, is produced
in the body and is also found in protein-rich foods such as fish, red meat, poultry, soy, whole
grains, beans and dairy products. Arginine methylation is a prevalent post-translational
modification that is understood to be involved in a number of cellular processes, including
transcriptional regulation, RNA metabolism and DNA damage repair
1-2
.
Post-translational modifications (PTMs) are occurrences by which a new functional group
is introduced to amino acids, which make it possible to have different combinations, leading to the
formation of 25,000+ proteins in our bodies, from just 20 amino acids. Moreover, PTMs are the
key to signaling transduction pathways in protein function. Methylation of arginine residues is a
PTM widely understood to play a significant role from RNA processing to modulation in protein
function. Arginine methylation can lead to three different species: monomethylated arginine
(MMA), asymmetric dimethylated arginine (ADMA) or symmetric dimethylated arginine
(SDMA), as shown in Figure A - 1.
117

Figure A - 1. Arginine methylation by protein arginine methlytransferases (PRMTs).
Methylation is performed by protein arginine methyltransferases (PRMTs) and methyl
addition from s-adenosyl methionine. In medical diagnostics, the ability to detect changes in
proteins is critical. For example, systemic lupus erythematosus is an autoimmune disease with a
wide range of symptoms. Currently, a common way to diagnose this disease is to look for
antibodies against specific protein modifications such as dimethylated arginine residues in
spliceosome Sm proteins
3
. While antibodies are known for their specificity, contamination is quite
common
4
. Additionally, the test, which typically implements fluorescence labeling techniques like
ELISA, is time-consuming and may cause error and contamination. Despite multiple studies on
lupus detection
5
, a reliable tool beyond ELISA is missing, since there is no single biomarker
responsible for diagnosing the illness. Additionally, adequate amounts of a biomarker must be
detected before a test achieves positive results. Developing an improved diagnostic tool with lower
limits of detection (LOD) is the motivation behind this project.
118

Appendix A-2. Materials and Methods
To accomplish this project, whispering gallery mode microresonators (WGM), shown
below in Figure A - 2, which are label free biosensors, will be used to detect the methylated
arginine molecule ADMA. Whispering gallery mode microresonators are fabricated from silica,
and have demonstrated the capacity to detect biomolecules, such as interleukin-2, at low
concentrations due to their high sensitivity
6
. Examples of such experiments, using liposomic
vesicles, have been previously described in Chapters 3 and 4 of this dissertation.
Briefly, by changing the functional group of a single molecule, and thus changing its
symmetry and size, adsorption to the proposed optical device induces a refractive index change
resulting in a distinct resonant wavelength shift. Such measurements were collected using
commercially available amino acids such as arginine.

Figure A - 2. Optical microscope image of the detection platform, a silica whispering gallery mode
resonator.
Appendix A-3. Results and Discussion
Commercially available L-arginine and ADMA solutions were prepared in ultra pure
MilliQ water, then sonicated for 5 minutes to degas the solution. WGM microcavities were
fabricated in the shape of a sphere by reflowing with a CO2 laser, and then treated with O2 plasma
to obtain a hydrophilic surface. The WGM microcavity was placed into a small aqueous chamber
and immobilized above an injection port. Solutions of varying concentrations of either arginine or
119

ADMA were introduced through the injection port at a constant flowrate, optimized at 60µL/min.
Due to the inconsistency of the pH of the solutions, as shown in Table A - 1, initial results
conducted in water were repeated in a buffer solution.
Table A - 1. Varying pH values of L-Arginine and ADMA solutions in both MilliQ water and
PBS.
L-Arginine ADMA
pH in water 9 3
pH in PBS ~6 ~6

In the following experiments, solutions of varying concentrations of ADMA were
introduced into the detection device surroundings while it was immersed in water. In Figure A - 3
below, the real-time detection of arginine and ADMA were conducted in ultra-pure MilliQ water.
Arginine concentrations as low as 1x10
-16
M were detected.

Figure A - 3. The real-time detection of arginine at varying concentrations to determine the limit of
detection, where (a) shows raw resonant wavelength shifts, (b) shows the line angle diagram of L-
arginine, and (c) summarizes the overall shift at each concentration.
In Figure A - 4 below, the same experimental setup was repeated with ADMA, across a
wider concentration range. There was a lack of reproducibility with increasing concentrations and
120

resonant wavelength shift. This was suspected to be due to pH fluctuations in the surrounding
environment. As increasing concentrations of ADMA are introduced into the detection chamber,
the surrounding water becomes more acidic, due to the presence of HCl. In order to distinguish the
resonator’s capacity to sense methylation from pH, this was addressed by using phosphate buffer
saline to stabilize pH effects.

Figure A - 4. Real-time detection of ADMA at varying concentrations where (a) shows the raw
resonant wavelength shifts in real-time, and (b) summarizes the overall shift at each concentration.
In the results below, both arginine and ADMA solutions were prepared in PBS instead of
water, ranging across the same concentrations as explored previously. Figure A - 5 and Figure A
- 6 reveal an improved stability for both compounds of interest.

(a) (b)
121

Figure A - 6. Real-time detection of ADMA in PBS across a range of concentrations.
Both L-arginine and ADMA solutions, while dissolved in PBS, appear to resolve the pH
fluctuations throughout the experiment as the noise appears to improve in comparison to the water
samples. As both solution concentrations increase, a concurrent increase in resonant wavelength
shift is detected, indicating the potential to use the present device as a methylation sensor. Further,
the desorption of arginine was detected, suggesting the recyclability of the detection platform. In
the figure below, the presence of arginine was detected by an increase in resonant wavelength until
the chamber was saturated, at which δλ plateaued. Once the wavelength shift stabilized, the
(a) (b)
Figure A - 5. Real-time detection of arginine in PBS, where: (a) depicts raw resonant wavelength shifts,
and (b) summarizes the overall shift at each concentration.
122

detection chamber was rinsed with water, removing the arginine, resulting in a blue shift that nearly
returns to the original resonant wavelength when the device is saturated in water only. This was
repeated twice, and shown by Figure A - 7.

Figure A - 7. Desorption measurement of arginine, where (a) overall resonant wavelength shifts,
and (b)indicates reproducibilty of the detection system
Since all experiments were previously conducted on bare silica devices, the resonant
wavelength shift detected was due only to adsorption, which is neither specific nor highly
selective. In the following experiments, surface-modification of the microresonators was explored,
as a means of evaluating whether or not affinity for the amino acid molecules could be enhanced
further than just adsorption. The mechanism by which a modified substrate can be conjugated to
amine groups found in L-argine and ADMA are summarized below in Figure A - 8a
7-9
. The WGM
microcavities were placed under O2 plasma for 5 mins followed by vapor deposition of (3-
aminopropyl)triethoxysilane (GOPTS) for one hour to functionalize the surface of the device. The
concentration of arginine was chosen to be 0.55mM to have a consistent pH with water, similar to
previous experiments. A 765nm laser was used and the solutions were introduced at a 40µL/min
rate. Reactive surfaces composed of GOPTS, a silane coupling agent, enhanced the sensitivity of
our device which resulted in greater adsorption of arginine as indicated in Figure A - 8b. The
123

wavelength shifts appear cleaner with improved resolution and reproducibility. Additional testing
of all arginine derivatives is required.

Figure A - 8. Desorption measurements of arginine with surface-modified detection platform,
where (a) illustrates the mechanism by which a silane group conjugates to amine groups, and (b)
shows the real-time detection of 0.55mM arginine solutions.
By employing silane chemistry to modify the surface of the optical device, L-arginine
molecules were found to be more accurately detectable with demonstrated reproducibility, due to
binding of the molecule to the sensor surface. Further, water rinses did not result in a complete
blue shift to the original wavelength as previously shown in Figure A - 7, indicating binding of
arginine to the surface.
Appendix A-4. Future Work
Beyond demonstrating the reproducibility of this detection system, further work is required
to determine the limit of detection for small molecules using the surface-modified optical detection
platform. The LOD must be lower than that of other presently used detection systems in order to
establish its potential as an improved sensor.
124

On a broader scale, unexpected results from this project indicated that the optical detection
platform also makes an effective pH sensor. Therefore, further investigations of systems in which
pH detection can be observed may be exploited for further diagnostic purposes.
125

Appendix A-5. Appendix A References
1. Bedford, M. T., Arginine methylation at a glance. Journal of Cell Science 2007, 120 (24),
4243-4246.
2. Bedford, M. T.; Richard, S., Arginine methylation: an emerging regulatorof protein
function. Molecular cell 2005, 18 (3), 263-272.
3. Pera, I.; Fritz, J., Sensing Lipid Bilayer Formation and Expansion with a Microfabricated
Cantilever Array. Langmuir 2007, 23 (3), 1543-1547.
4. Glasmästar, K.; Larsson, C.; Höök, F.; Kasemo, B., Protein adsorption on supported
phospholipid bilayers. Journal of colloid and interface science 2002, 246 (1), 40-47.
5. Keller, C.; Glasmästar, K.; Zhdanov, V.; Kasemo, B., Formation of supported membranes
from vesicles. Physical Review Letters 2000, 84 (23), 5443.
6. Nikitin, P. I.; Gorshkov, B.; Nikitin, E.; Ksenevich, T., Picoscope, a new label-free
biosensor. Sensors and Actuators B: Chemical 2005, 111, 500-504.
7. Piehler, J.; Brecht, A.; Valiokas, R.; Liedberg, B.; Gauglitz, G., A high-density poly
(ethylene glycol) polymer brush for immobilization on glass-type surfaces. Biosensors and
Bioelectronics 2000, 15 (9-10), 473-481.
8. Bañuls, M.-J.; Puchades, R.; Maquieira, Á., Chemical surface modifications for the
development of silicon-based label-free integrated optical (IO) biosensors: A review. Analytica
chimica acta 2013, 777, 1-16.
9. Funk, C.; Dietrich, P. M.; Gross, T.; Min, H.; Unger, W. E.; Weigel, W., Epoxy‐
functionalized surfaces for microarray applications: surface chemical analysis and fluorescence
labeling of surface species. Surface and interface analysis 2012, 44 (8), 890-894.

126

Appendix B. Porous Silica Sol-gel Microresonators
Appendix B-1. Introduction
Efforts on this project were focused on depositing porous thin films onto microtoroid
resonators for sensing applications. Nanoporous materials are proving to be an interesting platform
across a variety of fields and applications, including heterogeneous catalysis, molecule-selective
separations, optics, electronics, and sensing. Of great interest is the development of an optical
resonator containing a porous material deposited on its sensing region, thus mimicking a
biologically active scaffold. Nanoporous materials have demonstrated a remarkable increase in
surface-to-volume ratio that enhances the signals corresponding to the interaction between
biomolecular reactions. For example, the pores can be exploited as a means for protein/DNA
encapsulation, lowering the limit of detection in medical diagnostics. Due to their demonstrated
nanostructured matrix, sol-gel derived bioactive glass scaffolds have also been used in guided bone
regeneration
1-3
. This type of label-free, highly selective sensor finds applications in many fields
including medical diagnostics, environment, defense, and food security. Liquid crystal templating
is a method employed to form such pores, by exploiting the arrangement of micellar structures of
surfactant molecules
4-6
. Porous materials can be classified into three categories based on pore
diameter, as shown below in Table B - 1
7
.
Table B - 1. Classification of Porous Materials Based on Pore Diameters
Classification Pore diameter range
(nm)
Pore diameter range
(μm)
Pore diameter range
(Å)
Micropores < 2 < 0.002 < 20
Mesopores 2 – 50 0.002 – 0.05 20 – 500
Macropores > 50 > 0.05 > 500

Efforts in the Armani Lab have been focused on establishing a method to deposit
microporous sol-gel thin films onto silicon wafers with pore diameters ranging from 2-4 µm. More
127

recently, a method was developed to achieve pore diameters of ~100 nm and smaller, deposited
on silica pads photolithographically etched on silicon wafers. Preparing solutions of sol-gels in the
presence of a surfactant, dodecyltrimethylammonium bromide (DTAB), yields nanopores due to
the hydrophobicity of the long carbon chains of surfactant molecules, which aggregate together in
micelles when added to an aqueous solution. Many parameters affect pore formation, including:
(i) solvent, (ii) temperature, (iii) pH, (iv) aging time, and (v) initial precursor : water : catalyst
ratio. By varying two parameters, the concentration of surfactant and the annealing time, the pore
size and density of sol-gel films were optimized.
Sol-gel synthesis comprises of hydrolysis and subsequent condensation reactions.
Hydrochloric acid acts as a catalyst, initiating the hydrolysis of a silica precursor, which forms a
silica network from the suspended colloidal particles. These reactions frequently occur under
ambient conditions, allowing for facile production. Solutions are prepared at room temperature
and charged with an acid catalyst, upon which the silica network is allowed to age, followed by
the propagation of polymerization, during which the polymer matrix cross-links and forms the
network. The overall reaction is shown below:
Hydrolysis: Si(OC2H5)4 + 4H2O  Si(OH)4 + 4C2H5OH
Condensation: Si(OH)4  SiO2 + 2H2O
Overall Reaction: Si(OC2H5)4 + 2H2O  SiO2 + 4C2H5OH
Appendix B-2. Materials and Methods
The sol-gel composite thin films were synthesized using tetraethyl orthosilicate (TEOS)
(Alfa Aesar, 99.999 + %) as the silica sol-gel precursor. The reaction was carried out in an ethanol
solvent and catalyzed by hydrochloric acid (HCl) (EMD, 36.5-38.0%). Sol-gel solutions were
prepared holding constant molar ratios between solvent, precursor, catalyst and water (1:4:0.1:2).
128

The silane reagent TEOS was slowly added to ethanol, followed by H2O with surfactant. Lastly, a
solution of HCl was added dropwise to the solution in order to hydrolyze the TEOS.
Two surfactants of varying hydrocarbon chain lengths were investigated, DTAB and
cetyltrimethylammonium bromide (CTAB), as illustrated by Figure B - 1. The surfactant was
diluted to a 0.405 nmol concentration in deionized water. After stirring at room temperature for 2
hours, the sol-gels were aged for 1 week to allow for propagation of the polymer matrix. Sol-gel
films were spun coated onto toroid devices at 7000 rpm for 30 seconds to apply a homogeneous
thin film. Residual solvent and water were removed by heating on a hot plate at 140˚C for 1 hour,
followed by a subsequent bake at 210˚C for 1 additional hour to decompose and remove the
surfactant template.
The samples were then thermally annealed using a Lindberg/Blue M Tube Furnace from
Thermo Scientific. Samples were heated from 25˚C at a ramp rate of 1˚C/minute up to 450˚C and
held at this temperature for 12 hours to remove any remaining organic components, then cooled
down to 25˚C at a ramp rate of -5˚C/minute. Scanning electron microscopy (SEM) was used to
image the material surface for porosity. No additional sputtering of the wafer was necessary.

Figure B - 1. Two surfactants of varying hydrocarbon tail lengths were investigated.

H
3
C
N CH
3
H
3
C
CH
3
H
3
C
N
H
3
C
CH
3
CH
3
Br
Br
129

Appendix B-3. Results and Discussion
Prior to depositing the films onto the microtoroids, various sol-gel solutions were prepared
and characterized on bare silica wafers. The two parameters investigated included: TEOS:DTAB
ratio and anneal temperature. At temperatures above 200 ⁰C, cracking and sintering of the sol-gel
film was observed, as indicated by the table below. Millimolar concentrations of surfactant also
yielded larger pore sizes. Since the films needed to be nm-scale to remain within the sensing region
of the device, smaller pore sizes were necessary. By lowering the concentration of DTAB to 1 nM,
pore sizes and annealing at 140 ⁰C, pore sizes between 50-100nm in diameter were achieved. All
the prepared solutions and their optimized parameters were characterized with optical microscopy
and scanning electron microscopy, and summarized in Table B-2.

130

Table B - 2. Sol-gel solutions at varying parameters, and their characterized pore sizes.
Surfactant
DTAB DTAB DTAB DTAB DTAB
TEOS: DTAB

1:1/4 mM 1:1/4 mM 1:1mM 1:3 nM 1:1 nM
Anneal temp
(⁰C)

100 200 140 140 140
Anneal time
(h)
12 12 12 12 12
Diameter (nm)
280 – 570 100 – 500 125 – 625 50 – 100 50 – 100
Optical
Microscope

SEM image

131

The customized sol-gel protocol was then deposited onto circular silica pads. Optical
microscopy images of the spun-coated silica pads showed that a uniform layer of surfactant-loaded
sol-gels was successfully deposited on the surface with minimum cracking on the silica surface,
as indicated by Figure B - 2. The optimal spin coating parameters were found to be at 7000 rpm
for 3 minutes.

Figure B - 2. Optical micrograph images of the silica pads after spin-coating.
The presence of uniformly distributed pores over the device was further verified using
scanning electron microscopy (SEM). The white arrow in Figure B - 3 indicates that at the silica-
silicon interface, no substantial cracking is observed, despite a slightly raised (2 µm) platform.

Figure B - 3. Scanning electron microscope images of the spun-coated sol-gels on silica pads on
silicon. (b) interface of silica pads on the silcon wafer is indicated by the white arrow.
132

The same procedure was applied to a new geometry that was no longer planar – silica pads
on silicon wafers. Spin-coating onto silica disks yielded pores only on the silicon wafer, and not
on the silica, indicating surface tension effects. As shown in Figure B - 4, the sol-gel slipped off
the surface of the silica disks, and was only visible on the wafer. Further, evidence of significant
cracking was observed on the wafer, as further verified on the microtoroid devices in Figure B -
5.

Figure B - 4. Scanning electron microscopy images of the spun-coated silica disks elevated on a
silicon wafer.
133

Figure B - 5. Scanning electron microscope images indicate cracking of the toroid device and wafer,
due to the annealing procedure. Images taken at varying objectives: (a) X 75, (b) X 250, (c) X 650,
and (d) X 1,200.
Pore densities of ~100 nm and below, achieved using DTAB, are indicated by the SEM
images shown in Figure 6. The films were uniform, with no visible cracking.
134

Figure B - 6. Scanning electron microscopy images taken of the porous sol-gel thin films at varying
objectives: (a) X 5,000, (b) X 10,000, (c) X 20,000, and (d) X 30,000
Upon spin-coating onto the microtoroid sensing device, pores were observed only within
the inner edge of the torus, as shown in Figure B - 7.
135

Figure B - 7. Scanning electron microscopy image of porous sol-gels on a microtoroid via spin-coating
technology. Pores are observed only on the inner edge of the toroid, as well as on the silicon pillar
and substrate.
Appendix B-4. Conclusions/Future Work
Nanoporous materials can be deposited onto optical devices such as silica microresonators.
By altering the concentration of surfactant and the annealing temperature, pore diameters can be
varied. Methods that have been explored include oxygen plasma treatment of the device prior to
spin-coating, which makes the device surface more hydrophilic, attracting more surfactant
micelles. The device geometry can also be explored, including deposition of the thin films on
microspheres, rather than microtoroids, to minimize surface tension effects (Figure B - 8).
136

Figure B - 8. Scanning electron microscopy image of a microsphere on a silicon wafer at: (a) X 1,600,
and (b) X 7,000.
Students continuing this work must focus efforts on further optimization of sol-gel age-
time, sol-gel viscosity, anneal temperature, and deposition in order to completely coat the toroid
edge with pores. Future studies will investigate what types of sensing applications are appropriate
for this device. Gas detection, for example, would be of great interest. Gas molecule interaction
with a greater surface area due to a porous substrate may contribute to a lower limit of detection
(LOD). By functionalizing the surface of the sol-gel pores with an agent that is known to have a
high affinity for a certain gas, the LOD can be further lowered, yielding a highly sensitive detection
device. Efforts should be also focused on selecting an appropriate gas that will interact with the
device.

137

Appendix B-5. Appendix B References
1. Arcos, D.; Vallet-Regí, M., Sol–gel silica-based biomaterials and bone tissue regeneration.
Acta Biomaterialia 2010, 6 (8), 2874-2888.
2. Chung, J. J.; Li, S.; Stevens, M. M.; Georgiou, T. K.; Jones, J. R., Tailoring Mechanical
Properties of Sol–Gel Hybrids for Bone Regeneration through Polymer Structure. Chemistry of
Materials 2016, 28 (17), 6127-6135.
3. Mahony, O.; Jones, J. R., Porous bioactive nanostructured scaffolds for bone regeneration:
a sol-gel solution. Nanomedicine 2008, 3 (2), 233-245.
4. Nakanishi, K., Pore Structure Control of Silica Gels Based on Phase Separation. Journal
of Porous Materials 1997, 4 (2), 67-112.
5. Raman, N. K.; Anderson, M. T.; Brinker, C. J., Template-Based Approaches to the
Preparation of Amorphous, Nanoporous Silicas. Chemistry of Materials 1996, 8 (8), 1682-1701.
6. Barton, T. J.; Bull, L. M.; Klemperer, W. G.; Loy, D. A.; McEnaney, B.; Misono, M.;
Monson, P. A.; Pez, G.; Scherer, G. W.; Vartuli, J. C.; Yaghi, O. M., Tailored Porous Materials.
Chemistry of Materials 1999, 11 (10), 2633-2656.
7. Lee, K. Y.; Peters, M. C.; Anderson, K. W.; Mooney, D. J., Controlled growth factor
release from synthetic extracellular matrices. Nature 2000, 408 (6815), 998-1000.

Abstract (if available)

Abstract Biomaterials are substances engineered to interact with biological systems for a medical purpose. The study of biomaterials and biomimetic systems spans many fields including medicine, biology, chemistry, tissue engineering, and materials science, and they find a range of applications like optical diagnostics, imaging, and therapeutics. These applications rely heavily on self‐assembly, biocompatibility, and surface modification of polymers. During the course of this PhD, technical and scientific contributions were made to all three areas. Specifically, novel methods for modifying the surface of optical sensors to understand cell membrane behavior were developed, and biodegradable materials which permit the prolonged release of a therapeutic over an extended period of time were demonstrated. ❧ Over the course of my graduate study, this body of work has carried me through a variety of seemingly disparate experiments, from controlled environments outside of living organisms, or in vitro, to experimentation using whole living organisms, or in vivo. In whole however, it has yielded insight into different methods of performing experiments in both artificial environments as well as in cell cultures and animals, transitioning from fundamental basic science research to applied science. As such, these projects share a unifying goal of engineering solutions to clinical challenges such as the mechanisms of drug delivery and finding treatments for human disease and disorders such as cancer. ❧ In the first part of the dissertation, optical microresonators with lipid membrane coatings are designed to observe cell membrane behavior on a nanoscale, in vitro platform. The self‐assembly of a lipid bilayer from a solution of micelles is observed in real‐time on an optical platform, demonstrating its capacity to operate as a biosensor. Observing interactions on the nanoscale is best performed in parallel with studies in living systems, or in vivo. This is explored in the second part of the dissertation with animal studies, where the administration of a biomaterial, a hydrogel, was utilized to deliver protein therapeutics. This was complemented by a live imaging study, observing the trajectory of the implantable material in vivo using dye‐encapsulated hydrogels, demonstrating biocompatibility as well as its potential as a candidate in future clinical studies.

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Asset Metadata

Creator Sun, Victoria (author)

Core Title Engineering solutions for biomaterials: self-assembly and surface-modification of polymers for clinical applications

School Viterbi School of Engineering

Degree Doctor of Philosophy

Degree Program Chemical Engineering

Publication Date 10/10/2018

Defense Date 07/30/2018

Publisher University of Southern California (original), University of Southern California. Libraries (digital)

Tag biomaterials,CAR-T cell therapy,CD3⁺ T cells,CD8⁺ T cells,cell membrane,drug delivery,hydrogel,injectable polymers,interleukin 15,Lipids,OAI-PMH Harvest,optical microresonators,optical sensors,self-assembly

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Contributor Electronically uploaded by the author (provenance)

Advisor Armani, Andrea M. (committee chair), Graham, Nicholas (committee member), Ravichandran, Jayakanth (committee member), Wu, Wei (committee member)

Creator Email sun.victoria.l@gmail.com,victoria.sun@usc.edu

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