DEVELOPMENT AND APPLICATION OF MODIFIED HYLAURONIC ACID
DERIVED NANOPARTICLES FOR IMAGE GUIDED SURGERY AND DRUG
DELIVERY
BY
TANNER KINKADE HILL
A Dissertation Submitted to the Graduate Faculty of
WAKE FOREST UNIVERSITY GRADUATE SCHOOL OF ARTS AND
SCIENCES
in Partial Fulfillment of the Requirements
for the Degree of
DOCTOR OF PHILOSOPHY
Biomedical Engineering
May, 2015
Winston-Salem, North Carolina
Approved By:
Aaron M. Mohs, PhD, Advisor and Chair
Lissett R. Bickford, PhD
Nicole H. Levi, PhD
Frank C. Marini, PhD,
Thaddeus J. Wadas, PhD
DEDICATION AND ACKNOWLEDGEMENTS
I would first like to thank my advisor Dr. Aaron M. Mohs for the extensive support, guidance, and encouragement he has provided, without which this work would not have been possible. Throughout the time spent in his lab, Dr. Mohs has had an incredibly supportive, innovative, and driving role as my advisor and this has been critical in terms of learning, productivity, and well-being. I would also like to thank the other members of my committee, Dr. Frank C. Marini, Dr. Nicole H.
Levi, Dr. Lissett R. Bickford, and Dr. Thaddeus J. Wadas for their thoughtful critiques and advice how to improve this research. Amanda Davis, Dr. Steve
Kridel, and Dr. Todd Lowther also deserve special thanks for their work on the
Nano-ORL collaboration, as without their contribution the work would not have come to completion. Dr. Colleen Webb and Dr. Joseph von Fischer both deserve thanks for their incredible influence on my choice of a scientific career and the research training they provided me with during my time at Colorado State
University. I would also like to give a very special thanks to Dr. Kenneth
Klopfenstein of the Colorado State University Department of Mathematics, whose influence on me personally and scientifically cannot be overstated. Dr. K. was a wonderful mentor to me, his classes and teaching not only taught me calculus but about the world and myself as well. I would also like to thank the Mike and
Lucy Robbins family for their generous financial support of my graduate career through their fellowship. Finally, I would like to thank my parents, whose support has been a significant boon to me during my graduate career.
i
TABLE OF CONTENTS
DEDICATION AND ACKNOWLEDGEMENTS...... i
TABLE OF CONTENTS...... ii
LIST OF FIGURES AND TABLES...... viii
LIST OF ABBREVIATIONS...... x
ABSTRACT...... xiii
CHAPTER 1: INTRODUCTION…………………………………………………. 1
1.0 Nanoparticles in Medicine and Biomedical Research…….……………… 1
Liposomal Nanoparticles ……………………………………………………….... 3
Metallic Nanoparticles …………………………………………………………… 5
Quantum Dots ……………………………………………………………………. 7
Polymeric Nanoparticles ………………………………………………………… 8
Hyaluronic Acid Derived Nanoparticles ……………………………………….. 9
2.0 Biodistribution and Toxicity of Nanoparticles…………………………….. 16
The Reticuloendothelial System ……………………………………………….. 17
Tumor Structure and the Enhanced Permeability and Retention Effect …… 17
The Mechanical Environment ………………………………………………….. 19
Nanoparticle Properties Affecting Delivery ……………………………………. 19
Size ………………………………………………………………………………… 20
Shape……………………………………………………………………………... 22
Charge…………………………………………………………………………….. 23
Surface Modification...... 24
Toxicity……………………………………………………………………………. 26
ii
3.0 Clinical Imaging in Oncology………………………………………………. 27
X-ray and Computed Tomography…………………………………………….. 28
PET and SPECT………………………………………………………………… 30
Magnetic Resonance Imaging…………………………………………………. 31
Non-fluorescence Optical Imaging…………………………………………….. 32
Non-fluorescence Intraoperative Tumor Detection………………………… ... 33
4.0 Fluorescence Imaging and Image Guided Surgery……………………… 34
The Clinical Need for Image Guided Surgery………………………………… 35
Physics and Chemistry of Fluorescence……………………………………… 37
Fluorescence Imaging in Cancer……………………………………………… 39
Fluorescence Image Guided Surgery Instrumentation……………………… 45
5.0 Fatty Acid Synthase in Cancer……………………………………………. 47
Biology of FASN………………………………………………………………… 47
FASN Inhibitors and ORL………………………………………………………. 49
6.0 Conclusions…………………………………………………………………. 51
CHAPTER 2: SYNTHESIS OF NOVEL HYALURONIC ACID DERIVED
NANOPARTICLES……………………………………………………………… 53
1.0 Introduction…………………………………………………………………. 53
2.0 Materials and Methods…………………………………………………….. 54
Materials…………………………………………………………………………. 54
Aminopropyl-1-pyrenebutanamide Synthesis……………………………….. 54
Aminopropyl-5β-cholanamide Synthesis…………………………………….. 56
Conjugation of 5βCA or PBA to HLA…………………………………………. 56
iii
Conjugation of ODA to HLA…………………………………………………… 57
Hydrophobic Conjugate Characterization……………………………………. 58
Confirmation of NP Formation………………………………………………… 58
3.0 Results and Discussion…………………………………………………… 59
Hydrophobic Moiety Synthesis and Conjugation to HLA…………………... 59
4.0 Conclusions………………………………………………………………… 65
CHAPTER 3: IN VITRO CHARACTERIZATION OF INDOCYANINE
GREEN-LOADED NANOPARTICLES……………………………………….. 67
1.0 Introduction…………………………………………………………………. 67
2.0 Materials and Methods…………………………………………………….. 68
ICG Loading into Nanoparticles………………………………………………. 68
Nanoparticle Characterization………………………………………………… 69
NanoICG Serum Stability……………………………………………………… 69
Photostability……………………………………………………………………. 70
NanoICG Interaction with Serum Proteins…………………………………… 70
In Vitro Cellular Uptake of NanoICG………………………………………….. 71
Cytotoxicity………………………………………………………………………. 72
Statistical Analysis……………………………………………………………… 72
3.0 Results and Discussion……………………………………………………. 73
Physical, Chemical, and Optical Characterization…………………………... 76
NanoICG Interaction with Serum Proteins…………………………………… 82
In Vitro Cellular Uptake of NanoICG………………………………………….. 85
Cytotoxicity………………………………………………………………………. 87
iv
3.0 Conclusions………………………………………………………………… 89
CHAPTER 4: TUMOR UPTAKE, BIODISTRIBUTION, AND POTENTIAL
FOR IMAGE GUIDED SURGERY OF NANO-ICG IN XENOGRAFT
TUMOR...... 91
1.0 Introduction…………………………………………………………………. 91
2.0 Materials and Methods…………………………………………………….. 92 iRFP Transfection of MDA-MB-231 Cells……………………………………. 92
Initial NIR Fluorophore Enhanced IGS……………………………………….. 93
Extension of In Vivo Biodistribution and IGS……………………………….... 94
Tissue Phantoms……………………………………………………………….. 95
Statistical Analysis……………………………………………………………... 96
3.0 Results and Discussion…………………………………………………... 96
Initial In Vivo Investigation Using Subcutaneous Flank Tumors………….. 96
Image Guided Surgery………………………………………………………… 99
Further In Vivo Investigation with Orthotopic Breast Tumors and an
Additional Contrast Agent……………………………………………………… 102
Analysis with Tissue Mimicking Phantoms …………………………………... 109
4.0 Conclusions…………………………………………………………………. 113
CHAPTER 5: APPLICATION OF HYALURONIC ACID DERIVED
NANOPARTICLES TO DRUG DELIVERY…………………………………... 114
1.0 Introduction………………………………………………………………….. 114
2.0 Materials and Methods……………………………………………………... 115
Materials………………………………………………………………………….. 115
v
Drug Loading and Quantification……………………………………………… 116
Physical Characterization………………………………………………………. 118
FASN Inhibition………………………………………………………………….. 118
Inhibition of Lipid Synthesis……………………………………………………. 118
Cytotoxicity………………………………………………………………………. 119
Chemical Stability………………………………………………………………. 120
Mitochondrial Respiration Assays…………………………………………….. 121
Statistical Analysis ……………………………………………………………… 122
3.0 Results and Discussion…………………………………………………… 123
Drug Loading and Physical Characterization………………………………... 123
Inhibition of FASN and Lipid Synthesis………………………………………. 125
Cytotoxicity of Nano-ORL……………………………………………………… 127
Chemical Stability………………………………………………………………. 132
Mitochondrial Respiration Assays…………………………………………….. 134
4.0 Conclusions ………………………………………………………………… 140
CHAPTER 6: CONCLUSIONS AND FUTURE DIRECTIONS……………... 141
Summary of NanoICG Studies……………………………………………….... 141
Biodistribution and ICG Release to Serum Proteins……………………….... 142
An IGS-Specific Animal Model is Needed……………………………………. 145
Extension of NanoICG to SLN Mapping……………………………………… 146
Summary of Nano-ORL Studies……………………………………………….. 147
Diameter of Nano-ORL…………………………………………………………. 148
In Vivo Testing of Nano-ORL………………………………………………….. 149
vi
HLA Remains a Suitable Biopolymer for NP Research……………………… 150
Conclusions………………………………………………………………………. 150
WORKS CITED…………………………………………………………………... 152
APPENDIX………………………………………………………………………... 180
CURRICULUM VITAE…………………………………………………………… 186
vii
LIST OF ILLUSTRATIONS AND TABLES
Figure 1-1…………………………………………………………………………….2
Figure 1-2……………………………………………………………………………10
Table 1-1……………………………………………………………………………..14
Figure 1-3…………………………………………………………………………….38
Figure 1-4…………………………………………………………………………….41
Figure 1-5…………………………………………………………………………….45
Figure 1-6…………………………………………………………………………….49
Figure 1-7…………………………………………………………………………….51
Scheme 2-1…………………………………………………………………………..55
Figure 2-1…………………………………………………………………………….61
Figure 2-2…………………………………………………………………………….62
Table 2-1……………………………………………………………………………...63
Figure 2-3…………………………………………………………………………….65
Figure 3-1…………………………………………………………………………….74
Table 3-1……………………………………………………………………………...75
Figure 3-2…………………………………………………………………………….77
Figure 3-3…………………………………………………………………………….78
Figure 3-4…………………………………………………………………………….79
Figure 3-5…………………………………………………………………………….80
Figure 3-6…………………………………………………………………………….81
Figure 3-7…………………………………………………………………………….83
viii
Table 3-2……………………………………………………………………………...84
Figure 3-8…………………………………………………………………………….85
Figure 3-9…………………………………………………………………………….87
Figure 3-10…………………………………………………………………………...88
Figure 4-1…………………………………………………………………………….98
Figure 4-2…………………………………………………………………………….99
Figure 4-3…………………………………………………………………………….101
Figure 4-4…………………………………………………………………………….104
Figure 4-5…………………………………………………………………………….106
Figure 4-6…………………………………………………………………………….108
Figure 4-7…………………………………………………………………………….110
Figure 5-1…………………………………………………………………………….124
Figure 5-2…………………………………………………………………………….126
Figure 5-3…………………………………………………………………………….129
Figure 5-4…………………………………………………………………………….131
Figure 5-5…………………………………………………………………………….133
Figure 5-6…………………………………………………………………………….135
ix
LIST OF ABBREVIATIONS
HLA: hyaluronan/hyaluronic acid
NIR: near infrared
ICG: indocyanine green
IGS: image guided surgery
ORL: Orlistat
NP: nanoparticle
DLS: dynamic light scattering
AFM: atomic force microscopy
BSA: bovine serum albumin
PM: positive margin
FASN: fatty acid synthase
DOX: doxorubicin
PEG: poly(ethylene glycol)
RES: reticuloendothelial system
AuNP: gold nanoparticle
CT: computed tomography
MRI: magnetic resonance imaging
SPIO: superparamagnetic iron oxide
QD: quantum dot
ECM: extracellular matrix
x
MW: molecular weight
EPR: enhanced permeability and retention
ROS: reactive oxygen species
PET: positron emission tomography
SPECT: single photon emission computed tomography
US: ultrasound
FDG: fluorodeoxyglucose
ROLL: radioguided occult lesion localization
RSL: radioguided seed localization
RGL: radioguided localization
SLN: sentinel lymph node
DOT: diffuse optical tomography
ACLY: ATP citrate lyase
ACACA: CoA carboxylase
PBA: aminopropyl-1-pyrenebutanamide
5βCA: aminopropyl-5β-cholanamide
ODA: octadecylamine
NMR: nuclear magnetic resonance
EDC: 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide
NHS: N-hydroxy succinimide
FBS: fetal bovine serum
DMF: N,N-dimethylformamide
RBF: round bottom flask
xi
HCl: hydrochloric acid
MeOH: methanol
EtOH: ethanol
DMSO: dimethylsulfoxide
MWCO: molecular weight cutoff
AOI: area of interest
SNR: signal to noise ratio
CNR: contrast to noise ratio
OCR: oxygen consumption rate
WSS: wall shear stress
xii
ABSTRACT
This work presents research using the natural carbohydrate glycosaminoglycan hyaluronic acid (HLA) as a hydrophilic polymer backbone for the synthesis of nanoscale polymeric micelles loaded with either the near infrared (NIR) fluorophore indocyanine green (ICG) for image guided surgery (IGS), or the fatty acid synthase (FASN) inhibitor Orlistat (ORL) for chemotherapy. Hydrophobic conjugates were synthesized and conjugated to HLA in order to drive self- assembly into nanoparticles (NPs) and load their respective payloads. Chemical composition was confirmed by mass spectrometry and NMR. NP self-assembly was confirmed by dynamic light scattering (DLS) and atomic force microscopy.
Optical properties of ICG-loaded NPs, termed NanoICG, were examined and ICG loading was found to be dependent on the structure of the hydrophobic ligand.
Fluorescence quenching was observed in aqueous solution, and fluorescence could be reactivated by dissolution in a H 2O:DMSO mixture, or in PBS with bovine serum albumin (BSA). ICG release to BSA was observed in vitro . Nano-
ICG was found to have negligible cytotoxicity at physiologically relevant concentrations. In vivo results in a xenograft mouse model demonstrated that
NanoICG provided superior contrast between tumor and surrounding muscle tissue, and NanoICG successfully identified positive margins (PMs). ORL was successfully loaded into NPs, termed Nano-ORL, at a loading efficiency of approximately 98%. Hydrodynamic diameter was found to be dependent on ORL
xiii content. ORL extracted from Nano-ORL successfully inhibited FASN similarly to free ORL, and Nano-ORL inhibited lipid synthesis in cells at approximately the same level as free ORL. Nano-ORL was found to be more cytotoxic than free
ORL, and it was found that pre-incubation of both Nano-ORL and free ORL for 24 hours resulted in significantly reduced cytotoxicity of free ORL, while Nano-ORL remained just as cytotoxic. Metabolic analysis showed that Nano-ORL has a similar, negative impact on cellular metabolism as free ORL. These results present novel contributions in the form of the comparison of hydrophobic ligand structure in the loading of ICG into NPs, and the improved tumor accumulation and contrast achieved in vivo . Additionally, these results present the first HLA- derived NP formulation of ORL for use as a FASN inhibitor chemotherapeutic, and show that Nano-ORL has equal or greater activity against cancer cell lines than free ORL.
xiv
CHAPTER 1
INTRODUCTION
1.0 Nanoparticles in Medicine and Biomedical Research
Research into the properties of nano-scale materials has seen a significant advancement in recent years. These materials, despite differences in molecular and atomic composition, scale, shape, and purpose, are broadly grouped into the category of nanoparticles (NPs). NPs are generally considered to be synthetic or biological constructs that exist in the size range of 1-1000 nm in diameter.1 The term is often used to describe synthetic constructs but it is also useful to describe small proteinaceous particles such as drug-bound albumin as
NPs.2 Of direct interest to human health are those NPs that are engineered to provide some therapeutic or diagnostic effect for human disease. Material types range from various polymers and lipids to silicone, metals, and fullerenes. Each material type offers advantages as well as disadvantages and must be chosen according to the specific application. The applications of NPs in health and medicine broadly fall into either therapeutic or diagnostic categories, with a hybrid branch commonly termed “theranostic” that aims to build NPs with both diagnostic and therapeutic activity. The materials chosen for a specific NP generally belie part of its function, such as liposomal and polymeric NPs being used to solubilize drugs for therapeutic delivery, and superparamagnetic iron
xv oxide NPs used for magnetic resonance imaging. Here we briefly review the range of NP materials used in research and clinical application in order to better grasp the full range of potential. We then focus on polymeric NPs, and NPs that utilize the natural carbohydrate polymer hyaluronic acid (HLA) in particular. A set of examples of NP types is shown in Figure 1-1.
Figure 1-1. Examples of NP types. (A) Liposomal NPs showing self-assembly and drug loading into the aqueous core. (B) Gold NPs can be synthesized to form various shapes by changing the growth conditions. (C) Quantum dots typically have a polymer coating to solubilize them and prevent interaction with proteins and cells. Polymeric NPs can be formed from block copolymers (D), or polymeric chains with conjugated hydrophobic groups to drive self-assembly (E).
2
Many HLA-derived NPs are of this type (E). (A, D&E) adapted with permission from Macmillan Publishers Ltd: Davis, M. et al., Nanoparticle therapeutics: an emerging treatment modality for cancer . Nature Reviews Drug Discovery 2008,
7(9). (B) adapted with permission from Mieszawska, A.J. et al., Multifunctional
Gold Nanoparticles for Diagnosis and Therapy of Disease. Molecular
Pharmaceutics 2013, 10. (C) adapted from Advanced Drug Delivery Reviews,
60(11), Smith, A.M. et al., Bioconjugated quantum dots for in vivo molecular and cellular imaging , 1226-40, 2008.
Liposomal Nanoparticles
Liposomes may be composed of a number of different lipid combinations and form at least one spherical lipid bilayer in solution. The structure of this type of NP creates one or more hydrophobic zones in the form of bilayers, which surround an aqueous region. Both the bilayers and aqueous regions may be used to entrap drugs for controlled release and targeting. 3 Liposomes were one of the first NP formulations to be researched and extended to the clinic in the form of Doxil ®.1, 3 Doxorubicin (DOX) is an anthracyline with dose limiting cardiotoxicity. 4 Doxil ®, first approved by the FDA in 1995, is a formulation of DOX within a poly(ethylene glycol) (PEG) coated liposome. 5 The combination of liposomal entrapment and PEGylation allows Doxil ® to be more effectively delivered to tumors due to a decreased uptake by the reticuloendothelial system
(RES) in the liver, and has reduced cardiotoxicity compared to the non-liposomal
3 formulation. A number of additional liposomal drug formulations have since become approved or are under clinical investigation. 3
Lipids, such as phosphatidylcholines and phosphatidylglycerols, are used to form the lipid bilayers of liposomes. Cholesterol appears to make bilayers less permeable to entrapped drugs, and may be added to decrease the rate of release.6 Both PEGylated and non-PEGylated liposomes have been approved by the FDA.3 Added to these formulations there may be additional moieties such as protein ligands or monoclonal antibodies used to target certain cells or tissues.6
Attachment of additional molecules or atoms such as radionuclides for imaging may also be done through addition of chelators or entrapment.6, 7
Currently approved liposomal formulations are available for a number of indications. 3 The largest group of approved liposomes, and the largest group of new formulations in clinical trials, is directed toward chemotherapeutic delivery to various solid or non-solid cancers. Additional formulations have been approved for treatment of fungal infections, macular degeneration, influenza, meningitis, pain, and hepatitis A.
Current clinical studies with liposomal drug carriers are primarily focused on chemotherapy for cancer.3 In pre-clinical studies, research is focused on combinations of active targeting, nucleic acid delivery, and multi-functional formulations. 6 In targeted delivery, the targeting ligand does not appear to cause an increase in the total amount of NPs delivered to target tissues, though it may have an effect on the uptake of those nanoparticles by target cells. 6 Delivery of nucleic acids has a great deal of potential in treating numerous human diseases,
4 however the barriers remain equally significant. A number of investigations using nucleic acid-bearing liposomes have been performed, but clearance, targeted tissue delivery, and appropriate activity of delivered nucleic acids remain significant issues in this field. 6
Metallic Nanoparticles
A diversity of metallic NPs has been investigated in pre-clinical and clinical studies. Of these, major groups include gold NPs (AuNPs) and iron oxide NPs.
Both of these materials allow for extensive modification through physical or chemical attachment of coatings, targeting moieties, or drugs, which improve circulation time, delivery, and treatment. Additionally, both gold and iron oxide
NPs have proven useful as contrast agents for various imaging modalities, giving these materials an incredibly diverse potential for use in medicine.
AuNPs have been investigated extensively for multiple diagnostic and therapeutic methods. 8-10 The breadth of potential of AuNPs results from a number of useful properties, including easy synthesis, modification, and both diagnostic and therapeutic potential. AuNPs can be synthesized easily in the lab and control of shape makes it possible to produce spheres, cubes, rods, stars, and plates. While gold is largely physiologically inert, thiols can be easily conjugated to gold surfaces for improved functionality and biocompatibility.
PEGylation is a common method, 11-13 but addition of proteins or other polymers is also used. 9, 14 AuNPs have been used for X-ray contrast agents, fluorescence imaging, surface enhanced Raman spectroscopy, delivery of drugs and nucleic
5 acids, and photothermal therapy. 9 The density of AuNPs makes them well suited for X-ray imaging, and combination with Gadolinium allows for multimodality
CT/MRI. 15 Combination drug and photothermal therapy using AuNPs was investigated by Park et al., who found a synergistic effect between photothermal therapy and delivery of DOX.16 Not only was DOX release improved when
AuNPs were treated with near infrared (NIR) light, but the total cytotoxic effect was significantly more pronounced when both photothermal therapy and DOX were present than when only one or the other were used. Fluorescence of
AuNPs allows for combined imaging and therapy by combining the fluorescence imaging with photothermal ablation of tumors. 17-19 Jang et al. combined fluorescent AuNPs with photosensitizers for combined imaging, photodynamic therapy, and photothermal therapy. Fluorescent imaging of tumors gave a maximum tumor-to-background ratio of only 3-4, but the combined treatment of photodynamic therapy to produce reactive oxygen species and photothermal therapy for thermal ablation resulted in significantly smaller tumors than groups receiving other treatments eight days after treatment. Given the ease of use and these encouraging results it seems likely that AuNPs will find their way to human treatment.
Magnetic NPs are generally composed of various forms and alloys of iron.
Magnetite (Fe 3O4) and maghemite (Fe 2O3) NPs, as well as alloys with Mn, Co, or
Ni have been investigated for MRI signal enhancement, drug delivery, and hyperthermia. 20 Superparamagnetic iron oxide (SPIO) NPs have received much of this investigation due to the low toxicity and surface that can be modified for
6 biocompatibility, targeting, and therapy. 21 A number of these have been clinically approved, including GastroMARK ® (ferumoxsil), Combidex ® and Feridex IV ®.21
Those SPIOs that have been approved have been used as MRI contrast agents.
The usefulness of SPIO NPs lies in their ability to be magnetized but possess no net magnetization, which allows them to retain colloidal stability. 20 This allows for injection and target tissue accumulation, which can then be imaged with MR with either T1 or T2 weighting. Several polymer coatings have been investigated to improved pharmacokinetics of SPIO NPs in vivo . Commonly used among NPs is
PEG, which reduces recognition by the RES and thereby increases circulation time. Dextran and chitosan polysaccharides have also been used as hydrophilic, biocompatible, and biodegradable coatings for SPIO NPs. 21 SPIO NPs may also be targeted to specific tissues and cells by addition of targeting moieties such as antibodies, proteins, and small organic molecules such as folate. 21-24 A number of strategies have also been investigated for chemotherapeutic delivery with magnetic NPs, which has primarily focused on cancer therapy. Investigated drugs include paclitaxel, DOX, and methotrexate, and strategies involve physical complexation, entrapment within a hydrophobic coating, or attaching cleavable linkers for specific release upon delivery to target tissues. 21 Magnetic NPs may also be used as part of a larger NP complex. Ito et al. embedded magnetite NPs within liposomes, which were targeted with anti-HER2 antibodies, allowing for combination antibody and hyperthermic treatment. 25
Quantum Dots
7
Quantum dots (QDs) are nanoscale particles made from semiconducting materials such as cadmium sulfide, cadmium selenide, or cadmium telluride, which have size-dependent fluorescence that can be tailored to wavelengths in the NIR. Due to their high quantum yields and small size QDs have been investigated for their potential in both in vitro and in vivo imaging. 26, 27 However, the synthesis processes and materials required necessitate further solubilization and improvements in biocompatibility, such as conjugation of PEG or loading into liposomes or polymeric NPs. Nonetheless, QDs have shown pre-clinical successes in imaging cellular processes and lymph nodes. Jaiswal et al. used endocytic uptake and surface labeling with antibodies to specifically label cells for imaging over the course of a full week. 28 Kim et al. showed that NIR fluorescent
QDs coated with oligomeric phosphine resulted in soluble NPs approximately 10 nm in diameter, 29 which were capable of tracking the lymphatic system in both healthy mice and pigs, indicating the potential for QD tracking of sentinel lymph nodes (SLNs) in cancer patients. Further work is needed to prevent toxic side effects of QDs, as the materials that compose QD cores may contain relatively large amounts of Cd, Te, Pb, or As.
Polymeric Nanoparticles
An enormous variety of polymeric materials have been investigated as materials in NPs for use in medicine and cancer in particular. Combinations of random and block co-polymers utilizing poly(glutamic acid), poly(lactide), poly(caprolactone), N-(2-hydroxypropyl)methacrylamide, numerous hydrogel
8 materials, and polysaccharides such as chitosan, dextran, and HLA have all shown potential as materials for drug delivery or imaging. 30-35 Some of these drugs have found and are finding their way into the clinic, such as Cremophor- paclitaxel, a micellar formulation of castor oil and PEG, and next generation drugs such as Genexol ®-PM, a polymeric micelle formulation composed of diblock co-polymers of PEG-poly(D,L-lactic acid) and paclitaxel. 36-38
The advantages of polymeric NP formulations results from several attributes. 1, 20, 34, 39-41 First, the wide variety of polymers and molecular weights to choose from allows for a high degree of control over the properties of the NPs.
The ability to synthesize random or block co-polymers, dendrimers, polymer- conjugates, and composite materials allows for synthesis of a huge variety of
NPs. Many of the NP types previously discussed such as liposomal and metallic
NPs have been investigated as composite materials utilizing one or more polymers. Polymeric NPs are capable of loading a variety of drugs, contrast agents, and nucleic acids, both non-covalently and covalently. This can improve solubility and lessen dose-limiting toxicity. Surface modification with ligands and antibodies may improve targetability, and many polymers are biodegradable and biocompatible, allowing for higher levels of exposure without associated toxicity.
Of particular interest to this work are polymeric nanoparticles containing hyaluronic acid, a naturally occurring biopolymer with a wide range of physiological functions.
Hyaluronic Acid Derived Nanoparticles
9
Hyaluronic acid (HLA), or hyaluronan, is a linear, anionic polysaccharide composed of β(1,4) D-glucuronic acid β(1,3) N-acetyl-D-glucosamine ( Figure
1.2 ). HLA is present ubiquitously throughout the human body, primarily in the
ECM of soft tissues, and acts to lubricate joints and to support connective tissue. 42 High levels (> 200 µg/ml) of HLA are found in the umbilical cord, synovial fluid, and dermis. 43, 44 HLA is very hydrophilic as it contains a relatively large number of hydrogen bond donors (5) and acceptors (12) compared to the number of carbons (14) per repeat unit. This primary structure causes HLA to retain water very well, which contributes to its function in the ECM. 45 The presence of the carboxylic acid on D-glucuronic acid further contributes to its usefulness in material applications as this moiety may be used for crosslinking and conjugation.
Figure 1-2. Hyaluronic acid is a polymer of repeating β(1,4) D-glucuronic acid
(left) and β(1,3) N-acetyl-D-glucosamine (right). The natural MW of these polymers are in the range of 10 5 – 10 7 Da and can be tens of microns in length.
Degradation by acid catalyzed hydrolysis or using hyaluronidases can occur to further reduce the MW.
10
In addition to these roles, HLA serves as the ligand for several receptors including CD44 and RHAMM. In cancer HLA has been found at high levels in multiple tumors types, including breast, ovarian, non-small cell lung adenocarcinomas, and prostate cancer, and high levels have been associated with increased invasiveness. 42 Additionally, some tumors also express high levels of hyaluronidases HYAL-1 and HYAL-2, a case that may be involved in increased malignancy. 42, 46, 47 CD44 and RHAMM signal transduction are likely pathways that link HLA to tumor behavior, and CD44 splice variants have been associated with more malignant cancers. 48
HLA has been used previously for NP formulations and NP delivery to tumors. Because of its hydrophilic nature, HLA must be conjugated with other materials to form discrete NPs. Choi et al. showed that conjugation of HLA with a
5β-cholanic acid moiety triggers self-assembly of ≈200 kDa MW HLA into NPs, and that these NPs accumulate in tumors to a higher degree than non- conjugated HLA polymers. 49 This work was continued by introducing PEGylation of NPs, which resulted in improved delivery to tumors by avoidance of RES uptake, and further continued by loading of the hydrophobic chemotherapeutic drugs DOX and camptothecin into these PEGylated NPs.50, 51 In vivo delivery of camptothecin using these NPs showed a reduction in tumor growth over the course of 35 days when compared to free camptothecin, suggesting improved delivery to tumors. More recently, these materials have been used as nucleic acid delivery systems for RNA interference and have shown preclinical success in gene silencing. 52 Additional studies have investigated the potential for inclusion
11 of photosensitizers for photodynamic therapy. 53 Yoon et al. utilized the loading of chlorin e6 within PEGylated HLA NP cores, which was delivered to tumors in vivo and showed improved tumor size reduction after treatment with NIR light compared to delivery of free chlorin e6. Zhang et al. loaded Cy5.5 conjugated
HLA NPs with CuS for combined photoacoustic imaging and photothermal therapy.54 Both studies delivered these NPs in vivo and found that treatment reduced tumor volumes compared to controls. Similar studies have used additional hydrophobic ligands such as the amino acid histidine.55
Other studies have examined HLA composites and co-polymers in the formation of NPs. Studies by Cho et al. examined NPs formed by first conjugating ceramide to HLA, followed by mixing with Pluronic p85 in an attempt to improve NP stability and overcome multidrug resistance, and showed cellular toxicity and uptake within tumors. 56 Further work utilized PEGylation of HLA- ceramide NPs loaded with DOX, which resulted in smaller tumor volumes than non-PEGylated NPs loaded with DOX, presumably due to increased circulation time and lower uptake by the RES. 57 Co-polymers of HLA have also been examined using chitosan or poly(D,L-lactide-co-glycolide) for gene delivery and drug delivery respectively.58, 59 Composite NPs of gold-HLA and SPIO-HLA have also been investigated. Dakdouki et al. utilized HLA coated SPIO NPs to improve uptake into cancer cells, which allowed for MR imaging as well as drug delivery by conjugation of DOX.60 These NPs resulted in a greater decrease in T2* compared to Feridex, a clinically approved SPIO NP, and greater cytotoxicity at equivalent concentrations to free DOX. Another study utilized direct conjugation
12 of HLA to AuNPs in addition to the antibody drug Tocilizumab for the treatment of rheumatoid arthritis. 61 These NPs resulted in greater reduction (improved treatment) of mean clinical scores for arthritis induced in mouse feet compared to either AuNPs without Tocilizumab, or free Tocilizumab, indicating that HLA was a determining factor in treatment improvement.
In cancer imaging, HLA has been investigated preclinically for SLN mapping and image guided surgery. While not utilized in NP formation, HLA was mixed with patent blue or SPIO NPs and injected into the gastrointestinal tract of pigs, which resulted in prolonged and more discrete staining of lymph nodes. 62
However, this resulted primarily from the increased viscosity of the HLA-dye solution rather than NP-mediated delivery. HLA-indocyanine green conjugates have been examined for use in image guided surgery by direct conjugation of
ICG-OSU, a derivative of indocyanine green (ICG) that can be conjugated directly to HLA. This formulation successfully accumulated in xenograft tumors after systemic injection, however, while ICG is an FDA approved fluorophore,
ICG-OSU is not.
The extensive application of HLA, either as hydrophilic backbone, co- polymer component, surface coating, and conjugate in NP formulations clearly demonstrates its versatility. HLA is hydrophilic, biodegradable, biocompatible, and possesses numerous sites for simple conjugation. It also has potential to serve an active targeting function either through CD44 targeting or degradation by hyaluronidases expressed within tumors. A summary of the NP formulations utilizing HLA is found in Table 1-1.
13
Use of HLA NP type Indication Action Reference
Backbone Polymeric micelle Solid tumor Pre-Drug delivery 49 , 50 , 63 (5β-cholanic acid conjugate) Drug delivery 51
siRNA/miRNA delivery 52
Phyotodynamic 53 , 54 therapy/imaging Polymeric micelle Drug delivery 55 (Histidine conjugate) Polymeric micelle Drug delivery 56 , 57 , 64 , 65 (Ceramide conjugate) Photo X-linked HLA Drug delivery 66
Chemically X-linked Drug delivery 67 HLA Polymeric micelle Asthma Dexamethasone 68 (Flt-1 conjugate) delivery Co-polymer Polymeric micelle Solid tumor Drug delivery 58 , 69 Poly(DL-lactide-co- glycolide) Polymeric micelle siRNA delivery 70 Poly(ethyleneimine) HLA-chitosan Various Nucleic acid delivery 59 , 71 -73
Asthma Heparin delivery 74
Table 1-1. Summary of recent literature utilizing HLA in NP formulations.
14
Use of HLA NP type Indication Action Reference
Surface Moiety SPION Solid tumor MRI, drug delivery 60
Macrophages Imaging 75
AuNP Rheumatoid Immunotherapy 61 arthritis Hepatitis C Interferon α delivery 76
Diagnostics Nanoprobe 77
Mesoporous silica within Solid tumor Drug delivery 78 liposome Lipid NP cluster Solid tumor Drug delivery 79 , 80
Calcium phosphate NP Solid tumor siRNA delivery 81
Table 1-1. (continued)
15
2.0 Biodistribution and Toxicity of Nanoparticles
NP formulations are advantageous due to their size and chemical characteristics. The volume of NPs ranging from 10-500 nm in diameter is large enough to provide both interior hydrophobic environments in which drugs and small molecules can be loaded, while retaining a hydrophilic exterior to allow for solubility in blood. Additionally, the high surface area to volume ratio allows for chemical modifications that actively target specific cells and tissues, or provide immuno/metabolic evasion. 82, 83 Perhaps most importantly, NP diameters can often be tailored to fall within the critical size range to take advantage of the enhanced permeability and retention (EPR) effect displayed by many tumor types. Nonetheless, efficient delivery to target tumors remains difficult, with only a small percentage of injected dose being delivered. The barriers to effective delivery of NPs are numerous. 84 The size range must be greater than approximately 6 nm to avoid renal filtration, but sufficiently small to allow for extravasation into the tumor stroma. 85 Off-target delivery, particularly uptake by the RES, results in the majority of NP loss. For those particles that do extravasate into the tumor, they must then diffuse through the tumor stroma, avoid diffusion back into circulation or lymphatic drainage, and be taken up by target cells, should uptake be necessary for their function. For NPs that do not find their destination, the toxicity from off-target delivery or accumulation represents a significant concern, particularly for drug delivery formulations that require continued dosing for efficacy. These difficulties have been met by
16 research into the effects of surface modification, charge, size, shape, material composition, and conjugation of targeting and stealth ligands.
The Reticuloendothelial System (RES)
The RES is a system of macrophages found in the liver, spleen, and bone marrow that is capable of selectively removing particles from the blood which have bound opsonin proteins. 84 RES clearance is most prominent in the liver and spleen, and does not result from specific recognition of NPs by the macrophages, but rather from serum proteins bound to the NPs. These proteins are called opsonins and primarily include complement proteins and antibodies. 86 Opsonin binding can result in rapid clearance of NPs from the body and thus drastically reduce circulation and delivery. Thus, reduction of RES uptake is a key attribute of good NP formulations. However, even for NP formulations that guard against recognition by use of stealth mechanisms such as PEGylation, RES uptake remains significant, often at more than 50% of injected dose. 84
Tumor Structure and the Enhanced Permeability and Retention (EPR) Effect
For those NPs that are not taken up by macrophages, penetration into the tumor and delivery to cells provides an additional barrier. Tumor biology is complex and varies between tumor types, location, and patients. For these reasons, the vascular and stromal structures of any one tumor will be unique, and will thus result in different NP distributions. The extent of NP penetration into and retention within tumors depends on the vascular fenestration size, density
17 and composition of the ECM, pressure, and lymphatic drainage. These features are likely to be so variable that different NP formulations may be required for different tumor types. Nonetheless, similarities exist between all tumors exhibiting enhanced permeability and retention.
The EPR effect results from the poor vascular formation in solid tumors, causing preferential retention of macromolecules and NPs in solid tumors. 82, 87-90
When tumors reach a sufficient size, approximately 2-3 mm in diameter, oxygen and nutrients cannot diffuse rapidly enough to support cells in the interior of the tumor. When this occurs, growth factors are released that trigger angiogenesis in the surrounding tissue, forming new blood vessels. The vasculature formed however, is not well structured. Endothelial cell linings are generally not as close as in healthy vasculature and there is often a poorly formed or complete lack of smooth muscle cells lining the basement membrane. These properties lead to the increased extravasation of all molecules that can fit between the fenestrations in the endothelial layer, ranging from 200-2000 nm. After extravasation, larger particles are preferentially retained in the tumor due to the lack of appropriate lymphatic drainage. While ions and small molecules can diffuse back into the vasculature, macromolecules and NPs do not diffuse back as rapidly and thus tend to accumulate within the tumor interstitium. Thus, NPs with diameters smaller than the fenestrations can passively target those tumors without additional motivating forces. However, due to variability in the structure and extent of vasculature, tumor stroma, and lymphatic drainage, the EPR effect
18 varies by tumor type and location, making it a useful but not ubiquitous targeting method.
The Mechanical Environment
A theme that is frequently absent from the discussion in NP delivery is the physical environment experienced by NPs in flow conditions. Wall shear stress
(WSS) is the tangential force applied to a surface by a fluid moving across it. At a molecular level, WSS is essentially a measure of lateral contact forces between molecules moving at different velocities. In an ideal cylinder, shear stress is highest at the solid-liquid interface (WSS), and decreases toward the center of the cylinder. Physiological WSS values range from 50 dyne/cm 2 in the aorta, 91
2.8-100 dyne/cm 2 in capillaries, 92 and 17-211 dyne/cm 2 in arterioles. 93 In mice aortas, WSS values are predicted to range from 60-220 dyne/cm 2, and can be as high as 600 dyne/cm 2.94 For comparison, water flowing at 1 ml/s in a 3.175 mm
(1/8”) diameter tube experiences WSS of around 2.5 dyne/cm 2, and solutions standing stationary experience essentially zero shear stress. Contact between
NPs and vascular walls, as well as collisions with serum proteins, cells, and other molecules can thus be significantly higher than what is normally experienced in laboratory practice.
Nanoparticle Properties Affecting Delivery
Every aspect of the NP affects the interaction its biological environment, which includes cells, ECM and serum proteins, and physical factors such as flow
19 and fenestration size. NP size has physical effects on renal clearance, extravasation, diffusivity, and influences cellular interaction. NP shape also influences cellular interaction and cellular uptake by both cancer and macrophage cells. Surface charge, either positive or negative, may lead to interaction with charged groups in the extracellular matrix, while surface polarity may influence solubility and protein binding. Some of these effects may be hidden by surface modification. As concerns net biodistribution, the benefit of adding active targeting ligands has shown mixed results, but appears to produce increased cellular internalization, which may improve efficacy. These issues result in an incredibly complex set of parameters that must be considered when designing new NP formulations for a given purpose.
Size
Nanomaterials are by definition on the order of nanometers in at least one dimension, and in the case of NPs, three dimensions. These scales are chosen due to the unique physical and chemical properties that arise from particles at this scale, one of which is delivery and accumulation within tumors. However, even the effect of size on NP delivery is multifaceted. NPs above 6 nm but smaller than the 200-2000 nm vascular fenestration size of tumors will avoid renal clearance and penetrate into the tumor stroma to some extent.82 However, a further hindrance to delivery is the 50-100 nm fenestrations in the liver and interaction with hepatocytes that results in RES uptake.84 And what is optimal for avoiding RES uptake may not be optimal for tumor uptake. Cabral et al. found
20 that 30 nm drug loaded polymeric micelles, with a size small enough to pass through liver fenestrations, had the best penetration into poorly permeable tumors, compared to 50, 70, and 100 nm micelles, though all sizes penetrated well into highly permeable tumors. 95 Complicating the matter more, what is optimal for tumor uptake may not be optimal for tumor retention. Torchilin et al. showed that small 10 nm micelles accumulated in the tumor within 30 minutes of administration, but the total dose of taxol within the tumor was substantially lower at 2 hours compared to 30 minutes, unless a tumor targeting antibody was present. 96 While this observation possibly indicates a retention effect due to targeting, it is complicated by the addition of large (>100 kDa) monoclonal antibodies to an already small micelle, making it difficult to separate the effects of active targeting and decreased diffusion out of the tumor. Other work was performed strictly examining the MW on the effect of tumor penetration and retention. Dreher found that increasing the MW of fluorophore-linked dextran from 3.3 kDa to 2 MDa resulted in significant and meaningfully lower vascular permeability, a 100-fold reduction overall for the 2 MDa construct.97 However, total tumor accumulation was highest for 40-70 kDa molecules, as higher MWs resulted in lower penetration but longer circulation time, while lower MW resulted in higher penetration but rapid diffusion back into circulation. Another study found a size dependency of AuNP size in relation to uptake by HeLa cells, with 50 nm
AuNPs being taken up to a higher extent than 14, 30, 74, or 100 nm NPs. 98
21
Shape
NP shape also influences biodistribution. Champion and Mitragotri found that differently shaped polystyrene particles had different rates of phagocytosis. 99
These “worms”, elongated polystyrene particles, were multiple micrometers in length and had volumes equivalent to 3 µm diameter spheres. It was observed that the worms had significantly lower uptake by macrophages than spheres of equivalent volume. However, these worms had lengths of approximately the same size as the macrophages and thus may not represent the behavior of macrophages interacting with nanoscale materials. Chauhan et al. examined the effect of shape with nanoscale materials using PEGylated 35 nm diameter
CdSe/CdS quantum dot core with silica shell spheres, and PEGylated 54 nm length CdSe quantum dot core, CdS elongated rods. 100 The results showed that the transvascular flux and intratumor distribution of the rods were both significantly higher than the spheres. However, the mass difference of these particles was not given, and a volumetric calculation shows a much higher volume of the spheres than the rods, which may be a contributing factor to the differences observed. Chithrani also found a shape dependency on cellular uptake by HeLa cells, with lower aspect ratio AuNPs being taken up more than high aspect ratio AuNPs. 98 However, this effect could be two-sided, as increased uptake could benefit intracellular delivery of drugs, but hinder delivery by increasing uptake by macrophages.
22
Charge
NP charge has a significant effect on uptake by the RES as well as penetration through the tumor stroma. It is generally agreed that NPs with a moderate zeta potential (ξ) between ±10 mV have lower uptake by the RES. 84, 101
Levchenko found that the rate of liposome clearance was dependent on ξ, with more highly charged liposomes being cleared more rapidly. 102 Similar results have been produced in other studies utilizing different materials, suggesting that this effect does result from differences in ξ. 103, 104 This effect of surface charge likely results from alterations in NP interaction with plasma proteins, cell surface proteins, and the stroma. The presence of high surface charge on NPs increases opsonization and thus uptake by the RES. 86 However, while net charge leads to uptake, the surface must also be sufficiently polar, as uncharged but hydrophobic surfaces also result in higher levels of opsonization. 105 Charge effects interaction with both serum and stromal proteins, and thus also influences the penetration of
NPs into tumors and other tissues. With regard to stromal proteins, again we observe that particles with close to neutral ξ result in higher tumor penetration. 82,
84 These results have been explained theoretically by modeling NP interactions with extracellular matrix molecules such as collagen and HLA.106 The model agreed that particles with higher surface charge density should experience lower diffusion into the stroma due to interactions naturally occurring between positively and negatively charged molecules, and thus near neutral surface charge is ideal for tumor penetration after extravasation. This trend is confirmed by the effect of
23 surface modification of NPs, particularly the use of PEG, which provides a surface barrier of uncharged hydrophilic polymers.
Surface Modification
In order to combat the effects of opsonization or non-specific interaction with stromal proteins, and potentially allow for targeted delivery, a number of strategies for conjugation of surface moieties have been developed. Perhaps the most common methodology for making “stealth” NPs is PEGylation. 86 PEGylation works because PEG is uncharged, non-specific, and has low protein association.
This provides a physical buffer that reduces protein interaction with the underlying material, which reduces binding of opsonins, decreases uptake by immune cells, and lowers interaction with the extracellular matrix. However, the strategy of PEGylation is not without issues. Treatment with PEGylated NPs may result in development of hypersensitivity to PEG, resulting in decreased efficacy of future treatments. Judge et al. found that treatment of mice with PEGylated plasmid-containing liposomes resulted in the production of an anti-PEG antibody that decreased delivery to tumors after repeated exposure. 107 In addition to adaptive immune response, PEG may also trigger a response of the innate immune system through complement activation. A study in humans using Doxil ®
(PEGylated liposomal DOX), found hypersensitivity reactions and complement activation following administration in 45% of patients. 108 However, it is not clear from this study what component of Doxil triggered the response. The quantity and molecular weight of PEG also influences biodistribution. In micellar and
24 liposomal formulations, the total quantity of PEG may also affect the stability of the NPs themselves. 109 It has been reported that adding a sheddable layer of
PEG could provide the benefits of increased circulation time while allowing for reduced toxicity and sensitivity. 109 Additional methods of making stealth NPs have been investigated as well, including alternative polymers 110 or conjugation of “self” peptides that cause immune recognition of NPs as belonging to the host.
Active targeting refers to the strategy of linking proteins, peptides, antibodies, or any molecule that is specific to some aspect of tumor biology in order to improve tumor or cancer cell uptake. While active targeting does not appear to influence the physical factors determining NP deposition within tumors, evidence is building that it does affect receptor mediated endocytosis and thus may serve as a complementary strategy in tumor accumulation.82 The large number of FDA approved immunoconjugated drugs and imaging agents suggests that there is indeed a basis for active targeting with NPs. 111 However, studies have suggested that the net effect on tumor accumulation may be small, and the improved efficacy of targeted NPs results from increased cellular internalization. 112, 113 It must also be realized that the addition of active targeting moieties to the surface of NPs may alter the biodistribution of these NPs, further complicating the issue. Nonetheless, a number of targeted NP formulations are currently in clinical trials, the majority of which are targeted to solid tumors, and a large number of preclinical studies have taken place. 114
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Toxicity
The study of NP toxicity is a complex issue, due in large part to the vast array of materials and associated drugs and other molecules that compose NP formulations. Currently however, NPs approved for biomedical use have a good record in this area, particularly those formulations designed to deliver chemotherapeutics. Doxil provides one of the best examples, as the liposomal formulation of DOX prevents diffusion across healthy endothelial barriers and thereby greatly reduces the cardiotoxicity observed with free DOX.115 However, the toxicity profile of each new NP formulation will need to be determined independently. Broadly, the concerns in NP toxicity are the generation of reactive oxygen species (ROS), size and shape dependent accumulation, and material chemistry. 116, 117 Generation of ROS has been observed across multiple NP types. Damage occurs from oxidative stress through reactions with critical cellular components such as DNA and the cell membrane. Accumulation due to
NP size and shape may occur in a number of locations, including the liver, spleen, pulmonary system, and macrophages. 116 This may itself result in a range of complications, including inflammation, direct interaction and disruption of protein and nucleic acid function, and further generation of ROS. Direct material toxicities have been reported through environmental studies of metal oxide NPs, but this may depend on the both the route of administration and the total dose. 118
Unfortunately, there may also be poor overlap between in vitro and in vivo toxicological assays. Attempts to improve overlap have utilized 3D cell culture and cell co-culture techniques, but because of the vastly different dynamics
26 experienced by NPs in vivo and the much larger array of cell types and cell interactions, this overlap will likely never be entirely complete. 116, 117 Careful analysis will be required for each new NP formulation, as differences in physicochemical properties of the materials will result in different biodistributions as previously discussed, and subsequent cellular, tissue, and organ interaction will depend on NP material reactivity, biodegradability, and therapeutic payload.
3.0 Clinical Imaging in Oncology
Clinical cancer imaging utilizes a number of different modalities, including computed tomography (CT), magnetic resonance imaging (MRI), positron emission tomography (PET), single photon emission computed tomography
(SPECT), ultrasound (US), and optical imaging. While this may appear to provide a number of viable options for imaging of various cancers, in themselves all of these modalities lack specific targeting of neoplastic tissue. Molecular probes, such as fluorodeoxyglucose ( 18 F) (FDG), are capable of specifically targeting cancer cells, and molecular targeting is available clinically for use with
PET, SPECT, and MRI. 119 Molecular targeting with x-ray (CT) imaging is unlikely to occur due to the high molar quantities required to attentuate x-rays. 120 The use of microbubbles as contrast agents in US remains outside the realm of oncological imaging as the size of the these bubbles prevents their extravasation into the tumor stroma. 120 Targeted optical imaging however has been used both pre-clinically and clinically, and though the techniques, materials, and
27 instrumentation are still either new or experimental, this methodology holds significant potential, particularly in the field of image guided surgery. 121, 122 Here we review the basic medical imaging modalities and their uses in cancer diagnosis and treatment, with a focus on the potential for intraoperative image guidance.
X-ray and Computed Tomography
X-rays are produced by the acceleration of electrons through a potential difference to a target anode where a small percentage of them come close enough to target nuclei that the electrons lose kinetic energy and subsequently release an x-ray photon. 123 These photons are then either scattered or absorbed by molecules in tissue, causing attenuation depending on tissue density and composition. X-ray photons that are not absorbed by the tissue then reach either a screen-film or charge coupled device that records their location, providing information on electron flux through the tissue of interest.
Perhaps one of the most pervasive and common medical imaging procedures is the mammogram. This technique utilizes x-rays to detect neoplasms in the breasts and has become a standard of care in screening over the last several decades. Mammography has clearly proven successful in reducing the number of late stage diagnoses and deaths from breast cancer, however specific contrast agents are not available for standard mammographies and recent analyses of long term clinical utilization of mammography screening have shown that overdiagnosis is common. 124, 125 Bleyer et al. have estimated
28 that approximately 70,000 women were overdiagnosed with breast cancer in
2008, approximately one-third of all breast cancer cases. This overdiagnosis undoubtedly results in unnecessary medical expenditures, procedures, and patient duress. Thus, new methods of breast cancer screening are underway, including those utilizing dedicated CT. 126 Breast cancer screening with CT may improve resolution and thus potentially serve to both decrease incidence and overdiagnosis, but additional work is needed to determine the degree of improvement over standard mammography. 126, 127
Virtual colonoscopy by CT imaging remains a potential screening route to replace traditional colonoscopy. Traditional colonoscopy is an effective procedure but remains highly invasive. Earlier studies showed that virtual colonoscopy may not provide sufficient sensitivities for detection of cancers 128 , however recent studies have suggested that the sensitivities are adequate for detection of colorectal cancer and may even exceed that of optical colonoscopies when appropriate techniques such as fecal tagging are used. 129
Clinical CT provides anatomic imaging, and can easily distinguish between soft and hard tissues such as bone and surrounding muscle. CT also has comparatively good resolution (~1 mm 3) compared to nuclear imaging modalities, which have resolutions of approximately 8-12 mm 3.120 However, CT contrast agents must be injected at molar quantities and quickly become diluted, thereby precluding targeted imaging. 120 CT therefore lacks functional imaging that can be found with nuclear imaging modalities such as PET and SPECT.
29
Thus, dual modality imaging has become a subject of intense research and clinical application in oncology.
PET and SPECT
PET and SPECT operate by detection of gamma rays emitted from radionuclides delivered to the tissue of interest. In the case of PET, the radionuclide first releases a positron, which then collides with a nearby electron to release two oppositely directed gamma rays. 130 PET is well known for its use of FDG, a glucose analog that is taken up by cancer cells exhibiting high levels of aerobic glycolysis. 131 Additional functional imaging agents are also in use clinically for both PET and SPECT. 119, 132 The advantage of PET lies in the ability to synthesize radiolabelled probes that mimic molecules used in cells’ molecular machinary, thereby allowing for imaging of functional aspects of cancers.
Additional tracers in use include analogues of thymidine, choline, methionine, tyrosine, and several others. 133 This versatility of PET allows for its use in both pre and post-treatment of cancers, through diagnosis, staging, and restaging.
Combined imaging modalities are extremely powerful in their ability to distinguish both anatomic and functional aspects of tissues. The combination of
PET/CT allows for high resolution anatomic imaging with combined functional molecular imaging analysis, allowing for differentiation of healthy and diseased tissue by physical and biochemical attributes. 134 Combined PET/CT scanners are preferable for this methodology in order to avoid discrepancies in the physical
30 position of organs and patient. Similar attributes are available with SPECT/CT systems. 135
Combined imaging with PET/MRI may also be done to image cancers for multiple aspects of molecular and physical properties. For example, PET imaging can acquire data regarding expression of cell surface proteins and changes in those levels during treatment, while MRI can acquire physical data such as cell density and perfusion. 136 Thus, combined PET/MRI imaging can acquire data on biochemical and physical characteristics of tumors for diagnosis, staging, and progression during treatment that is not available through other modalities.136
Magnetic Resonance Imaging
MRI is regarded as the most useful imaging modality for imaging of soft tissue, with resolution at approximatley 1 mm 3 being comparable to CT and
US.120, 132, 136 This imaging modality relies on the alignment of the magnetic spin of nuclei in a strong magnetic field and subsequent perturbations of the spin axis by a secondary magnetic field. Protons of nuclei in different physical environments will return to their original alignment at different rates, thus allowing tissue contrast and imaging. The ability to precisely image soft tissues at high resolution at any depth makes MRI particularly useful for cancer imaging, and this modality is not limited to simple anatomical imaging. MRI can also be used to perform functional imaging using perfusion, diffusion, and contrast agents such as ferumoxide, ferumoxtran, and ferumoxytol, all of which are iron oxide particles. 119 Gadolinium ions, which are paramagnetic and shorten T1 relaxation
31 times of nearby protons, are also a commonly used MRI contrast agents. 137, 138 A number of studies show the effectiveness of MRI imaging in breast and prostate cancer. MR imaging in breast cancer was estimated to diagnose an additional
16% of patients with a positive predictive value between 52 – 77%, indicating approximately twice as many true positives as false positives. 139 However, screening and staging with MRI has also been shown to result in changes to decision making regarding surgical care, tending towards more radical surgeries, and the improvement in patient outcomes may not be improved.139, 140
Multiparametric MRI has also shown to be useful in imaging of prostate cancer. 141-143 Prostate cancer is typically diagnosed through non-imaging means such as detection of prostate specific antigen and biopsy, however the potential for anatomical and functional MR imaging has the potential to supplement these methods by biopsy guidance and preoperative characterization, and detection of local recurrance. The introduction of targeted contrast agents may further the potential for diagnosis and screening as well. 136
Non-fluorescence Optical Imaging
Imaging with visible light remains one of the most widely utilized forms of imaging in medicine. Endoscopy of the upper and lower gastrointestinal tract, and laparoscopic surgery, are the the major fields where optical imaging technology is utilized. Upper GI tract endoscopy has proven successful in the identification of numerous cancers, including esophageal and gastric cancer. 144
Similarly, the small intestines and lower GI tract can also be imaged
32 endoscopically, either through placement of handheld instrumentation or by swallowing a small capsule that images the gut during its passage. 145
Related to optical imaging are other imaging methodologies that utilize visible or NIR light to diagnose cancer. Optical coherence tomography, photoacoustic imaging, and raman spectroscopy are less widely used or remain under investigation for oncologic imaging, but may play a role in the future.
Optical coherence tomography, which utilizes near infrared (NIR) light reflections from tissue, is currently being investigated for imaging of both cervical and bladder cancers. 146 Photoacoustic imaging utilizes heating of tissue by photo- irradiation, which results in thermal expansion and the production of a mechanical wave that can then be measured by ultrasound. 147, 148 Raman spectroscopy is also under investigation for diagnosis of breast cancer, and has proven successful in identifying cancer in tissue samples, though not yet in patients. 149, 150
Non-fluorescence Intraoperative Tumor Detection
Radiological techniques are also used in the intraoperative treatment of cancer. Radioguided occult lesion localization (ROLL), and radioguided seed localization (RSL), together termed radioguided localization (RGL) for the intraoperative identification of non-palpable breast cancer. 151, 152 These techniques utilize the injection or placement of a radioactive tracer into the tumor prior to surgery. This is in contrast to the current gold standard of wire-guided localization, which utilizes the pre-operative placement of a wire into the non-
33 palpable lesion. ROLL utilizes Tc-99 while RSL utilizes a titanium seed containing Iodine-125. 151 Gamma tracers are then used to localize the radiation signal in the operating room and thereby improve resection of otherwise difficult to detect tumors.
A technique which has become recently available which does not use ionizing radiation or rely on any exogenous agents is margin detection by electrical properties of tissue. A number of recent studies have shown that a device called the MarginProbe ® can differentiate cancerous from healthy tissue so as to allow the detection of positive margins. 153, 154 This technology relies on differences in the membrane potential, cellular connectivity, and vascularity of cancer cells and tumors. Results showed that more positive margins (PMs) were detected using this technique than with visual inspection and palpation alone, resulting in few resections. However, false positive rates were higher among patients who’s lumpectomies were examined with the device. 153, 154 Nonetheless, the potential identify PMs without relying on delivery exogenous contrast agents may prove extremely useful alone or in addition to additional methods of PM detection such as fluorescence imaging.
4.0 Fluorescence Imaging and Image Guided Sugery
Fluorescence based imaging in oncology is a growing field in both research and clinical practice. Cancer diagnosis using endogenous and exogenous porphyrins date back to 1924 and 1942 respectively. 132 The use of
34 fluorescence in image guided surgery dates back to 1948, when fluorescein was observed to improve the detection of brain tumors. 155 This methodology requires three broad aspects for use: fluorescent contrast agents that accumulate in target tissues to sufficiently high levels for both detection and contrast, an excitation light source, and a light collection device capable of capturing both emmitted fluorescence and reflected visible light for anatomic localization. Of particular interest is NIR fluorescence imaging, which exhibits lower autofluorescence and tissue interference compared to other wavelengths, and is already in clinical use. 132 The methodologies, technologies, and recent experimental advances of oncologic fluorescence imaging are discussed here in detail.
The Clinical Need for Image Guided Surgery
Surgery is one of the most common and effective forms of cancer treatment, being used in the majority of solid tumor cases. Between 63-98% of either lung, breast, bladder, or colorectal cancer patients will receive surgery, depending on type, grade, and age at diagnosis. 156 Prostate cancer patients are also commonly treated with surgery, with surgery being chosen as part of intervention for between 6-57% of patients, depending on age, with younger patients being more likely to receive surgery. Surgery also offers the only option for curing colorectal metastases to the liver. 157 These cancers are among the most commonly occurring forms, and therefore it is not difficult to see that the majority of cancer patients in the US receive some form of surgical treatment. It is commonly understood that in most types of cancer it is not the primary tumor
35 that is the ultimate cause of death, but rather metastases that lodge in numerous critical areas which then grow and shut down necessary physiological functions.
It is therefore critical that the primary tumor, local metastases, and metastatic lymph nodes are removed prior to distant metastasis. However, much of the intraoperative detection of tumors relies on visual and tactile differentiation of healthy and diseased tissue, resulting in significant levels of residual cancer cells and undetected metastases.
PMs are residual cancer found at the borders of surgically removed tissue, and pose a significant health risk. Approximately 30% of breast cancer patients will experience recurrence of disease locally or systemically. 158, 159 Rates of PM occurrence are estimated to be between 20-40% for patients receiving breast conserving surgery. 160 For hepatic resection of colorectal metastases the intrahepatic recurrence rate is 11-37.5%, overall recurrance is as high as 40-
50%, and PMs are observed in 2-5% of patients. 161, 162 Importantly, in patients receiving both chemotherapy and surgical treatment for colorectal metastases to the liver, PMs have been identified as the only risk factor for patient survival that is associated with the surgery itself.163, 164 PMs also occur in 11-38% of patients receiving radical prostatectomy and are a strong negative predictor of recurrence and patient survival. 165-167 While some new imaging modalities such as
Cherenkov luminescence imaging may provide ways to detect PMs more effectively,168 current methods of detecting tumor margins intraoperatively involve palpation and visual inspection. These are characterized by poor tissue contrast and spatial resolution and lack detection nonpalpable lesions. 151, 152 Traditional
36 imaging modalities such as radiography, MRI, PET, SPECT, and US, typically involve instrumentation that is too cumbersome or dangerous to use in the surgical theater, or are incapable of providing tumor-specific information with high contrast or resolution. 120 To combat these limitations, a number of image guided modalities are currently under investigation. These include US, MRI, and optical fluorescence using both endogenous and exogenous fluorophores. 169-171
Physics and Chemistry of Fluorescence
Photon emission from molecules and atoms occur when electrons are excited above the ground state energy level and subsequently transition back to the ground state by emission of a photon of energy equal to the difference between the excited and grounds states, as shown in Equation 1.1,
[1.1] = ℎ = − where ℎ is Plank’s constant and is the photon frequency. This representation is inexact because both the ground and excited energy levels have additional thermal or vibrational energy levels within them. A Jablonski diagram is shown in
Figure 1.3 illustrating the absorbance excitation and fluorescence emmision of photons by electrons.
37
Figure 1-3. Jablonski diagram showing photon absorbance and fluorescence emission, designated and respectively. Absorbance occurs when a ℎ ℎ photon of energy is absorbed by an electron. The electron then undergoes ℎ vibrational relaxation as designated by the green dashed arrow, and emits a photon of energy to another vibrational energy level within the ground state. ℎ
Fluorescence is differentiated from phosphorescence specifically by the electron spins in the excited state. 172 In fluorescence, the electron is in a singlet state, with opposite spin to the paired electron in the ground state orbital. With both electrons having opposite spin, return to the ground state from the excited state is allowed and occurs with a rate of approximately 10 8 s-1. In contrast, phosphorescence occurs when the excited state electron has the same spin as the paired electron in the ground state orbital, thus preventing rapid photon emission and return to the ground state. Phosphorescence excited state lifetimes occur on the order of 10 3 to 1 s -1, many millions of times slower than fluorescence. It should be noticed that the emitted photon is generally of lower
38 energy, and thus higher wavelength, than the absorbed photon. This phenomon is called the Stokes shift, and results from thermal relaxation of the electrons within a single energy level without photon emission. Additionally, the electrons generally do not transition to the lowest thermal energy level of the ground state during emission.
Photon emission is usually independent of the excitation wavelength, a phonomenon known as the Kasha-Vavilov rule. This results from the rapid relaxation of electrons in excited states through non-radiative thermal relaxation within the excited state. Thus, all excited electrons transition to the ground state from identical engergy levels in the excited state. The emission spectra is often a mirror image of the absorption spectra. This phenomenon results from the similarity between the thermal energy levels within both excited and ground states, for example the spacing between the thermal energy levels E 1 and E 0 in
Figure 1-3. Photon absorption occurs from the lowest E 0 level to various thermal levels within E 1, and photon emission occurs from the lowest E 1 level to various thermal levels within E 0. Thus, as the thermal spacings are similar between E 1 and E 0, energy transitions of photon absorption and emission must be equally similar, resulting spectral symmetry.
Fluorescence Imaging in Cancer
Biomedical fluorescence imaging is concerned with wavelengths in the visible spectrum, from approximately 400-700 nm, and the NIR spectrum from approximately 700-900 nm. Commercially available fluorophores are available for
39 a large number of excitation and emission peaks. For research purposes, most are found in the visible spectrum. However, due to tissue autofluorescence and high absorbance of visible spectrum light, in vivo imaging is better suited to NIR fluorescence. Wavelengths below 700 nm are strongly absorbed by endogenous molecules such as hemoglobin and myoglobin, while water and lipids absorb strongly above 900 nm.173, 174 These phenomena create the “NIR window” from
700-900 nm in which photon absorbance is minimized and scattering is maximized, allowing for greater depth penetration.
There are currently two FDA approved NIR fluorophores: indocyanine green (ICG) (ex. 785 nm; em. 820 nm) and methylene blue (ex. 665 nm; em. 700 nm). ICG is of greater interest than methylene blue as the excitation and emission wavelengths show less tissue absorption and it has a greater quantum yield. 171, 175 Outside of image guided surgery, ICG is used clinically to measure cardiac output, hepatic function, and retinal angiography. 171 Within image guided surgery, ICG has been used for multiple oncological indications, including SLN mapping and identification of solid tumors. A number of additional contrast agents are currently being investigated, including several other cyanines such as
Cy5.5, Cy7, Cy7.5, IR-dyes (LI-COR, Lincoln, NE.), nanparticle formulations, and visible spectrum dyes. 176 Some examples are shown in Figure 1-4.
40
Figure 1-4. Examples of NIR fluorescing dyes include the FDA approved ICG
(ex. 785 nm, em. 820 nm) and methylene blue (ex. 665 nm, em. 700 nm), as well as investigational cyanines Cy5.5 (ex. 673 nm, em. 707 nm) and Cy7.5 (ex. 788 nm, em. 808 nm).
SLN mapping and biopsy is often used in an effort to detect metastatic disease, but proper identification of lymph nodes draining from the tumor, SLNs, must be achieved with high precision. Current methods utilize the visual dye
Patent blue or gamma probes, or a combination of both. 177 The use of ICG in
SLN mapping is recent but becoming well established. Use of ICG as a color dye has shown SLN detection rates of approximately 74%, while use of ICG fluorescence has resulted in detection of nearly 99% of SLN in breast cancer patients. 171 Additional work is required to identify the false negative rates in these patients, with a current best estimate of 7.7%. A number of clinical studies have
41 also shown success in SLN mapping with ICG in cervical, vulvar, and endometrial cancers. 170 Similar success rates of SLN detection occur for skin and gastric cancers. However, gastric and colorectal cancers may exhibit higher false negative rates when ICG is used, depending on tumor grade and location.
Studies in patients with breast, skin, gastro-intestinal, lung, cervical, esophageal, and other cancers have shown high rates of SLN detection. 171, 178-180 Studies that compare ICG fluorescence to the gold standard of radioisotope identification show that ICG fluorescence alone is comparable to radioisotope identification, and is just as effective when used in combination with radioisotope identification. 178, 180 While further work is needed, these initial studies suggest that IGC could provide equivalent, lower cost, and safer SLN detection than radioisotope administration.
Breast cancer imaging using ICG fluorescence has proven successful using diffuse optical tomography (DOT). Studies have shown good overlap in tumor localization between Gadolinium-enhanced MRI, palpation, and ICG, and
ICG may also be able to differentiate tissue pathologies.181, 182 DOT imaging using pharmacokinetic rate images over time may also allow for improved diagnoses. 183 Another study has shown it may be possible to utilize ICG in fluorescence mammograms. This study showed that in 21 patients fluorescence ratio mammograms gave sensitivity and specificity of 84-100% and 59-91% respectively. 184 In contrast, conventional mammography has 100% and 9-41% sensitivity and specificity. These results suggest that ICG fluorescence gives similar tissue differentiation potential to some conventional imaging modalities
42 and could serve to reduce false positive diagnoses. However, DOT and fluorescence mammography provide low contrast due to the depth of the imaging process, suggesting that fluorescence imaging may yield even better results when performed during surgical resection when the imaging depth is decreased.
Image guided surgery (IGS) using ICG to delineate tumors in humans has shown a number of preliminary successes, with most investigations examining the procedure in hepatic carcinomas. Gotoh et al. used intravenous injections of
0.5 mg/kg ICG between 1-8 days prior to surgery and found that ICG fluorescence successfully identified 100% of primary tumors and in 40% of the cases also identified additional hepatocellular carcinomas that went undetected either preoperatively or by intraoperative US. 185 These additional nodules were small, ranging from 3-6 mm in diameter. Another study examined the detection capacity of combined imaging with ICG fluorescence and intraoperative US with microbubbles compared to contrast enhanced CT/MRI, and found that the combined fluorescence-US method had higher sensitivity, accuracy, and positive predictive value than contrast enhanced CT/MRI. 186 Similar to these studies, IGS using ICG has also been shown to identify both hepatocellular carcinomas and colorectal liver metastases, and to differentiate the two by fluorescence pattern. 187 This study showed that ICG had 100% success in identifying lesions, some of which were not evident unless imaged by ICG, and hepatocellular carcinomas showed uniform fluorescence while colorectal liver metastases showed only rim fluorescence. The mechanism of ICG fluorescence localization within hepatocellular carcinomas has recently been investigated and appears to
43 result from a combination of features including cancer cell differentiation and expression of proteins that contribute to portal uptake. 188 ICG has also been used clinically to image hemangioblastomas, where it has been found to be useful in identifying vasculature feeding the tumors as well as hidden vessels. 189
The use of other fluorophores in image guided surgery for additional cancers has also met with clinical success. A phase III multicenter clinical trial investigating the use of fluorescence IGS for glioma patients showed significantly higher levels of complete tumor resection (90% vs 36%) and 6 month progression-free survival (41% vs 21%) when compared to conventional surgery conducted only with white light. 190 The use of 5-aminolevulinic acid in this study is interesting, as the molecule itself is not fluorescent but leads to accumulation of endogenous fluorescent porphyrins. This method requires lower wavelength
(violet-blue) excitation and thus both excitation and emmission exhibit higher absorption than if NIR fluorescence was used. In another study, folate-FITC was used as a targeted contrast agent for ovarian cancer, and resulted in significantly more tumors being identified than using white light alone. 122 Additionally, fluorescence intensity was qualitatively correlated with folate receptor-α expression and tumor grade, indicating successful active targeting.122 These studies both show the exceptional ability of fluorescence IGS to assist surgeons in diseased tissue identification. However, FITC (ex. 495 nm; em. 520 nm) and endogenous porphyrins both have absorptions and emissions outside of the optimal NIR window. The potential for delivery of NIR fluorophores to these solid tumors would likely improve the potential for complete resection even further.
44
Fluorescence Image Guided Surgery Instrumentation
A number of image guided surgery systems are currently available, in clinical trials, or under investigation.170, 191 These instruments are composed of an excitation source, generally either a laser or LED system, for excitation of fluorophores, and light collection system that preferably provides both white and
NIR image capture for overlay of the fluorescent signal onto the anatomical field, as shown in Figure 1-5. Additionally, the use of a handheld excitation source may also allow for spectral collection, providing a secondary means of signal detection.
Figure 1-5. Schematic diagram of image guided surgery with instrumentation showing channels for white (anatomical), laser, and NIR fluorescence. Image overlay allows for simultaneous viewing of fluorescent signal on the anatomical field for precise localization of cancer tissue. In this system, excitation is provided
45 by a handheld laser that also features fluorescent and Raman spectral collection, which may provide a further means of signal detection that is independent of the filtering parameters used in the camera system. Used with permission from Mohs et al. IEEE Biomedical Engineering , 2015.
Excitation light sources and wavelengths will depend on the contrast agents used, highlighting the interdependency of the contrast agents and instrumentation used. The light source must be powerful enough to give strong fluorescence signal in the surgical field, while maintaining a working distance for the surgery itself. 192 LED sources are capable of illuminating the entire field of view, but heat dissipation and adequate fluence are a concern. Laser excitation is another option, which allows for greater surgical control but requires personal protective goggles and there are concerns regarding patient exposure. 176 Field of view must be sufficient and adjustable for the surgeon to adequately capture the surgical field, and must be illuminated by a light source that does not overlap with the fluorescence signal. The collection optics and camera system are important, as signal loss by cameras due to low quantum efficiency and resolution loss will significantly affect the imaging potential. 176 Currently there is no gold standard in
NIR fluorescence imaging, and further studies are required to determine optimal system requirements and contrast agents.
46
5.0 Fatty Acid Synthase in Cancer
Biology of FASN
Lipid synthesis is a crucial part of cell growth and operation, as lipids are required for cellular membranes, modifying certain proteins, and energy storage.
Two sources are available for cells acquiring new lipids: circulating dietary lipids and de novo synthesis from precursor molecules. Most healthy cells utilize circulating lipids for structural lipids. However, nearly all cancers show increased de novo fatty acid synthesis, regardless of the level of circulating dietary lipids.193
The key protein responsible for this process is the homodimeric, multienzyme fatty acid synthase (FASN), which cyclically produces 16-18 carbon fatty acids from multiple catalytic sites, using malonyl-CoA, acetyl-CoA, and NADPH. 194
The primary product of FASN is the 16-carbon fatty acid palmitate. In healthy liver and other lipogenic tissues, palmitate is typically produced only to store excess energy in the form of triglycerides, and is regulated by dietary intake. 193, 195 However, the activity of FASN and fate of palmitate is different in cancerous tissue, where the predominant product is phospholipids, the primary class of lipids found in cell membranes. Regulation of FASN in these cells is typically independent of the level of circulating fatty acids. 193, 195 Evidence is accumulating that this activity of FASN confers a survival advantage to cancers and their precursor lesions, indicating its status as an oncogene and possibly a hallmark of cancer. 193, 196, 197 Increased expression of FASN has been shown in
47 every major cancer type and it has been recognized for some time as a potential target for therapy. 195, 197
Utilization of FASN in de novo fatty acid synthesis may be related to another hallmark of cancer, the Warburg effect. 193 This is the increased aerobic glycolysis, which many cancers show, that results in the conversion of glucose to lactate. 198 Both pyruvate and glucose-6-phosphate are upstream in the pathway of FASN catalyzed lipid synthesis, as shown in Figure 1-6. Pyruvate is converted to citrate by the citric acid cycle, and further to acetyl-CoA by ATP citrate lyase
(ACLY), then to malonyl-CoA by CoA carboxylase (ACACA). Both of these enzymes contribute to the lipogenic phenotype that ultimately results in FASN catalyzed production of lipids, which are eventually converted into membrane lipids for increased proliferation.
48
Figure 1-6. FASN primarily synthesizes palmitate from malonyl-CoA and acetyle-
CoA. Upstream contributors of FASN metabolites result from the increased uptake of glucose as a result of the Warburg effect. Citrate is converted to acetyl-
CoA by the enzyme ACLY and further to malonyl-CoA by ACACA, both of which may be upregulated in the lipogenic phenotype. Used with Macmillan Publishers
Ltd: Menendez et al., Fatty acid synthase and the lipogenic phenotype in cancer pathogenesis , Nature Reviews Cancer, 2007, 7(10).
FASN Inhibitors and ORL
A number of molecules that target FASN have been investigated, among which are Orlistat (ORL), cerulenin, C75, epigallocatechin gallate, and triclosan.197, 199-202 However, cerulenin is chemically unstable and C75 possesses side effects including severe weight loss and anorexia. 195, 201 Thus, new
49 molecules and formulations are needed that cause less collateral damage. As has been seen previously with other drugs with dose limiting toxicity, a NP formulation may help to reduce toxicity and improve delivery to tumors.
ORL, shown in Figure 1-7, will be the drug of focus for CHAPTER 5 of this body of work. ORL is an FDA approved weight loss drug that, when taken orally, binds specifically to gastric and pancreatic lipases. 203 In addition to the lipase activity however, ORL has been shown to be a potent inhibitor of FASN, where
ORL binds to the thioesterase within the enzyme, leading to the hydrolysis of the
β-lactone group within ORL and formation of a covalent bond between FASN and
ORL.204-206 In cellular assays ORL has been shown to effectively inhibit FASN by inhibiting 14 C-acetate uptake into fatty acids, reducing cancer cell proliferation as well as endothelial cell proliferation and angiogenesis. 206-208 In addition to FASN inhibition, there is also evidence supporting additional effects such as cell cycle disruption resulting in apoptosis. 209, 210 In vivo models have shown reductions in tumor growth, metastasis, and angiogenesis using a number of tumor models. 211-
214 However, due to the high hydrophobicity of ORL, all in vivo models to date have required administration either orally or via intraperitoneal injection. The structure of ORL shows a very hydrophobic molecule, and indeed its log(P) is predicted to be 7.3.215 Oral administration of ORL shows very low absorption, with enteric elimination accounting for approximately 96% of the total dose and urinary recovery accounting for 1.1%. 203, 216 Additionally, ORL is degraded after absorption resulting in two major metabolites that account for 42% of systemic
ORL content. 216 Due to this high hydrophobicity, low absorption, and poor
50 metabolic stability, a means of stabilizing and solubilizing ORL is needed to improve its biodistribution and delivery to solid tumors.
Figure 1-7. Orlistat (MW: 495.74) is FDA approved for oral administration as a weight loss drug that acts by inhibition of gastric lipases. ORL is also an inhibitor of FASN, and has been shown to reduce tumor growth when administered by intraperitoneal injection in mice. Improving solubility and metabolic stability may allow for intravenous injection and improved delivery for use as a cancer- targeting chemotherapeutic.
6.0 Conclusions
The body of research into nanoscale materials for biomedical application clearly indicates that there is a current and future place for this technology in medicine. Drug and contrast agent delivery using NPs has already found a place in clinical medicine, and the share of these pharmaceuticals in the market is likely to increase with the current formulations in pre-clinical and clinical investigations.
51
Among the many possible uses, drug and fluorescent contrast agent delivery to solid tumors are of high importance. It has already been seen that NP formulations can improve the total quantity of drug delivered to tumors, and perhaps equally important, they can also reduce associated toxicities and thereby lessen the already high burden placed on patients dealing with chemotherapy. Delivery of fluorescent contrast agents is a new but urgently needed area. Current surgical practices do not adequately identify the extent of solid tumors, local metastases, and sentinal lymph nodes, resulting in disease recurrence. It is therefore crucial that new technologies are developed and moved into the clinic that assist surgeons in the identification of these tissues.
In the following chapters, we will examine research into nanoscale polymeric micelles derived from the natural glycosaminoglycan HLA, and the application of these NPs to the delivery of indocyanine green for image guided surgery, and ORL for chemotherapeutic delivery. We first outline the chemistry behind the synthesis of these NPs, followed by the development and application of ICG-loaded NPs both in vitro and in vivo . We then turn our attention to the loading of the FASN inhibitor ORL into these HLA NPs to create a new therapeutic formulation and examine in detail the in vitro efficacy against a number of cancer cell lines.
52
CHAPTER 2
SYNTHESIS OF NOVEL HYALURONIC ACID DERIVED NANOPARTICLES
Text and the following figures: Scheme 2-1, Figure 2-1, Figure 2-2, and data in
Table-2-1, have been reproduced with permission from Hill, T.K. et al.,
Indocyanine Green-Loaded Nanoparticles for Image-Guided Tumor Surgery .
Bioconjugate Chemistry 26(2), 2015. AFM analysis was performed with the assistance of Dr. Adam Hall and Osama Zahid.
1.0 Introduction
In order to develop self-assembled NPs that use HLA as a hydrophilic backbone material it is necessary to conjugate hydrophobic ligands directly to
HLA in order to drive self-assembly. Hydrophobic moieties bearing carboxylic acid groups (1-pyrenebutyric acid, 5β-cholanic acid) were used as starting products for synthesis of hydrophobic ligands. Addition of 1,3-diaminopropane to these was performed in order to produce a product bearing a primary amine for conjugation to HLA, resulting in the products aminopropyl-1-pyrenebutanamide
(PBA), and aminopropyl-5β-cholanamide (5βCA). Octadecylamine (ODA) was also used to gain insight into the behavior of conjugates with varying hydrophobic ligand structures. Synthesized products were analyzed by mass spectrometry and NMR. Hyaluronic acid polymers were conjugated to one of these three different hydrophobic ligands at different ratios in order to drive self-assembly
53 into NPs. Conjugates were then analyzed by NMR to confirm conjugation, and subsequently by dynamic light scattering (DLS) and AFM to confirm presence and size distributions of NPs.
2.0 Materials and Methods
Materials
1-Pyrenebutyric acid, octadecylamine, 5-beta cholanic acid, 1,3- diaminopropane, N-hydroxy succinimide, and 1-ethyl-3-(3- dimethylaminopropyl)carbodiimide (EDC), ICG, and fetal bovine serum (FBS) were obtained from Sigma-Aldrich (St. Louis, MO). Unless otherwise noted, all water was obtained from a Barnstead NANOpure Diamond (Thermo Scientific;
Waltham, MA) system producing 18.2 MΩ water. Sodium hyaluronate was purchased from Lifecore Biomedical (Chaska MN). Methanol, and N, N dimethylformamide (DMF) were purchased from Fisher Scientific (Pittsburgh,
PA). Ethanol was purchased from the Warner-Graham Company (Cockeysville,
MD). Matrigel was purchased from BD Biosciences (San Jose, CA).
Penicillin/streptomycin and other media reagents were purchased from ATCC
(Manassas, VA).
Aminopropyl-1-pyrenebutanamide Synthesis
1-Pyrenebutyric acid (1.93 mmol, 558 mg) was dissolved in 5.5 ml MeOH with 3% concentrated HCl and refluxed for six hours at 60-65 °C in a 25 ml round
54 bottom flask (RBF). The reaction solution produced two layers; the top layer was clear and light yellow, while the bottom layer was dark and oil-like. The top layer was removed and the product, 1-pyrenebutyric methyl ester, was confirmed in the bottom layer and dried under vacuum. 1-pyrenebutyric methyl ester (0.881 mmol, 266.4 mg) was then dissolved into 6 ml 1,3-diaminopropane and refluxed at 130 °C for six hours to produce a clear, brown liquid. This solution was then cooled to 0 °C and PBA was precipitated by cold water, washed in cold water, and dried under vacuum. This process produced 0.52 mmol (180 mg, 59% yield).
A schematic of all the synthesis and conjugation used is shown in Scheme 2-1.
Scheme 2-1. (A) Synthesis of hydrophobic amides using 1-pyrenebutyric acid 5-
β-cholanic acid involves addition of MeOH followed by substitution with 1,3
55 diaminopropane. (B) Conjugation of hydrophobic moieties, PBA, 5βCA, or ODA, to HLA via EDC/NHS reaction.
Aminopropyl-5β-Cholanamide Synthesis
5-beta cholanic methyl ester was prepared according to the methods described in the literature. 217 Briefly, in a 25 ml RBF, 5-beta cholanic acid (2.8 mmol, 1.0 mg) was dissolved in 5 ml methanol. To the RBF, 180 µl of concentrated HCl was then added under stirring and this solution was allowed to reflux at 60-65 °C for 6 hours. The product was cooled to 0 °C and appeared as a white precipitate. The precipitate was collected under vacuum filtration, washed with cold methanol, and then dried under vacuum. 0.80 mg of 5-beta cholanic methyl ester was recovered (2.1 mmol, 75% yield). 5βCA was synthesized by dissolving 5-beta cholanic methyl ester (1.36 mmol, .51 mg) in 132 mmol (~11 ml) of 1,3-diaminopropane. This solution was refluxed at 130 °C for 6 hours and then allowed to cool to 0 °C. 5βCA was obtained from crystallization with water.
The white precipitate was then washed with 200 ml cold nanopure water under vacuum filtration and dried under vacuum to produce 0.952 mmol (.397 mg, 70% yield).
Conjugation of 5βCA or PBA to HLA
Sodium hyaluronate (90-95 mg, M N = 10-20 kDa), was dissolved in 25 ml of water. PBA at 5 or 10 wt% (14 or 28 mmol, 5 or 10 mg), or 5βCA at 5 or 10 wt% (12 or 24 mmol, 5 or 10 mg), was dissolved into 25 ml DMF under stirring at
56
30-40 °C. NHS and EDC, 78-154 mmol (10X molar ratio to PBA or 5βCA), were then dissolved into the HLA solution. The PBA or 5βCA DMF solution was then added dropwise to the activated HLA solution under constant stirring. The reaction was allowed to stir until completion at room temperature, 24-36 h. The reaction contents were then transferred to dialysis tubing (material: MWCO =
3500 Da, Spectrum Labs) and allowed to dialyze against 1:1 EtOH:H 2O for 24 h followed by H 2O for an additional 48 h. The polymer-hydrophobic moiety conjugate was removed from the dialysis tubing and lyophilized for storage at -20
ºC. This process produced High-PBA-HLA (10 wt% PBA), Low-PBA-HLA (5 wt%
PBA), High-5βCA-HLA (10 wt% 5βCA), and HLow-5βCA-HLA (5 wt% 5βCA).
Conjugation of ODA to HLA
ODA (9.3 mmol, 2.5 mg) was dissolved into a 70% EtOH:H 2O solution.
HLA, 97.5 mg, was dissolved into a separate solution of 70% EtOH:H 2O and 10- fold molar excess of EDC and NHS were then added. The ODA solution was then added slowly into the HLA solution under vigorous stirring and the reaction was allowed to proceed for 24-36 h at room temperature. The reaction mixture was then dialyzed (material, MWCO = 3500 Da, Spectrum Labs) against 1:1
EtOH: H 2O for 24 hours and water alone for 48 h. This material, ODA-HLA
(2.5wt% ODA), was then lyophilized and stored at -20 ºC.
57
Hydrophobic Conjugate Characterization
The PBA and 5βCA intermediate methyl ester products were analyzed by
GC-MS with a Thermo Scientific Trace Ultra gas chromatograph interfaced with a
Thermo Scientific TSQ Quantum XLS (Thermo Scientific; Waltham, MA) using splitless injection. Separation was performed on a 10.4 m DB-1 WCOT column
(Agilent Technologies Inc.; Santa Clara, CA). PBA and 5βCA were analyzed on a
Waters Q-Tof API-US mass spectrometer with Advion TriVerso NanoMate
(Advion Biosystems; Ithaca, NY) using positive ion electrospray. 5βCA-HLA,
PBA-HLA, and ODA-HLA conjugates were analyzed on a Bruker Avance 600
MHz NMR spectrometer with TXI Cryoprobe at 25º C, with 64 to 128 scans, 8192 to 16384 data points, and 10-12 second relaxation delay. 5βCA-HLA and PBA-
HLA were analyzed in 50/50 D 2O/DMSO-D6 (D 2O from Acros Organics “100.0%”
D, Fisher Scientific, DMSO-D6 from Cambridge Isotope Labs, 99.9% D). Data was processed in Mnova NMR (Escondido, CA).
Confirmation of NP Formation
Crosslinking of NPs with diaminopropane was performed here strictly to form a material that did not disassemble upon drying, thus making analysis by
AFM possible. NPs were not crosslinked in any other part of this experiment work. High-PBA-HLA NPs were exposed to 0.30 mol fraction diaminopropane, relative to the repeat unit of HLA, in presence of 10-fold molar excess of
EDC/NHS, and this solution was incubated at room temperature overnight in order to crosslink the self-assembled NPs, and thereby prevent disassembly. The
58 hydrodynamic diameter of this material was then analyzed by DLS using a
ZetaPlus system with an onboard dynamic light scattering (DLS) analyzer
(Brookhaven Instruments Corporation; Holtsville, NY). For AFM analysis, NPs were dispersed into water at 1 mg/ml. 20 µl of this solution was then placed on the center of a mica plate which had previously been cleaned by repeated application of standard scotch tape. Solution was then dried under light air flow and examined using MFP3D AFM (Asylum Research, Santa Barbara, CA.) operated in tapping mode .
3.0 Results and Discussion
Hydrophobic Moiety Synthesis and Conjugation to HLA
Three structurally distinct hydrophobic moieties were examined to determine their effect on driving self-assembly and ICG loading. 5β-cholanic acid and ODA have previously been reported to successfully drive self-assembly when conjugated to HLA, 50, 51, 218 while PBA has not been previously conjugated to HLA. Scheme 2-1 shows the reaction. 1-Pyrenebutyric acid and 5β-cholanic acid were refluxed in methanol to give 1 and 2, respectively. Next, methyl esters,
1 and 2, were converted to amide products, PBA and 5βCA, respectively, by refluxing 1,3-diaminopropane as shown in Scheme 2-1A. PBA and 5βCA each possess a primary amine moiety for conjugation to HLA. Since PBA is previously unreported, mass spectra and 1H-NMR of this product is shown in Figure 2-
1A&B. The structural characterization of 5βCA is shown in Figure 2-2A . ODA
59 was used as provided for conjugation to HLA. Conjugation of the primary amine group on PBA, 5βCA, or ODA to the carboxylic acid group of the glucuronic acid moiety was performed using 1-ethyl-3-[3-dimethylaminopropyl]carbodiimide hydrochloride (EDC) and N-hydroxysuccinimide (NHS) chemistry to facilitate amide bond formation ( Scheme 2-1B ).
60
Figure 2-1. Confirmation of synthesis of novel compounds. Mass spectrometry
(A) and 1H-NMR (B) confirmed the synthesis of PBA. (C) Quantification of conjugation of PBA to hyaluronic acid was achieved by calculating the ratio of
61 aromatic pyrene protons relative to the N-acetyl protons (a) on N-acetyl-D- glucosamine of hyaluronic acid.
62
Figure 2-2. Confirmation of previously synthesized compounds. (A) MS of 5βCA.
1H-NMR of (B) HLA-5βCA and (C) HLA-ODA.
Conjugation of the hydrophobic groups to HLA was performed using loading ratios of 5 wt% (Low) and 10 wt% (High) for PBA-HLA and 5βCA-HLA, and 2.5wt% for ODA-HLA. Conjugation ratios with greater than 2.5 wt% ODA had poor solubility and did not form NPs efficiently, while reducing the wt% of ODA lowers the ligand-to-polymer ratio below 1:1. Successful conjugation was observed for all products in the NMR spectra ( Figure 2.1C, Figure 2.2B-C). This synthesis produced amphiphilic HLA polymers with varying degrees of substitution of hydrophobic moieties. Conjugation levels were quantified by NMR and results are shown in Table 2.1.
Sample Theoretical Actual hydrophobic hydrophobic moiety moiety conjugation conjugation mol% a mol% b
Low-PBA-HLA 5.8 5.67
High-PBA-HLA 12.2 10.1
Low-5βCA-HLA 4.8 3.88
High-5βCA-HLA 10.1 6.0
ODA-HLA 2.8 2.77
Table 2-1. Conjugation levels of hydrophobic conjugates to HLA. A) Molar loading ratio, b) Calculated molar loading ratio determined by NMR integration
63
In order to determine the diameter of these NPs, crosslinking of High-
PBA-HLA NPs was performed in order to maintain their structure upon drying.
AFM and DLS was then performed to determine the presence of NPs as shown in Figure 2-3. It should be noted here that the crosslinking performed in Figure
2-3 was strictly used to determine NP diameter, and no other NPs presented in this work were crosslinked. NPs at the scale predicted were clearly present in both AFM and DLS analysis, and DLS analysis gave hydrodynamic diameters consistent with those observed for other NP types, as will be discussed in
CHAPTER 3. NP diameters as measured by AFM using the cross-sectional length were consistent with expected NP diameters from DLS. However, using height analysis of the AFM data shows heights much smaller than expected, ranging from 10-20 µm. This suggests that the NPs had undergone a significant degree of dehydration that affected their total volume, and thus the diameters obtained by AFM are not expected to conform with those observed in solution, but rather this data should be interpreted only to confirm the existence of NPs formed with these materials.
64
Figure 2-3. AFM of crosslinked High-PBA-HLA NPs. NPs were crosslinked in order to prevent disassembly upon drying, which had been previously observed.
Presence of NPs were confirmed by AFM and similar DLS distributions were obtained for both crosslinked and non-crosslinked formulations, with effective diameters of approximately 270 nm.
4.0 Conclusions
In this chapter we have examined the synthesis and conjugation of hydrophobic conjugates to HLA. PBA is a novel conjugate not previously examined and possesses several properties that have proved useful for future studies with these materials. These properties are the aromatic, planar nature of part of the molecule, which may allow for pi-pi stacking with other aromatic molecules such as fluorophores, the higher wavelength absorbances from the aromatic region that can allow for optical quantification of material, and more
65 rapid solubility of material compared to 5βCA, which may provide more complete conjugation to HLA. Other conjugates that have previously been examined were also used, including 5βCA which was synthesized in house, and ODA which was used as ordered. The conjugation of these materials to HLA was confirmed using
NMR, formation of NPs was confirmed using DLS and AFM, and further confirmation of NP formation will be presented in CHAPTER 3. These materials will form the basis of future investigations into NPs for image guided surgery and drug delivery that will be examined in the following chapters.
66
CHAPTER 3
IN VITRO CHARACTERIZATION OF INDOCYANINE GREEN-LOADED
NANOPARTICLES
Text and the following figures: Figure 3-1, data in Table 3-1, Figure 3-2, Figure 3-
3, Figure 3-4, Figure 3-5, Figure 3-6, and Figure 3-10, have been reproduced with permission from Hill, T.K. et al., Indocyanine Green-Loaded Nanoparticles for Image-Guided Tumor Surgery . Bioconjugate Chemistry 26(2), 2015.
1.0 Introduction
Detecting positive tumor margins and local malignant masses during surgery is critical for long-term patient survival. The use of image-guided surgery for tumor removal, particularly with near infrared fluorescent imaging, is a potential method to facilitate removing all neoplastic tissue at the surgical site. In this study we demonstrate a series of hyaluronic acid (HLA)-derived nanoparticles that entrap the near infrared dye ICG, termed NanoICG, for improved delivery of the dye to tumors. Self-assembly of the nanoparticles was driven by conjugation of one of three hydrophobic moieties: PBA, 5βCA, or ODA.
Nanoparticle self-assembly, colloidal stability, dye loading, optical properties, ICG release, cellular uptake, and cytotoxicity were characterized. ICG was found to load into all NP formulations and was influenced by hydrophobic ligand structure.
Fluorescence was quenched in aqueous solution but could be reactivated by
67 addition of an equal volume of DMSO or 37 mg/ml bovine serum albumin (BSA).
NPs were observed using DLS both with and without the presence of fetal bovine serum, and ICG appeared to release from NanoICG when serum proteins were present. Incubation of NanoICG with MDA-MB-231 cells resulted in cell- associated fluorescence, but this appeared to be independent of whether the ICG was localized in NPs. Cytotoxicity was found to be negligible at physiologically relevant concentrations.
2.0 Materials and Methods
ICG Loading into Nanoparticles
Amphiphilic HLA polymer conjugate, 8-16 mg, was dissolved in 5 ml of water, while ICG (.0026-.0052 mmol, 2.0-4.0 mg) was dissolved in 5 ml DMSO and added to HLA solution. This solution was vortexed and then dialyzed against
H2O for 24-36 h to remove DMSO and drive self-assembly of ICG-entrapped
NPs. Resulting NPs were then filtered through a PD-10 desalting column (GE
Lifesciences; Pittsburgh, PA) to remove free ICG. The clear green solution was then lyophilized and frozen at -20 °C until resuspension of NPs. Materials produced by this process are termed High-PBA-NanoICG, Low-PBA-NanoICG,
High-5βCA-NanoICG, Low-5βCA-NanoICG, and ODA-NanoICG.
68
Nanoparticle Characterization
NanoICG or empty NPs (empty NPs underwent same dialysis treatment, but without ICG) were diluted to a concentration of 0.1 mg/ml in PBS and filtered through a 0.45 um filter (Fisher Scientific; Pittsburgh,PA). The hydrodynamic diameter (HD) of the NP-containing solutions was determined on ZetaPlus system with an onboard dynamic light scattering (DLS) analyzer (Brookhaven
Instruments Corporation; Holtsville, NY). Absorbance (extinction) spectra were obtained on a UV-2600 (Shimadzu Scientific Instruments; Columbia MD).
Fluorescence spectra were obtained on a FluoroMax-4 fluorescence spectrometer equipped with a NIR extended range PMT (Horiba Jobin Yvon;
Edison, NJ). Spectra were obtained on aqueous solutions containing NPs and disassembled NP contents by addition of an equal volume of DMSO. ICG-loading content was determined with absorbance spectroscopy with a standard curve of
ICG in 1:1 H 2O/DMSO.
NanoICG Serum Stability
Solutions of 3% v/v FBS, 3% FBS + 0.35 mg/ml High-PBA-HLA NanoICG, or 0.35 mg/ml High-PBA-HLA NanoICG were made in PBS. Solutions were incubated at room temperature for 10 minutes prior to start of analysis. Dynamic light scattering using a Zeta-Plus instrument was performed as described previously to determine if hydrodynamic diameter of NanoICG or serum proteins changed in a mixed solution.
69
Photostability
ICG and NanoICG were dissolved separately in pure water and aliquoted into wells of an opaque 96 well plate (n = 3 each). ICG concentration of free or
NanoICG was approximately 0.0001 mg/ml (0.13 µM). Using the portable fluorescence spectrometer with a fiber-coupled handheld (SciApps; Laramie,
WY), 785 nm, 300 mW laser light was directed into the wells for three iterations of 10 seconds each. The fluorescence spectrum was taken after 0, 10, 20, and
30 seconds of cumulative exposure. The spectra were integrated from 797-977 nm to acquire the total signal present. Results were then averaged and normalized to the initial pre-exposure spectra.
NanoICG Interaction with Serum Proteins
BSA was dissolved in PBS to 37 mg/ml, and NanoICG was dissolved into the same solution at 1 mg/ml and either incubated at room temperature for 0 or 7 hours, or exposed to flow conditions. Exposure to flow conditions utilized a circularly connected piece of 1/8 X 3/16 Tygon 3350 silicon tubing pumped by an
Ismatec Reglo peristaltic pump for continuous flow. Solutions were mixed and then added immediately to the pump system, and flowed at rates of 0.1, 0.5, and
0.9 ml/s for one hour before stopping and immediate measurement of absorbance from 500-900 nm.
Elution of ICG with serum proteins was performed by mixing either
NanoICG or free ICG with 37 mg/ml BSA in PBS, followed by incubation at room temperature for either 0, 4, 8, or 24 hours. At the appropriate timepoint, solutions
70 were taken and eluted through 100 kDa MWCO Amicon Ultra 2ml centrifugal filter units at 3184 RCF for 30 min. Eluates were collected and examined using absorbance spectroscopy from 500-900 nm. Original solutions were also quantified for ICG and BSA content. Analysis was then performed by taking the relative fractions of ICG and BSA that eluted through the filter, as shown in
Equation 3.1, followed by simple linear regression.
[3.1] = /
In Vitro Cellular Uptake of NanoICG
500,000 MDA-MB-231 cells were seeded into each well of 6-well tissue culture dishes and allowed to adhere overnight. Cells were split into 5 groups: (1) unstained cells that received no treatment, (2) NanoICG-treated cells that received 1 hour pre-incubation with 10 mg/ml free 10 kDa HLA as a blocker prior to application of High-PBA-HLA NanoICG at 0.5 mg/ml (blocked), (3) NanoICG- treated cells that were not exposed to free HLA (no blocker), (4) free ICG-treated cells that received 1 hour pre-incubation with 10 mg/ml free 10 kDa HLA as a blocker prior to application of 0.1 mg/ml free ICG, (5) free ICG-treated cells that were not exposed to free HLA (no blocker). Free HLA was maintained in groups
(2) and (4) at 10 mg/ml over the course of the 1.5 hour incubation with either
NanoICG or free ICG. HLA pre-incubation and incubation with either NanoICG or free ICG were performed in serum free media. After 1.5 hours of incubation with either NanoICG or free ICG, cells were trypsinized from the plates and centrifuged twice in PBS to wash. Groups were then analyzed on a FACSAria
71 flow cytometer using a 760/80 nm bandpass filter with cell counts of
20,000/group.
Cytotoxicity
Cytotoxicity was evaluated by CCK-8 assay (Dojido, Japan). 2,000 MS1 mouse endothelial cells were seeded into 96 well plates and allowed to adhere for 24 hours in EMEM (Gibco), 10% FBS (Sigma Aldrich), 1% P/S (Gibco). Cells were then mixed with normal media containing High-PBA-HLA, High-PBA-
NanoICG, Low-PBA-HLA, Low-PBA-NanoICG, 5βCA-HLA, 5βCA-NanoICG,
ODA-HLA, or ODA-NanoICG at concentrations of 50 and 5 µg/ml. Cells were incubated for 24 hours followed by analysis with CCK-8 assay. Proliferation of
MDA-MB-231 cells was analyzed using a FluoReporter Blue DNA quantification assay (Life Technologies; Grand Island, NY). 2,000 MDA-MB-231 cells were seeded into wells of tissue treated 96 well plates. The cells were exposed to
High-PBA-HLA or High-PBA-NanoICG at 5 µg/ml, for 24 hours. After the 24 hour exposure standard media was applied. DNA was quantified at 0, 24, 48, 72, and
96 hours.
Statistical Analysis
Analyses of average NP size, ICG loading, photoprotective effect, and cytotoxicity were performed with one-way ANOVA. Analysis of ICG-BSA elution through 100 kDa filter membranes was analyzed with linear regression. Analysis
72 of cellular uptake of ICG was performed in FlowJo software, while all other analyses were performed using GraphPad Prism and Excel.
3.0 Results and Discussion
Such amphiphilic conjugates were necessary to drive self-assembly and provide a hydrophobic domain for entrapment of ICG, as schematically depicted in Figure 3-1A. Complete characterization of NanoICG formulations are shown in
Table 3-1, along with hydrophobic ligand conjugations as discussed in
CHAPTER 2 for reference. It was hypothesized that PBA could provide increased loading efficiency for ICG compared to 5βCA or ODA, because of the potential for π-π stacking between PBA and ICG. 219 ODA is capable of efficient self- association, but is not structurally similar to ICG. Each wt% of PBA and 5βCA conjugates were found to load more ICG than ODA-HLA; ODA-HLA loading was
13-35% less compared to the other hydrophobic moieties. ICG loading into PBA-
HLA NPs was equal or higher than 5βCA-HLA. Loading efficiency (33-50%) and
ICG content (6-10 wt%) of NanoICG was moderate compared to previous ICG- loaded NP formulations, 220, 221 which range from very low loading efficiency (1-
10%) and content (0.16-0.21 wt%), to moderate, 222, 223 and relatively high efficiencies of 34-97% and content up to 23%. 224 NanoICG with maximum ICG content was used in order to minimize the total mass of nanoparticle required for injection.
73
Figure 3-1. (A) Hydrophobic moieties conjugated to HLA drive self-assembly into nanoparticles that can entrap ICG. (B)
NP diameters range from 80-260 nm, with smaller diameters observed in NanoICG formulations. (C) NanoICG formulations increase ICG solubility in aqueous salt solution by providing a favorable electrostatic environment, while ICG alone or not entrapped in NPs has poor solubility and precipitates rapidly.
74
Sample Theoretical Actual ICG loading ICG wt% Fluorescence hydrophobic hydrophobic efficiency c activation d moiety moiety conjugation conjugation mol% a mol% b Low-PBA- 5.8 5.67 0.501 ± .004 10.0 ± .1% 31.7 ± 8.8 HLA High-PBA- 12.2 10.1 0.470 ± .004 9.4 ± .1% 28.2 ± 0.56 HLA Low-5βCA- 4.8 3.88 0.370 ± .003 7.4 ± .1% 22.3 ± 3.7 HLA High-5βCA- 10.1 6.0 0.475 ± .004 9.5 ± .1% 33.0 ± 1.2 HLA ODA-HLA 2.8 2.77 0.327 ± .003 6.5 ± .1% 21.2 ± 3.1 Table 3-1. Characterization and ICG loading of HLA conjugates. a) Molar loading ratio, b) Calculated molar loading ratio determined by NMR integration, c) ICG loading efficiency = (calculated ICG concentration measured by absorbance spectroscopy)/(theoretical concentration based on loading amount), d) Fluorescence activation = [integrated fluorescence intensity (790 – 950 nm) of NanoICG in 1:1 DMSO:H 2O]/[integrated fluorescence intensity (790 – 950 nm) of NanoICG in
H2O]
75
Physical, chemical, and optical characterization
Hydrodynamic diameters of empty NPs (i.e., those not loaded with ICG) ranged from 150 to 260 nm, whereas NanoICG formulations ranged from 80 to
150 nm depending on the hydrophobic group ( Figure 3-1B). Representative histograms of DLS analyses are shown in Figure 3-2. NPs were observed at all concentrations within the detection limits of the instruments used in this study.
NanoICG formulations increased the solubility of ICG in PBS, as was observed by an optically transparent green color throughout the PBS solution; centrifugation could not remove ICG from solution. The same quantity of ICG alone or ICG mixed with previously assembled empty NPs in PBS was insoluble and could be pelleted by centrifugation ( Figure 3-1C). The size of NanoICG remained consistent in FBS ( Figure 3-3).
76
Figure 3-2. Representative intensity-weighted (left) and number-weighted (right) dynamic light scatter results from each HLA-hydrophobic moiety conjugate. Black bars are nanoparticles that do not have ICG (empty nanoparticles); green bars represent nanoparticles that have ICG entrapped. 77
Figure 3-3. DLS histograms of NanoICG (top), FBS (bottom), and NanoICG mixed with FBS. NanoICG and FBS has distinct distributions centered at approximately 120 nm (hydrodynamic diameter) and 14 nm, respectively. Mixing
NanoICG in FBS resulted in a biomodal distribution with peaks occurring at 15 nm and 100 nm, representative of independent FBS and NanoICG populations and a stable nanoparticle diameter in serum.
The absorbance and fluorescence spectra of ICG in DMSO showed peaks near 800 and 825 nm, respectively ( Figure 3-4A). The extinction spectra of
NanoICG showed strong scattering in pure water and PBS, likely due to aggregation and electronic effects of ICG, 219 suggesting the ICG was self- associated within the NPs ( Figure 3-4B). NanoICG fluorescence was quenched in aqueous solution ( Figure 3-4B). NanoICG disassembly was induced by the addition of DMSO. The resulting absorbance and fluorescent spectra of ICG
78 closely resembled its characteristic shape ( Figure 3-4C), suggesting dye release. The fluorescence signal after DMSO addition over the quenched state for each NanoICG formulation is shown in Table 1. All NPs investigated showed quenching and activation ( Figure 3-5). Change in the absorption peak of ICG, indicative of scattering by NPs is consistent with previous reports. 221, 224
However, the fluorescence quenching observed in NanoICG ( Figure 3-5) is observed in only some formulations of encapsulated ICG previously reported in the literature,221, 225 but not others. 226 Nanoparticle entrapment of ICG did provide a modest, yet significant increase in the photostability of ICG compared to free ICG ( Figure 3-6). The photoprotective effect is consistent with dyes closely associated in a matrix. 227
Figure 3-4. (A) ICG dissolved in DMSO shows clear extinction and fluorescence peaks near 800 and 825 nm respectively. (B) Nanoparticle formulations of ICG show broadened absorbance spectra with a less discernable peak due to scattering by NP-entrapped ICG, and ICG fluorescence is quenched. (C) NPs disassemble in 50:50 DMSO:H 2O, decreasing the scattering in ICG absorbance
79 and fluorescence activation, which demonstrates the potential for optical property control.
Figure 3-5. Absorbance (left) and fluorescence emission (right) spectra of each nanoparticle when dissolved in either water (blue) or 1:1 H 2O:DMSO.
80
Figure 3-6. Exposure of ICG (black circle) and High-PBA-HLA NanoICG (red square) to a focused laser beam (785 nm, 300 mW) for 30 s. A modest, but significant increase in photostability is observed with NanoICG compared to free
ICG. It should be noted that this laser power is 10 to 40 fold higher than what is expected during intraoperative imaging, where the excitation is at a lower power and not directly at the focal point of the laser.
Interestingly, when ICG was entrapped, the NP diameter decreased, regardless of the hydrophobic conjugate. These data suggests that the hydrophobic groups of amphiphilic HLA facilitates close association of ICG.
Close packing of ICG in NPs is further supported by the optical and colloidal properties of NanoICG. First, the absorbance peaks of ICG, when loaded into
NPs, showed a decreased signal at ~795 nm and an increased scattering contribution to extinction spectra ( Figure 3-4B) when compared to ICG in DMSO
(Figure 3-4A). These finding are consistent with previously reported results. 218,
81
228 Second, fluorescence quenching was also observed in NanoICG formulations.
Finally, ICG becomes stable in high salt conditions when associated with amphiphilic HLA conjugates. Taken together, the findings of the broad absorbance of NanoICG, self-quenched fluorescence, and increased ICG salt stability strongly suggest that ICG is being closely associated within NPs via sequestration in hydrophobic domains. It is important to note that precise control of quenching and activation may also prove useful for image-guided surgery if high levels of fluorescence activation can be achieved specifically in the tumor environment. 224, 229
NanoICG Interaction with Serum Proteins
NanoICG was next analyzed for its stability in serum-like conditions. BSA was dissolved to 37 mg/ml in PBS and High-PBA-HLA was dissolved in this solution to a concentration of 1 mg/ml. Solutions were then exposed to a series of flow conditions to simulate shear stress that would be experienced in vivo .
Absorbances from 500-900 nm were taken in order to observe the shape and presence or absence of scattering around the ICG peak, and thereby gain insight into the physical state of ICG. The data in Figure 3-7 shows the standard
NanoICG peak in PBS (black), with a decreased abs max and scattering indicative of ICG present in NPs. Upon dissolution of NanoICG into BSA-PBS solution
(red), there is an immediate shift in the shape of the ICG spectrum, showing both narrowing and increasing height of abs max . This trend intensifies with increasing incubation time without flow (brown, 7 hours at room temperature), and with
82 increasing volumetric flow rate with corresponding increases in WSS (green, purple, blue). Quantitative analysis of this is shown in Table 3-2. Absorbance peaks of various NanoICG formulations were compared when dissolved in either
DMSO-H2O or PBS-BSA solutions after flow, and results indicate that the absorbance of ICG changes in an almost identical fashion when NanoICG is dissolved in DMSO-H2O or PBS-BSA. This changing ICG absorbance spectrum is indicative of a physical change in the proximity of ICG molecules to one another. As ICG can bind reversibly to the hydrophobic pockets of serum albumin, this trend likely results from changes in the physical location of ICG, moving from a closely associated state within the NP core to a BSA-bound state.
Figure 3-7. Analysis of ICG absorbance change from NanoICG in flow conditions with 37 mg/ml BSA. The increase in height and narrowness indicate a change in the physical state of ICG, likely due to the release of ICG from the NPs to BSA. 83
NP Type Ratio, NP in Ratio, NP in Ratio, Column 3 DMSO/H2O to PBS/BSA to PBS to Column 2 H2O (Ratio = 1 is perfect match) High-5βCA 2.014433 1.984559 0.98517
Low-PBA 2.110204 2.218076 1.051119
High-PBA 1.823256 1.946606 1.067654
Table 3-2. Analysis of relative absorbance peak increases of NanoICG in either
DMSO-H2O or PBS-BSA solutions. Column 2 shows the ratio of peak heights between NanoICG dissolved in DMSO-H2O or pure H 2O. Column 3 shows the ratio of peak heights of NanoICG dissolved in 37 mg/ml BSA in PBS, or PBS without BSA. Column 3 shows the ratio of Column 2 to Column 3, where a value of 1 indicates an equal change in absorbance between NanoICG in DMSO-H2O or PBS-BSA. Similar optical features are observed with NanoICG is dissolved in
PBS containing physiological concentrations of BSA, indicating ICG is being released to BSA.
In order to further test if ICG was releasing from NanoICG into BSA, we attempted to separate BSA from the NPs by size filtration. NanoICG or ICG was mixed with 36 mg/ml BSA in PBS and filtered at different timepoints through a
100 kDa MWCO protein filter in order to separate the BSA from the larger NP complexes. The quantity of BSA and ICG present in the initial solutions were measured with absorbance spectroscopy and the relative fractions of both ICG and BSA that moved through the filter were quantified and compared. Thus, if the
84 same fraction of ICG as BSA moved through the filter, then the relative fraction of
ICG:BSA would be 1.0, as shown in Figure 3-8. As the data indicates, when either NanoICG or ICG are exposed to BSA solution, approximately the same amount of ICG elutes with BSA. This data, combined with the absorbance spectra changes observed when NanoICG is exposed to flow conditions, indicate that NanoICG is likely releasing a substantial amount of ICG to bind serum proteins. Further work is required to elucidate the exact dynamics of release.
Figure 3-8. ICG and NanoICG elute similarly in presence of BSA, indicating that
ICG loaded into NPs is likely being released to serum proteins in vitro .
In Vitro Cellular Uptake of NanoICG
Cellular uptake of NanoICG vs free ICG was analyzed by in vitro incubation of MDA-MB-231 cells with either High-PBA-HLA NanoICG ( Figure 3-
85
9A ) or free ICG ( Figure 3-9B ), either with or without 10 mg/ml free HLA present as a blocking agent (Blue- with HLA, Red-without HLA). Results indicate that the effect of blocking by 10 kDa HLA on the uptake of NanoICG was negligible, as both blocked and unblocked NanoICG was taken up in cells at similar levels. The same results were observed with free ICG, as no difference was observed between cells that were exposed to free HLA. These results indicate that the uptake of NanoICG by cells is not dependent on the presence of free 10 kDa
HLA. While much larger molecular weight HLA polymers are found in vivo , these results indicate that the active targeting ability of NanoICG may be minimal.
However, active targeting of individual cells is not necessary for image guided surgery purposes due to resolutions being too low to detect individual cells. It has previously been shown that delivery of the majority of NP material to tumors relies on passive targeting via the EPR effect, and thus the lack of active targeting exhibited here is not likely to affect net delivery of ICG to the tumor bulk. 82
86
A B
Figure 3-9. Flow cytometry of MDA-MB-231 cells stained with either NanoICG
(A) or free ICG (B) with or without 10 mg/ml free HLA as a blocking agent (blue).
Results indicate there may be a slight be not meaningful blocking effect of free
HLA on the uptake of NanoICG, while uptake of free ICG by cells is not associated with the presence of HLA in solution.
Cytotoxicity
MS-1 mouse endothelial cells were chosen to represent mouse vasculature upon introduction of NPs. All NPs and NanoICG formulations did not affect the metabolic activity of these cells at physiologically relevant concentrations (i.e., 5-50 µg/ml of ICG) using a CCK-8 assay, as shown in
Figure 3-10A. Because High-PBA-HLA NanoICG was found to have the smallest effective diameter and approximately equivalent or higher ICG loading than other
NPs, these were further investigated for their cytotoxicity to MDA-MB-231 breast cancer cells. MDA-MB-231 breast cancer cells were chosen to determine the
87 effect on proliferation of target cancer cells. Proliferation of MDA-MB-231 cells was not affected after a 24 h exposure to High-PBA-HLA NanoICG, suggesting that the NPs do not have an effect on growth of target cell populations ( Figure 3-
10B). The results are analogous to the cytotoxicity profiles of other nanoparticle formulations of ICG. 221, 226
Figure 3-10. (A) A cell metabolic assay showed that NPs with and without ICG have cytotoxicity that is comparable relative to control cells. (B) Treatment of
MDA-MB-231 cells with 5 µg/ml High-PBA-HLA NP did not impact subsequent proliferation as determined by a DNA quantification assay.
88
3.0 Conclusions
The NPs developed here readily entrapped ICG through a self-assembly process, to a maximum of approximately 10 wt%. Differences in both NP formation and ICG loading potential were found to be dependent on the structure of hydrophobic ligand chosen, as both PBA and 5βCA ligands had similar ICG loading, while ODA has lower loading and also a lower requirement for molar conjugation to HLA needed for self-assembly. These results are important, as it shows that the choice of hydrophobic ligand plays a role in the loading of small molecules as well as the self-assembly process, thus suggesting that choice of hydrophobic ligand must be considered when investigating new formulations for loading small molecules. Differently structured hydrophobic conjugates will lead to both different NP stability, loading efficiency, and release, and thus the quantity and type must be considered in addition to the hydrophilic section.
Optical properties of NanoICG formulations universally showed reversible fluorescence quenching in aqueous solutions, and this fluorescence could be reversed either by dissolution in 50/50 DMSO/H 2O, where the reactivation occurred immediately, or over time by mixing in 37 mg/ml BSA. The fluorescence reactivation in the presence of serum proteins suggests a release from NanoICG, and that the rate of this release is dependent on flow conditions experience by the NPs. While fluorescence reactivation is a requirement for IGS, too quick a release may negate the benefit of using a NP formulation. Further work is needed to control the release of ICG in order to optimize the contrast between
89 healthy and diseased tissues. A modest but significant photoprotective effect was observed, which may indicate the potential for improved shelf life and stability in clinical practice. NanoICG was not observed to actively target CD44(+) cancer cells, though cells did show fluorescence. Active targeting is likely not necessary for this type of contrast agent however, as it is the tumor bulk that requires identification and delineation, so the intratumoral distribution of ICG may not play a meaningful role. NanoICG was not observed to decrease metabolism or proliferation of cancer cells in vitro , suggesting that low toxicities will also likely be observed in vivo . Additionally, as NanoICG is intended as a surgical contrast agent, it is unlikely that repeated exposure would be necessary.
90
CHAPTER 4
TUMOR UPTAKE, BIODISTRIBUTION, AND POTENTIAL FOR IMAGE GUIDED
SURGERY OF NANO-ICG IN XENOGRAFT MOUSE TUMOR MODELS
Text and the following figures: Figure 4-1, Figure 4-2, and Figure 4-3, have been reproduced with permission from Hill, T.K. et al., Indocyanine Green-Loaded
Nanoparticles for Image-Guided Tumor Surgery . Bioconjugate Chemistry 26(2),
2015.
1.0 Introduction
Fluorescence IGS in oncology has potential to better identify SLNs, positive tumor margins, and local metastases. In this chapter we investigate the potential of NanoICG formulations to identify tumor bulk and PMs in mice bearing solid tumors. Our preliminary in vivo experiment utilized the MDA-MB-231 human breast cancer cell line to establish tumors subcutaneously in the dorsal hip of female nude mice. Strong fluorescence enhancement in tumors was observed with NanoICG using a fluorescence image-guided surgery system and a whole- animal imaging system. Tumor-muscle contrast using NanoICG was significantly higher than ICG alone, and PMs were identified in one mouse. A second study was performed to verify and extend the results of the preliminary study using tissue phantoms to provide feedback on maximum potential imaging depths of tumors in non-fluorescing tissue. This second study confirmed the improved
91 signal and contrast observed in the initial study, and provided insight into deep tissue imaging of solid tumors.
2.0 Materials and Methods
iRFP transfection of MDA-MB-231 cells
Amplification of iRFP plasmid (piRFP 31857, Addgene, Cambridge, MA.) was performed by PCR using Platinum High Fidelity PCR SuperMix (Invitrogen cat#12532016) and custom designed primers. Primers were designed to include
15bp of homology to the LentiX Bicistronic Expression System (puro) from
Clontech (Cat #632183) in order to perform ligation utilizing In-Fusion HD cloning
(Clontech #639645). Agarose gel electrophoresis of PCR product was resolved, excised, and purified using Infusion spin column PCR purification kit. The LentiX
Bicistronic vector was linearized using EcoR1, resolved by agarose, excised and gel purified. For ligation of iRFP into the lentix vector, the In-fusion Cloning was carried out using 5x Infusion HD Enzyme mix, transformed, plated, and colonies selected. Restriction Enzyme Digest was performed to check for orientation of the product. Positive clones were verified and used to produce lentiviral supernatants. To make lentivirus, the Lenti-X HTX packaging system (Clontech) was used with Lenti-X 293T cells, and virus production was verified using Lenti-X
GoStix. MDA-MB-231 cells were transduced using 0.5ml of lentiviral supernatant with 8 µg/ml polybrene in DMEM media for 48 hours, virus removed, and drug selected for stable cell lines using 5 µg/ml puromycin for 5 days.
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Initial Near Infrared Fluorophore Enhanced Image-Guided Surgery
Breast tumor xenografts were introduced into 12-14 week old female athymic nude mice (Jackson Labs; Bar Harbor, ME) by injection of 2 x 10 6 iRFP
MDA-MB-231 cells in 50/50 media/matrigel subcutaneously into the right hip.
When the tumors reached 500-1000 mm 3, mice were injected with an intravenous infusion via a tail vein of either ICG or High-PBA-NanoICG (10 nmol
ICG total) in 200 µl water ( N = 5 mice/group). The quantities of ICG in both the free ICG and High-PBA-NanoICG solutions were determined by dissolving a fraction in 1:1 H 2O/DMSO and acquiring the absorbance spectrum, followed by making the appropriate dilutions to achieve 10 nmol/200 µl as determined by an
ICG standard curve. Absorbance spectra of both free ICG and High-PBA-
NanoICG were analyzed and diluted to equal absorbance values prior to injection. Mice were euthanized 24 hours after injection of NIR fluorescent contrast agents and the mice were imaged using a Pearl ® Impulse Small Animal
Imaging System (LI-COR Biosciences; Lincoln, NE). Tumors were then removed using a separate image-guided surgery system that has been previously described. 191, 230 In brief, skin was removed to visually inspect identifiable tumor tissue. Next the portable fluorescence spectrometer with a fiber-coupled handheld (SciApps; Laramie, WY) unit was used to detect NIR emission from the tumor. The laser also served as the directed excitation source for a wide-field
NIR imaging system (SpectroPath Image-Guided Surgery System; Atlanta, GA) that merges NIR and color channels and provides real-time video feedback for the surgeon. Using the IGS system, tumor boundaries were first highlighted and
93 then underwent a debulking procedure. Image-guided resection of enhancing regions was iteratively performed until either no fluorescence signal was observed at the tumor margin or until debulking was no longer possible. Mice were then re-imaged on the LI-COR whole-animal imaging system for comparison of iRFP fluorescence emission from the MDA-MB-231 tumor cells
(700 nm channel) with fluorescence emission from ICG (800 nm channel).
Tumor, muscle, kidney, liver, and spleen were harvested and underwent further fluorescence imaging. Using LI-COR software, tumor to muscle contrast was determined by generating an area of interest (AOI) around each tumor and corresponding muscle sample. The average signal intensities of these AOIs were then used to calculate the signal to noise and contrast to noise ratios as shown in
Equations 3.2 and 3.2. Histological analysis was performed with H&E staining.
[4.1] = = [4.2] = = −
[4.3] = ∗
Extension of In Vivo Biodistribution and Image Guided Surgery
Materials were identical to those used in the initial in vivo experiment as outlined above. Two million MDA-MB-231 cells previously transfected with an
NIR fluorescing protein iRFP were mixed in 50/50 media/matrigel and injected orthotopically into the inguinal mammary fat pads on both sides of 10-week old female nude mice (5 mice, up to 10 tumors per group). Tumors were allowed to
94 establish for 10-21 days prior to contrast agent administration. Presence of tumors and approximate tumor size was confirmed by live imaging using a Pearl ®
Impulse Small Animal Imaging System (LI-COR Biosciences; Lincoln, NE). Three contrast agents groups were used: (1) free ICG, (2) High-PBA-HLA NanoICG, and (3) High-5βCA-HLA NanoICG. NP formulations were filtered through a 0.45
µm filter to remove large aggregates. Each contained 10 nmol of ICG injected per mouse dissolved in pure sterile water. As determined by absorbance spectroscopy. Contrast agents were administered by tail vein injection and mice were sacrificed either 4 or 24 hours later. Mice were the dissected and imaged again using the Pearl ® Impulse to compare tumor and tissue signals. LI-COR software was used for image analysis to determine biodistribution, SNR, and
CNR. One mouse per contrast agent group was also examined using the IGS equipment as outlined above.
Tissue Phantoms
Tissue phantoms, which possess optical properties similar to that of natural tissue, were produced as described elsewhere. 231, 232 Briefly, 10 w/v% gelatin was added to a solution of 50 mM Tris-buffered saline, pH 7.4, 150 mM
NaCl, 0.1% NaN 3. This solution was warmed to 45 °C under constant stirring until gelatin was dissolved, followed by cooling to 25 °C and addition of 1 v/v%
Intralipid, and 1 w/w% bovine hemoglobin. This mixture was then pipetted into 6- well polystyrene plates to give multiple phantoms with thickness of 6 – 2 mm.
Plates were then covered and stored at 4 °C until used.
95
After tumor excision via IGS, tumors were placed underneath phantoms and examined using both medium-low (30 W), and medium (80 W) power.
Phantoms were placed over tumors to determine the maximum depth at which tumors could be observed under both power levels. Phantoms were also used in the LI-COR system to determine signal loss from tumors at increasing phantom depths up to 6 mm. Tumors were covered with phantoms made to depths ranging from 2-6 mm in 1 mm increments, and imaged. Signal loss as a function of thickness was measured.
Statistical Analysis
Organ biodistribution, tumor contrast, and size distribution were performed with two-way ANOVA using multiple comparisons with Tukey’s test. Total signal and mean signal as a function of tumor area were analyzed with simple linear regression. All statistical analyses were done in Prism 6.0 (GraphPad Software;
La Jolla, CA) and Excel.
4.0 Results and Discussion
Initial In Vivo Investigation Using Subcutaneous Flank Tumors
iRFP-labeled MDA-MB-231 human tumor xenografts were grown subcutaneously in the back right hip of female nude mice. When tumors were sufficiently large, mice were injected with either ICG in water (i.e., the standard clinical method) or the High-PBA-HLA NanoICG formulation. Mice were
96 euthanized 24 hours after injection and imaged with a LI-COR Pearl Impulse small animal imaging system and an image-guided surgery system. 191, 230 Figure
4-1A shows the signal to noise ratios (SNRs) of resected tissues analyzed using the small animal imaging system. Compared to ICG, NanoICG resulted in substantially higher signal in tumor, although the difference was not statistically significant (p = 0.088). The fluorescence due to ICG did not significantly differ in other tissues (p ≥ 0.78). The combination of higher signal in tumor, but similar background in surrounding tissue resulted in a significantly higher contrast to noise ratio (CNR) in mice that were administered NanoICG compared to ICG alone (p = 0.037) as shown in Figure 4-1B. The average CNR in NanoICG- treated mice was more than twice that of mice treated with free ICG (91 vs. 40).
This is an important distinction, as it is CNR, rather than SNR, that allows for more robust discrimination of tumor from normal tissue using IGS. The relative biodistribution of NanoICG at 24 hours was consistent with previous results of nanoparticle encapsulated ICG; i.e. showing the highest signal in liver and kidneys with higher signal in the tumor compared to surrounding tissues. 222, 226,
233-235 The majority of ICG accumulated in the liver for both NanoICG and free
ICG formulations, which is consistent with ICG clearance. 236, 237 While no significant difference was observed between the SNR of similar tissues,
NanoICG exhibited lower kidney and muscle accumulation, and higher liver and spleen accumulation than free ICG. Representative NIR emission images from
ICG and NanoICG are shown in Figure 4-2. The comparable overall SNR from
ICG and NanoICG are indicative of some degree of ICG dissociation from the
97
NPs. The higher liver and spleen signal from NanoICG may indicate some interaction with resident macrophage. The statistically significant CNR values between ICG and NanoICG in tumor indicate that NanoICG has distinctly different biodistribution compared to free ICG. These results are likely independent of NanoICG dose. A recent report demonstrates that ICG that was directly conjugated to PEG-grafted HLA, where ICG served as the hydrophobic ligand, resulted in tumor signal, both fluorescence and photoacoustic, that was more than double that of femur. 238 However, the dose of ICG used in this study,
50 nmol per mouse, was 5 times greater than the dose used in our study.
Figure 4-1. (A) Biodistribution of ICG 24 hours after injection as measured by
SNR. Liver accumulated the majority of ICG from both formulations, followed by kidney. NanoICG and ICG SNR were not found to be significantly different in the tested tissues, although the NanoICG signal was higher in tumor tissue (p = .088;
N = 5 mice/group). (B) The contrast-to-noise ratio (CNR) between tumor and surrounding muscle was significantly higher in NanoICG-treated tumors,
98 suggesting that the NanoICG formulation improved the potential for distinguishing between tumor and normal tissue. *p ≤ 0.05 was considered significant.
Figure 4-2. Representative NIR imaging of dissected organs from mice treated with free ICG (A) or NanoICG (B). Results show higher uptake in the tumor (t), liver (li), and spleen (s) of NanoICG treated mice. Intensities of both images have been set to the same scale. Order of organs displayed from left to right: tumor (t), muscle (m), heart (h), lung (lu), pancreas (p), kidney (k), liver (li), spleen (s).
Image Guided Surgery
Small animal florescence imaging, provided by LI-COR Pearl Impulse system, indicated iRFP fluorescence (700 nm channel; false-colored red) that was consistent with the location of the iRFP-labeled MDA-MB-231 cells; a representative example is shown in Figure 4-3A. Using the separate IGS
99 system, 191, 230 directed laser excitation (785 nm) of the tumor produced strong fluorescence enhancement due to the presence of ICG ( Figure 4-3B; false- colored cyan), which corresponded to the boundaries of iRFP-labeled tumor cells observed in Figure 4-3A . Laser excitation of surrounding normal skeletal muscle showed no fluorescent signal based on a predetermined threshold ( Figure 4-
3C). 191 Subsequently, the tumor was debulked under IGS using fluorescence enhancement from NanoICG. Figure 4-3D shows the presence of ICG in the resected tumor mass. A region of tissue that fluorescently enhanced at the tumor/normal tissue interface during IGS could not be excised because removal was restricted by the presence of bone ( Figure 4-3E). However, subsequent 700 nm imaging (in the whole animal imaging system) and histological analysis after animal necropsy revealed that this fluorescence-enhancing region was due to tumor infiltration into normal tissue ( Figure 4-3E&F). Further investigations are planned to determine NanoICG’s ability to enhance specifically at the tumor margin. Analysis of the debulked tissue and mouse carcass with whole animal fluorescence imaging at 700 nm to detect iRFP-labeled MDA-MB-231 cells
(shown in Figure 4-3F) confirmed the fluorescence enhancement detected in tumor by IGS.
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Figure 4-3. In vivo analysis of NanoICG tumor accumulation. (A) Preoperative imaging of iRFP shows the location of MDA-MB-231 breast tumor xenograft; the signal in the abdomen is due to rodent chow in the intestinal tract. (B) Image- guided surgery (left) and H&E histological staining (right) confirm high fluorescent signal in the bulk of the tumor, (C) while only low fluorescent signal is observed in adjacent normal tissue. (D) The resected tumor exhibits fluorescent signal;
101 presence of tumor was confirmed at histology. Tumor margins were then inspected for fluorescent enhancement to confirm tumor removal. (E)
Fluorescence was detected at the tumor margin, suggesting remaining tumor tissue. (F) Positive margins were confirmed by observing iRFP signal at the location indicated by the presence of NanoICG.
Further In Vivo Investigation with Orthotopic Breast Tumors and an Additional
Contrast Agent
This study differs from the previous section in that tumors were injected into the inguinal mammary fat pads of female nude mice, with one tumor on each side of each mouse. Cells used were again MDA-MB-231 breast cancer cells expressing iRFP protein. Both 4 and 24 hours timepoints were examined, and excised tumors were further analyzed using tissue-mimicking phantoms.
Mice treated with NanoICG showed lower SNR compared to free ICG at 4 hours, and higher SNR compared to those treated with free ICG at 24 hours, as shown in Figure 4-4A&B . Total ICG signal was higher at 4 hours for all organs and all contrast agents, suggesting that a significant amount of metabolism and excretion occurs between 4 and 24 hours. Free ICG shows high levels of fluorescence in the kidneys at 4 hours, which is consistent with the clearance of small molecules. SNRs of NanoICG formulations are significantly lower in the kidney but not significantly different in the liver at 4 hours, which is consistent with clearance of NPs. Biodistribution at 24 hours was also consistent with that expected of NP materials, with higher accumulation in the liver and spleen, and
102 approximately equal distribution in other organs. CNRs at 4 hours were highly variable ( Figure 4-4C ), with some CNRs from individual mice being negative
(higher signal in muscle). However, at 24 hours High-PBA-HLA NanoICG [95%
CI: 15.4-2.6] and High-5βCA-HLA NanoICG [95% CI: 12.62-0.65] showed higher tumor to muscle CNR than ICG ( Figure 4-4D). These results confirm the previously observed in vivo results that NanoICG formulations increase ICG accumulation within xenograft tumors and improve differentiation between diseased and healthy tissue at 24 hours. Interestingly, the 4 hour data gives insight into the metabolism and clearance of these contrast agents, and further suggest that NanoICG retains a NP structure in vivo . The kidney and liver signals are consistent with NP clearance at both 4 and 24 hours for NanoICG.
Additionally, total ICG administered was the same across all contrast agent groups, but the fluorescence signal is lower in NanoICG mice in all organs at 4 hours. It is thus possible that the fluorescence signal is being quenched due to tight packing of ICG within the NPs. Tumors were similarly sized across contrast agent groups and thus the high tumor SNR and CNR likely result from improved accumulation due to the NP formulation of ICG. While the observed loss of ICG from NanoICG to serum proteins previously observed in vitro is likely to be occurring, it appears that NanoICG still allows for improved tumor accumulation.
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Figure 4-4. NanoICG formulations show lower total signal in tumors at 4 hours relative to ICG (A), but higher signal relative to ICG at 24 hours (B). At four hours the ICG signal is higher in most tissues compared to NanoICG, particularly the kidney where small molecules are cleared (A). Contrast between tumor and muscle was highly variable at 4 hours, and no formulation is significantly different from another (C). The biodistribution of NanoICG at 24 hours reflects that expected of NP formulations, with higher signal in the liver and spleen and similar distributions in other tissues (B). The total signal present at 24 hours was less
104 than at 4 hours in all tissues, suggesting that a high amount of metabolism or clearance has occurred to reduce to total amount of ICG present. (D) Contrast between tumors and muscle at 24 hours showed that uptake into the tumor was significantly higher for NanoICG formulations.
Analysis of the size distribution of tumors in each treatment group is shown for 4 hours ( Figure 4-5A ) and 24 hours ( Figure 4-5B ). Tumors analyzed at 4 hours deviated slightly between groups, with the ICG group (9.45 ± 4.79 mm 2) being slightly larger than the High-5βCA-NanoICG group (5.56 ± 1.60 mm 2). At 24 hours, all groups showed similar tumor sizes. Total signal in the tumors were proportional to tumor cross-sectional area for both 4 hours (Figure
4-5C ) and 24 hours ( Figure 4-5D ), with 95% CI for these values as [ICG: 3125-
9660 (4 hrs), 984.7-1490 (24 hrs)], [High-PBA-NanoICG: -1476-5135 (4 hrs),
603.2-2772 (24 hrs)], and [High-5βCA-NanoICG: 1130-3633 (4 hrs), 1669-4557
(24 hrs)]. Slopes from Figure 4-5C&D are in units of (total pixel intensity)/mm 2.
However, mean signal per unit area, shown in Figure 4-5E (4 hours) and Figure
4-5F (24 hours), was not found to give a trend for increasing volumetric signal intensity with change in tumor size, as no slopes were found to be significantly different from zero. In other words, the signal intensity from equal tumor areas is not dependent on the size of the tumor (E&F ), but the total signal intensity from a tumor is dependent on tumor size ( C&D ). This indicates that contrast agent accumulation is approximately the same in a given volume of tumor, regardless of the size of that tumor, and thus tumor detection across the range tested is
105 limited by instrument sensitivity and contrast, rather than tumor size. As can be seen in Figure 4-5, even very small tumors with cross sections less than 1 mm 2 were detectable after resection, and thus could potentially be visualized in vivo .
Figure 4-5. (A) The average size of tumors in each contrast agent group at 4 hours was between 5-10 mm 2, and ICG tumors were slightly larger than High-
5βCA-NanoICG tumors. No difference in tumor size was observed at 24 hours
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(B). (C-D) Total signal present in a tumor is dependent on tumor size, with larger tumors having more total signal. However, the average signal intensity per unit area does not correlate with tumor size (E-F).
Tumors were examined in situ using the IGS system by first resecting overlying skin and examining the fluorescent signal at different locations of the mouse ventral surface. Images were captured before and after tumor removal.
As is seen in Figure 4-6, the signal on the surface at the tumor location is strong
(Figure 4-6A). Most tumors were clearly visible to the naked eye and the majority of the tumor bulk could be resected without the aid of the IGS system. Isolated tumors also showed strong fluorescent signal ( Figure 4-6B). However, in most cases the surgical margins could not be examined using the IGS system due to strong underlying fluorescent signal from the visceral organs ( Figure 4-6C). As seen in Figure 4-4, the fluorescent signal from the kidneys, liver, and spleen are very high, and the intestines appeared fluorescent as well. Thus, due to the hepatobiliary and renal accumulation, surgical margins are generally not discernable using this tumor model in mice. In the previous study ( Figure 4-3), the tumor location was in the back right hip on the dorsal surface and this location afforded protection of the signal from the visceral organs, allowing for effective detection of PMs. Future animal models will have to take this into account, to ensure that tumors are located in clinically relevant tissue while ensuring there is a sufficient distance from liver, kidney, and spleen to prevent
107 background signal from these organs. Thus, a larger animal model may be necessary.
Figure 4-6. (A) Administration of High-PBA-HLA NanoICG showed strong fluorescent signal in the tumor region prior to excision. (B) Surgical removal of the tumor shows a strong fluorescent signal when separated from the body, however the background signal from the internal organs remained high after excision (C). Laser is highlighting the previous position of the tumor, however no
108 tumor is present. Excess signal results from high levels of ICG in visceral organs in the background, indicating that this tumor model is inadequate for IGS.
Analysis with Tissue Mimicking Phantoms
Detection of local metastases is necessary for the complete removal of diseased tissue during surgery. However, these local metastatic sites will not always be located at the tumor surface, making detection at depth a critical feature of IGS. Tissue mimicking phantoms were used to determine the theoretical maximum detection depth of ex vivo tumors using ICG and NanoICG contrast agents. Signal loss of xenograft tumors was measured from 0-6 mm below phantoms. As shown in Figure 4-7, the average signal decreases with increasing tissue depth. Tumors were still visible up to 6 mm beneath tissue phantoms, and the 95% CIs for predicted depths of zero signal are 6.8 – 8.8 mm
(ICG), 6.7 – 9.8 mm (High-PBA-HLA), and 6.7 – 10.7 mm (High-5βCA-HLA).
Representative tumors at these depths are shown in Figure 4-7C . These phantoms did not contain any ICG, and therefore represent idealized tissues in which contrast agent accumulation does not occur, or from which it has cleared completely. In a real tissue, some level of contrast agent will be present, and this will likely decrease the sensitivity to tumors found at depth. Nonetheless, these results indicate that NIR fluorescence imaging using NanoICG provides sufficient signal to identify tumors below the tissue surface and may thus improve the identification of local metastases or SLNs in larger animal models or humans.
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Figure 4-7. Tumor signal decreases as a function of depth beneath tissue- mimicking phantoms at both 4 hours (A) and 24 hours (B). Relative SNR. (C)
Representative images from the 24 hour time point of tumors beneath phantoms ranging from 6-0 mm, showing an increase in scattering and decrease in photon penetration with increasing phantom thickness.
The image-guided tumor surgery study reported here, removal of a breast tumor xenograft in mice with image-guidance, demonstrates the potential of
NanoICG to depict tumor margins in the operating room. The colocalization of
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ICG signal and iRFP-labeled MDA-MB-231 tumor cells establishes that ICG delivery to tumors using NanoICG was successful and indicated the tumor boundaries. The strong contrast enhancement in the tumor in both studies showed that more ICG was delivered to tumors by NPs with lower background signal in surrounding tissue, which could provide meaningful data to surgeons in the operating room. In combination, these results suggest that use of NanoICG could potentially serve as a contrast agent for IGS.
Ultimately, one could envision using NanoICG with image-guided surgery to determine surgical margins. However, determining the tumor margin is complex. Currently, a positive margin is determined by the presence of neoplastic cells at the edge of the excised tumor. These can be detected by intraoperative pathology and definitively by IHC after the surgery. There is no method currently available to detect tumor cells remaining in the surgical cavity.
Using xenograft tumor models in mice, while useful to evaluate overall SNR and tumor contrast, has physical and biological limitations. With the resolution of optical imaging methodologies being approximately 0.1 x 0.1 mm, 230, 239 this gives the level of contrast agent distribution and signal necessary for effective surgical removal. Further investigation will be required to determine how the intratumoral biodistributions affects the surgeon’s ability to detect surgical margins by image-guided surgery.
The mechanism of ICG release to transition from a quenched to an unquenched state is likely the result of release of ICG from NPs to serum proteins and lipids or degradation via hyaluronidases. The comparable overall
111 biodistribution of NanoICG to ICG in non-tumor tissues at 24 hours indicates that
ICG exchanges from NanoICG to serum proteins to some extent. Intratumoral degradation by hyaluronidases provides an additional mechanism for tumor- specific fluorescence activation, but differentiating these effects will require additional investigation. The high kidney signal of free ICG at 4 hours suggests that there is a high level of renal clearance immediately after administration. As
NanoICG formulations do not appear to be cleared as rapidly by this mechanism, this suggests that NanoICG-treated mice may show higher tumor SNRs because more ICG remains in circulation. The fluorescent signal of free ICG-treated mice was higher on average in all organs at 4 hours. We also observe similar or greater levels of fluorescence in all tissues from animals treated with NanoICG formulations at 24 hours. As the same quantity of ICG was injected in all groups, this suggests that the NanoICG formulations are exhibiting a level of fluorescence quenching at 4 hours, which is causing the reduced signal.
Combined, the high signal of free ICG in all organs at 4 hours, equal or higher signal of NanoICG at 24 hours, and fluorescence quenching effect observed in vitro , suggest that NanoICG remains intact during circulation. This results in fluorescence quenching, and an increased circulation time, resulting in higher tumor CNR than ICG at 24 hours.
Analysis with tissue mimicking phantoms provided signal up to 6 mm deep and predicted maximum detection depths of approximately 8-9 mm. This represents an ideal situation in which the surrounding tissue provides minimal contrast. Studies using similar phantoms detected ICG-loaded agarose
112 inclusions up to 15 mm deep, with signal declining rapidly between 7-11 mm.232
Other studies have found that tissue penetration is dependent on tissue type, with fat having the higher penetration than liver or lung, and a maximum ICG detection depth of approximately 5 mm. 191 The potential maximum detection depth of tumors underlying healthy tissue is likely to be between 5-10 mm. 191, 230,
239 However, a number of factors will influence this depth. Volume and fluid distribution within the tumor will strongly affect the level of contrast agent uptake and thus total signal and imaging depth. 240 The number of cells in a tumor may influence fluorescence signal indirectly by affecting tumor size and vascularity, and therefore total contrast agent uptake.
4.0 Conclusions
The potential to use IGS in treatment of operable tumors is clear: improved detection of tumors and their boundaries could minimize recurrent disease. In this study NPs derived from HLA were used to deliver the NIR fluorophore ICG in an attempt to improve contrast enhancement. Tumor contrast was observed in both studies, providing strong evidence for a NP-mediated delivery of ICG. However, it has become clear that the choice of animal model and tumor location are of high importance when investigating the efficacy of contrast agents with IGS. Ultimately, the efficacy of NanoICG will be depend on whether the increased contrast enhancement in IGS results in decreased tumor recurrence, and further studies in larger animal models are required to test this.
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CHAPTER 5
APPLICATION OF HYALURONIC ACID DERIVED NANOPARTICLES TO
DRUG DELIVERY
This work was performed in conjunction with the Steven J. Kridel and W. Todd
Lowther labs. Empty and drug loaded NPs were provided by our lab for the following assays. Jill Clodfelter performed the direct FASN inhibition assay
(Figure 5-2A), Frances B. Wheeler performed the 14 C-acetate uptake assays
(Figure 5-2B and 5-4), and Amanda L. Davis performed the Seahorse metabolic assays (Figure 5-7). This work is currently in preparation for submission to
Molecular Pharmaceutics.
1.0 Introduction
ORL is an effective, irreversible inhibitor of FASN. However, it is highly hydrophobic and does not distribute well when administered orally. Thus, a method is needed to improve its distribution and activity when administered intravenously. In order to improve delivery of active ORL to tumors we have developed a novel ORL formulation, termed Nano-ORL, in which ORL is loaded into the hydrophobic regions of self-assembled polymeric NPs derived from
HLA.45, 241 In addition to being a natural, biodegradable polymer, HLA may also serve to improve uptake into tumor cells through interaction with the cell surface receptor CD44, which is overexpressed in a number of tumor types. 242-245
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Increased activity of hyaluronidases found in breast and prostate cancers may also serve to release therapeutic payloads upon delivery to tumors. 46, 47 In this chapter we modified HLA polymers with a hydrophobic moiety, PBA, to drive self- assembly and load ORL into Nano-ORL. Nano-ORL was analyzed for physical properties and ORL loading capability. FASN was confirmed as a target molecule for Nano-ORL using direct molecular inhibition and decreased lipid synthesis was confirmed by measurement of 14 C-acetate incorporation. Nano-ORL was demonstrated to be more toxic than ORL to cancer cell lines over a 48 hour period. This increase in toxicity is shown to result from an increased stability of the Nano-ORL formulation over free ORL when both are dissolved in aqueous solution. Finally, Seahorse XF24 analysis demonstrated decreased O 2 consumption and mitochondrial metabolism after treatment with Nano-ORL.
2.0 Materials and Methods
ORL was obtained from 120 mg Xenical capsules (Roche; Basel, CH) or in powder form (Alfa Aesar; Ward Hill, MA). Sodium hyaluronate was purchased from Lifecore Biomedical (Chaska MN). 1-Pyrenebutyric acid, 1,3- diaminopropane, N-hydroxy succinimide, and 1-ethyl-3-(3- dimethylaminopropyl)carbodiimide (EDC) were obtained from Sigma-Aldrich (St.
Louis, MO). Unless otherwise noted, all water was obtained from a Barnstead
NANOpure Diamond (Thermo Scientific; Waltham, MA) system producing 18.2
MΩcm water. Methanol, N, N dimethylformamide (DMF), 96-well tissue culture
115 plates (Falcon), were purchased from Fisher Scientific (Pittsburgh, PA). Ethanol was purchased from the Warner-Graham Company (Cockeysville, MD).
Penicillin/streptomycin and other media reagents (Gibco) were purchased from the Cell and Viral Vector Core Laboratory of the Comprehensive Cancer Center of Wake Forest School of Medicine. Cell lines were obtained from American Type
Culture Collection (Manassas, VA) and were grown in RPMI-1640 or EMEM with
10% fetal bovine serum, 1% penicillin/streptomycin.
Drug Loading and Quantification
ORL was isolated from Xenical capsules by mixing capsule contents into
10 ml ethanol or methanol, followed by vortexing and centrifugation at 1500 RPM for 5 minutes. Solution was then separated from the precipitated filler material, followed by a second separation and centrifugation. Stock 12 mg/ml (24.2 mM)
ORL solution was then aliquoted and stored at -20 °C. Stock solution was also prepared from powdered ORL (Alfa Aesar) by dissolution into ethanol or methanol to a concentration of 12 mg/ml followed by storage at -20 °C. ORL was loaded into NPs by dissolution of 18.0 mg HLA into 10 ml H 2O, followed by addition of 10 ml ethanol. ORL (2.0 mg, 4.0 µmol, 20wt% loading) was then added and the solution was dialyzed against pure water for 24 hours (MWCO =
3500 Da, Spectrum Labs). Material was then filtered through PD-10 columns (GE
Lifesciences; Pittsburgh, PA) to remove free ORL, followed by lyophilization to produce a low density, white, fibrous material, and was stored at -20 °C. This material is termed Nano-ORL.
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ORL loading efficiency was determined by extraction from Nano-ORL followed by HPLC quantification. Extraction of ORL was performed following the protocol for extraction of ORL from human plasma as described in the literature. 246 A standard curve of ORL extraction from NPs was performed by spiking solutions of empty NPs with known quantities of ORL from stock solution and performing the following extraction protocol. Empty NPs were dispersed into pure H 2O to a concentration of 1.0, 0.5, or 0.25 mg/ml. Stock ORL was then added to the empty NP solutions to a concentration of 0.20, 0.10, or 0.05. mg/ml.
1.0 ml aliquots of these solutions were then mixed with 1.0 ml acetonitrile, to which 5 ml n-hexane was added. This solution was then stirred for 30 minutes, at which time the hexane layer was removed and evaporated under vacuum.
Extracted ORL was dissolved in 100% HPLC grade MeOH and run on a
Beckman Ultrasphere UDS 4.6 x 250 mm column with 5 µm particle size using a mobile phase of 82.5:17.5 methanol:acetonitrile-.01% trifluoroacetic acid with a
Waters 510 pump, 717+ autosampler, and 2998 PDA UV/Vis detector.247, 248
Separation was isocratic and ORL was measured at 200-205 nm. Each ORL concentration was performed in quadruplicate. ORL content in Nano-ORL was then measured by performing these extraction methods and comparing to the standard curve. Loading efficiency of Nano-ORL was then determined by re- suspension of Nano-ORL into pure H 2O at a concentration of 0.50 mg/ml and extracted and analyzed in quadruplicate as described above. Extracted values were then analyzed using the standard curve.
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Physical Characterization
In addition to 20wt% ORL loading, 5wt% and 10wt% loading Nano-ORL formulations were also synthesized using identical conditions in order to test the effect of ORL loading on NP size. 0, 5, 10, and 20wt% ORL formulations were dissolved in pure H 2O to a concentration of 1.0 mg/ml. Hydrodynamic diameter was then determined using a ZetaPlus system with onboard dynamic light scattering (DLS) analyzer (Brookhaven Instruments Corporation; Holtsville, NY).
FASN Inhibition
The C-terminal thioesterase domain of FASN (FASN-TE, residues 2200-
2510) was expressed and purified as previously reported. 249 The inhibition assay using 4-methylumbelliferyl heptanoate (MUH) as the substrate was slightly modified from what was previously reported.250, 251 The 100 μl reactions contained 250 nM FASN-TE, 100 mM Tris pH 7.4, 50 mM NaCl, 1 mM EDTA,
120 μM MUH, 1% DMSO, and Orlistat 0.1-10 μM. Orlistat was preincubated with
FASN-TE for 30 min at 37 °C prior to starting the reaction with MUH. The samples within a 96-well plate were read, at 30 s intervals for 30 min, on a Tecan
Safire 2 instrument (excitation, 350 nm; emission, 450 nm).
Inhibition of Lipid Synthesis
For analysis of the inhibition of lipid synthesis by Nano-ORL, PC3 cells were seeded in two 24-well plates, 7 x 10 4 cells per well, and treated two days later with inhibitors or vehicle controls, in quadruplicate for 17-18 hours. [2-C14 ]
118 acetate (0.037MBq, 1 m Ci; PerkinElmer, Boston, MA) was added to wells in one plate for two hours. The cells were trypsinized, washed, and lysed in hypotonic buffer (20 mM Tris-HCl, pH 7.5, 1mM EDTA, 1mM DTT). Lipids were extracted in chloroform :methanol (2 :1). The organic fraction was washed twice with PBS, then transferred to vials for scintillation counting (Beckman Coulter LS 6500).
Protein concentration was determined by the Lowry method (Bio-Rad, Hercules,
CA) in cells treated in the duplicate plate.
Cytotoxicity
PC3, MDA-MB-231, U87mg, or RKO cell lines were seeded separately into 96-well plates at concentrations of 2000 cells/well and allowed to adhere overnight. Empty NPs and Nano-ORL were resuspended in culture media
(EMEM with 10% FBS and 1% P/S), and stock ORL was diluted into normal media from ethanol. Cells were then treated with either normal media, empty
NPs (0.062 mg/ml), ORL (0.0124 mg/ml, 25 µM), or Nano-ORL (0.062 mg/ml, theoretical [ORL] = 25 µM, actual [ORL] = 24 µM due to loading efficiency) and allowed to incubate at 37 °C for 48 hours. After 48 h incubation, cells were analyzed by CCK-8 assay (Dojindo, Japan) and results were normalized to the standard media group.
Similar methods were used to study the effect of pre-incubation on ORL and Nano-ORL. PC3, MDA-MB-231, U87mg, or RKO cell lines were seeded separately into 96-well plates at concentrations of 2000 cells/well and allowed to adhere overnight. ORL and Nano-ORL were first dispersed in serum-free media
119 at concentrations of 0.124 mg/ml (250 µM) and 0.62 mg/ml, respectively. These solutions were then incubated for 24 hours at 37 °C. ORL and Nano-ORL solutions were then diluted to 0.0124 mg/ml (25 µM) and 0.062 mg/ml, respectively, in serum-containing media. Empty NPs were also treated in a manner identical to Nano-ORL. Non-incubated ORL and Nano-ORL solutions were also formed at identical concentrations to pre-incubated ORL and Nano-
ORL. All non-incubated solutions, including standard media, were diluted to the same fetal bovine serum and penicillin/streptomycin concentrations as pre- incubated solutions. Solutions were then applied to cell lines and allowed to incubate for 24 hours. At 24 h, solutions were re-applied (both pre-incubated and non-incubated) in order to maintain concentrations of free ORL and Nano-ORL.
CCK-8 assay was performed at 48 hours and all groups were normalized to the standard media group.
Chemical Stability
Nano-ORL (0.5 mg/ml) and ORL (0.1 mg/ml) were separately dispersed into PBS and incubated at room temperature for 96 hours. At 96 hours 1 ml aliquots of each were removed and Orlistat was extracted as previously described. Samples were then dried under argon and redissolved in 82.5:17.5 methanol:acetonitrile with 0.1% trifluoroacetic acid and infused into a Quantum
Discovery Max (Thermo Scientific) quadrupole mass spectrometer at 10 µl/min for m/z 450-600 Da.
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Mitochondrial Respiration Assays
All cellular oxygen consumption rates (OCR) were measured using the
XF24 Extracellular Flux Analyzer (Seahorse Bioscience; Massachusetts, United
States) per manufacturer’s instructions. LnCap cells were plated on 22.4 μg/ml
Cell-Tak (Corning; Corning, New York, United States) at 40,000 cells/well in
RPMI-1640 with 10% FBS/1% PS. MDA-MB-231 cells were plated at 40,000 cells/well in RPMI-1640 10% FBS/1% PS. Cells were treated with identical media containing either vehicle or test compounds for 16 hours. Before OCR measurements were taken, cells were washed with Seahorse assay media supplemented with fresh sodium pyruvate and glucose per published Seahorse protocol and incubated at 37 °C without CO 2 for 1 hour. After equilibration, three
3-minute measurements were recorded. Oligomycin (1uM), FCCP (0.5uM), and
Rotenone/Antimycin A (1uM each) were injected into each well sequentially with three 3-minute measurements after each injection. The following metrics were used in accordance with published Seahorse protocols and previous publications 252 to analyze results:
[5.1] = 3
[5.2] =
[5.3] =
121
[5.4] =
[5.5] = 1 −
After completion of OCR measurements, cells from each treatment condition were pooled and counted via trypan blue exclusion. Total live cells counts were used to calculate the average number of live cells per well for each treatment group. The average number of live cells per well per treatment was then used to normalize OCR readings to live cell number.
Statistical Analysis
Extracted stock ORL measured by HPLC was plotted and analyzed using linear regression to form a standard curve. Extracted Nano-ORL was quantified using this standard curve. Average hydrodynamic diameters of Nano-ORL with different loading wt% ORL were plotted and analyzed using linear regression.
FASN inhibition, 14 C-acetate incorporation, cytotoxicity, and mitochondrial respiration assays were analyzed separately using two-way ANOVA with multiple comparisons using Tukey’s test.
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3.0 Results and Discussion
Drug Loading and Physical Characterization
ORL was loaded into NPs by dissolution of ORL and empty NPs into
50:50 H2O:EtOH followed by dialysis into pure H2O, PD-10 filtration, and lyophilization to produce Nano-ORL. A standard curve of ORL extraction from
NPs was created by spiking empty NPs with known quantities of ORL followed by extraction and quantification. Figure 5-1A shows the standard curve and extracted Nano-ORL values. An R 2 = .98 was achieved over the range from 0.05 mg to 0.20 mg/ml [ORL] for the standard curve. Nano-ORL was then re- suspended in pure H 2O to a concentration of 0.50 mg/ml, (theoretical [ORL] = 0.1 mg/ml) extracted and compared to the standard curve. Analysis of the extracted
Nano-ORL gave a loaded wt% of 19.3wt% ± 0.4wt%, corresponding to a 96.7% ±
2.4% loading efficiency. Thus, nearly all of the ORL loaded (20wt%) is retained
(19.3wt%) through the loading process and lyophilization.
The effect of ORL loading on NP size was investigated by loading different levels of ORL into NPs. NPs were loaded with 0, 5, 10, and 20wt% ORL using identical methods. These Nano-ORL formulations were then re-suspended in
H2O and hydrodynamic diameters were acquired using dynamic light scattering.
Figure 5-1B shows the linear effect on effective diameter of ORL loading, with an
R2 = .99 indicating that ORL loading has a direct influence on the size of Nano-
ORL. Hydrodynamic diameters range from approximately 290-580 nm. As the
Nano-ORL formulation with 20wt% loading has approximately 97% loading
123 efficiency, and formulations with lower ORL loading respond in a nearly perfect linear fashion to ORL, it is very likely that these lower ORL content Nano-ORL formulations have approximately the same loading efficiency.
Figure 5-1. (A) Extraction of ORL from empty NPs spiked with stock ORL resulted in a standard curve that is linear across the range of 0.05 to 0.20 mg/ml
(R 2 = 0.98). ORL extracted from Nano-ORL was found to be at 96.7% ± 2.4% of the theoretical value, giving Nano-ORL a composition of approximately 19.3wt%
ORL and 80.7wt% PBA-HLA. (B) Effective diameters, as determined by dynamic light scattering, were found to increase with increasing loading amounts of ORL. 124
Empty NPs were found to be in the range of 290 nm, increasing linearly up to
580 nm with approximately 20wt% ORL. This linear increase in diameter suggests that Nano-ORL size can be controlled using ORL content, and may allow us to optimize NP diameter for delivery to tumors in future studies.
Inhibition of FASN and Lipid Synthesis
Results in Figure 5-2A indicate that stock 10 µM extracted stock ORL and extracted Nano-ORL both inhibit approximately 93% of FASN activity. At 1 µM extracted stock ORL shows approximately 31% inhibition, and extracted Nano-
ORL shows 36% inhibition. At 0.1 µM extracted stock ORL shows 0% inhibition, and extracted Nano-ORL shows 7% inhibition. Extracted stock ORL and extracted Nano-ORL show similar FASN inhibtion over all concentrations measured, indicating that the Nano-ORL formulation is just as effective in inhibiting FASN activity as pure ORL. These results suggest that the process of loading and storing ORL in the Nano-ORL formulation is not causing degradation or resulting in lower ORL FASN inhibition.
Inhibition of lipid synthesis by FASN inhibition was then analyzed using
14 C-acetate uptake analysis. Acetate is metabolized upstream of FASN and eventually incorporated into the major FASN product palmitate, which is further metabolized into other lipids. Figure 5-2B shows measurement of this FASN activity in cells by treatment of cells with either empty NPs, free ORL, or Nano-
ORL, followed by addition of 14 C-acetate, which is incorporated into lipids.
Results indicate that lipid synthesis is reduced for cells treated with either free
125
ORL or Nano-ORL (ORL = 38%, Nano-ORL = 49% of the control). Neither free
ORL or Nano-ORL were found to be significantly different from one another [95%
CI: (-)5476 – 13862]. Empty NP vehicle was not found to be significantly different from control media [95% CI: (-)9192 – 10145], suggesting that the observed effect of Nano-ORL results solely from ORL incorporated into the NPs. All other groups were found to be significantly different from one another at α = 0.05.
Figure 5-2. (A) FASN was directly inhibited using ORL extracted from Nano-ORL or stock ORL that underwent the extraction protocol. The FASN inhibition of ORL extracted from Nano-ORL was similar to extracted stock ORL, indicating that
126
ORL activity is retained through NP loading and lyophilization. (B) Nano-ORL was tested for its ability to inhibit lipid synthesis in living cells. 14 C-acetate incorporation in lipids was measured when cells were exposed to empty NPs,
Nano-ORL, or ORL. Cells treated with empty NPs did not show a decrease in lipid synthesis, while both Nano-ORL and ORL inhibited lipid synthesis to a similar degree.
Cytotoxicity of Nano-ORL
Cytotoxicity of Nano-ORL was tested against human cancer cell lines PC3
(prostate), MDA-MB-23 (breast), U87mg (malignant glioma), and RKO
(colorectal) using CCK-8 assay. Cells were exposed to 25 µM ORL, 0.062 mg/ml
Nano-ORL (25 µM ORL theoretical concentration), 0.062 mg/ml empty NP, or normal media and allowed to incubate for 48 hours. As seen in Figure 5-3A-D both ORL and Nano-ORL show significant levels of toxicity, with relative viabilities ranging from 40-80% for ORL, and 21-57% for Nano-ORL. Nano-ORL was significantly more toxic than ORL in all cell lines tested (p < 0.0001). It is unlikely that this increase in toxicity is due to a toxic effect of the NP scaffold, as empty NPs show relative viabilities similar to normal media. Rather, it is possible that the increase in cytotoxicity of Nano-ORL is due to the increased solubility or stability of ORL due to its localization within NPs. ORL is very hydrophobic, and it was thought that prolonged exposure to cell growth media, even in the presence of fetal bovine serum, would potentially cause ORL to aggregate or hydrolyze and thus become less cytotoxic. To test this, we pre-incubated ORL, Nano-ORL,
127 and empty NPs for 24 hours at 37 °C prior to application to cells (designated with
PI superscript). Media was re-applied at 24 hours (again after 24 hours pre- incubation) to ensure a consistent level of pre-incubated drugs and materials, and plates were tested with CCK-8 assay. In Figure 5E-H we again observed that the Nano-ORL formulation is significantly more cytotoxic than the ORL formulation (p ≤ 0.0001 for MDA-MB-231, p ≤ 0.01 for PC and U87mg). ORL and
Nano-ORL were equally cytotoxic in RKO cells. As can be seen in the ORL PI groups however, ORL becomes much less cytotoxic after 24 hour pre-incubation in cell growth media, with relative viability of ORL PI increasing by 20-50% of the control group over ORL (p ≤ 0.0001 for all cell lines). In comparison, there was no significant difference between Nano-ORL and the pre-incubated Nano-ORL PI groups in any of the cell lines. No cytotoxicity was observed to result from empty
NPs. These results confirm that the Nano-ORL formulation results in significantly higher stability of ORL and thus may improve chemotherapeutic efficacy in vivo .
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129
Figure 5-3. 48 hour cytotoxicity analysis shows that the Nano-ORL formulation results in significantly lower relative viability of cancer cell lines PC3 (A), MDA-MB-231 (B), U87mg (C), and RKO (D) compared with free ORL. This result is not due to a toxic effect of PBA-HLA, as empty NPs exhibited no increase in cytotoxicity. Cells were then treated with ORL and Nano-ORL that had been pre-incubated for 24 hours prior cell exposure. It was observed that ORL was significantly less toxic after pre-incubation, while Nano-ORL did not lose efficacy, and both pre-incubated and non-incubated Nano-
ORL exhibited higher toxicities than free ORL and pre-incubated ORL in PC3 (E), MDA-MB-231 (F), and U87mg (G). RKO cells exhibited extremely high sensitivity to ORL (H). A schematic diagram of the experiment is shown at the bottom indicating when pre-incubations and media changes occur for each set of experiments.
130
Inhibition of 14 C-acetate uptake into lipids was then further examined using pre-incubated ORL and Nano-ORL in order to determine if pre-incubation of ORL had an effect on lipid synthesis. Conditions were identical to the previous study with the exception that two groups were added, free ORL or Nano-ORL receiving a 24 hour pre-incubation. Figure 5-4 shows that all groups containing free ORL or Nano-ORL inhibit lipid synthesis to a similar degree within each cell line, regardless of whether the formulation had been pre-incubated or not. Between groups containing a formulation of ORL, only Nano-ORL (pre-incubated) was found to be slightly higher than free ORL (not incubated), (95% CI: 1.8 – 39.8%).
However, these data result from level of lipid synthesis after a 17-18 hour exposure to Nano-ORL, indicating that the toxicity previously observed over 48 hours may be resulting from the additional 30 hours of exposure time.
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Figure 5-4. Pre-incubation of free ORL and Nano-ORL showed similar levels of lipid inhibition as shown by 14 C-acetate uptake. Groups containing either pre- incubated or non-incubated ORL or Nano-ORL were not found to be significantly different from each other, with the exception of pre-incubated Nano-ORL vs free
ORL (not pre-incubated), [95% CI: 1.764-39.84]. These results generally suggest that pre-incubation of ORL or Nano-ORL does not have an effect on lipid syntheses over the 17-18 hour exposure time tested here. This indicates that the previously observed toxicity over 48 hours may be resulting from the additional
30 hours of exposure, or from effects other than FASN inhibition.
Chemical Stability
In order to determine whether the maintenance of cytotoxicity after pre- incubation is a result of improved chemical stability or improved colloidal stability,
Nano-ORL and free ORL were incubated in PBS at room temperature for 96 hours, followed by extraction and analysis by mass spectrometry. It was hypothesized that exposure of ORL to aqueous solution may result in hydrolysis of the β-lactone ring, thereby reducing the FASN inhibition, while the Nano-ORL formulation may prevent this hydrolysis. However, as Figure 5-5 shows, the expected hydrolysis product (MW: 513.76) was not observed in either group.
Instead the mass spectrum shows ORL either protonated or associated with a positively charged ion (Na + or K +). This indicates that the increase in cytotoxicity of Nano-ORL is likely the result of increased solubility and availability to cells.
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133
Figure 5-5. MS analysis of extracted Nano-ORL (A) and free ORL (B) after 96 hour incubation in PBS. Identical results between (A) and (B), and the absence of a hydrolysis product at 513.76 indicate that ORL is not chemically degraded in either formulation in these conditions.
Mitochondrial Respiration Assays
To examine the consequences of ORL and Nano-ORL treatment on cellular metabolic function, MDA-MB-231 and LNCaP cells were treated with 50
μM ORL or 50 μM Nano-ORL (plus appropriate vehicles) for 16 hours then subjected to a Seahorse Mitochondrial Stress Assay ( Figure 5-6). Nano-ORL was able to reduce basal mitochondrial respiration by approximately 65% when compared to treatment with empty NPs ( Figure 5-6A, orange and purple lines prior to oligomycin injection, and Figure 5-6B). Basal respiration is a basic measure of O 2 use by cellular mitochondria. Likewise, Nano-ORL reduced coupling efficiency, which measures the amount of ATP turnover, by 58% when compared to treatment with empty NPs ( Figure 5-6C&G , orange and purple lines between FCCP and RTN/AA injections). The ability of Nano-ORL to significantly reduce mitochondrial function in MDA-MB-231 cells indicates that the cytotoxic effects observed mentioned previously are due, at least in part, to metabolic disruption.
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Figure 5-6. 40,000 MDA-MB-231 cells (A-D) or LNCaP cells (E-H) were plated in an XF24 well plate and allowed to adhere before treatment with Vehicle
135
(DMSO/EtOH), empty NPs, (NP), 50 µM free ORL, or 50 µM Nano-ORL for 16 hours. Oxygen consumption rates were measured (A&E) and values representing basal respiration (B&F), coupling efficiency (C&G), and spare respiratory capacity (D&H) were quantified with respect to cell number. Nano-
ORL reduced basal mitochondrial respiration (B) and coupling efficiency (C) in
MDA-MB-231 cell, and reduced coupling efficiency in LNCaP cells (G). ORL reduced coupling efficiency in both MDA-MB-231 and LNCaP cells. Nano-ORL was superior to ORL at reducing basal respiration in MDA-MB-231 cells, and overall showed a similar metabolic inhibition as ORL by decreasing basal metabolism and coupling efficiency by 50% or more. This suggests that Nano-
ORL has a strong ability to perturb mitochondrial function and reduce cellular metabolism in both breast and prostate cancer cell lines.
Previous in vitro and in vivo results have shown ORL to inhibit FASN, trigger apoptosis, and inhibit metastasis and angiogenesis. 206, 211, 214 Of particular interest in the investigation of Nano-ORL is the potentially broad applicability of
FASN inhibitors in the treatment of cancer. A growing body of evidence suggests that reliance on FASN in cancer is nearly ubiquitous and thus presents a valuable target for therapy. 193, 195, 197 Indeed, ORL has been used in pre-clinical investigations against a number of cancer types, including prostate, gastric, oral, skin, and leukemia. 206, 212-214, 253 In combination, the low bioavailability through enteric delivery, high degree of hydrophobicity, and potential for off-target effects
136 present significant barriers to clinical use of ORL and require novel developments in new ORL formulations.
Nano-ORL, as demonstrated here, is effectively solubilized by loading into the hydrophobic core of HLA NPs. The high loading efficiency observed with
Nano-ORL (96.7% ± 2.4%) shows a strong association between ORL and the NP core. While the observed effective diameters (400-600 nm) are above the ideal range for delivery through fenestrations in tumor vasculature, the increased solubility of Nano-ORL will likely increase total bioavailability and thus improve overall delivery to tumors. Additionally, modification of the polymer molecular weight or hydrophobic ligand content may allow for further decrease of NP diameter. The ability to control NP size based on ORL content will likely prove useful in optimizing biodistribution, circulation time, and ultimately delivery of
ORL to tumors.
The direct inhibition of FASN and decreased lipid synthesis ( Figure 5-2) demonstrate that Nano-ORL has the same molecular activity as free ORL. Nano-
ORL had similar or better FASN-TE inhibition compared to free ORL (Figure 5-
2A). This indicates that the molecular action of Nano-ORL is the same as that of free ORL, and the observed cytotoxicity results from FASN inhibition. Inhibition of lipid synthesis as indicated by 14 C-acetate uptake was also found to be similar between free ORL and Nano-ORL, further demonstrating that Nano-ORL has the same molecular action as free ORL and that Nano-ORL effectively inhibits new membrane synthesis in cancer cells.
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Analysis of the cytotoxic effect of Nano-ORL on prostate (PC3), breast
(MDA-MB-231), glioma (U87mg), and colorectal (RKO) cancer cell lines demonstrates the improved stability of the Nano-ORL formulation over free ORL
(Figure 5-3). 48-hour cytotoxicity analysis resulted in higher levels of toxicity from Nano-ORL as compared to free ORL ( Figure 5-3A). In order to show that this effect was resulting from improved stability of Nano-ORL and not a concentration effect, both ORL and Nano-ORL were pre-incubated for 24 hours prior to exposure to cells ( Figure 5-3B). This pre-incubation profoundly demonstrates the maintenance of the cytotoxic effect of Nano-ORL vs. free ORL, with the relative viability of free ORL-treated cells increasing by 20-30% relative to the media-only control. The relative viability of Nano-ORL-treated cells was not observed to change at all, and Nano-ORL was again found to be more cytotoxic than even non-pre-incubated ORL. This effect clearly demonstrates the potential advantages of NP encapsulation of ORL and lends further evidence towards the in vivo potential of Nano-ORL.
Interestingly, inhibition of lipid synthesis was not observed to be different over a 17 hour incubation when cells were treated with either pre-incubated or non-pre-incubated ORL formulations ( Figure 5-4). While this appears to be in contrast to the increased relative viability of cells when treated with pre-incubated free ORL, the 17 hour time point was chosen specifically to avoid complicating cytotoxic effects, thereby uncoupling the effects of cytotoxicity and lipid synthesis inhibition. Additionally, the inhibition assay was performed over 17 hours, while the toxicity assay was performed over 48 hours, which may provide additional
138 time for the effects of pre-incubation to manifest. Detailed metabolic analysis also shows that Nano-ORL decreases basal respiration and coupling efficiency equal to, or better than, free ORL ( Figure 5-6). This indicates that both the total O 2 consumed, and ATP turnover, are reduced in cells treated with Nano-ORL or free
ORL. Thus, both ORL and Nano-ORL exhibit a combination of reduced lipid synthesis and metabolic activity, which results in a cytotoxic effect against cancer cell lines. Nano-ORL, however, shows superior stability and solubility in aqueous solution, which results in higher cytotoxicity over extended periods of time.
Overall, these results show that Nano-ORL has the same molecular action as free ORL, but provides improved bioavailability due to the soluble NP formulation.
NP based drug formulations have previously demonstrated improved efficacy in the clinic, most notably in the forms of Doxil and Abraxane, and numerous pre-clinical studies are underway. 254 Our group and others have recently shown the potential for delivery of small hydrophobic molecules to tumors by physical entrapment within nano-scale polymeric micelles made with chemically modified HLA. 255, 256 Delivery with these materials have resulted in relatively high intra-tumoral concentrations of the selected payload, primarily due to the enhanced permeability and retention effect and improved drug solubility. In addition to the passive targeting of solid tumors, active targeting against cancer cells bearing HLA receptors such as CD44 remains a possibility. CD44 has been recognized as a potential target for acute myeloid leukemia, multiple myeloma, and chronic lymphocytic leukemia. 257-260 ORL has already been shown to be both
139 toxic against chronic lymphocytic leukemia primary cells, and substantially more toxic to these cells than healthy primary B cells. 253 More work is thus needed to investigate the potential for combination CD44 targeting with ORL delivery formulations for treatment of both solid and non-solid cancers.
Conclusions
This works represents the first investigation into a nanoparticle formulation of Orlistat. We have shown that this formulation, Nano-ORL, retains the FASN inhibition of free ORL through direct FASN inhibition and that this results in decreased lipid synthesis in cells. Nano-ORL also provides longer lasting cytotoxic efficacy in vitro , likely through improved serum stability and availability.
This also results in modulation of the metabolic activity of cells as observed by mitchondrial stress test analysis. These results should encourage further investigations into NP formulations of ORL, and for NP formulations of other
FASN inhibitors as well.
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CHAPTER 6
CONCLUSIONS AND FUTURE DIRECTIONS
Summary of NanoICG Studies
In this work we have shown that nanoscale self-assembled polymeric micelle formulations of modified HLA polymers can be used as delivery systems for ICG and ORL. A novel hydrophobic conjugate, PBA, was synthesized and conjugated to HLA, resulting in self-assembly in aqueous solution and entrapment of both ICG and ORL. Comparison to previously examined hydrophobic conjugates 5βCA and ODA showed that hydrophobic conjugates with roughly similar structures (PBA and 5βCA) provided similar levels of ICG loading, whereas octadecylamine conjugates exhibited lower ICG loading and also required lower conjugation ratios for soluble micelle formation. Loading of
ICG within the hydrophobic regions of NPs resulted in improved solubility in salt solution, and this process was dependent on the ICG loading step. These formulations were non-toxic at physiologically relevant concentrations. NanoICG was not found to actively target CD44+ tumor cells, suggesting that any increase in signal seen in vivo would be the result of passive targeting resulting from NP accumulation. In vivo results using NanoICG to highlight tumors suggested that
NanoICG provides approximately twice the contrast between tumor tissue and muscle at 24 hours compared to injection of free ICG making this formulation a candidate for further investigations into its usefulness as an image guided surgery contrast agent.
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Biodistribution and ICG Release to Serum Proteins
Further work is necessary to elucidate the details regarding the release profile of ICG in serum-mimicking flow conditions (Figure 3-7). Dissolution of
NanoICG into 37 mg/ml BSA-PBS solution (an average physiological concentration) resulted in rapid release of ICG from NPs to serum proteins. This release occurs to a significant extent immediately upon dissolution in BSA-PBS solution, and increasing levels of volumetric flow result in increasing ICG release to serum proteins. The WSS in an idealized cylindrical vessel at the liquid-solid boundary, where shear stress is highest, is shown in Equation 6.1:
[6.1] 32 =
2 where τ ω = wall shear stress (dyne/cm ), Q = volumetric flow rate, µ = viscosity,
D = diameter. Theoretical values calculated using Equation 6.1 for the flow rates in this experiment range from 0.287-2.58 dyne/cm 2. Using these values will result in underestimation of the actual shear stress experience by the NPs as the peristaltic pump action will result in transiently higher shear values as the fluid passes the pump head. Thus, a safe approximation for maximum shear stress experienced in this experiment is on the order of 10 dyne/cm 2. However, physiological shear stress values range from 50 dyne/cm 2 in the aorta, 2.8-100 dyne/cm 2 in capillaries, and 17-211 dyne/cm 2 in arterioles, and even as high as
600 dyne/cm 2 in mouse aortas. 91-94 Thus, the shear stress experienced in this system is a gross underestimate of shear stress experienced in humans or the in vivo model used here, meaning ICG release should actually release more rapidly. Additionally, we found that dissolution of either ICG or NanoICG in BSA- 142
PBS followed by elution through a 100 kDa centrifugal filter kit resulted in similar elutions of ICG and BSA, indicating that elution was not dependent upon the initial presence of ICG in the NP formulation (Figure 3-8). These results suggest that ICG is being released from NanoICG and should thus behave similarly to free ICG in vivo .
Contrary to the in vitro results, NanoICG appears to behave very much like a NP in its biodistribution and clearance in mice (Figure 4-1 & 4-4). In vivo results show that two different NanoICG formulations both result in improved
SNR and CNR compared to free ICG at 24 hours. Kidney signal is higher for ICG at 4 hours, and similar to NanoICG at 24 hours. These results are consistent with small molecule clearance for ICG and NP clearance for NanoICG. Additionally, this work was replicated in two separate animal experiments and with consistent results. It thus appears that ICG within NPs, while it does release to serum proteins, still remains within the NPs for some time. This is shown by the gradual increase in the release profile and likely causes the difference in biodistribution. It has also been shown that smaller molecules or polymers diffuse more easily from tumors back into circulation, so that the larger NanoICG particles would be retained in tumors longer than ICG. 97 One possible method to examine this in vitro would be to use tumors grown in decellularized tissues perfused with peristaltic pumps. This would allow for the combination of natural-like tumors in controlled flow conditions, and would afford precise control of solution composition such as NP and protein concentration. This approach would prevent
143 the complications of metabolism and excretion, and could thus focus solely on release, quenching, flow, and diffusion from tumors.
An additional complicating factor is the total ICG signal in mice at 4 hours, which shows ICG as being present in higher quantities in all tissues. ICG content of all contrast agents was measured with absorbance spectrometry prior to administration, so this is not likely to be a result of higher ICG administration levels. A possible explanation for this is the quenching phenomenon previously observed when ICG is closely bound within NanoICG. As the biodistribution indicates that ICG is retained within NP cores it follows that fluorescence quenching would also occur, which would reduce the observed signal. A future pharmacokinetic in vivo study is necessary in which ICG is extracted from tissues at various timepoints and more rigorously quantified in order to prevent systematic error due to quenching of NanoICG. The ICG release to serum proteins should also be further examined to determine to what extent ICG is released from NanoICG at various timepoints, and what effect solution composition has on this release.
Though NanoICG shows significantly higher CNR than free ICG, it may be possible to improve this further by controlling ICG release from the NPs. In order to prevent the loss of ICG to serum proteins, chemical modifications to NanoICG will be necessary. The most obvioius potential modification is PEGylation, which as been clearly demonstrated to reduce interactions with serum proteins in numerous studies. 86 This type of modification could be performed using the same chemistry required for conjugation of the hydrophobic groups to HLA, namely
144 amide bond formation between a primary amine on a PEG molecule and the carboxylic acid on HLA by EDC/NHS reaction. This method has been used previously with Cy5.5 labeled HLA, where PEGylation resulted in increased retention within the tumor and within the entire body, as well as lower relative uptake by the liver. 50 Another strategy to prevent loss of the fluorophore from the
NPs is through direct conjugation to HLA. Previous studies have investigated direct conjugation of ICG-OSU, an N-hydroxysuccinimide activated form of ICG, as well as direct conjugation of Cy5.5. 50, 228 Furthermore, chemical crosslinking of the HLA polymers could prevent dissociation and ensure that NPs remain intact during tumor accumulation. However, direct conjugation of fluorophores to HLA requires the use of non-FDA approved materials, significantly reducing the ease and speed of clinical translation. An alternative strategy to avoid this complication would be to test alternative NP formulations by utilizing different polymers. Other groups have successfully loaded ICG into PLGA or PLA based NPs, and a number of polymeric or liposomal formulations may need to be investigated in order to acquire optimal release kinetics and biodistribution. 220, 233, 261-263
An IGS-Specific Animal Model is Needed
Further addressing the in vivo issues in IGS is the choice of animal model.
New tumor models specifically considering the experimental requirements of IGS are needed. Mice provide high sample sizes but their physical dimensions prevent adequate separation of metabolic and excretory organs from target tissues (Figure 4-7). Additionally, for breast cancer models in particular, mice
145 provide little breast tissue with which to examine contrast between tumors and healthy tissue. Subcutaneous tumor models in which the tumor is placed in the back hip, as in Figure 4-3, provide the best optical protection of the tumor signal from the signal of visceral organs such as the liver and kidneys, but this location does not provide tumor orthotopic growth as would be found by injection into breast tissue. The studies demonstrated here have demonstrated the superior tumor distribution of NanoICG, but future in vivo studies will require survival surgery and direct comparison of contrast between target tissues. Thus, a different animal model, using a larger animal with orthotopic tumors, is necessary to gain a complete understanding of expected contrast between tumors and various tissues in a surgical setting. Some investigations into this have been performed in spontaneously occurring tumors in canines, which provides advantages in both animal size as well as natural tumor formation that provides realistic contrast agent distribution. 230 However, this method is not reasonable for routine work due to limitations on animal patient populations. One appropriate animal may be rats, which are substantially larger than mice while smaller than many other research species.
Extension of NanoICG to SLN Mapping
The current gold standard for SLN mapping involves radioisotope probes or the visual dye Patent blue. 177 ICG alone has previously shown success as both a color and fluorescent dye in SLN mapping, with rates comparable to radioisotope probes. HLA has also shown success in improving SLN mapping by
146 mixing with Patent blue or SPIO NPs, as free HLA increased the viscosity of the solution, which prolonged the duration of stain within the SLN. 62 NanoICG clearly shows increased retention within the tumor over free ICG, and would likely serve equally well or better than free ICG to image SLNs. Additionally, the ratio of ICG to HLA could easily be tailored to result in a higher viscosity solution and possibly decrease the quenching effect, which could further improve the retention and signal present for quick imaging. Using NIR imaging with ICG would serve both to provide surgeons with visual feedback as opposed to lymphoscintigraphy, and would decrease the radiation exposure of both patients and operating room staff.
Summary of Nano-ORL Studies
The Nano-ORL work presented here represents the first NP formulation of
ORL. ORL was efficiently loaded into the NP cores (Figure 5-1A), which demonstrates its highly hydrophobic nature. Nano-ORL was also found to inhibit
FASN and block lipid synthesis in vitro , indicating that the molecular action of
Nano-ORL is the same as that of ORL, and that the NP loading process does not cause ORL degradation ( Figure 5-2A&B ). Toxicity against cancer cell lines was found with both Nano-ORL and ORL, except that Nano-ORL appeared to be more toxic than free ORL over 48 hours ( Figure 5-3). This additional toxicity was not due to negative effects of the NP material itself, as empty NPs were not found to be toxic. Thus, the greater cytotoxicity must result from a synergistic effect of the Nano-ORL formulation. It was hypothesized that the Nano-ORL formulation provides better aqueous stability of ORL, thus increasing the total
147 exposure to the cells. Pre-incubation of both ORL and Nano-ORL resulted in greatly reduced cytotoxicity of ORL, while the toxicity of Nano-ORL was not effected. This indicates that the aqueous stability of ORL is increased when loaded into the hydrophobic cores of NPs. Further analysis indicated that the toxic effect of both ORL and Nano-ORL resulted in decreased metabolic activity of the cells due to perturbed mitochondrial function (Figure 5-7).
Diameter of Nano-ORL
Further work is needed with regard to the average NP size. Figure 5-1B indicates that the average diameter of Nano-ORL loaded with 20 wt% ORL is approximately 600 nm, while that of 5 wt% is approximately 400 nm. These diameters are too large to fit through the vascular fenestrations of many tumors.
Though many complicating and competing factors must be addressed in NP biodistribution, NPs of 100 nm in diameter have been found to be appropriate in many situations. 82 For delivery to solid tumors, it will thus be preferable to develop a formulation with signficantly smaller average diameters than what we have observed with 10 kDa High-PBA-HLA. Alternatives exist in the form of other polymeric or liposomal formulations, as these have commonly been used in NP drug delivery as outlined previously, however the exact composition will need to be determined experimentally.
Despite the non-optimal diameters, the increased solubility of Nano-ORL will likely increase total bioavailability and thus improve overall delivery to tumors.
Additionally, modification of the HLA molecular weight or hydrophobic ligand
148 content may allow for further decrease of NP diameter. Additionally, the ability to control NP size based on ORL content will likely prove useful in optimizing biodistribution, circulation time, and ultimately delivery of ORL to tumors.
In Vivo Testing of Nano-ORL
It is desirable to begin in vivo studies using Nano-ORL. The improved aqueous stability and prolonged activity observed are sufficient reason for initial investigations into the anti-tumor activity in living animals. No previous in vivo experiments have utilized systemic injection of ORL, 211-213 and thus both efficacy and toxicity studies are needed. Possible side effects of ORL administration are likely to be similar to those observed with C75, including weight loss and anorexia, 195, 201 , and ORL is known to be active against pancreatic lipases,
FASN, lipoprotein lipase, and potentially GADPH and β-tubulin as well. 253, 264
However, despite these potential complications, the potential beneficial outcomes of systemic administration of a solubilized form of ORL must be addressed. The in vitro results presented in this work demonstrate significant and meaningful improvements in the action of ORL against cancer cells. Additionally, previous
NP formulations of other drugs, such as Doxil ®, have resulted in lower off-target toxicities compared to the free drug, so it is reasonable to expect that a NP formulation of ORL could exhibit this behavior as well.
149
HLA Remains a Suitable Biopolymer for NP Research
HLA has been, and remains, a useful biopolymer for the synthesis of nanoscale polymeric micelles. However, HLA has both positive and negative aspects for use as a NP backbone. Benefits include the biodegradability, hydrophilic nature, potential for active targeting against CD44, multivalency for hydrophobic moiety, stealth moiety, drug, or targeting ligand conjugation.
Detriments include the molecular effects of different molecular weight HLA polymers, as large MW HLA is antiangiogenic and immunosuppressive, while small (20 kDa) MW HLA is angiogenic and inflammatory. 265 The polydispersity of both natural and commercial HLA, and the multivalency resulting in non-location- specific conjugation also result in a diversity of similar products formed in the same reaction. Simply, HLA is difficult to precisly control on a molecular level. As such, its limitations and variations should be recognized and addressed prior to utilizing HLA in new formulations. Nonetheless, it remains a highly useful polymer for NP formulations and the applications will doubtlessly expand.
Conclusion
The field of nanomedicine is expanding rapidly and for good reason.
Nanomedicines have shown numerous advantages over their free drug counterparts both pre-clinically and clinically, a number of which are evidenced here in the form of superior tumor contrast from ICG delivery and prolonged cytotoxic effects with ORL. Though there are significant hurdles in the delivery of
NPs to tumors, in many ways what remains is an optimization problem rather
150 than proof of efficacy. Optimization must occur in the form of NP size, stability, drug release rate, charge, and targeting. This will require novel contributions in the form of combinations and modifications of materials. Material choice will depend on the drugs or contrast agents being delivered, which in turn depend on the cellular and extracellular biochemical and biophysical environment that is being targeted. Related to the field of personalized genetic medicine, one can envision a field of personalized nanomedicine in which optimized NP formulations may be chosen based on an individual tumor’s expected attributes, such as vascularity, fenestration size, and predicted susceptibility to various chemotherapeutics. This could allow better outcomes at every stage of the cancer treatment process, from improved diagnosis, staging and restaging, optimized visualization during surgical resection, to improved chemotherapeutic outcomes.
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APPENDIX
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TANNER K. HILL CURRICULUM VITAE
Virginia Tech/Wake Forest University School of Biomedical Engineering and Sciences 1440 Brookwood Dr. Winston-Salem, NC. 27106 (336) 831-4301 [email protected] [email protected]
EDUCATION
2015 Ph.D. , Biomedical Engineering, Virginia Tech/Wake Forest University, Winston Salem, NC. (exp. May 2015)
2009 B.S. , Biological Sciences, Magna cum Laude, Colorado State University, Ft. Collins, CO. Minors: Biochemistry Physics, Mathematics
RESEARCH EXPERIENCE
2012-2015 Graduate Student Synthesis, characterization, in vitro and in vivo testing of polymeric hyaluronic acid nanoparticles for image guided surgery and drug delivery Virginia Tech/Wake Forest University Principal Investigator: Aaron M. Mohs, Ph.D.
2010-2012 Graduate Student Synthesis, characterization, in vitro and in vivo testing of electrospun collagen-poly-(caprolactone) tissue engineered vascular scaffolds for treatment of peripheral vascular disease Virginia Tech/Wake Forest University Principal Investigator: Sang Jin Lee, Ph.D.
2008-2009 Undergraduate Research Assistant Cloning of a recombinant plasmid for the purpose of expressing H4 histone with reversed amino acid sequence in tail domain for protein- protein interaction analysis Colorado State University Principal Investigator: Jeffrey Hansen, Ph.D.
2007-2008 Undergraduate Research Assistant
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Computational modeling of predator/prey interaction in two dimensional space. Computational modeling of Los Gatos Research Methane Analyzer using Matlab Colorado State University Principal Investigator: Colleen Webb, Ph.D., Joseph von Fischer, Ph.D.
SCIENTIFIC EXPERIENCE
Analytical Absorbance/fluorescence spectroscopy, dynamic light scattering, NMR, SEC, mass spec analysis, SEM
Synthetic Drug formulation chemistry, nanoparticle self-assembly, organic synthesis and purification, bioconjugation, electrospinning
Biological Mammalian cell culture, cytotoxicity assays, histology, fluorescence and light microscopy, flow cytometry, bioreactor conditioning, Southern/Western blots, PCR
Translational Design of experiments, data and statistical analysis, novel assay design and implementation, image guided surgery, small and large animal surgery, small animal tumor models, stress/strain testing, bioreactor design, detailed record keeping
SCHOLARSHIPS AND AWARDS
2014-2015 Mike and Lucy Robbins Fellowship Award Winner, Wake Forest University Graduate School of the Arts and Sciences
2010-2014 Graduate Research Assistant, Virginia Tech/Wake Forest University School for Biomedical Engineering, Summer
2008-2009 FEScUE Undergraduate Computational Modeling Fellowship, Colorado State University
2007-2008 First Year Undergraduate Physics Scholarship, Colorado State University
2006-2007 Colorado State University Undergraduate Scholarship
LEADERSHIP AND SERVICE
2013-2014 Biomedical Engineering Society Student Chapter President, Wake Forest University
2013-2014 Volunteer, Wake Forest Baptist Health
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2013-2014 Volunteer, Community Care Center of Winston-Salem
2011-2015 Undergraduate and Summer Student Mentor
2010-2015 Biomedical Engineering Society Member
AUTHORSHIPS AND PRESENTATIONS
Journal Publications
Indocyanine Green-Loaded Nanoparticles for Image Guided Tumor Surgery. Tanner K. Hill, Asem Abdulahad, Sneha S. Kelkar, Frank C. Marini, Timothy E. Long, James M. Provenzale, Aaron M. Mohs . Bioconjugate Chemistry 2015, 26(2), 294-303. http://pubs.acs.org/doi/abs/10.1021/bc5005679
Fabrication and Characterization of Medical Grade Polyurethane Composite Catheters for Near-Infrared Imaging. Andre T. Stevenson Jr., Laura M. Reese, Tanner K. Hill , Jeffrey McGuire, Aaron M. Mohs, Raj Shekhar, Lissett R. Bickford, Abby R. Whittington. Biomaterials 2015, 54, 168-176. http://www.sciencedirect.com/science/article/pii/S0142961215002902
Oral and Poster Presentations
Development of a Nanoparticle Formulation of Orlistat: Solubility, Stability, and Cytotoxicity. Tanner K. Hill , Frances B. Wheeler, Amanda L. Davis, Sneha Kelkar, Steven J. Kridel, Aaron M. Mohs . BMES Annual Meeting, poster presentation, 2014/10/24
Synthesis, Characterization, and Pre-Clinical Evaluation of Hyaluronan Based Nanoparticles for Image-Guided Surgery. Tanner K. Hill , Sneha Kelkar, Frank C. Marini, Edward A. Levine, James M. Provenzale, Aaron M. Mohs. Annual SBES Symposium, poster presentation, 2014/5/15
Synthesis, Characterization, and Pre-Clinical Evaluation of Hyaluronan Based Nanoparticles for Image Guided Surgery. Tanner K. Hill , Sneha Kelkar, Frank C. Marini, Edward A. Levine, James M. Provenzale, Aaron M. Mohs. 2014 Duke University Fitzpatrick Institute for Photonics Annual Symposium. Poster presentation, 2014/3/11
Near Infrared Fluorescent Polysaccharide-Based Nanoparticles for Image-Guided Tumor Surgery. Tanner K. Hill , Sneha Kelkar PhD, Frank C. Marini PhD, Aaron M. Mohs, PhD. 2013 NCI Alliance for Nanotechnology in Cancer Annual Principal Investigators Meeting. Podium presentation, 2013/9/19
Development of Hyaluronan-Based Nanoparticles for Intraoperative Breast Cancer Imaging. Tanner K. Hill , Sneha Kelkar, Asem Abdulahad, Timothy Long, Frank C. Marini, Aaron M. Mohs. 2013 Annual SBES Symposium. Poster presentation, 2013/5/16
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Near Infrared Image Guided Tumor Removal: Epidermal Growth Factor as a Target. Tanner K. Hill , Frank C. Marini, Aaron M. Mohs. 2012 BMES Annual Meeting. Podium presentation, 2012/10/27
An Innovative Approach to Building Clinically Relevant sized Tissues and Organs. Tanner K. Hill , Jaehyun Kim, Peter Masso, Sang Jin Lee, Karl-Erik Andersson, Anthony Atala, James J. Yoo. Poster presentation, 15 th Annual Hilton Head Workshop: Regnerative Medicine, Innovations for Clinical Applications, hosted by Georgia Tech. March 18, 2011
Book Chapters
Electrospun Nanofibers in Tissue Engineering. Mitchell R. Ladd, Tanner K. Hill , James J. Yoo, Sang Jin Lee. Chapter 16 of “Nanofibers – Production, Properties and Functional Applications”. ISBN 978-953-307-420-7, InTech Publishing. 2011
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