Smart Brachytherapy Spacers for Combined Chemo-Radiation Therapy: Local Delivery of Nanoparticles, Chemotherapeutics, and Molecular Inhibitors for Cancer Treatment

Smart Brachytherapy Spacers for Combined Chemo-Radiation Therapy: Local delivery of nanoparticles, chemotherapeutics, and molecular inhibitors for cancer treatment

A Dissertation Presented

Jodi Elizabeth Belz

The Department of Bioengineering

in partial fulfillment of the requirements for the degree of

Doctor of Philosophy

in the field of

Bioengineering

Northeastern University Boston, Massachusetts

May 2017

ABSTRACT

Prostate cancer remains the second leading cause of cancer related deaths in men with over 161,360 new cases and 26,730 associated deaths in the U.S. alone in 2017. While standardized screening has become routine and allowed for early stage detection, prostate cancer treatment options are scarce and often leave survivors with reduced quality of life due to off-target side effects. We have utilized technology in the standard clinical procedure Brachytherapy to deliver a new local chemotherapy implant with the standardized radiation treatment without the need for additional procedures. Brachytherapy uses plastic inert spacers to help clinicians place and separate the radioactive seeds. We have designed a biocompatible and biodegradable spacer that provides this spatial guidance, with the added therapeutic benefit of sustained chemotherapy and radiosensitization locally at the tumor site to sensitize the cancer throughout the course of brachytherapy rather than intermittent radiosensitization experienced with systemic chemotherapy administered every three weeks. These ‘smart brachytherapy spacers’ can be modified to tailor their release of docetaxel to coincide with varying half-lives of any radioactive seed used in brachytherapy.

In this work, I have developed, characterized, and extensively tested a docetaxel

loaded biodegradable implant for the treatment of prostate cancer. Our spacers were

fabricated with a docetaxel loaded Poly(lactic-co-glycolic) acid cylindrical implant for

intratumoral injection via an 18 gauge applicator needle for local, sustained therapy. Our

spacers exhibit diffusion driven release in vitro over 75 days, designed to sensitize I-125

(t 1/2 = 60 days) brachytherapy seeds most commonly used for treatment of prostate cancer.

The spacers were tested for therapeutic efficacy against clinically administered docetaxel iii and resulted in significant tumor inhibition and improved survival (median survival time

(MST) of spacers 52 days versus 26 with IV DTX, p<0.01). Next the docetaxel spacer was combined with fractionated radiation therapy at reduced doses, to determine the radiosensitization and synergistic therapeutic response. Mice treated with local combined chemo-radiation resulted in significant survival improvement (MST 209 days vs. 120 in radiation therapy alone and 85 in spacers alone, p<0.01) and tumor inhibition, with 33% of mice cured. These results combined with a full toxicity study were completed and prove the therapeutic potential for successful clinical translation and impact.

Acknowledgements

I would like to first and foremost acknowledge my parents and siblings for their

unconditional love and support. Mom and Dad, I have never gone a day unloved and for

that I am forever grateful. This dissertation is dedicated to you. Thank you for always

supporting me and encouraging me to be the best version of myself. I love you and you are

my dearest inspirations.

To my advisor, Dr. Srinivas Sridhar. I sincerely thank you. You have been patient,

giving, and supportive. Working with you has awarded me more opportunities than most

and I am grateful for all you have taught me. We have travelled across the world and back

and genuinely thank you for everything you have done for me.

To my second family, the Sridhar Lab. Paige and Rajiv, you have made this journey

an unbelievable and worthwhile experience. I have learned so much from both of you.

Thank you for your guidance, support, suggestions, and most importantly your laughter.

You made every day more enjoyable and I will miss not working alongside you every day.

You are not just colleagues, but true friends. Rita and Tim, the lab would not run without

either of you so thank you for keeping us going. Noelle, Anne, Yuan, and Tej, you have

been tremendous help in this work and I thank you for your contributions.

I would like to thank our numerous collaborators both at Dana Farber Cancer

Institute and Michigan State. Specifically, Robert Cormack for your large role, I appreciate all of your feedback as a member of my committee and of your enthusiasm and support on v this project. Your help has been instrumental in its success. I would also like to thank

Ravina, Dolla, Aniruddha, and Houari who have sacrificed many weeks and weekends to help make sure my radiation treatments were completed on time. Karen Liby, you were a wonderful and kind collaborator who I hope to work with again in the future.

Lastly, I would like to acknowledge the patients for whom we worked so hard for: those lost, surviving, and thriving. This past year, I received devastating news about a friend who lost their battle to prostate cancer. The implications of my work have never seemed so important. So I thank you and your families for sharing your battle and your continuous courage and fight. You are why I am so passionate about my work and you are the truest form of inspiration.

LIST OF TABLES

Table 1. TNM staging system for prostate cancer...... …………..……………………….8

Table 2. Commonly used radiation sources for brachytherapy ………………………...... 20

Table 3. Listed copolymer properties of PLGA available for purchase (Sigma)………... 59

Table 4 Design of preliminary in vivo LCRT experiment to determine appropriate dosing for synergistic therapeutic effects …………………...………………………………… 100

Table 5. Treatment plan for toxicity study showing time of sacrifice and number of animals per group ……………………………………………...……………………………….. 117 Table 6. Talazoparib implants decrease tumor size and extend survival in BRCA1-deficient mice………………………………………………………………………………..……140

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LIST OF FIGURES

Figure 1. Male Anatomy showing physiological location of prostate located below bladder, around urethra, and laterally to rectum...... ………………………………...…………….4

Figure 2. Zonal depiction of the prostate. The prostate is divided into four regions, the central zone, transition zone, peripheral zone, and anterior fibromuscular stroma (AFS); ED: ejaculatory ducts, SV: seminal vesicles…………………………...... ………….5

Figure 3. Gleason Grading System from 1 to 5, with grade 5 having the worst prognosis. The pathologist examines the biopsy sample of the prostate tissues microscopically, identifies any malignant areas, and assigns a score to each one. The first score is the predominant cancer type present, while the second number is the secondary pattern…... 10

Figure 4. Brachytherapy for prostate cancer using transrectal ultrasound guided imaging. Dozens of applicator needles are used to guide radioactive seeds into or near the tumor and placed using careful dose mapping of the tumor………………………………………… 15

Figure 5. Software imaging used to map out exact placement of applicator needles and brachytherapy seeds for complete dose mapping of the tumor while sparing radiation to healthy tissues prior to procedure……………………………………………………….. 21

Figure 6. Post-operative x-ray image showing permanent radiopaque brachytherapy seeds in prostate………………………………………………………………………………...22

Figure 7. Wientjes et al show spatial drug distribution in the prostate as a function of distance from the injection site after dogs were given intraprostatic infusions of doxorubicin (0.3m/150 AL over 150 minutes). Results show high drug concentrations are achievable, yet uneven distribution was attained…………………………………………23

Figure 8. Poly D,L-lactic-co-glycolic acid, where x and y represent the number of times each monomer repeats……………………………………………………………………24

Figure 9. PLGA hydrolysis results in lactic acid and glycolic acid oligomers…………..26 Figure 10. Taxanes stabilize microtubules, preventing mitosis, leading to eventual programmed cell death or apoptosis……………………………………………………...30 viii

Figure 11. Paclitaxel and Docetaxel are structurally identical other than a difference in two functional groups which are highlighted in gray………………………………………….31

Figure 12. Size distribution of SiNP acquired from A) DLS. Data shows Gaussian distribution of silica nanoparticles with hydrodynamic diameter of ~36nm for SiNP encapsulating either Cy7.5, Cy5, or DTX and B) TEM image shows uniform silica nanoparticles with diameter of 30nm……………………………………………………..46

Figure 13. Stability of Silica Nanoparticles with docetaxel after fabrication, dialysis, and filtering showed no significant change in size during storage for 7 days in 4°C………...47

Figure 14. Release profile of docetaxel from silica nanoparticles in buffer (37°C, pH 7.4) over 6 days………………………………………………………………………………..48 Figure 15. PC3 cell viability after incubation with DTX loaded SiNP for 24, 48, and 72 hours. IC50 values are reported in molar concentration…………………………………..49 Figure 16. Confocal image of PC3 cellular uptake of Cy5 loaded silica nanoparticles (red) and RNA (blue) after 2 hr and 48 hr incubation…………………………………………50

Figure 17. Surviving fractions of empty SiNP, free docetaxel (DTX), and docetaxel loaded silica nanoparticles versus radiation at doses from 0-8 Gy using a small animal radiation research platform…………………………………………………………………………51

Figure 18. Schematic of INCeRT dual release platform with PLGA matrix embedded with Silica NPs loaded with chemotherapeutic or imaging moieties…………………………..54

Figure 19. Schematic of INCeRT spacer applications in brachytherapy for prolonged radiosensitization. Spacer aimed to have slow sustained therapeutic release throughout course of radiation treatment (blue curve) at tumor site rather than intermittent radiosensitization currently experienced with intravenous treatment administered every three weeks (black)………………………………………………………………………56

Figure 20. Photograph of modified smart brachytherapy spacers INCeRT-1 and INCeRT- 2. INCeRT spacers are loaded with silica nanoparticles and near infrared dye Cy7.5 in PLGA matrix that is morphologically similar to commercial brachytherapy spacers for theranostic applications…………………………………………………………………..60

Figure 21. Cross sectional and lateral views of INCeRT spacer using SEM. Chemical analysis of homogenous spots confirm silica nanoparticle pockets embedded throughout spacer using EDS…………………………………………………………………………61

Figure 22. (a) Illustration of intradermal spacer diffusion. As the spacer degrades, fluorescent silica nanoparticles are released for a slow sustained diffusion. Further, drug encapsulated in the nanoparticle can be released. A TEM image of (b) 30 nm and (c) 200 nm silica nanoparticles. (d) A series of SEM images of a flash frozen and fractured spacer ix with uniform distribution of nanoparticles. As magnification of the cross sectional area is enhanced, silica pockets are observed……………………………………………………65

Figure 23. (a) Schematic and (b) photograph of the fluorescence imager used to acquire image sequences for this work (see text for details). M, mirror; LP, long pass filter; BP, band pass filter; WL, white light………………………………………………………….68 Figure 24. (a-d) Fluorescence image sequence of a representative phantom with AlexaFluor 750 dye diffusion at 1 minute, 15 minutes, 30 minutes, and 60 minutes respectively. (e-h) Fluorescence image sequence showing 30 nm nanoparticle phantom diffusion for 1 minute, 6 days, 11 days, and 15 days. (i-j) Fluorescence image sequence showing 200 nm nanoparticle phantom diffusion for 1 minute, 6 days, 11 days, and 15 days. All images are normalized to the maximum at minute 1………………………………………………….74 Figure 25. Diffusion (fluorescence) as a function of distance from the phantom center, averaged over all phantoms imaged (N=3) and normalized to the intensity at x=0 for each time point. Data for (a) free AF750 dye, (b) 30 nm nanoparticles, and (c) 200 nm nanoparticles are shown. (d) Average maximum fluorescence intensity normalized to initial time point per phantom type showing the continual decrease in maximum fluorescence over time. Inset time scale of free AF750 dye diffusion is on the order of minutes compared to days in part (d). (e) Diffusion coefficients found by fitting the diffusion profile curves to Eq. 1 on a logarithmic scale…………………………………..75

Figure 26. Example in vivo fluorescence image sequence of Cy 7.5 spacers. (a) White light image, (b) normalized fluorescence image, (c) normalized fluorescence image overlaid on white light image at minute 1, and (d) Overlaid fluorescence image acquired on day 15 after dissection of spacer from the mouse………………………………………………...76

Figure 27. Example in vivo image sequence of Cy 7.5 spacers in row 1, 30 nm NP spacers in row 2, and 200 nm NP spacers in row 3. Each row represents the changes over the 15 days of a specific spacer. The images in each row are normalized to minute 1……….77

Figure 28. Averaged fluorescence diffusion curves over all spacers imaged (N=8), normalized to the intensity at x=0 per time point for (a) Cy7.5 dye spacers, (b) 30 nm nanoparticle spacers, and (c) 200 nm nanoparticle spacers. (d) The average maximum spacer fluorescence intensity normalized to initial time point showing the continual decrease in maximum fluorescence over time. (e) The fluorescence area imaged (full width at half maximum) around the spacer as a function of time, averaged over all trials. As indicated, diffusion occurred continuously over time, with the maximum diffusion area observed on day 4 for 30 nm NPs, day 6 for 200 nm NPs, and continuously increasing area observed for free dye……………………………………………………………………..78

Figure 29. Docetaxel loaded implant a) schematic shows mechanism of intratumoral release as polymer slowly degrades. B) SEM images show smooth surfaces intercalated with nano-size pores, and C) release of docetaxel show short burst release (~30% in 6 days) followed by extended linear release for 75 days………………………………………….88

Figure 30. Release of docetaxel from PLGA implants of different lengths (n=3)………...90

Figure 31. Average tumor response in mice treated with DTX (18 mg/kg) implants and intravenously compared to untreated control and control vehicle. Individual animal tumor response is plotted against control mice for each group above……………………………93

Figure 32. Kaplan-Meyer survival curve of mice in LCT study……………………….... 94

Figure 33. Representative pictures of treated and untreated mice, with emphasis on the ulceration (week 4) and tumor-free mouse (week 8) post treatment with DTX spacers…..95

Figure 34. Weight change (%) in LCT groups after treatment for determination of gross toxicity…………………………………………………………………………………...97

Figure 35. Efficacy of combined chemo-radiation therapy using DTX-spacers implanted intratumorally in prostate cancer animal model. The graph shows the change in tumor volume when different treatment combinations were studied in tumored animals. DTX- spacers containing 250 µg of total DTX (~9 mg/kg body weight of mice) was implanted in mice (red arrow) and radiation dose of 5, 10 and 15 Gy were given 13 days post implantation of spacers (black arrow)…………………………………………………104

Figure 36. Toxicity assessment in prostate cancer tumored animals treated with combined chemo-radiation therapy using DTX-spacers and external beam radiation. The percent change in body weight was measured for each group of mice receiving different treatment combinations……………………………………………………………………………105

Figure 37: Combined chemo-radiation therapeutic efficacy studies with the DTX spacers implanted intratumorally in subcutaneous prostate cancer tumored animals (n=7 per group) and fractionated radiation dose of 10 Gy (4 Gy/ day for 5 days). The upper panel shows the individual mouse data of the different groups compared to the controls (no treatment). The lower panel shows the tumor volume fold change over a period of 100 days post spacer implantation…………………………………………………………………………….107

Figure 38: Box and whisker plots showing the geometric mean difference in tumor volume and comparative significance among various treatment groups after 135 days of treatment……………………………………………………………………………….. 109

Figure 39. Kaplan-Meyer survival curve shows median survival time of LCRT treatment groups…………………………………………………………………………………….94

Figure 40: Gross toxicity assessment by measuring the change in body weight in prostate cancer tumored animals treated with local chemotherapy using spacers……………..…112

Figure 41. Percentage change from geometric means of hematology counts in treatment groups to levels of healthy mice…………………………………………………………121

Figure 42. Percentage change from geometric means of hematology counts in treatment groups to levels of untreated tumored control mice……………………………………..122 Figure 43. Percentage change from geometric means of blood chemistry levels in treatment groups to levels of healthy mice…………………………………………………………124

Figure 44. Percentage change from geometric means of blood chemistry levels in treatment groups to levels of untreated tumored control mice……………………………………..125

Figure 45. Biodegradable PLGA implants for sustained intratumoral delivery of Talazoparib. a) Schematic depiction of PLGA implant degradation (blue) for localized delivery of Talazoparib (red). b) Scanning Electron Microscopy images of flash-frozen and fractured, drug-loaded PLGA implants a solid smooth surface with nanoscale pores homogenously spread throughout surface of implant. c) In vitro release profile of Talazoparib from 2 mm PLGA implants (n=3) in phosphate buffered saline (pH 6.0)……………………………………………………………………………………...138

Figure 46. In vitro analysis of different PARP inhibitors on W780 and W0069 breast cancer cell lines derived from Brca1 Co/Co ;MMTV-Cre;p53 +/- mice. Western Blot data testing protein expression after treatment with Olaparib, Niraparib, and Talazoparib on Brca1 - deficient cell line in a) W780 b) W0069 cells. C) Measured dose response curve and IC50 of Talazoparib in W780 and W0069 cells………………………………………………139

Figure 47. Treatment with 50 µg Talazoparib implants decreases tumor size and slows tumor growth in Brca1 -deficient mice. Implants (2 mm × 0.8 mm diameter) containing 0 µg (control) or 50 µg Talazoparib were injected into established mammary gland tumors in female Brca Co/Co ;MMTV-Cre;p53 +/- mice. Mice treated by oral gavage received 6 doses of Talazoparib (50 µg Talazoparib total). A) Average tumor volumes before and after treatment (n=5-8/group). *, P < 0.05 vs. initial tumor volume; ǂ, P < 0.05 vs. all other groups. B) Change in tumor volume over a period of 28 days. C) Average body weight before and after treatment (* P < 0.05 vs. initial weight for the treatment group)……….141

Figure 48. Immunohistochemistry studies showing Talazoparib implants decrease PCNA staining and increase γ-H2aX staining in tumors from Brca1 -deficient mice. A). Immunohistochemical staining for PCNA (top) and γ-H2AX (bottom) of tumors treated with 50 µg Talazoparib implants for 2 weeks. B). Quantification of the percentage of PCNA-positive or γ-H2AX-positive cells (mean ± SEM) from 4-5 tumors per group. (P < 0.001 versus control)……………………………………………………………………143

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TABLE OF CONTENTS

Abstract……………………………………………………………………………………ii Acknowledgements……………………………………………………………………….iv List of Tables……………………………………………………………………………...vi List of Figures……………………………………………………………………………vii Synopsis of dissertation…………………………………………………………………..xv

Chapter 1. Introduction…...………………………………………………………………..1 1.1 The Prostate………….…………….…………………………………………………..3 1.2 Overview of Prostate Cancer………………………….………………...... 5 1.2.1 Epidemiology of prostate cancer…………………….……...………...... 5 1.2.2 Diagnosis and staging of prostate cancer……………………….……………………7 1.2.3 Current treatment of prostate cancer……………………….……………………….10 1.2.3.1 Therapies for localized prostate cancer……………………….………………….10 1.2.3.2 Therapies for metastatic prostate cancer……………………….…………………17 1.3 Regional and local cancer treatment……………………….…………………………18 1.3.1 Low dose rate brachytherapy……………………….………………………………18 1.3.2 Regional cancer chemotherapy……………………….……………………….……22 1.4 Biodegradable drug delivery systems……………………….…………………..……23 1.5 Degradation of PLGA……………………….……………………………………..…25 1.6 Docetaxel (Taxotere) ………….……………………………………………………..29 1.7 Nanotechnology for oncology………….………………………………………….....31 1.7.1 Nanomedicine for chemotherapeutic delivery………….…………………………..31 1.7.2 Radiosensitization using nanotechnology………….………………………….…...33 1.8 Research Strategy and overview of dissertation………….………………………..…35

Chapter 2. Fabrication of Silica Nanoparticles…………………….…………………….40 2.1 Introduction…………………….………………….…………………….…………...40 2.2 Materials and Methods…………………….…………………………………………41 2.2.1 Synthesis and Characterization…………………….……………………………….41 2.2.2 In vitro Efficacy…………………….………………….…………………….……..43 2.3 Results and Discussion…………………….…………………………………………45 2.4 Conclusion……………………………………………………………………………52

Chapter 3. Fabrication of INCeRT Spacers………………………………………………53 3.1 Introduction…………………….…………………………………………………….54 3.2 Materials and Methods…………………….…………………………………………57 3.3 Results and Discussion………………….……………………………………………58

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Chapter 4. In vivo optimization of ‘Smart Brachytherapy Spacers’ using optical imaging………………………………………………………………………………….. 62 4.1 Introduction………………….…………………………………………………….....62 4.2 Methods and Materials………………….……………………………………………66 4.2.1 Synthesis, fabrication, and characterization of Silica nanoparticles and Spacers …..66 4.2.2 Fluorescence imaging system………………….…………………………………...67 4.2.3 In vitro Imaging of phantoms………………….……………………………………69 4.2.4 In vivo imaging of spacers………………….………………………………………70 4.2.5 Image Processing and Data Analysis………………….……………………………71 4.3 Results………………….…………………………………………………………….73 4.3.1 In Vitro Characterization of Nanoparticles, Spacers, and phantom imaging………..73 4.3.2 In Vivo Imaging………………….…………………………………………………76 4.4 Discussion and Conclusion………………….………………………………………..78

Chapter 5. Docetaxel loaded implants as a monotherapy for prostate cancer……………..81 5.1 Justification for implant reformulation………………….……………………………81 5.2 Materials and Methods………………….……………………………………………82 5.2.1 Docetaxel Implant Fabrication and characterization………………….……….……82 5.2.2 In vivo therapeutic efficacy…………………………………………………………84 5.3 Results and Discussion………………….……………………………………………86 5.3.1 Docetaxel Implant characterization………………….…………..…………………87 5.3.2 Docetaxel Implant therapeutic benefit………………….………………..…………91

6. Docetaxel loaded implants with radiation for combine local chemo-radiation………...99 6.1 Introduction 996.2 Materials and Methods…………………………………………...99 6.3 Results and Discussion……………………………………………………………...102 6.3.1 Determination of parameters in preliminary study………………………………...102 6.3.2 Local Chemo-Radiation Therapeutic efficacy…………………………………….106 6.4 Conclusion…………………………………………………………………………..112

Chapter 7. Toxicity of Local Chemotherapy and Local Chemo-Radiation Therapy…….114 7.1 Introduction…………………………………………………………………………114 7.2 Materials and Methods………………….…………………………………………..115 7.3 Results and Discussion………………….…………………………………………..118 7.4 Conclusion………………….……………………………………………………….126

Chapter 8. Sustained PARP inhibition for breast cancer treatment using spacer platform…………………………………………………………………………………128 8.1 Introduction………………….……………………………………………………...129 8.2 Materials and Methods………………….…………………………………………..132 8.2.1 Talazoparib Implant Fabrication and characterization………………….…………132 8.2.2 In vitro Therapeutic Efficacy………………….…………………………………..134 8.2.3 In vivo Therapeutic Efficacy……….……………………………………………...135 8.3 Results……….……………………………………………………………………...136 8.3.1 PLGA implants continuously release Talazoparib……….……………………….136 8.3.2 Sustained PARP inhibition produce apoptosis and DNA damage in vitro ………..138 xiv

8.3.3 Sustained PARP inhibition produces tumor shrinkage……………………………139 8.4 Discussion....………………………………………………………………………..143 8.5 Conclusion………………………………………………………………………….148

Chapter 9. Perspectives and Conclusions……………………………………………….149

References………………………………………………………………………………154 Appendix………………………………………………………………………………..167

SYNOPSIS OF DISSERTATION

Prostate cancer remains the second leading cause of cancer related deaths in men with over 161,360 new cases and 26,730 associated deaths in the U.S. alone in 2017. While early detection has caused a shift in staging, prostate cancer remains a worldwide concern.

Between 2006 and 2012, the 5-year survival rate for men diagnosed at a local stage was

100%. This number decreases to 29.3% if the tumor is diagnosed after metastasis. Overall, any man diagnosed with prostate cancer has a 98.7% 5-year survival rate, a 98% 10-year survival rate, and a 95% 15-year survival rate. Additionally, there is an 83% chance of mortality from prostate cancer for any man diagnosed before turning 65 years old. The incidence rate of prostate cancer is extremely high. Nearly one in six men will be diagnosed with prostate cancer during their lifetime in the United States with an estimated 2,850,139 men living with prostate cancer in 2013. These numbers are only expected to grow in the future. As the population of people living over the age of 65 increases, the proportionate number of prostate cancer cases is expected to quadruple by 2030, resulting in an immediate need for new treatments and cures.

Treatment plans depend on the age, health, PSA levels, Gleason score, and risk of metastasis, however all result in significant side-effects. Prostate cancer treatment options are scarce and often leave survivors with reduced quality of life due to off-target side effects. While doctors do as much planning to avoid toxicities, local therapies including prostatectomy, radiation therapy and cryotherapy can leave patients with bowel or urinary incontinence, erection problems including impotence or infertility, and lymphedema, if lymph nodes are damaged during therapy. We have utilized technology in the standard clinical procedure Brachytherapy to develop a new local chemotherapy implant with the xvi standardized radiation treatment without the need for additional procedures. Brachytherapy uses plastic inert spacers to help clinicians place and separate the radioactive seeds. We have designed and fabricated a biocompatible and biodegradable spacer that mimics this spatial guidance, with the added therapeutic benefit of sustained chemotherapy and radiosensitization locally at the tumor site to sensitize the cancer throughout the course of brachytherapy radiation rather than intermittent radiosensitization experienced with systemic chemotherapy administered every three weeks. These ‘smart brachytherapy spacers’ can be modified to tailor their release of docetaxel to coincide with varying half- lives of any radioactive seed used in brachytherapy.

In this dissertation I have developed, characterized, and extensively tested a docetaxel loaded biodegradable implant for the treatment of prostate cancer. Our initial goal was to load the spacers with silica nanoparticles which would release drugs providing a dual-release platform. Chapter 2 describes in detail the synthesis and characterization of these 30 nm silica particles which encapsulated either dye for imaging and tracking, or docetaxel for radiosensitization. Silica nanoparticles loaded with docetaxel had IC50 values of 4.633nM, and 2.218nM after treatment on human prostate cancer cell line PC3 after 48 and 72 hour incubation respectively. These are similar to the reported IC50 value of docetaxel on PC3 cells (2-60 nM). When combined with radiation, docetaxel loaded silica nanoparticles resulted in significant radiosensitization.

These particles were loaded into a PLGA matrix, as described and characterized in chapter 3. In order to optimize the release in vivo, docetaxel was replaced with near infrared dye Cy7.5 to act as a model drug. The in vitro and in vivo diffusion and release kinetics xvii were determined using a custom optical imaging platform designed for particle tracking in real time in live animals.

We used this platform in chapter 4 to optimize the formulation of the spacer. Results indicated that nanoparticles aggregated and behaved as microparticles rather than individual nanoparticles, making the release difficult to predict and reproduce. In contrast, free dye exhibited optimal release kinetics and free docetaxel embedded in PLGA matrix was selected for therapeutic studies.

Chapter 5 discusses docetaxel-loaded spacers as monotherapy compared to

clinically relevant Taxotere, the IV formulation of docetaxel. Our spacers were fabricated

with a docetaxel loaded Poly(lactic-co-glycolic) acid cylindrical implant for intratumoral

injection via an 18 gauge applicator needle for local, sustained therapy. Our spacers exhibit

diffusion driven release in vitro over 75 days, designed to sensitize I-125 (t 1/2 = 60 days) brachytherapy seeds most commonly used for treatment of prostate cancer. The spacers were tested for therapeutic efficacy against clinically administered docetaxel and resulted in significant tumor inhibition and improved survival from local chemotherapy (LCT)

(median survival time (MST) of spacers 52 days versus 26 with IV DTX, p<0.01) with no significant weight loss as seen in the systemic formulation.

Docetaxel spacers were combined with fractionated radiation therapy at reduced doses, mimicking brachytherapy, to determine the radiosensitization and synergistic therapeutic response in chapter 6. Mice treated with local combined chemo-radiation resulted in significant survival improvement (MST 209 days vs. 120 in radiation therapy alone and 85 in spacers alone, p<0.01) and tumor inhibition, with 33% of mice cured.

Chapter 7 discusses a full toxicity study to test hematology and blood chemistry of all xviii treatment groups. Blood work results showed an improved enzymatic levels in local chemotherapy compared to systemic chemotherapy. LCRT had improved hematology

(white blood cell counts, platelet levels, and lymphocytes) compared to all other treatment groups, however enzyme levels indicated a potential for kidney disease within the group.

All studies used reduced concentrations of docetaxel compared to clinical dosing. Small ulcerations at the tumor site in groups treated with docetaxel spacers suggest local toxicities rather than systemic toxicities will limit the maximum tolerated dose.

Chapter 8 shows the versatility of our biodegradable drug eluting platform with applications to breast cancer for local sustained delivery of PARP inhibitor Talazoparib in a Brca-1 deficient mouse model. As a whole, this dissertation describes the development and potential for a new local treatment in prostate cancer. This ‘smart spacer’ shows significant success as a monotherapy, as well as a radiosensitizer when combined locally with radiation therapy, suggesting great potential for successful clinical translation and impact.

Chapter 1. Introduction

Cancer is the second most common cause of death in the US, accounting for 1 in 4 deaths and continue to be a leading cause of death worldwide [1]. Cancer is defined as a group of diseases involving abnormal cell growth with the potential to invade or spread into other parts of the body. There are over 200 different types of cancer with varying symptoms dependent on the type and patient genetics, sex, diet, etc. with dozens of sub- types[2]. Tumors often exhibit heterogeneity; meaning cells within the tumor have different genetics, express different biomarkers, and respond differently to treatments, making it nearly impossible to successfully eradicate all cancer cells with a single approach[3].

Prostate cancer is the second leading cause of cancer related deaths in men. Nearly one in six men will be diagnosed with prostate cancer during their lifetime in the United

States with an estimated 2,850,139 men living with prostate cancer in 2013[1,4]. While five year survival rates are a high 98.9% in the last decade, prostate cancer is often slow growing with a 20- 30% chance of relapse and recurrence, which results in fewer treatment options, added toxicities, and a poorer prognosis[5–7]. Substantial improvements in diagnosis and staging have been made with digital rectal examination, serum levels of prostate specific antigen (PSA) measurements, and transrectal ultrasound, leading to a staging shift to organ confined and local disease. Over 80% of prostate cancer is diagnosed in local stages (stage I and II)[5,8]. When confined to the organ, such as in these early stages, men have a much higher chance of cure or survival. Because of its slow-growing nature, many clinicians suggest the popular “watch and wait” treatment modality over the 2 more aggressive treatment options, as the associated risks of treatment often outweigh the benefits. Significant morbidity can result from local tumor progression and migration.

Current treatment options for localized prostate cancer include prostatectomy, radiation therapy, and cryotherapy. These options are associated with low but significant morbidities such as impotence, urinary incontinence, rectal leakage, and instances of GI or bladder cancers. Androgen deprivation therapy (ADT) is used with significant success, however, tumor recurrence often results in hormone refractory tumors which are resistant to further

ADT and often more difficult to treat. Over 40 percent of all prostate cancers are diagnosed in men under the age of 65[7,9]. Therefore, a need exists for effective nonsurgical treatments that can eradicate localized prostate tumors without leaving significant side effects which greater affect the patient’s quality of life. A biodegradable implant which has a slow sustained release for weeks to months has the potential to offer such benefits. With such an implant, a controlled and designated concentration of drug could be tailored to release at the disease site each day to provide high local drug concentrations with the potential to minimize side effects and toxicity experienced with current treatment options.

Docetaxel has shown clinical efficacy in head and neck cancers, breast cancer, stomach cancer, non-small-cell lung cancer, and metastatic prostate cancer and can be used in combination with other chemo-agents[10]. While it has shown promise in pre-clinical models, docetaxel for localized prostate cancer is not used clinically due to the poor accumulation of drug at the disease site with high systemic side effects. The incorporation of docetaxel into a biodegradable implant offers the potential to increase the amount of drug available at the tumor while administering much lower concentrations to the patient and sparing toxicities to healthy organs. Here I discuss the development of this new 3 approach to docetaxel delivery for localized prostate cancer to enhance treatment efficacy.

The implant was designed as a modified brachytherapy spacer to be used in conjunction with permanent low dose brachytherapy as a radiosensitizer or alone as a chemo- monotherapy.

1.1 The Prostate

The prostate gland is the largest accessory gland of the male reproductive system and is composed of four lobes. The prostate is a small fibromuscular and glandular organ shaped like a walnut and located inferior to the urinary bladder and anterior to the rectum as shown in Figure 1 [11]. The posterior section of the prostate can be palpated through the anterior wall of the rectum during a digital rectal exam. The prostate has a dense, fibrous prostate capsule, is composed of 30-50 tubuloalveolar glands, and is surrounded by a fibrous prostatic sheath. The fibromuscular stroma is located between the tubuloalveolar glands. The prostate surrounds the urethra as it exits the bladder and merges with the ductus deferens at the ejaculatory duct [12]. The prostate is responsible for producing a secretion which makes up about 30% of the semen volume. These prostatic secretions are a milky white mixture of simple sugars (such as fructose and glucose), enzymes, and alkaline chemicals which act as nutrients for sperm to aid in motility and protect the sperm from degradation by neutralization of acidic vaginal secretions to promote survival in the female body[11]. Two prostate specific enzymes that are secreted by the secretory epithelial cells are prostatic acid phosphatase (PAP) and prostate specific antigen (PSA). Both are markers measured in serum for monitoring the progression and diagnosis of prostate cancer, 4 respectively. PSA, or kallikrein-3 (KLK3), is a glycoprotein enzyme secreted by the epithelial cells which liquefies semen in the seminal coagulum.

Figure 1. Male Anatomy showing physiological location of prostate located below bladder, around urethra, and laterally to rectum [adapted from 11] .

Scientists have divided the human prostate gland histologically into four zones according to their function as shown in Figure 2 [13]. The peripheral zone makes up about

70% of the glandular volume of the prostate and is the sub-capsular portion of the posterior aspect of the prostate gland that surrounds the distal urethra. It is from this portion of the gland that between 70-80% of all prostatic cancers originate. The central zone is about 25% of the glandular fraction and surrounds the ejaculatory ducts. This zone accounts for about 5

2.5% of prostate cancers which tend to be more aggressive and likely to invade the seminal vesicles.

The transition zone accounts for 5% of the glandular volume and surrounds the proximal urethra. While 10-20% of prostate cancers originate in this zone, it is the most frequent site of benign prostatic hyperplasia. The fourth zone or anterior fibromuscular zone is approximately 5% of the prostate volume and is devoid of any glandular components. This zone occupies the anterior surface of the prostate and is comprised only of smooth muscle cells and fibrous tissue [13].

1.2 Overview of Prostate Cancer

1.2.1 Epidemiology of prostate cancer

Prostate cancer represents a leading cause of death in men worldwide. In 2016,

prostate cancer diagnosis accounted for 10.7% of all cancer cases and 4.4% of cancer 6 related deaths in the United States. Other than skin cancer, prostate cancer is the most common cancer among American men [5]. The American Cancer Society estimated

161,360 new cases and 26,730 deaths from prostate cancer in 2017 in the United States alone with higher rates in developed nations and lower rates in third world country, a result of improper diagnosis. The incidence rate of prostate cancer increases with age: 60% of men who are diagnosed are over 65 and the average age at the time of diagnosis is 66.

Prostate cancer can be a serious disease but is not fatal for most diagnosed men. About 1 in 39 men will die of prostate cancer, however treatment options often lead to significant side effects and long term problems [14,15]. Studies suggest there are many risk factors including age, diet, race, genetics, familial factors, and environmental factors. Black

American men have the highest incidences of prostate cancer diagnosis and death. As the population of people living over the age of 65 increases, the proportionate number of prostate cancer cases is expected to quadruple by 2030, resulting in increased funding for research in the field for prevention and treatment [5,16].

Between 2006 and 2012, the 5-year survival rate for men diagnosed at a local stage was 100%. This number decreases to 29.3% if the tumor is diagnosed after metastasis.

Overall, any man diagnosed with prostate cancer has a 98.7% 5-year survival rate, a 98%

10-year survival rate, and a 95% 15-year survival rate [14]. Relative survival is the preferred metric for analyzing survival and prognosis for cancer patients in population- based studies. Relative survival compares the observed survival of a group of cancer patients to the survival time of members of the general population who share the same demographics (i.e. age, gender, race, and state of residence). While these statistics may seem positive, the incidence rate and frequency of diagnosis makes prostate cancer the 7 third leading cause of cancer related deaths in all men. The widespread use of PSA testing has resulted in a dramatic increase in the diagnosis and treatment of prostate cancer, however many men do not benefit from therapeutic intervention because the disease is either indolent or disseminated at the time of diagnosis. Prostate cancer often progresses slowly and many men die of other causes while living with the disease. Additionally, current therapeutic options for prostate cancer can leave patients with long-term impotence, or urinary and bowel problems. The survival and prognosis of patients diagnosed with prostate cancer is heavily influenced by the stage of tumor at the time of diagnosis. Ninety- two percent of all prostate cancer is diagnosed locally, where the tumor is confined to the primary site, or regionally where the tumor has spread to regional lymph nodes [17,18].

This early stage detection is attributed to standardized testing during physicals in men over the age of 50, or 45 if there is familial history of the disease. Once the disease has metastasized, the median survival time is between 1-3 years [6]. Therefore, treating the disease at localized stages is becoming increasingly demanded by patients and has warranted the development of new effective treatment modalities for the localized disease.

1.2.2 Diagnosis and staging of prostate cancer

PSA is present in small quantities in men with healthy prostates, but is often elevated in the presence of cancer or other prostate disorders. A PSA test in combination with a digital rectal examination of the prostate contributed to early diagnosis of prostate cancer, however high levels are indicative of abnormal prostate activity rather than the presence of a tumor resulting in only 30% of men with high levels having a diagnosis of prostate cancer [16,19]. The PSA test measures the quantity of protein in the blood in 8 nanograms per milliliter. If the tested levels are elevated or if the patient is presented with symptoms indicative of abnormal prostate activity (i.e. burning, frequent or interrupted urination, blood in urine or semen, or new onset of erectile dysfunction), a biopsy is carried out to confirm or exclude the tumor diagnosis [20].

Several staging systems are used to classify prostate cancer in the United States, the most common being the American Joint Committee on Cancer TNM classification system. This classification gives three key pieces of information and is summarized in

Table 1 [21,22]: (1) T refers to the extent of the primary (main) tumor.

Table 1: TNM staging system for prostate cancer [22]

There are two classifications of staging the tumor, clinical and pathological. The clinical stage is the doctor’s best estimate to the extent of the disease based on the results of a physical exam, a digital rectal exam, lab tests, prostate biopsy, and any imaging conducted.

After surgery, doctors can also stage the tumor via pathology, which uses the clinical tests in combination with the results of the surgery to more accurately determine any dissemination of the disease that was not seen in images [21]. Staging parameters for both clinical and pathological are the same, ranging from T1 to T4 with T1 describing localization to the prostate, and T4, a disseminated disease. (2) N categories describe whether the cancer has spread to regional lymph nodes. (3) M categories describe whether the cancer has spread to distant parts of the body [22]. The most common sites of prostate cancer metastases are the bones and distant lymph nodes, although it can also spread to other organs such as the liver and lungs [23].

Another common staging classification method is the Gleason score system, which is determined from histological sections taken from the prostate biopsy. A pathologist will grade the tumor histology from 1-5 based on differentiation of the majority of tumor cells in the biopsy as shown in Figure 3 and both scores are added. The more uniform and differentiated the cells, the lower the score and better the prognosis. The higher number scores, such as 8-10, indicate the cancer cells are poorly differentiated or undifferentiated and likely to grow rapidly [24]. The D’Amico risk category is not used to stage all prostate cancers like the TNM or Gleason scoring systems, but rather it is used to estimate the risk that a prostate cancer has spread outside of the prostate. The system uses the PSA level,

Gleason score, and the T stage of the cancer to divide men into three risk groups: low, intermediate, and high [25]. To properly stage a tumor, doctors will use all available 10 information, including the T, N, and M categories, Gleason score and PSA level, to determine the overall stage of the cancer. A reference guide provided by the American

Cancer Association for staging prostate cancer can be found in the Appendix. The stage is expressed in Roman numerals from I (the least advanced) to IV (the most advanced) and helps determine treatment options and overall prognosis of the patient.

1.2.3 Current treatment of prostate cancer

1.2.3.1 Therapies for localized prostate cancer

The course of treatment for prostate cancer varies greatly depending on the stage and grade of the disease along with the demographics of the patient. For patients with slow growing cancer, the preferred method is watchful waiting or active surveillance where 11 clinicians constantly monitor the growth and progression of the disease. This approach is likely not a good option when the patients presents with a fast-growing cancer (high

Gleason score), a cancer that is likely to spread outside the prostate (based on PSA levels), or if the patient is young and healthy, out of concern that the cancer might become a problem over the next 20 to 30 years [27]. When intervention is needed, surgery, radiotherapy, cryotherapy and some clinical trials including photodynamic therapy are used.

In 2010, a total number of 138,000 prostatectomies were performed in the US, nearly half of the diagnosed cases that year [28]. Until recently , radical prostatectomy was regarded as the first line of treatment for localized prostate cancer and a chance at cure if the cancer was localized inside the prostate gland. A radical prostatectomy is a surgery by which the entire prostate gland and some surrounding tissue (including the seminal vesicles) is removed from the patient. Recently, radical prostatectomy has moved from an open surgery to laparoscopic incisions to remove the prostate. This has resulted in faster recovery times, shorter hospital stays, and less pain and bleeding experienced by the patient. However, the rates of major side effects from the laparoscopic surgery, such as erection problems and incontinence, seem to be the same as the previously used open prostatectomy. While a prostatectomy can result in cure, results and outcomes of patients greatly vary depending on the demographics of the patient, the tumor grade, and the precision of the surgeon. While this procedure does offer the potential for complete tumor eradication, treatment-related mortality risk is substantial, reaching as high as 1% and increasing to over 2% for patients over the age of 75. Because of nearby nerve bundles adjacent to the prostate, the patient can be left impotent as a result of the surgery. There is 12 a wide reported range within the literature stating that 63-94% of patients experience some level of erectile dysfunction (ED) post-radical prostatectomy. Over 80% of patients report urinary incontinence, or bladder leakage, immediately after surgery, some with symptoms still present 18 months later. In all cases, the surgery results in a loss of patient fertility

[29]. In a retrospective analysis, nearly 2000 men who had undergone radical prostatectomy with a curative intent were followed for a mean of 5.3 years. Of these 2,000 men, 15% demonstrated an abnormal PSA of 0.2ng/ml or higher, which is considered evidence of a biochemical recurrence. Of these 315 men, 103 (34%) developed clinical evidence of recurrence with a median time to develop metastasis of 8 years. Once metastasis occurred, the median time to death was an additional 5 years. In this cohort, over 5% of men who underwent a radical prostatectomy developed recurrence and the metastatic disease [30].

Radiotherapy is commonly used and effective in patients presenting localized disease. Radiation therapy is used as a first treatment for men with stage I prostate cancer where the cancer is localized in the prostate gland. Cure rates for men with low grade cancer is typically about the same as men treated with radical prostatectomy [31]. It is also used in combination with other treatment options for cancers that have disseminated into nearby tissues. For late stage cancers, radiation is used to maintain control or relieve symptoms of the patient, and in such cases of recurrence, it is a second line of treatment after surgery. There are two major types of radiation therapy; external beam radiation therapy (EBRT) and brachytherapy (internal radiation).

External beam radiation therapy uses focused beams on radiation from a machine outside the body to specifically target the prostate gland. A team of radiation oncologists 13 will take careful measurements using MRI or CT scans to calculate the correct angles for aiming the radiation beams and the proper dose of radiation during treatment planning. Due to the physiological location of the prostate, careful planning must be done to assure minimization of radiation dosing to nearby healthy organs. To focus radiation more precisely on the tumor while sparing exposure to nearby tissues, newer EBRT techniques are currently clinically used including: (1) Three-dimensional conformal radiation therapy which uses special computers to precisely map the location of the prostate. Radiation beams can be shaped and aimed at the prostate from several different directions to minimize damage to normal tissues while administering the proper dose to the tumor site. (2)

Intensity modulated radiation therapy (IMRT) is an advanced form of 3D therapy and is the most common type of EBRT for prostate cancer treatment. It uses the technology of the 3D conformal therapy along with a computer driven machine that moves around the patient as it delivers radiation. Along with the ability to shape beams from different angles, the intensity of the beam can be adjusted to limit doses reaching nearby normal tissues.

Newer machines have imaging scanners built into them for advanced image guided radiation therapy , letting doctors take real time images of the prostate to make minor adjustments in aiming just before delivering the radiation dose. All of these technological improvements help deliver radiation more precisely with the aim to minimize patient toxicity and side effects. (3) Stereotactic body radiation therapy is a technique that utilizes image guidance to deliver large doses of radiation to a specific area. Because the radiation doses are so high, the entire course of treatment is given in days, rather than the weeks that it takes to deliver IMRT. (4) Proton beam therapy is a form of radiotherapy which uses focuses beams of protons instead of x-rays. X-rays release energy both before and after 14 they hit their target while protons release their energy after traveling a certain distance, causing little damage to tissues in which they pass. In theory this technique would allow the delivery of more radiation to the tumor while doing less damage to nearby normal tissues, however the machines required to make protons are expensive and not widely available in many treatment facilities [32]. While doctors do as much planning to avoid toxicities, EBRT can leave patients with bowel problems, or radiation proctitis, urinary problems, or radiation cystitis, erection problems including impotence, and lymphedema, if lymph nodes are damaged during therapy [33].

Brachytherapy, sometimes called seed implantation or interstitial radiation therapy, uses small radioactive pellets placed directly into the prostate. For patients with a risk of tumor migration, brachytherapy can be combined with external radiation or boosts , to decrease the likelihood of such an event. Like EBRT, imaging tests such as CT scans, MRI, and trans-rectal ultrasound are used to help map and guide the placement to accomplish the calculated required dose of radiation needed. There are two types of radiation therapy: low dose and high dose. High dose rate (HDR) brachytherapy is a temporary and less common form of brachytherapy. During the short procedure, hollow needles guide catheters into the prostate. Radioactive iridium-192 or cesium-137 is then placed in the catheters for 5-15 minutes for 3 brief treatments given over 2 days. HDR is often combined with EBRT given at low doses with the aim to concentrate most of the radiation in the prostate itself while sparing normal tissues [34]. Low dose rate (LDR) or permanent brachytherapy is more commonly used. In this approach, radioactive pellets (seeds) are placed inside thin 18- gauge needles (1.3mm outer diameter) which are inserted through the skin between the scrotum and anus and into the prostate under trans-rectal ultrasound guidance as seen in 15

Figure 4. Typically 70 to 100 pellets are left in place as the needles are removed, emitting low doses of radiation for weeks or months. Radiation from the seeds travels extremely short distances so the seeds can give high concentrated doses in a very small area which limits damage to nearby tissues, and in some cases EBRT can be added alongside brachytherapy if there is a higher risk for cancer spread. The side effects of brachytherapy are the same as EBRT, however studies report brachytherapy side effects are less common and severe, making it a desired option for young, healthy men [35]. Low dose brachytherapy will be discussed in further detail later in this chapter. 16

Cryotherapy is a secondary line of treatment, typically done after radiation therapy, in which freezing temperatures are used to ablate cancer cells using a metal probe placed under ultrasonic guidance. During this procedure, a catheter is placed in the urethra with warm salt water running through it to protect the urethra from freezing. Cryotherapy uses 17 argon gas infused through the probe into the prostate gland to preferentially target the cancer, limiting the damage to normal prostate tissue. However due to its new development and clinical use, no long-term studies have been conducted to validate the efficacy of the procedure. Additionally, rare but serious side effects can occur after cryotherapy with <1% of men developing a fistula between the rectum and bladder [37]. Another option to treat local or regional prostate cancer is clinical trials. There are over 90 Phase III drug trials and more than 400 Phase I/II trials in progress for prostate cancer in North America and

Europe. Those approved will join the five new drugs that have been approved for men with advanced metastatic disease since 2010 [38,39]. One of particular interest is a local

Padeliporfin vascular-targeted photodynamic therapy conducted in the UK to compare patient safety and outcome to active surveillance in men with low-risk prostate cancer, mentioning that most low risk prostate cancer patients often need therapeutic intervention at some point in their lifetime. For this focal therapy, fibers were positions at the tumor site using ultrasound imaging. When in place, 4mg/kg padeliporfin was administered intravenously over 10 minutes. The drug was then preferentially activated only at the treatment zone by laser light at 753 via the fibers for about 20 minutes. Results showed the procedure was safe and effective compared to the active surveillance cohort, with emphasis on early treatment to eradicate the cancer and prevent recurrence or metastasis [40].

1.2.3.2 Therapies for metastatic prostate cancer

Hormone therapy has been the first line of treatment for advanced prostate cancer via surgical castration or by medicinal castration to achieve androgen deprivation therapy

(ADT) since 1941. The goal of this therapy is to reduce levels of the male hormone, 18 androgen, in the body or to stop them from affecting the growth of prostate cancer cells as androgen is a stimulant prostate cancer cell growth [41]. The two main androgens, testosterone and dihydrotestosterone (DHT) are made by the testicles and in small quantities, the adrenal glands. While ADT slows the growth and progression, it alone is not a cure and nearly all patients develop a resistance to the therapy, leading to a widely debated best time and dose for treatment [42]. When resistance develops, recurring tumors are castrate-resistant or hormone-refractory and much more difficult to treat effectively.

When hormone therapy fails for metastatic disease, few options remain for patients.

A final line of treatment is chemotherapy. Most commonly for advanced metastatic prostate cancer, Docetaxel, or Taxotere, is administered intravenously in combination with steroid prednisone. If unresponsive, Cabazitaxel, then Mitoxantrone and Estramustine are used.

While both of these drugs have shown to help improve survival time and reduce symptoms, chemotherapy very rarely is curative. Due to their toxic nature on rabidly dividing cells, chemotherapy has to be administered in cycles to allow the body enough time to recover.

Systemic administration of anticancer agents commonly results in cytotoxic side effects such as severe allergic reaction requiring medicine prior to treatment, myelosuppresion, nausea, hair loss, and peripheral neuropathy or damage to nerves [43,44].

Metastatic prostate cancer nearly always spreads to the bone first, leading to high levels of pain, fractures or breaks, or high blood calcium levels which can be dangerous or life threatening. If the cancer has grown outside of the prostate, the clinical aim is to prevent or slow the spread of the disease. Once a patient presents bone metastasis, prognosis is poor, with a median survival time <1.5 years, and treatment is focused on controlling or relieving pain [41]. 19

Bisphosphonates and Denosumab are two drug options used to slow down osteoclasts, bone cells that break down hard mineral structure of bones which become overactive in the presence of prostate cancer metastasis. Radiopharmaceuticals are radioactive drugs which give off radiation that kills cancer cells. Unlike EBRT, radiopharmaceuticals, such as Radium-223, can reach and treat all affected bones at the same time, however they often result in low blood cell counts, which increases risk of infection of disease. A final approach for treatment is external radiation therapy, corticosteroids, and pain medicines to aid in pain management associated with the advanced disease [37]. The poor prognosis associated with metastasis of prostate cancer highlights the severity in necessity for treatment and cure at local and regional stages.

1.3 Regional and local cancer treatment

1.3.1 Low dose rate brachytherapy

Brachytherapy is a common form of treatment for prostate, cervical, breast, and skin cancers. Brachytherapy is a form of radiotherapy in which the radiation source is sealed in a seed and placed inside or next to the required treatment area. Results have demonstrated cure rates comparable to patients treated with surgery or EBRT or improved when used in combination with these techniques. Brachytherapy can be used alone or in combination with EBRT and chemotherapy. Brachytherapy seeds irradiate a highly localized area surrounding the sources and therefore reduce exposure of tissues further from the seeds, allowing for planned concentrated dosing directly to the tumor site with reduced probability of unnecessary damage to surrounding healthy tissue. Additionally, the seeds move with the patient or tumor so they maintain their relative position for treatment 20 which is difficult to achieve with EBRT. Brachytherapy is typically an outpatient procedure making it highly tolerable for patients with fewer required visits than other radiotherapies.

In 2013, the global market for brachytherapy reached $680 million with growth projected to reach $2.4 billion by 2030, attributed mainly to the short procedure and recovery time with fewer side effects than other treatment modalities [45].

Brachytherapy seeds are between 5-8mm in length and 0.8mm in diameter, which can be injected using a hollow 18 gauge applicator needle. Brachytherapy seeds are often separated physiologically with inert plastic spacers of the same size which provide clinicians with spatial and temporal guidance during the procedure and help separate the seeds and prevent ‘hot spots’ from overlap of adjacent radiation zones[46,47]. Different types of brachytherapy can be classified according to: (1) the placement of the radiation sources in the target treatment area, (2) the rate of intensity of the irradiation dose, and (3) the duration of the dose [35,47]. The rate of intensity and duration of the dose are directly related to the irradiation source of the brachytherapy seed used. Table 2 shows the different radiosources used during brachytherapy. The most common isotope used for LDR brachytherapy is iodine (I)-125 with a half-life of 60 days and a mean energy of 31.4keV at a rate of 2Gy/hr[48].

Table 2: Commonly used radiation sources for brachytherapy Radionuclide Type Half-life Energy Cesium-131 (131Cs) Electron Capture, ε 9.7 days 30.4 keV (mean) Cesium-137 (137Cs) β−- particles 30.17 years 0.662 MeV Cobalt-60 (60Co) β−- particles 5.26 years 1.17, 1.33 MeV Iridium-192 (192Ir) γ-rays 73.8 days 0.38 MeV (mean) Iodine-125 (125I) Electron Capture, ε 59.6 days 27.4, 31.4 and 35.5 keV Palladium-103 (103Pd) Electron Capture, ε 17.0 days 21 keV (mean)

Treatment planning is critical in brachytherapy patients not only for the placement of the seeds, but the insertion sites of the applicator needles. A range of images are collected using x-ray, ultrasound, CT scans, and MRI to create a three-dimensional visualization of the tumor size and shape and its relation to the surrounding tissue and organs. Initial planning helps ensure that ‘cold spots’ (too little radiation) and ‘hot spots’ (too much radiation) are avoided during treatment, which can result in either treatment failure or adverse toxicities, respectively. Software is used to identify the optimal spatial and temporal distribution of the radiation sources within the applicators of the implanted tissue or cavity. This approach of ‘dose-painting’ allows the brachytherapy team to graphically represent the distribution of irradiation of the treatment plan optimally tailored to the anatomy of each patient before the procedure begins Figure 5 [49–52].

During the surgery, applicators are inserted and correctly positioned in line with initial planner under trans-rectal ultrasound imaging guidance. After completion of the delivery of the radioactive sources, the applicators are carefully removed and the radiopaque seeds remain inside the patient permanently, which can be seen post-op via x-ray (Figure 6).

Between 2003 and 2012, brachytherapy had the largest decline in hospital stays post-OR procedures, dropping 24.4% in men age 45-64 and 27.3% aged 65-84, alluding to the quick recover time and safety of the procedure [54].

Figure 6. Post-operative x-ray image showing permanent radiopaque brachytherapy seeds in prostate [55]

1.3.2 Regional cancer chemotherapy

The major goal of regional chemotherapy administration to tumor-bearing organs is to achieve high drug concentrations for prolonged periods of time while sparing healthy tissue and organs from drug toxicity. Examples of successful regional chemotherapies include intraperitoneal treatment of advanced ovarian cancer, pelvic perfusion of rectal cancer, biodegradable implants for glioblastoma treatment, biodegradable films for the treatment of glial tumors of the brain, topical treatment of skin cancer, and many more [56–

60]. Each organ has different physiological properties that affect drug distribution that need to be considered.

Many research groups have conducted preclinical work on the development of intra-prostatic chemotherapy, with the aim to completely treat organ-confined prostate cancer and prevent local recurrence or metastasis [61–63]. Regional chemotherapy for prostate cancer can potentially lessen impotence and incontinence incidences compared to side effects of prostatectomy, radiation therapy, and cryotherapy. Additionally, the small 24 size of the drug molecules makes chemo-agents highly diffusive within an organ to hard- to-reach tumor cells.

Wientjes et al infused doxorubicin into the prostate to show high drug concentrations can be achieved. Although the diffusion within the prostatic tissue was uneven (Figure 7), the studies resulted in very low concentration of drug levels (6ng/mL) in plasma due to a natural barrier that blocks diffusion outside of the prostate via the fibromuscular stroma, which encases the prostate [62]. These studies demonstrate great potential and feasibility for further development of regional chemotherapies for prostate cancer with minimal toxicities outside of the target tissue.

1.4 Biodegradable drug delivery systems

The tunable properties and biocompatibility of biodegradable polymers have led to

their widespread use in controlled drug delivery. Currently, Zoladex, a PLGA implant for

delivery of Goserelin acetate in endometriosis and prostate cancer, and Gliadel, a BCNU

or carmustine loaded polyanhydride wafer for malignant gliomas. Zoladex is a subdermal 25 cylindrical implant injected subcutaneously in the upper abdomen. The Gliadel wafer implant is approximately 1.45cm diameter by 1mm thick and designed to pack the surgical cavity after resection to deliver carmustine directly to the remaining tumor cells [64,65].

Poly (lactic-co-glycolic) acid or PLGA is an FDA approved biodegradable and

biocompatible copolymer that has been widely studied for the use of drug delivery systems.

Its chemical formula is shown below in figure 8. PLGA has tunable mechanical and

degradation properties and is widely characterized for its extensive use in developing

devices for controlled delivery and release of many types of molecules, proteins, and drugs.

Many groups have shown that PLGA can be injected subcutaneously or intratumorally for

long-term controlled drug delivery without the need for surgery. The hydrophobicity of the

polymer can be altered by adjusting the ratio of the two polymers PLA and PGA, so that a

wide variety of payloads can be incorporated into nanoparticles, microspheres, or implants.

PLGA as an implant has been used intratumorally to deliver molecules, magnetic particles,

nanoparticles, drugs, and imaging agents to improve the bioavailability at the tumor site

and minimize any leaking of particles from the tumor [66–69]. Substantial research has

been conducted that supports the use of biodegradable PLGA implants in the use of cancer

treatment for prolonged drug delivery with minimal systemic side effects.

Figure 8. Poly D,L-lactic-co-glycolic acid, where x and y represent the number of times each monomer repeats [70] .

Preclinical studies have been conducted to fabricate biodegradable intraprostatic implants using copolymer PLGA for regional intraprostatic drug therapy to treat prostate confined cancer. Results showed release kinetics of doxorubicin from the PLGA cylinder relied heavily on the molecular weight of the polymer and percent loading of the drug, with burst release as high as 73% in twenty-four hours. In vivo canine models confirmed these results with rapid release of 80% of drug in 48 hours, however spatial drug distribution showed high concentrations confined to the lobule containing the implant and reaching concentrations 8 times higher than achieved with clinical intravenous injection. Results showed the implant caused necrosis within the lobule, sparing severe damage to the prostatic nerve bundles and urethra indicated feasibility in a PLGA-based implant for the delivery of regional prostatic chemotherapy [63].

1.5 Degradation of PLGA

Polyester PLGA is a copolymer of poly lactic acid (PLA) and poly glycolic acid

(PGA). It is the best defined biomaterial available for drug delivery with respect to design and performance. PLGA is generally an acronym for poly D,L-lactic-co-glycolic acid where D- and L- lactic acid, two enantiomeric forms to describe an asymmetrical α-carbon,

forms are in equal ratio. PGA is void of any methyl side groups and shows high crystalline

structure in contrast to PLA. PLGA can be processed into almost any shape and size, and

can encapsulate a wide range size of molecules. It is soluble in a wide range of common

solvents, including chlorinated solvents, tetrahydrofuran, acetone, or ethyl acetate [66].

PLGA biodegrades by the hydrolysis of its ester linkages into lactic acid and

glycolic acid as shown in figure 9. PLA is more hydrophobic than PGA due to its presence 27 of ethyl side groups. A PLGA copolymer lactide rich results in a less hydrophilic polymer which absorbs less water and subsequently degrades more slowly. The Tg (glass transition temperature) of PLGA copolymers is reported to be above the physiological temperate

37°C, hence they are glassy and rigid. Due to the hydrolysis of PLGA, glass transition temperature, moisture content, and molecular weight can change with time as degradation proceeds, resulting in a time-dependent and variant release of incorporated drug molecules.

Physical properties such as the size of the device, surface shape, exposure to water, and storage temperature, along with the type of drug incorporated into the polymer, all play a role in setting the release rate. Mechanical strength, swelling behavior, and capacity to undergo hydrolysis and subsequent biodegradation rate of the polymer are directly influenced by the degree of crystallinity of the PLGA, a parameter dependent on the type and molar ratio of PLA and PGA in the copolymer chain [71].

Figure 9. PLGA hydrolysis results in lactic acid and glycolic acid oligomers

PGA is crystalline in structure and when copolymerized with PLA, results in a

reduced crystalline structure and increases the rate of hydration and hydrolysis and

subsequently the degradation of the polymer. In general, higher content of PGA and lower

chain sizes lead to quicker rates of degradation with the exception of a 50:50 ratio of

PLA:PGA, which exhibits the fastest degradation due to its highly amorphous content. 28

PLGA copolymer undergoes degradation by hydrolysis or biodegradation through cleavage of its backbone ester linkages into oligomers, and then monomers. The degradation process for these polymers is mainly through uniform bulk degradation of the matrix where the rate of polymer degradation is slower than the rate of water penetration into the matrix. The increase of carboxylic end groups as degradation occurs autocatalyzes the process. The collective process of degradation is from bulk diffusion, surface diffusion, bulk erosion, and surface erosion, making the in vivo release rate pattern difficult to predict.

The biodegradation rate of the PLGA copolymer is influenced by molar ratio of the lactic and glycolic acids, the molecular weight of the polymer, the degree of crystallinity, and the

Tg of the polymer, resulting in a complicated release of drug from the degrading matrix

[72,73].

PLGA predominantly degrades by bulk erosion in two steps: (1) an initial burst of drug release related to drug type, concentration, and polymer hydrophobicity. Drug on the surface of the polymer is released as it comes into contact with the medium at a rate dependent on its solubility as well as water penetration into the polymer matrix. Initial burst of drug release is related to drug type, drug concentration and polymer hydrophobicity. (2) In the second phase, drug is released progressively through the thicker drug depleted layer. Water inside the polymer hydrolyzes the polymer into soluble oligomer and monomer products, which create channels for drug to be released by diffusion and erosion until the entire polymer solubilizes. Hydrophobicity of the drug can influence the rate at which the aqueous phase moves into the matrix. It is unclear how enzymatic activity in vivo plays a role in PLGA degradation, however some researchers suggest the 29 difference between in vitro and in vivo degradation rates can be attributed to enzymatic action during metabolic excretion of the foreign substance [72,74].

Overall, it is important to consider a number of parameters affecting degradation when designing a tunable PLGA drug matrix. (1) Effect of polymer composition : The

amount of glycolic acid is a critical parameter in tuning the hydrophilicity of the matrix

and thus the degradation and drug release rate. (2) Effect of crystallinity or glass transition

temperature (Tg) . Glass transition temperature varies depending on the composition of the

copolymer, the chemical structure of the drug, and the loading capacity of the drug in the

polymer. It is important to keep glass transition temperature over physiological conditions

in order to preserve the integrity of the implant. (3) Effect of average molecular weight .

Polymers with higher molecular weights have longer polymer chains, which require more time to degrade than small polymer chains. (4) Effect of the drug type . The presence of drug may change the degradation mechanism from bulk erosion to surface degradation.

The drug release profile is completely dependent on the drug type, and while the relationship is not easily predicted, it is clear that the chemical properties of the drug greatly affect the drug-release mechanism of a particular system. (5) Effect of the size and shape of the matrix . The ratio of surface area to volume is a significant factor for degradation where higher ratios result in higher degradation of the matrix and with drug release. (6)

Effect of pH. Both alkaline and strongly acidic media accelerate polymer degradation, however the difference between slightly acidic and neutral media is less pronounced due to autocatalysis of the carboxylic end groups. (7) Effect of enzymes. This effect is not entirely understood, however the difference between in vitro and in vivo degradation is hypothesized to be a result of enzymatic versus hydrolytic cleavage. (8) Amount of drug 30 loading . This plays a significant role on the rate and duration of release. Matrices having a

higher drug content tend to express a larger initial burst release than those with lower

content [66,71–74].

1.6 Docetaxel (Taxotere)

Docetaxel is part of the taxane family and a chemotoxic agent via mitotic inhibition.

Like paclitaxel, docetaxel’s mechanism of action works by reversibly binding to, and

stabilizing, microtubules. During cell separation, in the G2/M phase, microtubules are in

constant polymerization and disassembly due to dynamic instability, resulting in growing

and shrinking phases on microtubules. During polymerization, the tubulin dimers are

bound in a GTP bound state, where the α-tubulin is stable, however the GTP bound to β- tubulin can be hydrolyzed to GDP. GTP-bound tubulin continues to bind to the end of the microtubule causing growth, a term referred to as “rescue.” When hydrolysis catches up to the tip of the microtubule, “catastrophe” or rapid depolymerization occurs. This “search and rescue” model continues until the (+) ends of microtubules encounter kinetochores, a process required for chromosome separation and mitosis [75].

The principal mechanism of action of the taxane class of drugs is disruption of microtubule function, which is essential to cell division. Docetaxel stabilizes GDP-bound tubulin, preventing depolymerization and chromosome separation. This prevents the cell from continuing through mitosis and prolongs the time in the G2/M until eventual cellular apoptosis occurs [76]. This mechanism is shown in figure 10.

Figure 10. Taxanes stabilize microtubules, preventing mitosis, leading to eventual programmed cell death or apoptosis [adapted from 77] .

The G2/M phase is the most radiosensitive phase of the cell cycle, inherently making docetaxel a radiosensitizer. This theory is supported by two randomized clinical trials show docetaxel as a radiosensitizer [78,79]. Docetaxel is structurally similar to paclitaxel and was synthetically developed after Paclitaxel production proved too difficult and expensive for adequate drug supply. Docetaxel presents two chemical modifications that make it different from paclitaxel. It has a hydroxyl function group on carbon 10, whereas paclitaxel has an acetate ester, and a tert-butyl carbamate ester on the phenylpropionate side chain instead of the benzaminde in paclitaxel. These differences in structure are highlighted in Figure 11 The carbon 10 functional group causes docetaxel to be more water-soluble than paclitaxel, resulting in a more potent microtubule stabilizer with 3-fold better cellular uptake and 3-times slower efflux than that of paclitaxel [80].

Figure 11. Paclitaxel and Docetaxel are structurally identical other than a difference in two functional groups which are highlighted in gray.

The docetaxel formulation approved by the Food and Drug Administration (FDA) for human use is Taxotere, a formulation of lyophilized anhydrous docetaxel solubilized in polysorbate-80 and 14% ethyl alcohol then added to 0.9% Sodium Chloride for infusion.

For the treatment of prostate cancer, docetaxel is administered intravenously at 75mg/m 2 every 3 weeks for 6 administrations in the form of Taxotere [81]. The half-life of the drug is on the order of 84 hours (22 hours in mice) with poor bioavailability at the tumor site due to its high plasma protein binding affinity (>98%). Docetaxel undergoes hepatic metabolism via the cytochrome P450, CYP3A4, and CYP3A5 subfamilies of isoenzymes, meaning patients with liver disease do not qualify for this treatment [82].

1.7 Nanotechnology for oncology

1.7.1 Nanomedicine for chemotherapeutic delivery

Nanotechnology for medical applications, termed nanomedicine, has many advantages. Nanoparticles (typically 1-1000 µm) are on the same size scale as biomolecules such as receptors, antibodies, and nucleic acids [83–85]. The surface of 33 nanoparticles can be functionalized with biomolecules, enabling specific targeting to organelles or cells. Additionally, the surface can be modified for control over solubility, stability, and surface charge. Nanoparticles have a high surface area to volume ratio, allowing for a high payload (drugs, inhibitors, radioisotopes, siRNA, etc.) to be carried or encapsulated in the nanoparticle. Once at the target site, the high payloads can cause devastating damage to cancer cells. Nanoparticles can be targeted to tumors specifically using surface modifications and ligands, or passively via extravasation through leaky vasculature common in the tumor microenvironment [86,87].

Nanotechnology in cancer treatments is already a reality providing a wide range of new tools and possibilities, from earlier diagnostics and improved imaging to better, more efficient, and more targeted therapies. Nanomedicine offers the ability to deliver payloads specifically to tumor cells while masking the particle from the immune system for long circulation and fewer off-site toxicities.

Application of nanomaterials in biomedical field spurs the development of various formulations. For successful translation of these formulations, it is imperative to evaluate the design and properties of these nanoparticles. The applications of nanoparticles in oncology include enhanced drug delivery, efficient tumor targeting, treatment monitoring and diagnostics. The ‘theranostic properties’ associated with nanoparticles have shown enhanced delivery of chemotherapeutic drugs with superior imaging capabilities and minimal toxicities. While chemotherapy is frequently used for a first line of treatment, adverse toxicities and low bioavailability associated with the nonspecific systemic drugs often limit treatment timing and dosing in patients [88]. To overcome these toxicities and improve circulation time, many researchers are using nanoparticles as a means for 34 delivery[84,85]. Nanotechnology based delivery systems are making a significant impact on cancer treatment. Some important technological advantages of nanotherapeutic drug delivery systems include prolonged half-life, improved bio-distribution, increased circulation time of the drug, controlled and sustained release of the drug, versatility of route of administration, and increased cellular uptake cancer. A wide number of nanoparticles are being studies for medical applications including polymer-drug conjugates, ceramics, micelles, dendrimers, immunoconjugates, and liposomes, each of which offer unique properties, biocompatibility, targeting, and surface modifications [83].

1.7.2 Radiosensitization using nanotechnology

Several nanoparticle based formulations have been developed over years which have shown potential in either enhancing the effective radiation dose or reducing the normal tissue toxicities. There are several different mechanisms by which nanoparticles have shown promise in improving radiation therapy (RT) in different cancer models. The role of nanoparticles, in conjunction with improving the efficacy of RT, can be categorized into four major subcategories: (1) Increasing tumor response by radiosensitizing nanoparticle. As the name suggests radiosensitization is the use of agents, either drugs or

nanoparticles, to increase the sensitivity of the tumor cells to radiation [89,90].

Radiosensitization using nanoparticles is dependent on the type and design of nanoparticle.

These properties determine the mechanism by which nanoparticles will radiosensitize the

tumor tissue. Nanoparticles can themselves act as radiosensitizers (high atomic number

(Z)-nanoparticles) or can act as delivery vehicle for various known radiosensitizers

(metronidazole, tirapazamine, gemcitabine, docetaxel etc.) [91–93]. The efficacy of this 35 treatment is highly dependent on the efficacy of targeting the tumor cells. (2) Targeting molecular pathways to overcome radiation resistance. Despite advances in treatment planning, there are a number of patients who suffer local recurrence after radiation therapy.

Three main biological factors have been shown to be very critical in improving the outcome of radiation therapy, including the extent of hypoxia, ability of surviving cells to repopulate with in the treatment time frame and the development of resistance of tumor cells [94–97].

Modulation and targeting these factors by means of using inhibitors can efficiently improve outcome of treatment [53]. The use of nanoparticles can efficiently target the tumor site not only to modulate the microenvironment resulting in sensitization of hypoxic areas of tumor but nanoparticles can also be used to deliver inhibitors which target specific biological processes like DNA repair process. (3) Reducing non-specific radiation toxicity in normal tissue . The major concern of radiotherapy is the emergence of late-reacting tissue damage in healthy organs, which can take months to manifest. At the onset of radiation exposure, free radicals are formed though ionizing reactions that are highly reactive with

DNA and RNA, causing mutagenesis and cell death. Cellular oxidative damage is initiated by the generation of reactive oxygen species (ROS), which mechanistically change the oxidative state of cells and result in damage to mitochondrial function. While healthy cells have protective and repair mechanisms to combat these effects, they are not capable of blocking all of the damage which ultimately leads to normal tissue death. Because of these long-term effects, reducing radiation toxicity in patients treated with radiotherapy to protect normal tissues from radiation-induced damage is mandatory. Two major strategies to do so include 1) the use of radioprotectants to help defend healthy tissue from damage, and 2) the use of radiosensitizing agents at tumor sites to enhance the effects of radiation. 36

However, each of these treatment methods has side effects and toxicities of their own. The use of nanotechnology can be used to reduce toxicity of the radioprotectants and sensitizing agents, enhance targeting to selected tissue, and possibly reduce the radiation dose required for treatment. (4) Enhanced imaging efficacy . The success of radiation therapy relies immensely on the accuracy of imaging. The three basic steps involved in modern day radiotherapy is imaging, treatment planning and treatment delivery. The imaging steps provide patient’s structural and anatomical information using computed tomography (CT)

[98]. This information is transferred to treatment planning system in which tumor lesions are identified and target volume is defined. After a complex assessment of several factors, the patient is irradiated. Thus, both the steps of planning and delivery are entirely dependent on the quality and accuracy of the imaging. The role of nanoparticles in providing multimodal imaging platforms has been well studied and several promising platforms have been reported.

For the scope of our work, we will focus on increasing tumor response by means of radiosensitizing nanoparticles. Here we will explore the use of ultra-small silica nanoparticles to delivery radiosensitizing and chemo-toxic agent docetaxel for the treatment of prostate cancer. The aim of this approach is to increase the intracellular concentration of docetaxel in prostate cancer cells for enhanced radiosensitization during brachytherapy.

1.8 Research Strategy and overview of dissertation

The research strategy of this work is to fabricate a modified brachytherapy spacer

to deliver chemotherapeutic and radiosensitizer, Docetaxel over the course of weeks to 37 months for a sustained localized therapy. The overall goal for the project is to study the efficacy and toxicity of the localized treatment with and without radiation as compared to a conventional therapeutic regimen. The motivation and inspiration behind these spacers comes from an idea to modify and replace commercially used brachytherapy spacers which are commonly used in clinical applications via injection using an 18-gauge brachytherapy applicator needle. The project goals will be achieved through the following specific aims:

Aim 1: Fabrication, characterization, and optimization of biodegradable spacers.

Formulations using nanoparticles will be tested against free-drug embedded spacers. In

vitro and in vivo optical imaging will be used to evaluate the release kinetics and diffusion

properties of each spacer formulation to optimize the delivery profile. Silica Nanoparticles

(SiNPs) will be synthesized to encapsulate Docetaxel to aid in intracellular uptake at tumor

site. SiNPs will be chemically tagged with NIR-dye (model drug) for optical tracking.

Fluorescent nanoparticles and free dye will be embedded in a biodegradable polymer and

tested in vitro and in vivo for optimal release kinetics tracking fluorescent dye Cy7.5. The

distribution profile of the released “drug” in free form and in SiNPs from the doped spacers

will be estimated using non-invasive NIR fluorescence imaging of the live mice by IVIS

LUMINA II instrument. Each image will be converted to an intensity scale I(x,y,t) using a

colormap and pixel intensity will be plotted against space and time. The SiNPs distribution

estimated by fluorescence imaging will be correlated as an indirect evaluation of the drug

distribution subcutaneously and thus will provide an insight into the estimated distribution

pattern and behavior of NPs in the tumor matrix. Expected outcomes: Quantitative results

on pharmacokinetics, diffusion, and space-time DTX distribution will be obtained and will 38 set the stage for subsequent local chemotherapy (LCT) and local chemo-radiation therapy

(LCRT) experiments in Aim 2.

Aim 2: In vivo evaluation of docetaxel doped brachytherapy spacers versus conventional

chemotherapy to test as a monotherapy. In vivo evaluation of docetaxel (DTX) doped brachytherapy spacers’ synergistic effects as a radiosensitizer in combination with fractionated radiation to mimic brachytherapy dosing. I hypothesize that synergistic chemo-radiation therapy using localized delivery of DTX from brachytherapy spacers will provide improved therapeutic outcomes compared to RT or DTX spacer alone. This study will consist of two parts: (1) compare the efficacy and toxicity of local chemotherapeutic administration (LCT) with systemic IV administration of DTX of equivalent dose to determine the feasibility of an implant as a monotherapy. (2) Efficacy of the LCRT approach will be studied using fractionated external beam radiation therapy to mimic brachytherapy dosing in combination with the local drug spacer. The results of these studies will provide quantitative information on the localized treatment approach in animal models as a monotherapy and combinatorial radiosensitizer. The results will test the hypothesis that brachytherapy with localized DTX delivery leads to better outcomes than brachytherapy alone.

Aim 3: Determine associated toxicities from new treatment modality both with and without radiation in an immune competent mouse model. In vivo evaluation of toxicities associated with docetaxel eluding brachytherapy spacers compared to conventional chemotherapy and radiation will be assessed. Tumored immune-competent mice will be used for histological 39 and hematology studies at different times after treatments to study acute and chronic toxicities associated with each treatment group. I hypothesize that localized chemotherapy delivery of docetaxel from smart brachytherapy spacers will reduce systemic toxicities associated with conventional chemotherapy due to the localization in the tumor and generally lower drug doses. The study will compare the toxicity of local chemotherapeutic administration with systemic IV administration of DTX of equivalent dose, both with and without radiation. The toxicity of the local delivery is expected to remain in a local vicinity to the spacer and therefore local and systemic toxicities will be assessed. The results of these studies will provide necessary information of toxicities associated with this new treatment method in order to move into larger animal models for preclinical studies. It is expected that our local implant will be no more toxic than conventional chemotherapy.

A successful completion of these specific aims would set the foundation for moving forward into preclinical studies of a therapy in which a new biodegradable drug delivery system can be injected using current clinical procedures. Delivering chemotherapy and radiosensitizers directly into the tumor using our modified brachytherapy spacer allows for:

a. Tailored release profiles that achieves radiosensitization synchronized with the

radioactive decay rate for different sources such as 125 I, 131 Cs, and 103Pd during

image-guided brachytherapy. This ensures that the cancer remains continuously

radiosensitized during delivery of radiation, compared with intermittent

radiosensitization achieved with adjuvant systemic chemotherapy administered

every 3 weeks. 40 b. LCRT introduces dose painting of the tumor by controlling the sensitizing

parameters such as drug loading per spacer, release kinetics (by varying polymer

cross-linking) and number and spatial distribution of spacers, resulting in better

delivery of drug to the tumor. c. The synergistic effect of radiosensitization and radiation therapy could lead to

reduced radiation doses required, minimizing dosing to adjacent organs like the

rectum and to improved survival.

d. Significantly improved pharmacokinetics with minimal systemic side effects from

local intratumoral dosing thus in scaling up, the MTD in LCRT should be

determined by local (eg necrosis) rather than systemic toxicity. e. Spacers can be easily modified to include other drugs for other cancers (i.e. AuNP

for Pancreatic Cancer, Talazoparib for Breast Cancer)

Chapter 2. Fabrication of Silica Nanoparticles

2.1 Introduction

The applications of nanoparticles in oncology include enhanced drug delivery, efficient tumor targeting, treatment monitoring, and diagnostics. The “theranostic properties” associated with nanoparticles have shown enhanced delivery of chemotherapeutic drugs with superior imaging capabilities and minimal toxicities [83,86].

In conventional chemotherapy, only a fraction of the administered drug reaches the tumor site or cancer cells. For successful translation of these formulations, it is imperative to evaluate the design and properties of these nanoparticles. Organically modified siloxanes or ORMOSILS are organic/inorganic hybrid materials which offer unique combinations of properties unachievable by other materials. Many researchers have studied alkylsilanes and alkoxysilanes as components for hybrid materials because of their capability for surface modification, which has proved useful in adding targeting ligands, imaging moieties, and stealth components [99,100].

Here, we describe the design of ultra-small silica nanoparticles to encapsulate a radiosensitizing drug for combined chemoradiation therapy. The small size of nanoparticles allows for better dispersion and uptake of the drug within the highly vascularized tumor tissue. Unlike the common sol-gel method, ultra-small silica nanoparticles can be prepared using a hydrolysis and polycondensation reaction in a reverse micellar medium. Silica nanoparticles are synthesized using an oil-in-water microemulsion method. The microemulsion method provides a robust synthetic route in which the inner hydrophobic core is used to encapsulate chemotherapy drug, docetaxel while the outer hydrophilic region provides dispersibility of the synthesized nanoparticles 42 in an aqueous environment [101,102]. Docetaxel is commonly used for treatment of resistant or metastatic prostate cancer, and is known to have radiosensitizing properties. Here, we describe a systematic approach for synthesizing these theranostic nanoparticles for application in prostate cancer.

Silica nanoparticles can be formulated at various sizes, ranging from 10-250nm in diameter. The core is non-polar and can trap a variety of hydrophobic chemotherapeutics, small molecules, or dyes. The surface of the nanoparticles can be modified with PEG and/or targeting ligands for long term circulation and targeted cellular uptake. The

PEGylation of nanoparticles minimizes the opsonization process and allows for increased systemic circulation. The functional groups on the surface of the nanoparticles can be used for conjugating targeting molecules like peptides, antibodies or aptamers. Below, we describe the formulation of docetaxel loaded ultra-small silica nanoparticles for radiosensitization in prostate cancer. We conjugated a near-infrared fluorophore Cyanine

7.5 to the silica nanoparticles to track the nanoparticle in vitro and in vivo (Chapter 4) and load chemotherapeutic drug docetaxel into the hydrophobic regions of the nanoparticle to sensitize prostate cancer cells in vitro before radiation.

2.2 Materials and Methods

2.2.1 Synthesis and Characterization

Organically modified Silica Nanoparticles (ORMOSIL NPs) were synthesized using a previously reported method from Kumar et al. in the nonpolar core of an oil-in- water microemulsion [101,103]. Briefly, 2.2% surfactant AOT was dissolved in deionized water at room temperature under magnetic stirring at 1300rpm. Once dissolved completely, 43

300 µL Butanol-1 was added and stirred, followed by either 75 µL Docetaxel (10 mg/mL in DMSO), 50 µL Cy 5 (5mg/mL in DMSO), or 50 µL Cy7.5-silane precursor (Cy7.5 NHS ester (1.0 mg/mL), conjugated to aminopropyltriethoxy silane). After stirring, 100 µL

Triethoxyvinylsilane (VTES) was added to the reaction mixture and stirred for 45 minutes before a polymerization reaction was initiated with the addition of 10 µL ammonium hydroxide. The final reaction mixture was covered with foil to prevent photobleaching and allowed to stir overnight. Samples were dialyzed in cellulose membrane tubing with a cut- off size of 12-14kDa against deionized water at room temperature for 48 hours under gentle stirring. SiNPs were sterilized with a syringe filter (.45 µm) before use. To make larger particles of 200 nm in diameter size, the quantity of butanol-1 was increased to 700 µL and the stir speed decreased to 800 rpm.

Particle size was assessed using Dynamic Light Scattering (DLS) and confirmed using Transmission Electron Microscopy (TEM). Dynamic light scattering (DLS) measurements (90Plus zeta sizer, Brookhaven Inc., Holtsville, NY) measured the hydrodynamic diameter of the silica nanoparticles. TEM images (JEM- 100CX microscope, JEOL, Peabody, MA) were obtained using a at an acceleration voltage of 80 kV. The specimens were prepared by drop casting the sample dispersion onto an amorphous carbon coated 300 mesh copper grid. Drug loading was quantified using high performance liquid chromatography (HPLC). Quantification of docetaxel was determined via HPLC using an Agilent 1100 system and reverse phase C18 Zorbek column. A standard curve was prepared for free Docetaxel in the mobile at concentrations ranging from 1-1000

µg/mL. The drug was quantified using a 65:35 methanol:water mobile phase with a flow rate of 1.0 mL/min. Absorbance was detected at 254nm with t ret =5.97 minutes. The 44 standard curves were fit with a linear regression curve (R 2=1.00) and the curve was used to quantify the concentration of Docetaxel encapsulated in the silica nanoparticles. All solvents were HPLC grade and degassed under sonication before use. Before HPLC, silica nanoparticles (500 µL) were dissolved in dichloromethane (DCM) followed by extraction of Docetaxel using methanol. Particles were disrupted via sonication for 30 minutes then left over night in 4°C before evaporation of DCM under gentle air. Samples were diluted to 1 mL in mobile phase and spin filtered using 100kDa membranes to remove any particulates. Flow through containing the drug was collected and analyzed for drug loading.

All samples were run in triplicate.

Stability of the nanoparticle samples was measured daily for 7 days following

fabrication and dialysis using DLS for any size changes. Samples were stored at 4°C.

Release Studies were conducted in PBS (pH 7.4) at 37°C for 6 days. Silica nanoparticles

were places in a cassettes (membrane size 10kDa) and placed in buffer. Buffer was

removed at predetermined time points for HPLC analysis and replaced with fresh buffer to

maintain a concentration gradient favorable for release. All studies were conducted in

triplicates.

2.2.2 In vitro Efficacy

PC3 cell lines derived from human male metastatic prostate cancer were used for

all studies. All cells were grown in F-12K medium supplemented with 10% FBS and 1%

Penn/Strep. Confocal microscopy was used to image the uptake of silica nanoparticles

encapsulating fluorophore Cy 5 in PC3 cells. Cells were plated on 8-chamber cover slips

at 10,000 cells/well and left for 24 hours for attachment. A concentration of 0.1% silica 45 nanoparticles were incubated for 2 hours and 48 hour prior to imaging. Cells were washed twice with PBS and fixed with 10% formalin for 30 minutes. Formalin was removed and replaced with 0.1% Propidium Iodide (PI) in PBS and imaged using a Zeiss LSM 710 laser- scanning confocal microscope (PI Ex:533 nm, Em:617 nm; SiNP-Cy 5 Ex:684 nm, Em:710 nm) using a 20× or 40× objective.

IC50 values were calculated to determine the sensitivity of the cells to chemotherapeutic agent Docetaxel. PC3 cells were seeded into 96 well plates (Corning) at

10,000 cells per well. The following day cells were treated with Docetaxel concentrations ranging from 0 to 1000 nM in both free formulation and encapsulated in silica nanoparticles. Metabolic activity of the cell was measured 48 hours after seeding using an

MTS Assay (Promega) to determine cell viability. Dose response was plotted and fit using a variable slope four parameter logistic equation constrained at 0 to determine the IC50 value (Prism). All experiments were done in triplicate. Empty particles were tested for biocompatibility.

Clonogenic Assays were completed using DTX-SiNP, Empty Si-NP, Free DTX, and Control. Dosing was chosen to be the concentration of NPs which killed 25% of cells in the MTS assay so that a synergistic effect may be observed when combined with radiation. An equivalent dose of free DTX was used as comparison. For the clonogenic assay, cells were seeded in T25 and left to attach for 24 hours before treatment. Cells were treated for 48 hours with drug, SiNP with drug, or empty NPs before irradiation (0-8 Gy).

Cells were washed twice, media was replaced, and cells were left for 4 hours before replating at low density concentration (800-6000 cells) in 6-well plates. Cells were left untouched in a 37°C, 5% CO 2 incubator for 12 days to allow colony formation. Cells were 46 then washed twice with PBS, fixed in 10% formalin for 30 minutes, and stained with crystal violet. Colonies composed of 50 cells or greater were manually counted using a pen and used to estimate the surviving fraction as a function of the plating efficiency. To generate the radiation dose response curves, the data was fitted to a linear quadratic model using

CS-Cal software. For all radiation studies, cells received a single dose of 0 – 8Gy X-rays using a 220 kVp beam delivered at 13 mA (dose rate of 5.45 Gy/min) via a 0.15mm copper filter using a Precision Small Animal Radiation Research Platform (SARRP, XRad).

2.3 Results and Discussion

In conventional chemotherapy, only a fraction of administered drug reaches the tumor site or cancer cells. Here, ultra-small silica nanoparticles were designed to encapsulate a radiosensitizing drug for combined chemoradiation therapy. The small size of nanoparticles allows for better dispersion and uptake of the drug within the highly vascularized tumor tissue. Silica nanoparticles were synthesized with an oil-in-water micro-emulsion method. The microemulsion method provides a robust synthetic route in which the inner hydrophobic core is used to encapsulate chemotherapy drug Docetaxel while the outer hydrophilic region provides dispersibility of synthesized nanoparticles in aqueous environment. ORMASIL nanoparticles have been widely studied for applications in nanomedicine due to their tunable properties and surface modifications. In this work, we were able to covalently conjugate near-infrared fluorophore Cy7.5 to the silica nanoparticle for direct optical tracking of the particle. Near infra-red dyes have the advantage of high wavelengths and low energies, allowing for further penetration through tissue, making it a viable candidate for tracking particles for theranostics in vivo using optical imaging as 47 described in detail in Chapter 4. Dynamic Light Scattering (DLS) was used to determine the hydrodynamic diameter of the SiNP. DLS resulted in a Gaussian distribution of sizes centered about 36 nm as shown in Figure 12a. TEM confirmed a uniform and homogenous mixture of silica nanoparticles sized on average diameter 30nm (Figure 12b).

The size of the silica nanoparticle synthesized was independent of the molecule encapsulated (either Cy 5, Cy7.5, or docetaxel). The size stability testing over 7 days shows no significant size increase or decrease of the SiNP after 7 day storage in 4°C (Figure 13) which is consistent with the robust nature of ceramic materials. Silica nanoparticles as carriers for drug delivery are not easily degraded and therefore it is important to ensure the cytotoxic agent must be able to escape the nanoparticle in physiological conditions. 48

Figure 13. Stability of Silica Nanoparticles with docetaxel after fabrication, dialysis, and filtering showed no significant change in size during storage for 7 days in 4°C

Figure 14 shows after 24 hours only 23.8% of the docetaxel loaded into the particles is released, and after 48 hours 49.7% is released in buffer (37°C, pH 7.4). With an encapsulation of 32% after synthesis, dialysis, and sterilization, this is equivalent to the amounts of 5.712 µg/mL and 11.928 µg/mL. Between days 2-6, less than 5% of the docetaxel loaded is released which can be attributed to the stable nature of these nanoparticles. 49

Figure 14. Release profile of docetaxel from silica nanoparticles in buffer (37°C, pH 7.4) over 6 days.

Cell viability was assessed using an MTS assay on prostate cancer cell line PC3 cells for

24, 48, and 72 hours of incubation with docetaxel loaded silica nanoparticles to determine the sensitivity of the nanoparticles. Results showed IC50 values of 147nM, 4.633nM, and

2.218nM respectively (figure 15). The population doubling time for PC3 cells is 25hr, which may account for the large increase in cytotoxicity when incubating docetaxel loaded nanoparticles for 24 to 48 and 72 hours. Cells in different stages of the cell cycle may be undergoing mitosis until the nanoparticles are taken up and the docetaxel released, resulting in the delayed cytotoxic effects. It is also important to note that the release profile of the silica nanoparticles is a slow releasing with ~50% of the drug released after 48 hours which could account for the delayed cytotoxic effects in the IC50 results. 50

Figure 15. PC3 cell viability after incubation with DTX loaded SiNP for 24, 48, and 72 hours. IC50 values are reported in molar concentration.

Prostate cancer cell line PC3 showed high sensitivity to docetaxel loaded silica nanoparticles after 48 hours of incubation so cellular uptake was conducted after 2 hour and 48 hour incubation with Cy5 loaded SiNP to confirm the nanoparticles are translocated and remain intracellularly during this time period. RNA was stained with propidium iodide to determine the in vitro location of the particles after 2 hours and 48 hours of incubations.

Figure 16 is a confocal image of PC3 cells at 40x incubated with SiNP loaded with Cy5 for 2 hours and 48 hours. It is clear from the two images that very little difference exists in the uptake of the nanoparticles indicating either the nanoparticles are taken up and remain in the cytoplasm, or they are in constant recycling of uptake and efflux in the PC3 cells. It can be hypothesized that the SiNPs remain in the lysosome until the drug is released for the duration of the 48 hours, however it is important to mention that the dye is encapsulated 51 rather than chemically conjugated to the particle so it is unclear whether the nanoparticles remain intact during this duration. It is important to note that the ‘model drug’ or Cy 5 remains in the lysosomes and surrounds the nucleus until the cell passes through the G2/M phase where the nucleus dismantles and the nanoparticles can fill the cell as seen in the top left quadrant of the 48hr image.

Figure 16. Confocal image of PC3 cellular uptake of Cy5 loaded silica nanoparticles (red) and RNA (blue) after 2 hr and 48 hr incubation.

The radiosensitization properties of docetaxel makes it an attractive candidate for combined chemo-radiation therapy. The slow release and high cytotoxic effects of ultra- small docetaxel loaded silica nanoparticles give rise to their applications for prostate cancer treatment. A clonogenic assay was done to show the enhanced radiosensitivity of DTX loaded SiNPs compared to vehicle controls, and free docetaxel (figure 17). Vehicle controls showed no significant difference in surviving fraction from radiation controls at any radiation dose. The IC25, or the concentration that inhibits 25% of cell function, from docetaxel loaded silica nanoparticles after 48 hours on PC3 cells was chosen for dosing in all clonogenic studies. The IC25 of DTX SiNP, or 1.62nM, was selected for pretreatment 52 so that cells remained viable enough to allow for any enhanced cell sensitivity to radiation to be seen. Based on the IC50 (4.633e-009 M) and the hill slope of the curve (-1.044), this value was determined to be IC 25 = 1.62E-009 M. An equivalent dose was added for free

Docetaxel, and the same concentration of empty SiNP were added as a pretreatment for vehicle controls. While free docetaxel showed no radiation enhancement at any dose, docetaxel loaded silica nanoparticles incubated on PC3 cells for 48 hours had a significant improvement (P<0.05) from both controls and free docetaxel and radiation doses of 4, 6, and 8 Gy. Normalized surviving fractions were plotted using linear quadratic regression curves from the linear parameter α and the quadratic parameter β calculated separately for

each radiation dose–survival curve. Each experiment was done in triplicate for

significance.

Figure 17. Surviving fractions of empty SiNP, free docetaxel (DTX), and docetaxel loaded silica nanoparticles versus radiation at doses from 0-8 Gy using a small animal radiation research platform.

2.4 Conclusion

This work has shown enhanced therapeutic outcomes in vitro with encapsulated

docetaxel loaded silica nanoparticles of diameter 30nm when combined with radiation.

Silica NPs are robust and remain stable in 4°C storage, with slow degradation and release

(~50% of drug after 6 days) in buffer at physiological conditions. SiNP with DTX with

prolonged cellular uptake and low efflux after 48 hr incubation in prostate cancer cell line

PC3. Cell viability tests show high sensitivity to docetaxel in PC3 cells at 48 and 72 hours

with IC50 values of 4.633nM, and 2.218nM respectively. Clonogenic studies showed PC3

cells incubated with 1.62 nM docetaxel loaded silica nanoparticles for 48 hours had

significantly enhanced sensitivity to radiation doses or 4, 6, and 8 Gy to cells treated with

equivalent an equivalent dose of free docetaxel indicating a superior treatment using silica

nanoparticles for combined chemo-radiation therapy for the treatment of prostate cancer.

Chapter 3. Fabrication of INCeRT Spacers

In both brachytherapy and external beam radiation therapy, the use of inert and biocompatible spacers or fiducials are used to achieve high spatial accuracy during clinical procedure. These implantable devices are a critical technical component for accurate radiation delivery however they themselves provide no direct therapeutic benefit. Here lies an excellent opportunity to fabricate morphologically and physically similar spacers to not only provide the spatial and temporal accuracy during delivery of radiation therapy, but also incorporates a therapeutic component such as a chemotherapy drug, radiosensitizer, nanoparticles or imaging modality to provide in situ theranostics.

Here we aim to develop and fabricate a biodegradable modified smart spacer to be concomitantly delivered during brachytherapy for the treatment of prostate cancer. These modified smart spacers, termed INCeRT, or Implantable Nanoplatforms for Chemo-

Radiation Therapy, is a dual release platform implanted directly at the tumor site for enhanced and tunable control for the delivery of a number of molecular agents. Our platform consists primarily of a PLGA matrix which is extruded into the physical dimensions of current clinical brachytherapy spacers with the capability to degrade and release silica nanoparticles loaded with either chemotherapeutics, imaging agents, molecular inhibitors, etc. for enhanced cellular uptake and a secondary controlled release as shown in the schematic in figure 18.

. Figure 18. Schematic of INCeRT dual release platform with PLGA matrix embedded with Silica NPs loaded with chemotherapeutic or imaging moieties.

3.1 Introduction

Poly(lactic-co-glycolic) acid (PLGA) is an FDA approved polymer well known for its biocompatibility and biodegradability and is highly utilized and characterized for biomedical applications. For our application, PLGA was an ideal candidate due to the hyphobic flexibility it offers by altering the polymer ratios [72]. The copolymer is soluble in commonly used solvents and miscible with a number of molecular agents allowing for incorperation of a wide number of payloads, including hydrophobic chemotherapy drug docetaxel or silica nanoparticles loaded with dyes or drugs. PLGA has robust properties and can be fabricated into almost any size and shape as a drug delivery system [104]. Many groups have shown that PLGA can be injected subcutaneously or intratumorally for long- term controlled drug delivery without the need for surgery [63,69,104,105]. 56

It is critical to consider a number of parameters affecting degradation when designing a tunable PLGA drug matrix. (1) Effect of polymer composition : The amount of glycolic acid is a critical parameter in tuning the hydrophilicity of the matrix and thus the degradation and drug release rate. (2) Effect of crystallinity or glass transition temperature

(Tg) . Glass transition temperature varies depending on the composition of the copolymer,

the chemical structure of the drug, and the loading capacity of the drug in the polymer. It

is important to keep glass transition temperature over physiological conditions in order to

preserve the integrity of the implant. (3) Effect of average molecular weight . Polymers with

higher molecular weights have longer polymer chains, which require more time to degrade

than small polymer chains. (4) Effect of the drug type . The presence of drug may change the degradation mechanism from bulk erosion to surface degradation. The drug release profile is completely dependent on the drug type, and while the relationship is not easily predicted, it is clear that the chemical properties of the drug greatly affect the drug-release mechanism of a particular system. (5) Effect of the size and shape of the matrix . The ratio of surface area to volume is a significant factor for degradation where higher ratios result in higher degradation of the matrix and with drug release. (6) Effect of pH. Both alkaline and strongly acidic media accelerate polymer degradation, however the difference between slightly acidic and neutral media is less pronounced due to autocatalysis of the carboxylic end groups. (7) Effect of enzymes. This effect is not entirely understood, however the

difference between in vitro and in vivo degradation is hypothesized to be a result of

enzymatic versus hydrolytic cleavage. (8) Amount of drug loading . This plays a significant

role on the rate and duration of release. Matrices having a higher drug content tend to

express a larger initial burst release than those with lower content [72,104]. 57

The scope of this work utilizes past researchers’ experimental results to carefully select a copolymer with appropriate chain lengths and polymer ratios to fabricate a biodegradable implant for applications in brachytherapy during treatment of prostate cancer. The most common brachytherapy seeds are composed of I-125 with a radioactive half-life of 60 days [52]. The overall aim would be to sensitize the cancer cells throughout the duration of brachytherapy treatment locally using a biodegradable drug-eluding implant. This provides a substantial advantage over current clinical procedures in which patients receive a single dose of intravenous Taxotere once every three weeks resulting in intermittent periods of radiosensitization at the tumor site (figure 19) [82]. Here we utilize silica nanoparticles to load a NIR dye as a ‘model drug’ to optimize the PLGA matrix parameters and size of the nanoparticles which best achieve the degradation and diffusion we desire. These parameters can be tailored and altered to achieve degradation rates that align with other brachytherapy seeds for wide application and use.

3.2 Materials and Methods

The fabrication of silica nanoparticle doped PLGA spacers was carried out in two steps involving first the extraction of SiNP from aueous media and second, nanoparticle and polymer extrusion. Silica nanoparticles were prepared as directed in Chapter 2. Briefly, the synthesis of conjugated Cy7.5 silica nanoparticles (conCy7.5-NPs) was carried out using a slightly modified oil-in-water microemulsion method. To conjugate the Cy7.5-

NHS ester fluorophore into the NPs, Cy7.5-Silane was prepared by conjugating Cy7.5-

NHS ester to a silane precursor using (3-aminopropyl) triethoxysilane (APTES) under argon atmosphere. The NPs were prepared at room temperature with 2.2% AOT (wt/v) in

10mL HPLC water, followed by the addition of 1-butanol and Cy7.5-silane at 1200rpm stirring. After 15 minutes, vinyltrimethoxysilane (VTES) was added and the mixture continued to stir for 45 minutes, at which point, ammonium hydroxide was added and the reaction was covered and left to stir overnight. The nanoparticles suspension was then dialyzed against distilled water for 48h at room temperature, using a cellulose membrane with a cut-off size of 12-14 kDa. Following dialysis, the nanoparticles were sterile filtered

(0.45µm) and stored at 4 °C for future use.

After synthesis, dialysis, and filtering, nanoparticle formulations were transferred to a pear-shaped vial and placed in -80°C overnight before lyopholization. Following that,

250 mg of Poly(lactic-co-glycolic) acid or PLGA (50:50 PLA:PGA, 54-69 kDa) was dissolved in minimum amount of acetone. Lyophilized particles were collected, weighed, and resuspended in minimum amount of DMSO via sonication and vortex agitation then added dropwise to the PLGA solution. The entire mixture was mixed thoroughly to get a viscous paste, which was transferred to a 1ml syringe using a Luer stub adapter (0.5in; 59

18G). All air bubbles were removed and an infusion pump (Harvard Apparatus) was used to slowly inject the PLGA/SiNP mixture into a silion tube (inner diameter 0.8 mm) attached to the adapter. After infusion, the ends of the tube were clipped and dried overnight at

60°C. The dried spacers were cut into 3-5 mm length and stored at 4°C in the dark. To fabricate free Cy7.5 doped PLGA spacers, the same procedure was used substituting an equivalent concentration of Cy7.5 solution (5 mg/mL in DMSO) instead of the SiNP.

Transmission Electron Microscopy (TEM) images (JEM- 100CX microscope,

JEOL) were obtained using a at an acceleration voltage of 80 kV. The specimens were prepared by drop casting the sample dispersion onto an amorphous carbon coated 300 mesh copper grid. Dynamic light scattering (DLS) measurements (90Plus zeta sizer, Brookhaven

Inc.) measured the hydrodynamic diameter of the conCy7.5-NPs. To prepare samples for

Scanning Electron Microscopy (SEM), spacers were flash frozen in liquid nitrogen then fractured with a cooled razor. Fractured and non-fractured spacers were attached to a specimen mount using a conductive carbon adhesive tab and then coated with 10 – 15 nm of carbon using a vacuum evaporator (DV502, Denton Vacuum). Samples were imaged in both SE (secondary electrons) and BSE (back-scattered electrons) mode (S-4800 field emission SEM, Hitachi) at 5 kV. Energy-dispersive X-ray spectroscopy (EDS) analysis was carried out by selecting several region of interest (ROI) from the spacers and evaluating the elemental composition of the ROI.

3.3 Results and Discussion

PLGA has been fabricated into cylinder implants using solvent extruding methods for a number of biomedical applications. Zhou et al. used a 50:50 copolymer ratio of PLGA 60 to develop a multiple-drug delivery implant for intraocular management of proliferative vitreoretinopathy [68]. Zhang et al. studied the optimization of release properties via a biodegradable controlled antibiotic release devices for osteomyelitis [106]. Ortiz et al. fabricated a biodegradable intraprostatic implant slow release of doxorubicin [63]. Each of these investigators laid a heavy foundation studying the theoretical and experimented release kinetics of their PLGA-based drug delivery matrix. In each study, a solvent extrusion method was used similar to the method we use in our work, which involves dissolving the PLGA copolymer and incorperating the drug or antibiotic agent into the matrix before extrusion into a silicon tube. Results of each group impressed the importance of the molecule loaded, the loading capacity into the matrix, and the matrix parameters

(copolymer ratio, MW, inherint viscocity, etc.) on the resulting degradation and release kinetics of the molecule incorperated. From these results, we were able to select a polymer that would provide the desired degradation and release for our INCeRT platform. Table 3 depicts the commercially available PLGA copolymers and their properties [107]. Based on the results of others’ and the preliminary results aquired in our lab, we selected a 50:50 polymer blend with molecular weight 54-69 kDa. The expected degradation is less than 3 months, or 90 days, with a glass transition temperature above physiological conditions which meet the criteria set for use in brachytherapy.

Table 3. Listed copolymer properties of PLGA available for purchase (Sigma). lactide:glycolide Glass Transition ratio terminated PLGA MW (Da) Degradation Temp

50:50 acid 7,000-17,000 <3 months Tg 42-46 °C

50:50 acid 38,000-54,000 <3 months Tg 46-50 °C

50:50 ester 54,000-69,000 <3 months Tg 48-52 °C

65:35 acid 24,000-38,000 <5 months Tg 46-50 °C

75:25 acid 4,000-15,000 <6 months Tg 42-46 °C

Silica nanoparticles with near infrared dye Cy7.5 both encapsulated in (INCeRT-

1) and chemically conjugated to the particle (INCeRT-2) were prepared as described in

Chapter 2 and embedded into PLGA matrix as shown in figure 20. Two different sizes were synthesized: (1) average diameter of 30 nm and (2) average diameter of 200 nm to compare to release and diffusitivity of the nanoparticles from the matrix in vitro and in vivo and determine which had optimal degradation and release for the treatment of prostate cancer.

Figure 20. Photograph of modified smart brachytherapy spacers INCeRT-1 and INCeRT-2. INCeRT spacers are loaded with silica nanoparticles and near infrared dye Cy7.5 in PLGA matrix that is morphologically similar to commercial brachytherapy spacers for theranostic applications.

SEM images of flash frozen spacers show a smooth surface on the lateral view of the implant after fabrication. The topography of the cross-section from the top view shows a uniform distribution of silica nanoparticle ‘pockets’. The region of interest of one of these pockets is circled in red in figure 21. The molecular composition was analyzed using

Energy-dispersive X-ray spectroscopy (EDS) analysis was carried out by selecting several region of interest (ROI) from the spacers and evaluating the elemental composition of the

ROI. Regions outside of circular pockets showed no presence of silica, indicating the 62 nanoparticles formed large aggregate ‘macromolecule’ pockets between 3-5 µm in diameter. SEM confirms uniform distribution throughout the entirety of the PLGA matrix.

Figure 21. Cross sectional and lateral views of INCeRT spacer using SEM. Chemical analysis of homogenous spots confirm silica nanoparticle pockets embedded throughout spacer using EDS.

The degradation and release properties of INCeRT spacers were determined using optical imaging both in vitro and in vivo as described in chapter 4 for the purpose of optimization before testing their therapeutic efficacy. It is important to note that while in vitro release kinetics were studied, literature indicates enzymatic degradation of PLGA plays a role to an unknown extent in in vivo release kinetics. A complete literature search was done to estimate the polymer most suitable, however it is understood that the matrix components may need to be adjusted after in vivo assessment.

Chapter 4. In vivo optimization of ‘Smart Brachytherapy Spacers’ using optical imaging

Drug loaded implants are a new, versatile technology platform to deliver a localized payload of drugs for various disease models. One example is the Implantable Nanoplatform for Chemo-Radiation Therapy (INCeRT) where inert brachytherapy spacers are replaced by spacers doped with nanoparticles loaded with chemotherapeutics and placed directly at the disease site for long term localized drug delivery. However, it is difficult to directly validate and optimize the diffusion of these doped nanoparticles (NPs) in in vivo systems.

To better study this drug release and diffusion, we developed a custom macroscopic fluorescence imaging system to visualize and quantify fluorescent NP diffusion from spacers in vivo . To validate the platform, we studied the release of free fluorophores, and

30 nm and 200 nm NP conjugated with same fluorophores as a model drug, in agar gel

phantoms in vitro and in mice in vivo . Our data verified that the diffusion volume was NP

size-dependent in all cases. Our near-infrared imaging system provides a method for by

which NP diffusion from INCeRT spacers can be systematically optimized (e.g. particle

size or charge) thereby improving treatment efficacy of the platform.

4.1 Introduction

Implantable drug delivery systems placed at the site of disease represents an

appealing approach to delivering a number of different classes of drugs that cannot be

effectively delivered via more common routes such as intravenous, oral, or topical

administration. An implantable drug reservoir can increase the bioavailability of drug 64 without the need for multiple injections [108]. To prevent a secondary intervention for drug reservoir removal, many implantable devices have been made biodegradable. The degradation rate of the implant combined with the diffusivity of the drug makes understanding and predicting the release profile in vivo a challenge. For example, polymer-

Dexamethasone implants are used for intravitreal dosing for persistent macular edema, showing significant dexamethasone levels up to 6 months post implantation. However, detection of drug levels was performed using common techniques, such as liquid chromatography-tandem mass spectrometry in sacrificed animals, which greatly limits the amount of data collected to a single time point [109]. Similar post-sacrifice analysis techniques or complex biopsies were used in determining diffusion profiles of clinically used biodegradable implants in delivering insulin, steroids, chemotherapeutics, antibiotics, or analgesics [108]. In this work, we aim to develop an imaging platform which can be used to understand the mechanism and quantification of the drug diffusion from these implants. We also aim to evaluate the size-dependent diffusion properties using real time diffusion in live animals by studying different sized fluorescent silica nanoparticles (NPs) incorporated in these implants.

We have fabricated a modified brachytherapy spacer for application in localized drug delivery during the brachytherapy procedure. During brachytherapy, clinicians use identical but inert spacers to achieve spatial and temporal distribution of radioactive brachytherapy seeds for highly controlled radiation dosing of the tumor. However, these spacers themselves do not currently have any therapeutic function [110,111]. Since these spacers are frequently used as internal radiation therapy guides in clinics, this offers an opportunity for in-situ delivery of drugs by fabricating spacers which are doped with 65 chemotherapeutic drugs or radiosensitizers, without additional discomfort to the patient.

Our group leveraged this opportunity and fabricated “Implantable Nanoplatform for

Chemo-Radiation Therapy” (INCeRT) spacers for localized delivery of therapeutics [112].

While systemic chemotherapy administered intravenously is routinely used, both with and without radiation therapy in the management of cancer, poor tumor accumulation and systemic toxicity due to intravenous delivery are a major limitation when creating a personalized treatment plan [113–115]. Localized delivery of a chemotherapy agent would eliminate the high systemic toxicities while allowing for an effective and concentrated dose directly at the tumor site. The robust INCeRT platform can be tailored for a controlled and continuous release of the drug/radiosensitizer from the spacer for a sustained local delivery of a high therapeutic dose optimal for long term brachytherapy treatments. The size, shape and delivery procedure for INCeRT spacers are identical to spacers used in current clinical settings, with the added benefit of local chemotherapy. This will act as a slow-release drug reservoir for simultaneous combinatorial chemo and radiation therapy. However, to prove our concept of slow and sustained drug release from INCeRT implants, it is imperative to study the release of the drug and diffusion properties in real time using a non-invasive method. Thus, to demonstrate this, we have used NIR fluorophore, Cyanine 7.5 as a model drug to incorporate in the INCeRT spacers and study the diffusion in real time using optical imaging methods. We have further incorporated Cy7.5 in the NPs to evaluate the size dependent release and diffusion from the spacers. This release is conceptually illustrated in figure 22a , demonstrating the nanoparticle release from the spacer and the subsequent

release of the therapeutic agent from the nanoparticle. The use of Cy7.5 conjugated NPs in 66 the spacer allows for studying the diffusion properties of materials with different molecular weights or sizes when embedded in the implants or spacers.

Figure 22. (a) Illustration of intradermal spacer diffusion. As the spacer degrades, fluorescent silica nanoparticles are released for a slow sustained diffusion. Further, drug encapsulated in the nanoparticle can be released. A TEM image of (b) 30 nm and (c) 200 nm silica nanoparticles. (d) A series of SEM images of a flash frozen and fractured spacer with uniform distribution of nanoparticles. As magnification of the cross sectional area is enhanced, silica pockets are observed.

Fluorescence imaging has a number of advantages for this application, including

the ability to perform non-invasive serial imaging over time on the same animal(s), the

broad availability of biocompatible cell labeling techniques (red fluorescent proteins,

AlexaFluor dyes, and Cyanine dyes), the lack of ionizing radiation, and the relatively low

cost of the instrumentation involved [116–118]. To achieve this, we developed and

validated a new broad-field fluorescence imaging system to observe NP diffusion in bulk

biological tissue. This allowed us to systematically quantify diffusion both in phantoms in

vitro and in mice in vivo . Moreover, we chose to image at NIR “diagnostic window”

wavelength region (corresponding to approximately 650-850 nm), where light absorption

and scatter is minimal compared to the visible range [118]. We performed in vitro studies 67 with free AlexaFluor 750 dye and NPs of different sizes (30 nm and 200 nm) in agar gel phantoms in vitro and analyzed the diffusion over time. We also performed studies for 2 weeks in mice tracking and analyzing the diffusion in vivo . As we show, we were able to quantify size-based diffusion in both models. As such, this technology represents a new platform by which we can optimize INCeRT NP characteristics – for example with respect to size and surface chemistry - to achieve sustained chemotherapy release from the spacers over weeks or months in vivo.

4.2 Methods and Materials

4.2.1 Synthesis, fabrication, and characterization of Silica nanoparticles and Spacers

Synthesis of silica nanoparticles and INCeRT Spacers can be found in detailed description in Chapters 2 and 3. Transmission Electron Microscopy (TEM) images (JEM-

100CX microscope, JEOL, Peabody, MA) were obtained using a at an acceleration voltage of 80 kV. The specimens were prepared by drop casting the sample dispersion onto an amorphous carbon coated 300 mesh copper grid. Dynamic light scattering (DLS) measurements (90Plus zeta sizer, Brookhaven Inc., Holtsville, NY) measured the hydrodynamic diameter of the conCy7.5-NPs. To prepare samples for Scanning Electron

Microscopy (SEM), spacers were flash frozen in liquid nitrogen then fractured with a cooled razor. Fractured and non-fractured spacers were attached to a specimen mount using a conductive carbon adhesive tab and then coated with 10 – 15 nm of carbon using a vacuum evaporator (DV502, Denton Vacuum, Moorestown, NJ). Samples were imaged in both SE (secondary electrons) and BSE (back-scattered electrons) mode (S-4800 field emission SEM, Hitachi, Chiyoda, Tokyo, Japan) at 5 kV. Energy-dispersive X-ray 68 spectroscopy (EDS) analysis was carried out by selecting several region of interest (ROI) from the spacers and evaluating the elemental composition of the ROI.

4.2.2 Fluorescence imaging system

A schematic diagram and photograph of our custom designed fluorescence imaging system is shown in figures 23a and b, respectively. Samples (mice or agar phantoms) were placed on a platform and trans-illuminated with the output a tunable Titanium:Sapphire laser (Mai Tai, XF-1, Spectra-Physics, Inc., Santa Clara, CA). We chose to use a near- infrared fluorophore and laser combination because attenuation and autofluorescence of biological tissue is well known to be significantly reduced versus blue or green wavelength ranges [119]. The output wavelength of the laser was set to 740 nm (which matched the absorption spectra of the AlexaFluor750 and Cy7.5 dyes). The output of the beam was filtered with a 730 nm “clean up” filter with 30 nm bandpass (z730/30 nm, Chroma

Technology, Rockingham, VT) to remove any out-of-band near infrared (NIR) laser output. The 1 mm diameter beam was expanded to approximately 150 cm in diameter using a lens pair in simple telescope configuration (f = 50 mm and 200 mm, Edmund Optics,

Barrington, NJ) followed by a third expanding plano-concave lens (f = -25 mm, Edmund).

The power at the sample was adjusted using a set of neutral density filters and was approximately 15 mW (equivalent to 0.015 mW/cm 2) for these experiments. A ground glass diffuser plate was placed under the samples to create a relatively homogeneous illumination profile and remove any speckle or interference pattern from the illumination light.

Samples were imaged with a high-sensitivity electron-multiplied charge coupled device camera (EMCCD, iXon EM +855 Andor Technology, Belfast, Northern Ireland) fitted with a 35 mm lens (NT54-689, Edmund). The imaging field of view was 105 mm x 105 mm at a working distance of 47 cm. Each image was 1004 x 1006 pixels. Camera exposure time and image acquisition was controlled using a personal computer running Andor control software. Although the camera could be operated with intensifier gain, this was turned off for the experiments described below. The camera was fitted with either a 740 nm filter (to record “intrinsic” images at the laser wavelength), with a 40 nm bandpass

(et740/40m, Chroma) or a 780 nm longpass filter (et780lp, Chroma) to collect fluorescent light from the sample. The filters were mounted in a five-position filter wheel (LCFW5,

Thorlabs , Newton, NJ). White light images of the sample were acquired by moving the

filter wheel to an open position so that all wavelengths were detected by the camera. Mice

were top-illuminating with a white light source (Fiber Optic Illuminator Model 190, Dolan- 70

Jenner Industries, Lawrence, MA), whereas clear agar phantoms were back-illuminated with a LED ring (Digi-Slave L-Ring 3200, Edmund).

4.2.3 In vitro Imaging of phantoms

Agar phantoms were placed on the imaging platform and trans-illuminated as described above. The movement of free AlexaFluor (AF) 750 dye (Molecular Probes, Life

Technologies, Carlsbad, CA), 30 nm nanoparticles, and 200 nm nanoparticles were tested in separate experiments to observe the effect of particle size on diffusion. Free dye and nanoparticles were injected directly into the clear phantom with an insulin syringe. The phantoms were made in 250 mm diameter by 25 mm deep cell culture dishes (Corning Inc.,

Corning, NY). The phantom material was made with 1.5 grams of Agarose (Mr=-0.10,

Acros Organics, Geel, Belgium) and 250 mL of distilled water resulting in 0.6% agar gel.

The mixture was boiled for 10 minutes until the powdered agar was dissolved and then poured into the cell culture dish to cool and harden [120,121]. For experiments involving free dye, AF-750 was injected into the agar phantom and images were acquired at 15 minute intervals for a total of 3 hours (N=3). In this case, it was possible to continuously make measurements without removal of the phantom from the stage. Exposure times were

1 ms for intrinsic images (i.e. at the wavelength of the laser) and were between 0.5 and 4 s for fluorescence images, depending on the acquisition time point (later times required longer exposures due to dye diffusion and dilution). For experiments involving nanoparticles (30 nm and 200 nm diameter), the phantoms were imaged daily for 15 days after the injection of the nanoparticles (N=3). To prevent drying of the agar, these phantoms were refrigerated between experiments. A wire marker was placed in the upper left hand 71 corner to allow image co-registration between imaging sessions. For nanoparticle formulations, fluorescence images required between 0.5 and 30 s exposures (as above, longer exposures were required for later time points). White light images were taken of the phantoms by back-illumination with the LED ring during each imaging session.

4.2.4 In vivo imaging of spacers

Animal experiments were conducted in agreement with all relevant guidelines and regulations set by the Northeastern University, and with approved institutional protocols by the American Association of Laboratory Animal Care. Six-six week old male athymic nude mice (Hsd:Athymic Nude-Foxn1nu) were procured from Charles River lab, and were housed in a group of six in standard cages with free access to food and water and a 12 hrs light/dark cycle. All animals acclimated to the animal facility for at least 48 hrs before experimentation. All possible parameters that may cause social stress, like group size, type

(treated and non-treated) etc., among the experimental animals were carefully monitored and avoided. Animals were observed daily for any behavioral abnormalities and weighed weekly.

INCeRT spacers were implanted subcutaneously in mice using a sterile clinical brachytherapy applicator needle. Two nanoparticle or free-dye coated spacers were implanted in the left and right rear flanks of nude mice (N = 4 spacers per case). Three groups of spacers were studied: (1) PLGA Spacer doped with Cy7.5 free dye, (2) PLGA

Spacer doped with conCy7.5-NPs with average diameter 30nm, (3) PLGA Spacer doped with conCy7.5-NPs with average diameter 200nm. Each mouse was imaged once per day for 15 days following implantation. It is important to note that the mice were imaged in trans-illumination mode (as opposed to reflection mode). As we discuss, trans-illumination 72 imaging (combined with image normalization) is significantly more quantitatively accurate when imaging in deep tissue with significant optical inhomogeneities. During imaging, mice were placed on a custom mask to avoid saturation of the camera by direct laser illumination. Three sets of images were acquired during each imaging session as follows:

First, white light images were acquired with a 1 s exposure time. Second, intrinsic images

(at the wavelength of the laser) were acquired with the 740 nm filter in front of the camera and 5 ms exposure times. This “intrinsic” image was used for normalization of the fluorescence images to improve quantitative accuracy. 16 Third, fluorescence images were acquired with the 780 nm filter in front of the camera and exposure times was between 7.5 and 100 s, depending on the imaging time point.

4.2.5 Image Processing and Data Analysis

We first performed basic pre-processing and normalization as follows. First, a dark field image was acquired (by capping the lens) and then subtracted from each image. All images were normalized to exposure time and laser intensity (which was measured on each day) so that the units were in photon counts/s/mW. Next, fluorescence images were normalized by pixel-by-pixel division by the corresponding intrinsic image. Other authors have shown previously that this normalization operation improves the accuracy of fluorescence images

– particularly when imaged in trans-illumination geometry – since it accounts for the laser illumination profile and minor differences in illumination intensity. Normalization also improves quantitative accuracy in the presence of tissue optical property heterogeneities and improves imaging depth [122]. 73

Phantoms: Each of the data sets for agar phantoms injected with either free AF750 dye, 30 nm, or 200 nm nanoparticles was analyzed by first finding the center of the injection site using a custom written MATLAB (MathWorks, Natick, MA) routine. Images were then averaged radially from this center point for each time point. These intensity curves were then fit to a simple one dimensional diffusion equation with a point source:

 2   − x  A  4Dt  I(x,t) = o e  4πDt

where I(x,t) was the intensity at each time and position, A0 was the total amount of diffusing substance, D was the diffusion coefficient, t was the time following injection, and x was

the radial distance from the source. 19 The equation was fit to the averaged experimental data for each injection type to yield the diffusion coefficient D. The complete spatial and temporal profiles were used in each fit for each case.

In Vivo Data: The diffusion process in vivo is significantly more complex than in agar phantoms - for example, due to active transport of nanoparticles away from the site of injection, tissue boundaries and inhomogeneities, light diffusion – which prevented fitting of the simple diffusion equation as above. Therefore, diffusion of nanoparticles and AF750 dye from the implanted spaces was quantified using three additional metrics: i) diffusion area, ii) diffusion profile, and iii) spacer intensity. For the first (diffusion area), we considered the maximum fluorescent intensity for each time point and then computed the fluorescence area for pixels exceeding 50% of this value on each day (i.e. the “full width at half maximum area”). Pixel numbers were converted to area (in mm 2) using the 74 calibrated image pixel size of 0.011 mm 2. This area was calculated for each spacer and for

each time point (N=4 for each). Second, we computed intensity (diffusion) profiles for the

spacers at each time point. Since spacer orientation differed between mice, we first applied

a thresholding routine to automatically identify the long and short axes of the cylindrically

shaped spacer in the in vivo fluorescence images. Next, the fluorescence profiles were

computed by averaging 11 lines orthogonal to the long axis of each spacer. Third, the mean

fluorescence intensity of each spacer was computed as a function of time.

4.3 Results

4.3.1 In Vitro Characterization of Nanoparticles, Spacers, and phantom imaging

Dynamic Light Scattering (DLS) was used to determine the hydrodynamic diameter of the conj.Cy7.5-NP. Two batches were prepared: one with an average diameter of 30 nm and the other with an average diameter of 200 nm particles. These sizes were confirmed using TEM as shown in figure 22b and 22c respectively. Spacers were flash frozen then fractured and mounted for SEM imaging. The cross-sectional image in figure 22d shows an overall uniform distribution of the silica across the face of the spacer with additional panels showing the distribution magnified. Chemical analysis confirmed the pockets in the region of interest to be silica nanoparticles.

Example fluorescence images measured from agar phantoms injected with

AlexaFluor-750, 30 nm nanoparticles and 200 nm particles are shown in figures 24a-d,

24e-h and 24i-l, respectively.

Figure 24. (a-d) Fluorescence image sequence of a representative phantom with AlexaFluor 750 dye diffusion at 1 minute, 15 minutes, 30 minutes, and 60 minutes respectively. (e-h) Fluorescence image sequence showing 30 nm nanoparticle phantom diffusion for 1 minute, 6 days, 11 days, and 15 days. (i-j) Fluorescence image sequence showing 200 nm nanoparticle phantom diffusion for 1 minute, 6 days, 11 days, and 15 days. All images are normalized to the maximum at minute 1.

As is evident from these image sequences, the free dye diffused significantly more rapidly than the nanoparticles (i.e. on the order of minutes versus days) due to their larger size. As described above, normalized corrected fluorescent profiles were computed and are shown in figures 25a-c for the three sets of injections. By inspection of these figures, diffusion of the nanoparticles was small but significant over the 15 day period, whereas the diffusion of the free dye occurred rapidly over the span of a few hours. In all cases, the measured fluorescence intensity steadily decreased over time. This is shown explicitly in figure 25d, where the maximum intensity is shown as a function of measurement time (intensity data 76 for the free dye is shown in figure 25d inset). In combination, figures 25a-d indicate the anticipated size-dependent diffusion behavior, i.e. that the larger nanoparticles diffused less than either the smaller nanoparticles or the free dye. Likewise, the 1-dimensional diffusion equation was then fit to the data from figures 25 a-c. The resulting diffusion coefficients (D) for each of the three sets of phantoms are shown in figure 25e. As shown, the diffusion coefficient for the free dye was 24.92 mm 2/day, 0.02 mm2/day for 30 nm NPs, and 0.01 mm 2/day for 200 nm NPs. This agrees well with visual inspection of the diffusion behavior in figures 4a-d and the anticipated size-dependency.

Figure 25. Diffusion (fluorescence) as a function of distance from the phantom center, averaged over all phantoms imaged (N=3) and normalized to the intensity at x=0 for each time point. Data for (a) free AF750 dye, (b) 30 nm nanoparticles, and (c) 200 nm nanoparticles are shown. (d) Average maximum fluorescence intensity normalized to initial time point per phantom type showing the continual decrease in maximum fluorescence over time. Inset time scale of free AF750 dye diffusion is on the order of minutes compared to days in part (d). (e) Diffusion coefficients found by fitting the diffusion profile curves to Eq. 1 on a logarithmic scale.

4.3.2 In Vivo Imaging

Spacers coated with either free dye (Cy7.5), 30 nm or 200 nm nanoparticles were implanted in the rear flanks of nude mice (N = 4 for each) and then imaged daily for 15 days. Example representative image sets obtained for implanted spacers at a single time point in vivo is shown in figure 26. A white light image of the mouse is shown in figure

26a, with the position of the implanted spacers indicated by the red arrows. The normalized

fluorescence image is shown in figure 26b. For anatomical reference, an overlay of the

normalized fluorescence image over the white light image of the animal is shown in figure

26c. After the animals were euthanized, we also removed the skin layer and the spacers to

image the residual fluorophores and nanoparticles without light scatter due to the skin

tissue. As shown in figure 26d (in this case for Cy7.5 coated spacer), the residual

fluorescence was still clearly visible and demonstrates the continuous diffusion from the

spacer over time.

Time-series images for each of the 3 spacer types (representative images) are shown in figure 27. As is indicated, all of the spacers decreased in intensity over time due continuous diffusion of the fluorescence from the spacers, and decrease was observed for spacers coated with nanoparticles than for free dye.

Likewise, the fluorescence area appeared to increase over time due to this diffusion process. This is shown explicitly in figure 28. The diffusion profiles for the free dye, 30 nm and 200 nm nanoparticles as a function of distance from the spacer over time are shown in figures 28a-c, respectively. A small amount of diffusion (~1 mm) beyond the profile measured on day 1 is evident in all cases (we note that the diffusion profile observed on day 1 is principally due to light diffusion in scattering biological tissue, as opposed to fluorophores or nanoparticle diffusion from the spacer). The normalized fluorescence intensity for each case is shown in figure 28d, indicating continuous diffusion of the fluorophores and nanoparticles over time. Likewise, the full width at half maximum fluorescence area is shown in figure 28e. As shown, the area of diffusion reached its maximum on day 4 for the 30 nm nanoparticles and day 6 for the 200 nm particles, 79 indicating a size-dependence of the process (i.e. diffusion was slower with the larger nanoparticles). The fluorescence area decreased after this maxima, presumably due to active transport of the nanoparticles away from the site of the spacer, for example by blood vessels.

4.4 Discussion and Conclusion

INCeRT represents a new strategy for local tumor control by which previously inert spacers are doped with drug-delivering NPs that diffuse from the biodegradable implant continuously over time. Ideally, NPs will be optimized so that they diffuse several mm 80 from the implant site and are retained in the tissue for weeks, i.e. to treat a large tumor volume for an extended time period. Optimization of these diffusion properties is difficult to measure in vivo . Therefore, in this work, we developed and validated a new custom fluorescence imaging platform to directly visualize and quantify diffusion of NPs in gel phantoms in vitro and in mice in vivo . In both cases, we observed size-dependent diffusion profiles as expected. In phantoms, this was a straightforward relationship which was well modeled by a 1-D diffusion equation. In vivo , size dependent diffusion was also observed,

but was significantly more complex as we have noted above, due to both light propagation

effects in biological tissue and active transport mechanisms which carry away NPs in

tissue. However, in this proof-of-concept study we were able to confirm that NPs are

indeed observed in tissue for weeks after implanting INCeRT spacers and are therefore not

rapidly removed from the site.

This work provides a platform by which the NPs can be further optimized for

diffusion characteristics both in vitro and in vivo , for example, with respect to size, surface

chemistry or target tissue type. Future work involves diffusion from INCeRT spacers in

tumors, more specifically in prostate cancer models where brachytherapy is often used as

a salvage therapy in patients clinically. Understanding the release intratumorally will help

optimize the spacer parameters to deliver the chemotherapeutic/radiosensitizer at sufficient

doses for therapeutic efficacy for the entirety of the brachytherapy dose regimen. Our

research interests lie in treating early stage, slow progressing, non-metastatic prostate

cancer, and in salvage cases such as recurrence, where brachytherapy or internal radiation

therapy is a common mode of treatment. During brachytherapy, clinicians space

brachytherapy seeds with inert biocompatible spacers for highly controlled dosing of the 81 prostate [123,124]. Recurring tumor treatment is limited by previously given treatments due to the toxicity to the patient [78]. Adding an implant to brachytherapy can reduce the toxicity seen to healthy tissue while radiosensitizing the local prostate tumor for a greater chance of therapeutic effect, acting as a slow release drug depot for simultaneous combinatorial chemo and radiation therapy. The robust platform can be tailored for a controlled and continuous release of the drug/radiosensitizer from the spacer for a sustained local delivery of high therapeutic dose optimal for long term brachytherapy treatments.

As we have demonstrated, near-infrared imaging has distinct advantages in that it allows us to image the same animal serially over the course of weeks, non-invasively and at a relatively low cost. The imaging platform demonstrated in this work can be used to determine the release profile of these spacers to alter the parameters of the implant for an optimized release profile to mimic that of what is seen in brachytherapy for sustained combinatorial treatment.

Chapter 5. Docetaxel loaded implants as a monotherapy for prostate cancer

5.1 Justification for implant reformulation

Results from the in vivo degradation and release kinetics of INCeRT spacers using optical imaging resulted in the reformulation of the matrix. The diffusivity constant of the

30nm and 200 nm silica nanoparticles was very low compared to that of a free flowing dye.

When embedded in the PLGA matrix, the surface charge of the nanoparticles resulted in large aggregated pockets of particles which behaved as microparticles on the order of 2-5

µm in diameter as confirmed by the cross-sectional SEM images. Additionally, it was apparent from the in vivo imaging that there was a maximum diffusion of the nanoparticles

within the first week when the spacer was placed subcutaneously on the flank. This is most

likely a result of the complex biological processes such as the reticuloendothelial system

(RES) which work to capture and degrade foreign particles[125]. As a result, INCeRT

spacers were optimized to: (1) remove any use of silica nanoparticles. Silica nanoparticles

aggregate and undergo phagocytosis via macrophages. While the use of nanoparticles

could provide unique opportunities, silica nanoparticles have not been approved for

biological applications in humans and for the purpose of translation through the clinic, only

FDA approved materials should be used. Additionally, the tumor matrix is dense and

fibrous and literature suggests particles >12 nm will be unsuccessful in complete diffusion

and tumor dosing [126]. (2) The copolymer should be adjusted for a faster release. Previous

studies have shown that Doxorubicin embedded in PLGA (50:50, 7-19 kDa) degrades as a

function of its loading capacity [63]. A decrease in drug loading resulted in a prolonged 83 degradation with a lower burst release. While the drug appeared to degrade in a number of weeks, it is important to take into account the chemical properties of the drug loaded into the matrix. Doxorubicin hydrochloride has a molecular weight of 579.98 g/mol and is soluble in water at 10 mg/mL, DMSO at 100 mg/mL, and very poorly soluble in ethanol

[127]. Comparatively, Docetaxel has a molecular weight of 807.4 g/mol and in very poorly soluble in water, soluble in DMSO at 200 mg/mL and ethanol at 100 mg/mL [128].

Docetaxel is much more hydrophobic than doxorubicin. As the polymer begins to degrade, and water navigates through pores to interact with drug, doxorubicin should freely dissolve and clear way for further and faster diffusion. Docetaxel, however is expected to have a slower and longer release profile when incorporated into a matrix of the same parameters due to its limited interaction with water as the PLGA degrades [72]. As a result, we have fabricated a biodegradable implant incorporating free docetaxel in a PLGA (50:50 copolymer, MW 7-17 kDa) matrix.

5.2 Materials and Methods

5.2.1 Docetaxel Implant Fabrication and characterization

Docetaxel implants were fabricated using a polymer extrusion method as reported previously by our group with several modifications [112]. The Docetaxel implants are composed of a solid Poly(lactic-co-glycolic) acid (PLGA, 50:50 PLA:PGA and MW 7-

17kDa) matrix embedded with 14% (w/w) DTX. Briefly, docetaxel implants were fabricated by dissolving docetaxel (25 mg) in minimum amount of chloroform (~50 µL) and mixing with PLGA solution (180 mg) in chloroform. The polymer/drug slurry was mixed thoroughly and left at 4°C for 30 minutes or until the mixture became a viscous 84 paste. The paste was transferred to a 1 mL syringe and extruded into sterile silicone tubing

(i,d, 0.8mm) at a rate of 2.5 µL/min using an infusion pump (Harvard Apparatus). After extrusion, tubes were clipped at the ends and dried overnight at 60 °C. Implants were then cooled to room temperature and cut to appropriate size (2±0.3 mm in length) for storage in a closed vial at -20 °C. Drug loading per unit length (mm) spacer was determined from randomly selected implants (n=3). Implants were measured for length, dissolved, and

Docetaxel was quantified using high-performance liquid chromatography (HPLC).

The loading of the drug in the Docetaxel implants was quantified via HPLC method using an Agilent 1100 system and reverse phase C18 Zorbek column. A standard curve was prepared for free Docetaxel in both the mobile phase and PBS buffer at concentrations ranging from 1-1000 µg/mL. Docetaxel implants (n=3) were dissolved in dichloromethane followed by extraction of Docetaxel using methanol. The drug was quantified using a 65:35 methanol:water mobile phase with a flow rate of 1.0 mL/min. Absorbance was detected at

254nm with t ret =6.97 minutes. The standard curves were fit with a linear regression curve

(R 2=1) and the curve was used to quantify the concentration of Docetaxel in each spacer.

Docetaxel implants were flash frozen in liquid nitrogen and fractured using a cooled razor. Fractured and non-fractured implants were attached to a specimen mount using a conductive carbon adhesive tab and then sputter coated with 10 – 15 nm of platinum using a vacuum evaporator (DV502, Denton Vacuum, Moorestown, NJ). Spacer samples were imaged in both SE (secondary electrons) and BSE (back-scattered electrons) mode using aS-4800 field emission SEM (Hitachi, Chiyoda, Tokyo, Japan) at 5 kV.

The release profile of Docetaxel from the implants was studied using HPLC. To determine the release kinetics, Docetaxel implants (n=3) were cut to predetermined 85 lengths, measured, and placed in 1 ml of PBS (pH 6.0 adjusted with 1 N HCl) at 37°C with gentle shaking to mimic biological tumor conditions for 70 days to coincide with common half-life of brachytherapy seed radiation. At each predetermined time point, buffer was completely removed and replaced with fresh buffer to maintain the concentration gradient between drug depot and buffer. Each collected buffer fraction was subjected to HPLC analysis as described earlier and Docetaxel concentration released was quantified. Care was taken as to not agitate and disrupt the integrity of the spacer during the buffer transfer.

After 75 days, the spacer was dissolved and any remaining drug was extracted and quantified. The total drug released and remaining in the spacer was summed and compared to the estimated drug loading from HPLC analysis for the given length of spacer. The released drug from the implants was plotted as a function of time. Each experiment was done in triplicate for significance. Release profile of docetaxel implants was determined at three different lengths to determine if the release was dependent on the length of the implant to account for batch-to-batch variation in loading between experiments and to account for the scale up in length when moving into larger animal models and eventually humans.

5.2.2 In vivo therapeutic efficacy

All animal studies were performed in accordance with protocols approved by the

Institutional Animal Care and Use Committee (IACUC) at Northeastern University. For in vivo therapeutic efficacy evaluation, an equivalent dose of Docetaxel (Taxotere) was administered one-time intravenously and compared to a one-time intratumoral implantation of Docetaxel spacer. Empty spacers were used as a vehicle control. Drug 86 loading of the implant was determined to be 100 µg DTX/mm spacer and in order to get the required therapeutic dose (18mg/kg), 2 spacers of length 2.5 mm were injected laterally to one another intratumorally. The in vivo therapeutic efficacy of Docetaxel implants was carried out in athymic nude (nu/nu) mice with a subcutaneous PC3 xenograft on the right flank. PC3 cell lines derived from human male metastatic prostate cancer were used for all studies. All cells were grown in F-12K medium supplemented with 10% FBS and 1%

Penn/Strep. The engrafted mouse models were generated by subcutaneously injecting PC3 cells at a concentration of 1.5e6 cells/mouse in matrigel on the right flank of 6−7-week old male athymic nude mice using a 1 mL Monoject tuberculin syringe. Tumor growth was monitored every 24−48 h manually with calipers. Treatment in mice began when the tumor size reached approximately 200-300 mm 3 in volume.

The Docetaxel implant was administered intratumorally using clinical grade 18G brachytherapy applicator needles (Bard Medical, Covington GA). For intratumoral implantation, each Docetaxel implant was preloaded into the applicator needle and guided laterally into the left and right of the center of the tumor. The implant was ejected by pressing the stylet simultaneously as the outer needle (cannula) was withdrawn from the tumor. Animal skin glue was used to seal the small hole in the skin to avoid any leakage of the implant out of the tumor. Four groups of animals (n=7-9 animals per group) were used in the study, out of which a control arm received intratumoral implantation of blank PLGA implants (no docetaxel), one additional control arm was untreated, one Docetaxel implant group which were treated with a two 2.5 mm long implants for a total dose of 18 mg

DTX/kg, and an intravenous docetaxel group which received a one-time equivalent dose of docetaxel in Taxotere form (18 mg DTX/kg mouse in 1:1 Ethanol:Tween-80 diluted 1:9 87 in 5% glucose to a final concentration of 50 µL/mouse) via retro-orbital injection. Free docetaxel was given i.v. for comparison to the current clinical administration of treatment

[128]. Tumors size was monitored using calipers and mice were weighed thrice weekly to evaluate weight loss. Tumor volumes were calculated using the formula:

= ℎ ∗ ℎ 6

where length and width are perpendicular diameters of the tumor, length being the longer

of the two dimensions. Mice were sacrificed if the tumor volume exceeded 1 cm 3, the tumor

volume exceeded 10% of body weight, if the mouse lost 15% of its initial body weight, or

if a tumor ulcerated or interfered with mobility.

Results are expressed as mean ± standard error of mean (SEM) and were analyzed using the t test, χ2 test, or one-way ANOVA on ranks (Wilcoxon signed rank test) using

Prism7 or SigmaStat3.5 software. All P values were two-sided with a P value of <0.05 considered statistically significant.

5.3 Results and Discussion

Current brachytherapy procedures utilize inert spacers that are delivered intratumorally via an 18G needle. Our implants, of similar dimensions to brachytherapy spacers, can be easily and readily administered as an outpatient procedure using existing clinical protocols for prostate cancer treatment. In this study, we designed a degradable

PLGA matrix containing docetaxel for slow and continuous release directly into the tumor.

Localized sustained delivery of docetaxel reduces the amount of drug required to achieve therapeutic benefit while delivering highly concentrated dosing for several weeks directly to cancer cells. The use of a local intratumoral docetaxel implant offers the potential to 88 bypass toxicities and challenges associated with intravenous delivery, metabolic stability, and pharmacokinetics. Future studies will combine the implant with radiation to produce synergistic therapeutic outcomes, however the scope of this study is to determine the feasibility of therapeutic benefit using a slow releasing implant versus the conventional clinical administration.

5.3.1 Docetaxel Implant characterization

In this study, we designed a biodegradable Poly(lactic-co-glycolic) acid matrix containing embedded docetaxel for the treatment of prostate cancer. PLGA is an FDA- approved, biodegradable polymer with low glass transition temperature provides an excellent matrix for implant fabrication. The hydrophobicity of Docetaxel ensures a uniform distribution of drug in the PLGA matrix, which was confirmed by scanning electron microscopy. PLGA implants (0.8 mm diameter) containing 100 µg/mm Docetaxel were fabricated by extrusion followed by curing. A single batch of fabrication with the above mentioned PLGA/Docetaxel mixture yields approximately 25 implants of 2.5 mm length. The implant was designed to allow direct placement into solid tumors using an 18G brachytherapy needle, after which Docetaxel was slowly released by PLGA degradation and allowed to diffuse into the tumor (Fig 29a). 89

In these studies, we used implants of 2.5 mm length, equivalent to 250 µg Docetaxel per spacer, as measured by HPLC. Scanning electron microscopy images of flash-frozen and fractured PLGA implants showed a smooth surface with homogenously distributed nanoscale pores (Fig 29b). It is likely that these pores represent pockets of Docetaxel 90 distributed throughout the implant. Docetaxel implants showed continuous drug release in phosphate buffered saline (pH 6.0) at 37°C. Drug release was characterized by an initial burst release (13% by day 2), followed by brief quick linear release (32% by day 7), before a slowed linear release from day 7 (R 2 = 0.9961) until day 75 during which time the

remaining 70% of the docetaxel was released (Fig 29c). The initial burst release can be

attributed to desorption of drug molecules from the implant surface. The prolonged burst

(6 days) is likely due to the highly hydrophobic nature of docetaxel and its very low

solubility in water. Release kinetics studies carried out with different length implants (1.8

mm, 2.5 mm, 4.15 mm) showed that the percentage of drug release was independent of

implant length indicating that drug dosing to the tumor can be adjusted by changing the

length of the implant without affecting the kinetics of the drug released or by administering

multiple implants in the tumor (figure 30). For all three different size of spacers, the release

profile showed a slow and sustained behavior over the 70 days, with almost close to 100%

of the DTX was released by ~75 days. It is important to note that the release from the

different size of the spacers showed similar profile suggesting that the rate of docetaxel

diffusion from the PLGA matrix remains same and is independent of the length of the

spacers. This is a very important result as this directly correlates to the release profile of

the spacers which will be used for implantation in larger animals or human brachytherapy

applications. The slow and sustained release of docetaxel is very critical in pre-

radiosensitizing the prostate tumor during the entire time frame of brachytherapy. Standard

intravenous administration of docetaxel results in intermittent radiosensitization as

opposed to the sustained radiosensitization which is achieved with docetaxel spacers 91

(figure 19). The release kinetics studies clearly indicates the high potential of this platform in radiosensitizing the tumors during the entire time frame of radiation treatment.

Figure 30. Release of docetaxel from PLGA implants of different lengths (n=3)

The in vitro release kinetics showed a slow sustained release of the drug over 70 days, comparable to the 73.8 day half-life of the radiotope used in brachytherapy that we ultimately aim to sensitize. Since the release of the drug in buffer showed a diffusion driven release, it is highly unlikely that there is any interaction between the drug molecules and

PLGA matrix. Additionally, the HPLC analysis resulted in no absorbance shift of the retention peak or additional peaks indicating metabolites, suggesting the Docetaxel is stable and active in the PLGA matrix during and after fabrication. Typically, PLGA is degraded via the hydrolysis of the ester bonds of the polymer backbone resulting in formation of acidic oligomers and we anticipate a similar degradation mechanism for

Docetaxel implants. The sustained release of the drug can be attributed to the diffusional constraints and slow degradation of the PLGA matrix. The competing dynamics with the 92 continuous release of Docetaxel and simultaneous inhibitory response on the tumor cells will also alter the release profile of Docetaxel from the implants. However, the release profile in buffer indicates zero order kinetics suggesting a diffusion based release of

Docetaxel. Thus, we predict that inside the tumor matrix, the release of Docetaxel will be a combination of diffusion as well as enzymatic degradation and surface erosion of the implants. While we expect these release kinetics to be altered in vivo due to enzymatic degradation of the PLGA polymer, we did observe a therapeutic benefit for as long as >100 days, suggesting that Docetaxel release in vivo is slow and sustained.

5.3.2 Docetaxel Implant therapeutic benefit

Pilot studies were carried out previously with small batch of animals (n=3 per group), to determine in vivo experimental parameters (tumor size and treatment doses) and therefore it was imperative to carry out a more systematic experiment with large group size to have a statistical significance. Four groups of animals (n=7-9 animals per group) were used in the study: (1) a control arm receiving intratumoral implantation of blank PLGA implants (no docetaxel), (2) an untreated control arm was untreated, (3) a Docetaxel implant group which were treated with a two 2.5 mm long implants (total dose of 18 mg

DTX/kg), and (4) an intravenous docetaxel group which received a one-time equivalent dose of docetaxel in Taxotere form (18 mg DTX/kg). The cumulative drug doses used this study are half of the reported maximum tolerated dose of athymic nude mice treated with intravenous docetaxel. Mice were monitored for body condition and weight loss in addition to the tumor response three times a week after treatment for any improvement in survival. 93

Tumor response is shown in figure 31. Individual animal tumor volume plots are shown for each treatment group compared to untreated control mice (black curves). Mice treated with empty spacers (green) show no difference in tumor growth from untreated controls, indicating the PLGA implant provides no therapeutic benefit and any response observed in mice treated with DTX spacers is a result of the docetaxel. The control groups

(non-treated and blank spacer group) showed continuous growth of tumor and the animals were sacrificed approximately at day 30 because of the excessive tumor burden (tumor volume >1cm3) on mice. Mice treated with IV Docetaxel showed a heterogeneous therapeutic response within the treatment group (blue) with overall no significant change in tumor growth from the control groups. The mice treated with a single dose of free DTX via intravenous route showed an overall delay in tumor growth from the day of administration but the tumor progression increased after day 15 and by day 30, the tumor volume was ~1cm 3. The change in tumor volume for free DTX group was not statistically significant when compared to other control groups. Mice treated with docetaxel implants

(red) however showed a significant improvement in tumor response. The overall response of the group showed tumor growth was arrested and no progression was observed for 30 days compared to all other groups (P<0.01).

The measured tumor volume results show significant growth inhibition of tumors treated with local slow-releasing Docetaxel spacers as opposed to uninhibited and continuous growth of tumors treated in all other groups. The survival study, as presented in the Kaplan-Meyer curve in figure 32, clearly shows increased survival for the mice treated with DTX-spacers as compared to other treatment groups. The results observed show 89% survival for the mice treated with Docetaxel spacers as opposed to 0% in all other treatment groups 39 days after treatment. The local treatment group showed a delayed response in tumor growth with no weight loss over the course of the experiment. As the 95 spacer degraded, DTX was released into the tumor and the mice showed tumor inhibition or tumor free treatment for a number of weeks with minimal systemic toxicity to the mice.

The free DTX did show some therapeutic benefit but 0% mice survived beyond 40th day.

The median survival time of the group treated with IV DTX was 27.1 days as compared to the no treatment control median survival time of 15.7 day survival and compared to the blank spacer median survival time of 21.5 days. The DTX spacer group median survival increased to 58.9 days (P<0.0001 compared to all other groups) with 33% of the mice showing complete tumor response and disease free survival for up to 120 days.

Figure 32. Kaplan-Meyer survival curve of mice in LCT study.

In a few mice treated with DTX spacer slight ulceration of the tumor were observed.

While many healed after a few days, some of the mice were sacrificed due to the severity of the local toxicity. Figure 33 shows the representative pictures of treated and untreated mice, with emphasis on the tumor-free mouse 8 weeks post treatment with DTX spacers. 96

The control mice (no treatment) and mice with free DTX (intravenous) were sacrificed on day 30 and 40 respectively. The third and fourth panels shows the mice which were treated with DTX-spacers and the treatment efficacy can be clearly visualized with the tumor volume decreasing with time. The third panel shows the observed ulceration at the tumor site ~4 weeks post treatment. The ulceration in these mice may be a result of the high cell kill achieved with implanted DTX-spacers in combination with the common side effects of docetaxel which result is skin sensitivity. The fourth panel shows the ulceration healed and the mouse was tumor free ~8 weeks post treatment with DTX-spacers with a small scar where the tumor previously resided.

Figure 33. Representative pictures of treated and untreated mice, with emphasis on the ulceration (week 4) and tumor-free mouse (week 8) post treatment with DTX spacers.

Animal weight was recorded as a means of determining gross toxicity after treatment

in local chemotherapy (LCT). Figure 34 shows the percent change in body weight for

different animal treatment groups. The control groups with no treatment and blank spacers

did not showed any significant change in body weight. Similarly, the group of animals 97 treated with docetaxel spacers did not experience any acute depression in body weight after implantation. However, the mice treated with an equivalent dose of free DTX injected intravenously showed a significant ~6% decrease in body weight for close to 14 days post injection (P<0.05), most likely a result of systemic toxicity from the free drug. This observation clearly indicates the systemic toxicity associated with the free drug injected intravenously whereas almost no toxicity with locally implanted Docetaxel spacers. No gross toxicity was observed in any of the control mice treated with empty spacers, suggesting the method of fabrication completely absolves the implant of any trace amounts of toxic solvents. During fabrication, we dry the implants above the glass transition temperature of the polymer to induce the temperature dependent phase transition. This heating also allows for high vapor pressure and volatile solvents such as chloroform to completely evaporate during that drying time, which is evident in the lack of toxicity in mice treated with the implant.

Figure 34. Weight change (%) in LCT groups after treatment for determination of gross toxicity. 98

It is important to discuss the local toxicities of the docetaxel implant. While the empty implants caused no inflammation or ulceration after intratumoral injection, (8/9) mice treated with two docetaxel spacers showed tumor ulcerations located between the two spacers, indicating an overlap in dosing. Initially, mice with any ulcerations had to be sacrificed per the animal protocol, however an amendment was made to treat the mice with

Tylenol for 3 weeks to relieve the animal of any pain while the small ulcers healed. In some cases, the local toxicities became so great that there was evident muscle damage to the underlying flank, resulting in obstructed movement of the mouse and scarification. This may be a result of the rapid elimination of the tumor with the competing dynamic of the slow continuous release of docetaxel implants for weeks after the cancer was eradicated, resulting in drug diffusion to nearby tissue. While an important observation, nearly all mice recovered after Tylenol treatment and this phenomenon was limited to 2 mice which highlights a limitation of the small animal model we used. Literature reports evidence that the prostate lobes provide a natural barrier to prevent drug diffusion in large animal models and therefore the outcome may be limited to the subcutaneous tumor model we are using.

The result of the experiment impressed the importance of determining the diffusivity of the docetaxel implant and the concentration gradients it produces to be able to properly dose map the prostate and accurately plan treatments in men. This must be done in larger animal models such as canines where the physiology is more relevant to humans and would be an important step before translation into the clinic, and unfortunately is outside the scope of this work.

The high therapeutic index observed in mice treated with docetaxel spacers with minimal observable toxicity can be attributed to the slow and sustained release of docetaxel 99 from spacers. This prolonged release of drug from the spacers provided a constant supply of drug to the tumor and thus showed a superior survival as compared to other group which received free docetaxel intravenously. Systemic administration results in only a fraction of drug to accumulate in the tumor, whereas docetaxel spacers offer 100% of the drug directly to the tumor cells slowly over a number of weeks. Additionally, the mice treated with systemic docetaxel showed loss of body weight for 15 days post treatment whereas no weight loss was observed for mice treated locally. The tumor growth inhibition seen in our

Docetaxel implant group was evident months after treatment with no observable toxicities and attests to the long-term benefit of slow releasing Docetaxel implants over systemic administration. These results have clearly validated the feasibility of our local delivery platform and laid a strong foundation to carry out a combined chemo-radiation therapy study in vivo using these spacers in conjunction with radiation.

100

6. Docetaxel loaded implants with radiation for combine local chemo- radiation

6.1 Introduction

Combined chemo-radiation therapy (CRT) requires synchronization of the two modalities to achieve maximum benefit. Several drugs are being currently tested as a radiosensitizer with external beam radiation therapy (EBRT) in a large ongoing multinational randomized trial [129]. However systemic administration of these drugs have several drawbacks: (a) high systemic toxicity, (b) fast pharmacokinetics (c) low residence time in tumor (d) limited periods of radiosensitization that are not synchronized with the radiation dosage and (e) poor spatial distribution in tumor [130]. The use of modified smart brachytherapy spacers for localized drug delivery will provide several advantages over the conventional systemic delivery such as (a) local delivery of therapeutics in high dose to the tumor, (b) slow and sustained release of the drugs, (c) ability to fine tune release kinetics,

(d) minimal systemic drug toxicities to non-targeted organs and (e) greater spatial distribution of the drug in the tumor [131]. Docetaxel loaded implants showed great efficacy as a monotherapy when compared to the therapeutic benefit of an equivalent dose given intravenously and therefore laid a strong foundation for their use in combination with radiation for local chemo-radiation therapy (LCRT).

6.2 Materials and Methods

Docetaxel spacers were fabricated and characterized as described in chapter 6.

HPLC was used to confirm drug loading was 100 µg/mm. For the purpose of 101 radiosensitization, a single spacer of 2.5mm in length was used to pre-sensitize the tumor, half the dose administered previously (9 mg/kg), so that a synergistic effect could be seen when combined with radiation. A preliminary combination study was conducted to determine the optimal radiation dose before a larger experiment was conducted for significance.

All animal experiments were carried out following the guidelines and regulations set by the set by the Northeastern University, under the approved animal use protocols by the American Association of Laboratory Animal Care and ACURO. Animals were procured from Charles River Laboratories and were housed in a group of five in standard cages with free access to food and water and a 12 hrs light/dark cycle. All animals acclimated to the animal facility for at least 48 hrs before experimentation. All possible parameters that may cause social stress, like group size, type (treated and non-treated) etc., among the experimental animals were carefully monitored and avoided. Animals were observed daily for any behavioral abnormalities and weighed weekly. The same mouse model was used as previously described. Briefly, engrafted PC 3 tumor mouse models were generated by subcutaneously injecting PC3 cells at a concentration of 1.5e6 cells/mouse on the rear right flank of mice using a 1 mL Monoject tuberculin syringe.[11] Tumor growth was monitored every 24−48 h until a tumor size of approximately 5 mm in diameter was observed.

To evaluate the efficacy of combined chemo-radiation therapy, we used the same prostate cancer subcutaneous xenograft animal models. In order to assess the combined effect of two treatments and to study if these combination works in an additive or synergistic manner, it was critical to standardize several parameters before the start of 102 actual experiment. Thus, we initiated a pilot experiment to optimize the drug dose to be used in the spacers (determined from the therapeutic dose from the previous LCT experiment), the appropriate radiation dose, and the timing of irradiation after implantation of spacers in tumored animals. The design of the experiment is described in table 4.

Table 4: Design of preliminary in vivo LCRT experiment to determine appropriate dosing for synergistic therapeutic effects Radiation Radiation Radiation Groups No Radiation (5 Gy) (10 Gy) (15 Gy) Controls: No 3 3 3 3 treatment DTX-spacers 3 3 3 3 (9 mg/kg)

Results of the preliminary experiment provided the parameters required for a larger animal experiment. The tumored animals were randomly assigned in four groups (n=7 per group) and included: (1) an untreated control group, (2) a group was treated with docetaxel spacer

(9 mg/kg), implanted intratumorally, (3) a group treated with fractionated 10 Gy radiation dose (4 Gy/day for 5 consecutive days) and (4) a combination group treated with a

Docetaxel spacer and fractionated 10 Gy radiation using Small Animal Radiation Research

Platform (SARRP) (9 mg/kg + 4 Gy/day x 5). Irradiation experiments were carried out using Seifert ISOVOLT 225M2 X-ray source (Seifert X-ray, Lewistown, PA) on a SARRP platform operating at 220 kVp and 13 mA using a 0.15 mm copper filter at Dana Farber

Cancer Institute (Boston, MA). The DTX spacers were implanted intratumorally when the tumor size was between 200-250 mm 3 and radiation was given 13 days post spacer implantation. All treatments with docetaxel spacers were given with a single spacer ~2.5 mm in length for a dose of 9 mg/kg body weight of mice. Tumor size was monitored every 103 alternate day by physical measurements of the tumor size by calipers and body weight was recorded. As described previously, mice were sacrificed if the tumor volume exceeded 1 cm 3, the tumor volume exceeded 10% of body weight, if the mouse lost 15% of its initial body weight. In case of ulceration, mice were given Tylenol for 3 weeks until wound healed. Results are expressed as mean ± standard error of mean (SEM) and were analyzed using the t test, χ2 test, or one-way ANOVA on ranks (Wilcoxon signed rank test) using

Prism7 or SigmaStat3.5 software. All P values were two-sided with a P value of <0.05 considered statistically significant.

6.3 Results and Discussion

6.3.1 Determination of parameters in preliminary study

For estimating the combined efficacy, it was imperative to find the right balance between the DTX dose and radiation dose, so that the individual therapy alone shouldn’t show increased survival as we saw with the LCT experiment. Thus, for the LCRT experiment, we reduced the DTX concentration to half (~250 µg) and tested these spacers implanted intratumorally in mice by combining with 3 different radiation doses (5, 10 and 15 Gy) given in a single fraction using Small Animal Radiation Research Platform (SARRP). The docetaxel spacers were implanted intratumorraly when the tumor size was between 200-

250 mm 3 and radiation was given after 13 days post spacer implantation. A period of 13 days was picked in between the two treatments so that sufficient docetaxel be released from the spacers and diffuses into the tumor matrix prior to the tumor irradiations. We anticipated a higher degree of therapeutic benefit as released DTX will act as radiosensitizing molecule in presence of radiation. 104

Figure 35 shows the antitumor activity of different treatment combinations used in

tumored animals (n=3 per group). The control group of animals which received no

treatment showed a continuous increase in tumor volume and all mice were sacrificed by

day 56 due to large tumor volume (> 1cm 3). The group of mice which received only radiation (5, 10 and 15 Gy groups) showed a radiation dose dependent tumor growth arrest initially, for roughly up to 55 days, following which the tumor growth recurred and increased drastically. The tumor volume for all three groups which received 5, 10 and 15

Gy was in the range of 1.1-2.1 cm 3 on day 90 and these animals were sacrificed. It was interesting to note that the group of animals which received only DTX-spacer (half the dose as compared to LCT experiment), also delayed the tumor growth significantly for 40 days after which the tumor volume started increasing again. This is in contrary to the LCT experiment, where we observed the decrease in tumor volume and PFS for more than 120 days using double the dose. This clearly demonstrated the effect of DTX concentration which is required to inhibit the tumor growth and further validate our drug concentrations used in this experiment.

105

Figure 35. Efficacy of combined chemo-radiation therapy using DTX-spacers implanted intratumorally in prostate cancer animal model. The graph shows the change in tumor volume when different treatment combinations were studied in tumored animals. DTX-spacers containing 250 µg of total DTX (~9 mg/kg body weight of mice) was implanted in mice (red arrow) and radiation dose of 5, 10 and 15 Gy were given 13 days post implantation of spacers (black arrow).

The group of animals which were treated with both docetaxel spacers and radiation showed a dose radiation dependent therapeutic benefit. The batch of animals treated with docetaxel spacer and 5 Gy radiation showed tumor growth suppression for ~50 days after which the tumor volume started increasing again suggesting the radiation dose of 5 Gy was suboptimal to arrest tumor growth. This antitumor tumor effect was still pronounced and was much greater when compared to the batch of mice which received 5 Gy radiation only in which the tumor volume exceeds by ~0.5 cm 3 at day 90. The maximum tumor growth

inhibition was observed for batches of mice which received 10 and 15 Gy radiation along

with docetaxel spacers. These groups of mice showed continuous regression of tumor

volume and showed a progression free survival up to 90 days. 106

Gross toxicity was assessed in groups using body weight measurement as previously described and shown in figure 36. Since the docetaxel concentration in these groups of animals was reduced to half as compared to the animals treated in the previous

LCT experiment, there was no significant change in body weight in any of the experimental group. The only toxicity observed were groups in which the mice were treated with 15 Gy radiation (with and without DTX-spacers) which showed a decrease of ~12% of body weight. Since the weight loss was very similar in the two groups, it was quite apparent that this toxicity was induced as a result of high radiation dose. However, while these mice gained weight after 20 days post irradiation, 2/6 mice treated with 15 Gy had to be sacrificed from losing >15% of its body weight post treatment.

107

Due to similar therapeutic benefit, and minimal toxicity a combination of 9 mg/kg docetaxel spacer and 10 Gy radiation dose was selected for a safe and effective dose for scale-up.

6.3.2 Local Chemo-Radiation Therapeutic efficacy

The therapeutic studies for LCRT efficacy involved two animal experiments in which we measured the therapeutic response of external beam radiation therapy, in combination with docetaxel. The radiation dose was optimized from the experiment carried out in the preliminary study and determined to be 10 Gy, in which we observed a high therapeutic benefit with minimal radiation associated toxicity. While the docetaxel spacer is designed for ultimate use in brachytherapy, the co-implantation of brachytherapy seed and the drug eluding spacers in 100 mm 3 tumor size was not possible due to the small size of the tumor, so we adopted an alternative strategy by fractionating the radiation dose into

5 daily fractions of 4 Gy/day using EBRT approach to mimic the prolonged tumor exposure to radiation experienced in brachytherapy. Based on calculations from our collaborators in the Radiation Oncology group at Dana Farber Cancer Institute, 5 consecutive daily fractions of 4 Gy should have an equivalent tumor effect to a one time radiation dose of 10

Gy (assuming α/β = 3 for PC3 tumor) [111].

Figure 37 shows the antitumor activity of different treatment combinations used in

tumored animals. The graphs at the top show the individual growth curves of tumor volume

for mice treated with either docetaxel implant (9 mg/kg) in blue, fractionated radiation (4

Gy x 5) in green, and the combination of both DTX spacer (9 mg/kg) and fractionated 108 radiation (2 Gy x 5) in red each plotted against the untreated control mice in black

(n=7/group).

Mice were treated with DTX spacers on day 1, and radiation on days 13-17 to pre- sensitize the tumor to radiation. This dosing was chosen from previous experiments in which tumors showed significant growth inhibition beginning 7-12 days after the spacers were injected. The individual animal data depicts the varied response of the DTX-spacer 109 treatment group. While one mouse showed complete tumor regression, the group overall showed the least significant improvement from the control group (p<0.05). The radiation treated mice had a better response, with significant tumor regression from the controls however all tumors in the group showed tumor regrowth by day 90. The mice treated with the combination therapy, both DTX spacer and fractionated radiation therapy, had the most uniform and significant response in tumor inhibition (p<0.001 vs control, p<0.01 vs DTX spacer). At day 90, 40% of the mice in the combination group showed a complete treatment response and 40% continued to inhibit tumor growth for progression free survival (PFS).

The lower panel in figure 37 shows the mean tumor volume fold change of each treatment group (+/- SEM) normalized to the starting tumor volume at treatment. Controls (black) grew uninhibited with little variation in the group. DTX-Spacer (blue) and Radiation

(green) groups showed a significant inhibition in tumor volume from the controls (p<0.05), however both only delayed tumor growth. The combination of radiation and DTX-spacer

(red) was the only group to have an overall decrease in tumor volume (days 32-78) with significant tumor inhibition and reduction from all other groups. While the average trend is increasing, it is apparent from the individual animal curves that the majority of the mice within the group continued to show tumor inhibition and regression at day 100. This data successfully shows the radiosensitization achieved by pre-treating tumors with localized docetaxel before radiation therapy. The degree of tumor response seen with the combination treatment is unachieved with either monotherapy in the form of radiation or

DTX-spacer alone. These animals were monitored further and survival curves were plotted to evaluate the efficacy of treatment.

110

) ** 3 * Tumor Volume (mm Volume Tumor

) ) /kg /kg g g Control 9m 9m r ( er ( ace diation (2Gyx5) a R + Spac DTX Sp 5) yx 2G ( n io at Radi

Figure 38: Box and whisker plots showing the geometric mean difference in tumor volume and comparative significance among various treatment groups after 135 days of treatment.

Figure 38 shows the box and whisker plots to show the median difference in the tumor volume compared across the different treatment groups 90 days after treatment. The plots clearly indicate the significant improvement in therapeutic efficacy across the control

(no treatment) or DTX spacers vs combination treatment group. Figure 39 shows the

Kaplan Meier survival curves for untreated tumored mice (control) or tumored mice treated with radiation, docetaxel spacer, or the combination of radiation and DTX-spacer

(n=7/group). The data shows an enhanced survival for all treated mice as compared to the controls, in which 100% of the mice were sacrificed by day 50. Mice treated with DTX- spacers had a significantly increased median survival time of 85 days compared to 42 days in controls (p<0.01) with 28% of mice still alive 100 days after treatment. Mice treated with fractionated radiation show a better survival than DTX-spacer alone with 60% of the 111 mice alive at day 100 and a median survival time of 120 days (P<0.01). While the mice in treated with radiation had an improved median survival, all mice were sacrificed by day

168 as compared to the last mouse surviving 218 days after treatment with the DTX spacer.

It is worth mentioning here that previously when mice were treated with a single fraction of 10 Gy dose, the tumor growth was significantly inhibited during the first 40 days but showed a regression of the tumor by day 90. Thus, here a fractionated approach showed a better treatment response as opposed to the treatment done with single dose of radiation.

Mice treated with the combination therapy had the most significant response and therapeutic benefit, with all of the mice (100%) still alive at day 100. Median survival time was significantly enhanced from all other treatment groups (P<0.01) to 209 days with an overall survival time extending to 302 days. This significant therapeutic benefit in combination therapy is achieved because of the pre-sensitization of the tumor by the released docetaxel from the DTX spacers.

Figure 39. Kaplan-Meyer survival curve shows median survival time of LCRT treatment groups. 112

The gross toxicity assessment in case of combined chemoradiation therapy was carried out by analyzing the change in body weight of treated animals. Figure 40 shows the average weight change of each group after treatment. No significant weight loss was seen in any of the treatment groups. Preliminary studies indicated maximum toxicity post radiation treatment occurred within 7 days of therapy, especially for the groups treated with a single fraction of 15 Gy radiation dose. Thus, the approach of fractionated dosing not only mimics the actual brachytherapy procedure but also minimized the radiation associated toxicities. The fractioned dosing caused no significant weight loss when given as monotherapy or in combination with the docetaxel spacers. Also, it is worth mentioning here that since no significant weight loss was seen in the previous LCT study using DTX spacers with twice the dose of docetaxel (18 mg/kg previously) as we used in current experiment, we did not expected any gross toxicities at this low dose. Both treatment groups with docetaxel spacers showed no apparent adverse toxicities other than minor ulcerations (local toxicity) in a few mice. All ulcerations healed and no mice were sacrificed as a result of local toxicities. The combination group showed no weight loss and overall no change in behavior in the mice, indicating there were no treatment associated toxicities in any of the groups. 113

Figure 40: Gross toxicity assessment by measuring the change in body weight in prostate cancer tumored animals treated with local chemotherapy using spacers

6.4 Conclusion

The therapeutic efficacy studies established significantly improved outcomes using docetaxel implants when tumored mice were treated with a combination of radiation and local chemotherapy as opposed to either of the two monotherapies alone. The fractionated radiation treatment approach to mimic the brachytherapy procedure showed better therapeutic efficacy in combination with docetaxel spacers. The combination treatment using the fractionated radiation dose equivalent to 10 Gy and DTX spacers was the only treatment to achieve an overall reduction in tumor volume after treatment and showed a complete treatment response in 40% of the treated mice and progression free survival (PFS) in other 40%. The mice treated with either monotherapy had a delayed tumor growth as compared to control groups but showed an uninhibited progression of tumor over a period 114 of 100 days. The safety of combination therapy using the fractionated radiation treatment and docetaxel spacers was evaluated using by measuring the change in the mice body weight. No significant loss of body weight in mice were observed in any of the treatment groups. This clearly indicates that the local delivery of docetaxel using our spacers in combination with fractionated radiation (brachytherapy) provides high therapeutic benefits with minimal toxicities. The promising in vivo therapeutic results using sustained releasing

docetaxel spacers for localized chemotherapy or in combination with fractionated radiation

therapy laid a strong foundation to facilitate further preclinical studies with elaborated

safety studies which can help translate these modified brachytherapy spacers in clinical

phase 1 testing in near future.

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Chapter 7. Toxicity of Local Chemotherapy and Local Chemo-Radiation

Therapy

When developing a new drug delivery system, it is imperative to test any treatment associated toxicities. For an intra-prostatic implant, we must consider local toxicities as well as systemic toxicities with the understanding that maximum tolerated dosing will most likely be limited by local toxicities (i.e. necrosis) experienced at the site of implantation.

Chemotoxic agents often target rapidly dividing cells such as bone marrow, gastrointestinal lining, and lining of the mouth and cheek [10]. Because of this, the most common adverse effects associated with docetaxel is low white blood cell counts due to bone marrow’s fast proliferation rate, and liver toxicity due to the hepatometabolism of the drug [81,82,128].

7.1 Introduction

Toxicities associated with docetaxel and radiation both need to be explored as monotherapies and as combotherapies to assess the risk of combining these two therapies locally versus the conventional clinical treatment. Our PLGA implant loaded with docetaxel has shown significant therapeutic success over conventional treatments with limited gross toxicity as measured by body weight. The promising in vivo therapeutic results using sustained releasing docetaxel spacers for localized chemotherapy or in combination with fractionated radiation therapy laid a strong foundation to facilitate further preclinical studies with elaborated safety studies which can help translate these modified brachytherapy spacers in clinical phase 1 testing in near future. To do so, we implore the use of immune-competent mice with metastatic prostate cancer cell line TRAMP-C1 to 116 determine any and all toxicities associated with the new implant both with and without radiation compared to conventional chemotherapy treatment modalities.

7.2 Materials and Methods

In previous studies, humanized prostate cancer PC3 cell line was injected subcutaneously for a prostate tumor xenograft on the right flank of immune-compromised male athymic nude mice. In order to properly assess the toxicities associated with treatment, mice must have competent immune systems. For the purpose of this study, 7-8 week old C57BL/6 mice were purchased from Charles River Laboratory and were housed in a group of five per cage in standard cages with free access to food and water and a 12 hrs light/dark cycle. All animals acclimated to the animal facility for at least 48 hrs before experimentation. All possible parameters that may cause social stress, like group size, type

(treated and non-treated) etc., among the experimental animals were carefully monitored and avoided. Animals were observed daily for any behavioral abnormalities and weighed weekly. Animal experiments were conducted in agreement with all relevant guidelines and regulations set by the Northeastern University, and with approved institutional protocols by the American Association of Laboratory Animal Care.

TRAMP-C1, a mouse derived cell line was purchased from ATCC. TRAMP-C1 is

an androgen-dependent prostate cancer cell line which represents various stages of cellular

transformation and progression to androgen-independent metastatic disease was selected

to best test toxicities experienced by both tumor burden and treatment. TRAMP-C1 was

cultured in Dulbecco's modified Eagle's medium with 4 mM L-glutamine adjusted to

contain 1.5 g/L sodium bicarbonate and 4.5 g/L glucose supplemented with 0.005 mg/ml 117 bovine insulin and 10 nM dehydroisoandrosterone, 90%; fetal bovine serum, 5%; Nu-

Serum IV, 5% as suggested per ATCC and passaged at a ratio of 1:10. For tumor inoculation, cells were washed twice with PBS and tryspinized for 10 minutes in 37°C, 5%

CO 2. Trypsin was deactivated in media and cells were centrifuged at 600rpm for 5 minutes.

The media was removed and the cell pellet was washed with serum-free media before centrifugation. Cell pellets were collected, washed, and centrifuged three times. TRAMP-

C1 cells were re-suspended in Corning Matrigel Membrane Matrix, Phenol Red-Free, at a final concentration of 30e6 cells/mL. Each mouse was shaved and injected with 100uL of cell suspension on the rear right flank and observed for 2-4 weeks for tumor growth. Each mouse was randomly assigned to a treatment group shown in Table 5 and treated accordingly. Each group has five mice per timepoint to account for any errors in failed tumor inoculation, blood collection, or natural deaths.

The intravenous solution of docetaxel was formulated by dissolving Docetaxel in a

1:1 Tween-80: Ethanol solution and diluting 1:9 in 5% Glucose for a final concentration of 5.5 mg DTX/mL immediately before use. Roughly 50 µL solution was injected retro- orbitally into anesthetized mice for a dose of 9 mg DTX/kg mouse. Mice injected with spacers were treated under anesthesia. Spacers were implanted subcutaneously in mice using a sterile clinical brachytherapy applicator needle directly into the tumor. Animal glue was used to seal the small hole left from the 18 gauge applicator needle. All radiation was conducted using a small animal radiation research platform (SARRP) at Dana Farber

Cancer Institute. Mice were transferred at least 24 hours before and after treatment for acclimation. A 220kVp, 13mA small field irradiator with a 0.15mm copper filter using conformal radiation plans to specifically target the tumor was used to treat mice for 85 118 seconds or 4Gy daily for 5 consecutive days, a total equivalent dose of 10 Gy. Mice receiving both IV DTX and Radiation were treated with IV DTX 24 hours prior to the first radiation dose, to allow for drug accumulation at the tumor site. Mice treated with DTX

Spacers and radiation were pretreated with spacers for 12 days before the first radiation dose to allow for complete tumor dosing prior to radiation. All sacrifice timepoints commenced after the final day of each treatment.

Table 5. Treatment plan for toxicity study showing time of sacrifice and number of animals per group Timepoint 1 Timepoint 2 Timepoint 3 Group 24 hr post 1 week post 3 weeks post treatment treatment treatment DTX Spacer (10 mg DTX/kg) n= 5 n= 5 n= 5 IV DTX (10 mg DTX/kg) n= 5 n= 5 n= 5 Radiation (4Gyx5) n= 5 n= 5 n= 5 IV + Radiation (10 mg DTX/kg) + (4Gyx5) n= 5 n= 5 n= 5 Spacer + Radiation (10 mg DTX/kg) + n= 5 n= 5 n= 5 (4Gyx5) Control (Blank spacer) (0 mg DTX/kg) n= 5 n= 5 n= 5 Control (no treatment) n= 5

At the time of sacrifice, mice were anaesthetized using 2.5% isofluorane and

1ml/min O 2 before euthanasia via a cervical dislocation and cardiac puncture for blood

collection. Whole blood totaling 200 µL was collected in a lavender EDTA-

microcentrifuge tube and inverted 20 times to prevent clotting. A gold gel serum separator

tube was used to collect 500 µL blood. After clotting, the gold micro-centrifuge tubes were

centrifuged at 3000 rpm for 10 minutes and serum was separated for clinical analysis.

Blood chemistry was performed at VRL in Rockville, MD. Organs were collected, washed,

and fixed in formalin for histological staining and analysis.

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7.3 Results and Discussion

Docetaxel loaded spacers have been tested as a monotherapy against standard

chemotherapy Taxotere and in combination with radiation to show sensitization and

sustained therapeutic efficacy for the treatment of prostate cancer in our previous studies.

In order to progress into the clinic, it is imperative to determine any and all adverse side

effects from both the docetaxel implant as a monotherapy, and any toxicities it may cause

when administered concomitantly with radiation. To determine whether this treatment is

more toxic than current therapy, it is necessary to test all combinations with and without

radiation as is currently done in the clinic. It is also important to determine whether any

adverse effects are a result of the implant or procedure, or a result of the drug being released

from the implant.

Animals in each group were sacrificed at three different time-points post treatment.

Time-points were selected based on results of earlier studies and published data to test for

any adverse side effects and the most toxic times of each treatment to give a clear

comparison. For example, the half-life of docetaxel is published to be between 18-84 hours.

Previous studies have shown that significant weight loss occurred in mice immediately

after treatment with a two week recovery time, so if toxicity is observed as a result of the

treatment at this dose, we should expect to see it at both time-points 1 and 2 with a recovery

at time-point 3. Similar time-points for radiation toxicity were observed in pilot studies,

however we have reported no significant weight loss at these fractionated low doses in

previous experiments. The spacer is a slow releasing implant. If an adverse reaction occurs

due to the implantation of the foreign object, we should detect toxicities within the first 24

hours. Although these implants have resulted in no significant weight loss in previous 120 studies, it is important to test for both acute and chronic toxicities post injection, resulting in the decision of a 3 week time-point.

Complete blood chemistry and hematology was tested for any changes after treatment. Results were compared to levels obtained in tumored untreated mice and healthy mice to determine if changes were a result of the cancer or a result of the treatment.

Measurements were taken at different time-points to determine whether the observed toxicity changed over time. Twenty-two blood chemistry parameters and twenty-five hematology components were tested. Results for group were averaged and the percent change from the weighted geometric means were calculated from both the tumored controls and the normal healthy ranges for male C57BL/6 mice. The percent differences from healthy mice (figure 41) and tumored mice (figure 42) were plotted as a heat map to best represent any significant changes within and between treatment groups. From this data, it is apparent that mice treated with docetaxel spacers show the least change in white blood cell counts from healthy mice, while all other groups, including blank spacers show reduced level. White blood cells help the body fight off infection. Strong chemotherapy or radiation can lower white blood counts. Neutropenia, or the condition associated with low neutrophils (white blood cells) is a common side effect of intravenous docetaxel. It is clear that when administered locally using a slow release implant, mice treated with docetaxel spacers do not experience neutropenia. The mice treated with IV docetaxel did show a slight recovery of white blood cell counts from day 1 to day 7, which was expected as the pharmacokinetics of the drug clear from the system in 2-3 days. Similarly, mice treated with docetaxel spacers were the only treatment group to show normal lymphocyte counts of a healthy mouse (or increased from the tumored mice). All other mice have a 121 significantly reduced number of lymphocytes which is common with radiation or chemotherapy treatments which cause bone barrow damage. Another common result of cancer treatment is a drop in platelet numbers (PLT). If platelet numbers are too low, patients are no longer able to receive treatment as the condition is very serious. In mice treated with radiation and mice radiation with DTX spacers, platelet levels deviate the least from healthy mice and increased from the tumored control mice, which was unseen in any other group. Additionally, looking at the MPV values or mean platelet volume illustrates the average size of platelets found in blood which is indicative of the age of the platelets.

A high MPV value indicates a high number of larger, younger platelets in the blood which may be due to the bone marrow producing and releasing platelets rapidly into circulation.

Mice treated with radiation and combined radiation and docetaxel spacers had normal MPV values as compared to both the healthy and untreated tumored controls. The elevated MPV values from both healthy mice and controls in all other groups suggests treatment associated toxicity which is forcing the body to create new, larger platelets. 122

Figure 41. Percentage change from geometric means of hematology counts in treatment groups to levels of healthy mice. 123

Figure 42. Percentage change from geometric means of hematology counts in treatment groups to levels of untreated tumored control mice.

124

Blood chemistry results can indicate specific organ disease or toxicities so it is critical to test for a wide range of enzymes and metabolites. Serum was analyzed to determine the change in levels from healthy mice (Figure 43) and from tumored mice

(figure 44) for the various treatment groups. There was no significant difference between the docetaxel spacer group and the blank spacer group in the blood chemistry panel. No inflammation markers were observed within 24 hours as we might have expected as a result of the procedure. This further attests to the safety of docetaxel when delivered in a slow release implant, however the when combined with radiation, mice experienced slight changes in blood levels which need to be considered for toxicological evaluation. Blood urea nitrogen (BUN) levels significantly were elevated in 21 day time point from the combination DTX Spacer and Radiation group. Additionally, we see reduced levels of glucose within this group. Both of these can implicate abnormal kidney function. This is not seen in either of the monotherapies or any other group however we cannot rule out the possibility of kidney disease. Liver enzymes (other than GGT which returned to normal by day 21) within the combination treatment group show no deviation from the normal range of a healthy mouse. Cholesterol levels appeared low which is not unusual in the presence of cancer. Groups treated with IV DTX showed elevated phosphorus levels by day 7 with normal levels after 21 days. Phosphorus is associated with kidney and liver activity.

Docetaxel is metabolized in the liver after administration. These elevated levels within the first week may be evidence of acute toxicities in the metabolized organs, with a recovery in the following week. This result is consistent with the gross toxicity measurement we observed with weight loss, where groups treated with IV DTX expressed a maximum weight loss by week 1 and recovered by week 2. Uric acid levels were elevated in all 125 groups, often a result of the chemotherapy or radiation treatment due to the high cell turnover. Futher histological evaluation of healthy organs would provide better insight as to what is actually causing these blood chemistry changes and confirm any presence of toxicity.

Figure 43. Percentage change from geometric means of blood chemistry levels in treatment groups to levels of healthy mice.

126

Figure 44. Percentage change from geometric means of blood chemistry levels in treatment groups to levels of untreated tumored control mice.

When discussing toxicity, it is important to mention that during our previous experiments, we have seen small ulcerations of the skin in near proximity to the docetaxel loaded implant indicating local toxicity. We have not seen this side effect in any of the empty spacers, meaning it must be either a result of: (1) the drug is released and causing 127 skin irritation; (2) the drug is released, causing tumor death, and the immune system is clearing out the dead cells rapidly, leaving a hole behind where the tumor once was; or (3) the docetaxel release is resulting in necrosis rather than apoptosis. To determine this, docetaxel loaded implants were placed subcutaneously on the left flank and observed for any local toxicities. No ulcerations were observed (only skin redness), implicating the ulcers are a result of tumor response. Additionally, we have observed few instances in which the hind flank near the tumor has significant muscle damage in mice that were completely cured. As mentioned previously, this is most likely a result of the continuous docetaxel releasing into nearby tissue after the tumor has been eradicated. While this problem is limited to the small animal and subcutaneous tumor model we are using, it is important to understand the diffusivity in larger, more relevant animal models to prevent any necrosis around the implant before scaling up for human use.

7.4 Conclusion

We have completely a full blood chemistry and hematology study to determine any associated toxicities with our slow releasing docetaxel implant both with and without radiation as compared to current clinical procedures. Immune-competent mice were inoculated with a mouse derived androgen-dependent cell line TRAMP-C1 and treated with a number of combination and monotherapies in order to determine any deviation in toxicity due to the new administration route of docetaxel. Blood work analysis concludes that combined LCRT via docetaxel spacers with fractionated radiation result in improved white blood cell counts and decreased the presence of neutropenia currently experienced from systemic administration. Additionally, platelet counts were elevated compared to 128 other groups, with the platelets of a mature age compared to the rapid production of new larger platelets needed for recovery after treatment that the other groups expressed. It is important to highlight that while liver enzymes appeared to be in normal range, a concern associated with the metabolism of docetaxel after intravenous administration, the blood chemistry levels indicated abnormal kidney function in the combined LCRT group. To rule out acute kidney toxicities, histological analysis on organs should be evaluated further.

After a full toxicological analysis, local toxicity rather than systemic toxicity appears to be the limiting dosing factor. This work shows a successful completion of the development of an effective and safe new treatment modality, which can be used as a monotherapy or in combination with radiation therapy, for local and sustained therapy in prostate cancer.

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Chapter 8. Sustained PARP inhibition for breast cancer treatment using spacer platform

Talazoparib, a potent PARP inhibitor, has shown promising clinical and pre-clinical activity by inducing synthetic lethality in cancers with germline Brca1/2 mutations.

Conventional oral delivery of Talazoparib is associated with significant off-target effects, therefore we sought to develop new delivery systems in the form of an implant loaded with

Talazoparib for localized, slow and sustained release of the drug at the tumor site in Brca1 -

deficient breast cancer. Poly(lactic-co-glycolic acid) (PLGA) implants (0.8 mm diameter)

loaded with subclinical dose (25 or 50 µg) Talazoparib were fabricated and characterized.

In vitro studies with Brca1 -deficient W780 and W0069 breast cancer cells were conducted

to test sensitivity to PARP inhibition. The in vivo therapeutic efficacy of Talazoparib

implants was assessed following a one-time intratumoral injection in Brca1 Co/Co ;MMTV-

Cre;p53 +/- mice and compared to drug-free implants and oral gavage.

Immunohistochemistry studies were performed on tumor sections using PCNA and γ-

H2AX staining. Sustained release of Talazoparib was observed over 28 days in vitro . Mice treated with Talazoparib implants showed statistically significant tumor growth inhibition compared to those receiving drug-free implants or free Talazoparib orally. Talazoparib implants were well-tolerated at both drug doses and resulted in less weight loss than oral gavage. PARP inhibition in mice treated with Talazoparib implants significantly increased double-stranded DNA damage and decreased tumor cell proliferation as shown by PCNA and γ-H2AX staining as compared to controls. These results demonstrate that localized and sustained delivery of Talazoparib via implants has potential to provide superior treatment 130 outcomes at sub-clinical doses with minimal toxicity in patients with BRCA1 deficient tumors.

8.1 Introduction

The breast cancer-associated gene 1 (BRCA1 ) is the most frequently mutated tumor suppressor gene found in familial breast cancers. Women with this mutation have a 50-

80% risk of developing breast cancer in their lifetime. As many as 75% of these cases are diagnosed as triple-negative breast cancer (TNBC), a subtype known for its poor prognosis.[132–134] Mutations of the BRCA1 gene modulate many cellular functions including DNA damage and repair, homologous recombination, cell-cycle regulation, transcription, epigenetic modification and apoptosis. Typically, BRCA1 is required for homologous recombination, a commonly utilized double-stranded break DNA repair mechanism. In tumor cells which have a loss of both BRCA1 genes, homologous recombination is lost and cells must rely on non-homologous end joining, a more error prone pathway, for survival. [135–138]

Poly-ADP-Ribose Polymerase (PARP) plays a key role in the repair of single- stranded DNA breaks and is recruited to sites of DNA damage to initiate repair.

[137,139,140] PARP inhibition (PARPi) produces cell death via synthetic lethality in cancers with genetic predispositions for impaired DNA repair or impaired transcription, such as BRCA1 mutants. [132,141,142] Drugs that inhibit PARP1 activity prevent the repair of single-stranded DNA breaks, which propagate into double-stranded breaks after

DNA replication. PARPi in cells deficient in homologous recombination results in genomic instability and ultimately cell death by forcing cells to repair double-stranded breaks using 131 lower fidelity methods such as non-homologous end joining. [143–147] PARP inhibitors have shown promising results in clinical trials for breast cancers, particularly those due to either inherited mutation or somatic inactivation of BRCA1/2 genes which are present in

~15% of the patient population. [148–150] These drugs are also expected to have wide

applicability beyond BRCA mutations in breast cancers that exhibit “BRCAness” such as

sporadic tumors with homologous recombination defects. [151–153]

Talazoparib (BMN-673) is known for high PARP1 and PARP2 inhibition with

potency at much lower concentrations than earlier generations of PARP1/2 inhibitors like

Olaparib or Veliparib [141,154]. The cytotoxicity of Talazoparib at nanomolar

concentrations is believed to be related to the trapping of PARP-DNA complexes based on

studies in PARP1 -/- mouse models. [155–157] However, the hydrophobic nature of

Talazoparib, combined with its oral delivery route, limits drug bioavailability (~50% after first-pass metabolism), necessitating the use of high doses in patients and leading to off- site toxicity. [141] Thus, Talazoparib is an attractive drug candidate for breast cancer patients with Brca1 mutations, but its potency may limit long-term use. [145,158]

Local, sustained delivery of drugs has been achievable in the past few decades with the use of biodegradable polymers. Poly (lactic-co-glycolic) acid or PLGA is an FDA approved biodegradable and biocompatible copolymer that has been widely studied for the use of drug delivery systems. PLGA has tunable mechanical and degradation properties and is widely characterized for its extensive use in developing devices for controlled delivery and release of many types of molecules, proteins, and drugs. [66,73] Many groups have shown that PLGA can be injected subcutaneously or intratumorally for long-term controlled drug delivery without the need for surgery. The hydrophobicity of the polymer 132 can be altered by adjusting the ratio of the two polymers PLA and PGA, so that a wide variety of payloads can be incorporated into nanoparticles, microspheres, or implants.

[71,104] PLGA as an implant has been used intratumorally to deliver molecules, magnetic particles, nanoparticles, drugs, and imaging agents to improve the bioavailability at the tumor site and minimize any leaking of particles from the tumor. [63,67,69,105,112]

Substantial research has been conducted that supports the use of biodegradable PLGA implants in the use of cancer treatment for prolonged drug delivery with minimal systemic side effects.

Here, we report a novel, biodegradable implant for localized and sustained delivery of Talazoparib to treat Brca1 -deficient breast cancer with minimal treatment associated toxicities. The in vitro efficacy of Talazoparib was compared to that of other PARP inhibitors Olaparib and Niraparib using two breast cancer cell lines derived from Brca1 - deficient mice. For in vivo delivery, Talazoparib implants were fabricated using poly(lactic-co-glycolic) acid (PLGA) with tuned release of drug over several weeks.

Implants containing 25 or 50 µg Talazoparib were administered intratumorally into

Brca1 Co/Co ;MMTV-Cre;p53 +/- mice and their treatment efficacy was compared to drug-free control implants and Talazoparib treatment by oral gavage. Mice were monitored for signs of gross toxicity and immunohistochemistry was performed to assess DNA damage and cell proliferation in the tumors. The results showed that localized delivery of Talazoparib significantly delays tumor growth and prolongs survival time at sub-clinical doses. This new approach of localized treatment using Talazoparib implants is an attractive strategy to treat the 75% of TNBC patients and 33% of overall breast cancer patients which would benefit from PARPi therapy. [143,148,149] 133

8.2 Materials and Methods

All solvents used were HPLC grade and purchased from Sigma-Aldrich unless otherwise stated. Talazoparib, Olaparib and Niraparib were purchased from SelleckChem (Houston,

TX). Poly(lactic-co-glycolic) acid (PLGA) (Mw 7-17 kDa) was purchased from Sigma-

Aldrich. Syringes and silicon tubing (i.d. 0.8 mm) were purchased from Fisher Scientific.

Unless otherwise stated, all cell culture products (nutrient media, trypsin, Penstrep, fetal bovine serum) were obtained from Life Technologies and were used without further purification. Antibodies were procured from Abcam and Calbiochem. For cell viability assay CellTiter 96 ® AQueous One Solution Cell Proliferation Assay (MTS) reagent was

procured from Promega.

8.2.1 Talazoparib Implant Fabrication and characterization

Talazoparib implants were fabricated using a polymer extrusion method as reported

previously by our group with several modifications. [112] The Talazoparib implants are

composed of a solid Poly(lactic-co-glycolic) acid (PLGA, 50:50 PLA:PGA and MW 7-

17kDa) matrix embedded with 4% (w/w) Talazoparib. Briefly, Talazoparib implants were

fabricated by dissolving Talazoparib (9 mg) in chloroform and mixing with PLGA solution

(190 mg) in chloroform. The polymer/drug mixture was mixed thoroughly and left at 4°C

for 30 minutes or until the mixture became a viscous paste. The paste was transferred to a

1 mL syringe and extruded into sterile silicone tubing (i,d, 0.8mm) at a rate of 2.5 µL/min

using an infusion pump (Harvard Apparatus, Holliston, MA). After extrusion, tubes were

dried overnight at 60 °C. Implants were then cooled to room temperature and cut to 134 appropriate size (2 ± 0.3 mm in length) for storage in a closed vial at -20 °C. Drug loading per unit length (mm) spacer was determined from randomly selected implants (n=3).

Implants were measured for length, dissolved, and Talazoparib was quantified using high- performance liquid chromatography (HPLC).

The loading of the drug in the Talazoparib implants was quantified via HPLC method using an Agilent 1100 system and reverse phase C18 Zorbek column. A standard curve was prepared for free Talazoparib in both the mobile phase and PBS buffer at concentrations ranging from 1-100 ug/mL. Talazoparib implants (n=3) were dissolved in dichloromethane followed by extraction of Talazoparib using methanol. The drug was quantified using a 50:50 methanol (0.1% TFA):water (0.1%TFA) mobile phase with a flow rate of 0.8 mL/min. Absorbance was detected at 232nm with t ret =3.98 minutes. The standard curves were fit with a linear regression curve (R 2=1) and the curve was used to quantify the concentration of Talazoparib in each spacer.

Talazoparib implants were flash frozen in liquid nitrogen and fractured using a cooled razor. Fractured and non-fractured implants were attached to a specimen mount using a conductive carbon adhesive tab and then sputter coated with 10 – 15 nm of platinum using a vacuum evaporator (DV502, Denton Vacuum, Moorestown, NJ). Spacer samples were imaged in both SE (secondary electrons) and BSE (back-scattered electrons) mode using aS-4800 field emission SEM (Hitachi, Chiyoda, Tokyo, Japan) at 5 kV.

The release profile of Talazoparib from the implants was studied using HPLC. To determine the release kinetics Talazoparib implants were cut to predetermined lengths, measured, and placed in 1 ml of PBS (pH 6.0 adjusted with 1 N HCl) at 37°C with gentle shaking to mimic biological tumor conditions for 28 days. At each predetermined time 135 point, buffer was completely removed and replaced with fresh buffer to maintain the concentration gradient between drug depot and buffer. Each collected buffer fraction was subjected to HPLC analysis as described earlier and released Talazoparib was quantified.

Care was taken as to not agitate and disrupt the integrity of the spacer during the buffer transfer. After 28 days, the spacer was dissolved and any remaining drug was quantified.

The total drug released and remaining in the spacer was summed and compared to the estimated drug loading from HPLC analysis for the given length of spacer. The released drug from the implants was plotted as a function of time.

8.2.2 In vitro Therapeutic Efficacy

W0069 and W780 cell lines derived from mammary tumors from genetically engineered

Brca1 Co/Co ;MMTV-Cre;p53 +/- mice were used for all studies. All cells were grown in

Dulbecco’s Modified Eagles Medium (DMEM) supplemented with 10% FBS and 1%

Penn/Strep. To evaluate and compare the efficacy of various PARP inhibitors, W0069 and

W780 were treated with 10 µM Olaparib, Talazoparib, or Niraparib for 24-72 hours. [134]

Cell lysates were processed for Western blotting with antibodies against cleaved PARP

(Upstate), cleaved caspase 3, γ-H2aX (Abcam), or α-tubulin (Calbiochem). IC50 values were calculated to determine the sensitivity of the cell lines to PARP inhibitors

Talazoparib, Olaparib, and Niraparib. W0069 and W780 cells were seeded into 96 well plates (Corning) at 1000 cells per well. The following day cells were treated with

Talazoparib concentrations ranging from 0 to 100 nM. One week after seeding, cell viability was measured with an MTS assay (Promega) to measure the metabolic activity of the cells. Dose response was plotted and fit using a variable slope four parameter logistic 136 equation constrained at 0 to determine the IC50 value (Prism). All experiments were done in triplicate.

8.2.3 In vivo Therapeutic Efficacy

All animal studies were performed in accordance with protocols approved by the

Institutional Animal Care and Use Committee (IACUC) at Dartmouth College. For in vivo

therapeutic efficacy evaluation, two different lengths of Talazoparib implants were used.

Fabricated implants with the same unit loading of Talazoparib were cut in to two different

lengths of 1 mm and 2 mm. Since the drug loading in Talazoparib implants was 25 µg per

mm, 1 and 2 mm long implant contain a total of 25 and 50 µg of Talazoparib respectively.

The in vivo therapeutic efficacy of Talazoparib implants was carried out in a genetically

engineered Brca1 Co/Co ;MMTV-Cre;p53 +/- breast cancer mouse model. Treatment in mice was started when the tumor size was approx4-6 mm in diameter. The Talazoparib implant were administered intratumorally using 18G brachytherapy applicator needles (Bard

Medical, Covington GA). For intratumoral implantation, the Talazoparib implant was preloaded into the applicator needle and guided laterally into the center of the tumor. The implant was ejected by pressing the stylet simultaneously as the outer needle (cannula) was withdrawn from the tumor. Animal skin glue was used to seal the small hole in the skin to avoid any leakage of the implant out of the tumor. Four groups of animals (n=6-8 animals per group) were used in the study, out of which a control arm received intratumoral implantation of blank PLGA implants (no Talazoparib), one additional control arm was treated with free Talazoparib via oral gavage and two Talazoparib implant groups which were treated with a single 1 mm or 2 mm long implant. Free Talazoparib was given orally 137 for comparison to the current clinical administration of treatment. An equivalent dose of free Talazoparib (as in 2 mm long implant) was given a total of six times by oral gavage every other day over a period of 12 days (8.3 µg per gavage). Tumors size was monitored using calipers and mice were weighed twice weekly to evaluate weight loss. Tumor volumes were calculated using the formula: volume = length x width x height/2. Mice were sacrificed if the tumor diameter exceeded 1 cm, the tumor volume exceeded 10% of body weight, or if a tumor ulcerated or interfered with mobility. For immunohistochemistry, tumors were harvested, fixed in neutral-buffered formalin, embedded in paraffin and sectioned. Sections were stained with PCNA (Santa Cruz 1:200) or γ-H2aX (Abcam 1:100) antibodies as previously described [134]. To determine the

percentage of PCNA- or γ-H2aX-positive cells, two sections on the periphery of each tumor (n=4-5/treatment, ~1000 cells each) were counted. Slides were randomized and group labels removed prior to counting.

Results are expressed as mean ± standard error of mean (SEM) and were analyzed using the t test, χ2 test, or one-way ANOVA on ranks (Wilcoxon signed rank test) using

Prism5 or SigmaStat3.5 software. All P values were two-sided with a P value of <0.05 considered statistically significant.

8.3 Results

8.3.1 PLGA implants continuously release Talazoparib

Poly(lactic-co-glycolic) acid implants (0.8 mm diameter) containing 25 µg/mm

Talazoparib were fabricated by extrusion followed by curing. A single batch of fabrication with the above mentioned PLGA/Talazoparib mixture yields approximately 20 implants of 138

2 mm length. The implant was designed to allow direct placement into solid tumors using an 18G needle, after which Talazoparib was slowly released by PLGA degradation and allowed to diffuse into the tumor (Fig 45a). In these studies, we used implants of 1 and 2 mm length, equivalent to 25 µg and 50 µg of Talazoparib respectively, as measured by

HPLC. Scanning electron microscopy images of flash-frozen and fractured PLGA implants showed a smooth surface intercalated with homogenously distributed nanoscale pores (Fig

45b). It is likely that these pores represent pockets of Talazoparib distributed throughout the implant. Talazoparib implants showed continuous drug release in phosphate buffered saline (pH 6.0) at 37°C. Drug release was characterized by an initial burst release (10% in day 1), followed by a linear release (R 2 = 0.9983) until day 28. (Fig 45c). The initial burst

release can be attributed to desorption of drug molecules from the implant surface. Release

kinetics studies carried out with different length implants showed that the percentage of

drug release was independent of implant length indicating that drug dosing to the tumor

can be adjusted by changing the length of the implant without affecting the kinetics of the

drug released. 139

Figure 45. Biodegradable PLGA implants for sustained intratumoral delivery of Talazoparib . a) Schematic depiction of PLGA implant degradation (blue) for localized delivery of Talazoparib (red). b) Scanning Electron Microscopy images of flash-frozen and fractured, drug-loaded PLGA implants a solid smooth surface with nanoscale pores homogenously spread throughout surface of implant. c) In vitro release profile of Talazoparib from 2 mm PLGA implants (n=3) in phosphate buffered saline (pH 6.0).

8.3.2 Sustained PARP inhibition produce apoptosis and DNA damage in vitro

PARP inhibitors Olaparib, Niraparib, and Talazoparib were compared in BRCA1 -mutant

breast cancer cell lines W780 and W0069 derived from Brca1 Co/Co ;MMTV-Cre;p53 +/- genetically engineered mice to evaluate the potency of three PARP inhibitors in these cell 140 lines. Talazoparib was found to be the most potent inhibitor in the W780 cell line (Fig.

46a), as evidenced by an increase in cleaved PARP (c-PARP), cleaved caspase 3 (c-caspase

3), and γ-H2AX protein expression following 48-72 hours treatment. These proteins are markers of apoptosis and DNA damage, respectively. Similar results were obtained in

W0069 cells (Fig. 46b), with both Talazoparib and Niraparib observed to be more potent than Olaparib. Both W0069 and W780 were highly sensitive to growth inhibition by

Talazoparib, with IC50 values of 11 and 2.6 nM respectively (Fig. 46c).

Figure 46. In vitro analysis of different PARP inhibitors on W780 and W0069 breast cancer cell lines derived from Brca1 Co/Co ;MMTV-Cre;p53 +/- mice. Western Blot data testing protein expression after treatment with Olaparib, Niraparib, and Talazoparib on Brca1 -deficient cell line in a) W780 b) W0069 cells. C) Measured dose response curve and IC50 of Talazoparib in W780 and W0069 cells.

8.3.3 Sustained PARP inhibition produces tumor shrinkage

To determine the in vivo therapeutic efficacy of sustained Talazoparib delivery,

drug-loaded PLGA implants were locally injected into Brca1 Co/Co ;MMTV-Cre;p53 +/- spontaneous mammary gland tumors 4-6 mm in diameter. Mice received a one-time implantation of a single 1 or 2 mm length implant using an 18G brachytherapy needle.

Untreated mice showed no significant difference in tumor growth rate or mouse weight to mice treated with empty implants (data not shown). For the entirety of the study, controls 141 refer to mice treated with control implants. As shown in Table 6, tumors treated with 1 mm control implants (0 µg Talazoparib) grew exponentially and had an average life span of

11.8 ± 1.9 days. These tumors increased in volume more than 7-fold, from an initial average of 67 ± 28 mm 3 to a final average of 504 ± 67 mm 3 (P=0.002). In contrast, tumors treated

with 1 mm Talazoparib implants (25 µg drug) decreased in size by 50%, from an average

initial tumor volume of 130 ± 42 mm 3 to a final volume of 65 ± 14 mm 3 (p = 0.002). The average lifespan of the mice treated with 1 mm drug-loaded implants was more than twice as long as mice in the control group (29 vs. 12 days; P < 0.001).

Table 6. Talazoparib implants decrease tumor size and extend survival in BRCA1- deficient mice

Initial Tumor Final Tumor Volume Volume # of Lifespan after (mm 3) (mm 3) mice/group treatment (days) 0 µg control implant (1mm) 66.7 ± 27.5 504 ± 67.4 a 5 11.8 ± 1.9 25 µg Talazoparib implant 130.1 ± 42 65.4 ± 14.1 b 8 29 ± 3 g

0 µg control implant (2mm) 72.1 ± 24.9 484.3 ± 68.8 c 7 11.9 ± 1.4 50 µg Talazoparib implant 44.4 ± 5.2 14.9 ± 5.7 c,d,e 7 44.1 ± 1.8 h 537.1 ± 50 µg Talazoparib gavage 55.4 ± 16.2 244.2 f 6 39.8 ± 6.4 h

Values are mean ± SEM. a, P = 0.002 0 µg control implant (1mm) final volume vs. initial volume b, P = 0.002 25 µg Talazoparib implant final volume vs. 1mm control final volume c, P < 0.001 50 µg Talazoparib implant final volume vs. 2mm control final volume d, P < 0.05 50 µg Talazoparib implant final volume vs. 50 µg Talazoparib implant initial volume e, P < 0.001 50 µg Talazoparib implant final volume vs. 50 µg Talazoparib gavage final volume f, P = 0.026 50 µg Talazoparib gavage final volume vs. initial volume g, P < 0.001 25 µg Talazoparib implant vs. 0 µg control implant (1mm) h, P < 0.001 50 µg Talazoparib implant and 50 µg Talazoparib gavage vs. 0 µg control implant (2mm)

142

Since implants containing 25 µg Talazoparib were well-tolerated, we chose to test

2 mm implants containing 50 µg Talazoparib (Fig. 47). Measurements of tumor volume

(Fig. 47a), tumor growth rates (Fig. 47b), and mouse survival (Table 6) demonstrated that the higher dose of 50 µg Talazoparib was more effective. Tumors treated with 50 µg

Talazoparib implant decreased in volume by 67%, from 44 ± 5 mm 3 at the beginning of

treatment to 15 ± 6 mm 3 after treatment (P < 0.05). The average lifespan of mice receiving

Talazoparib implants was 44 days vs. 12 days (P < 0.001) for mice receiving control PLGA implants. An additional control arm, in which mice were treated with 50 µg Talazoparib by oral gavage over 12 days (8.3 µg x 6 doses), showed a slowed tumor growth rate (Fig.

47b) compared to the 0 µg control implant, but was not as effective as the 50 µg implant group. As a result, the final tumor volume in the oral gavage group was no better than the control group but the average lifespan was increased (40 days vs. 12 days, P < 0.001).

Figure 47 . Treatment with 50 µg Talazoparib implants decreases tumor size and slows tumor growth in Brca1 -deficient mice. Implants (2 mm × 0.8 mm diameter) containing 0 µg (control) or 50 µg Talazoparib were injected into established mammary gland tumors in female Brca Co/Co ;MMTV-Cre;p53 +/- mice. Mice treated by oral gavage received 6 doses of Talazoparib (50 µg Talazoparib total). A) Average tumor volumes before and after treatment (n=5-8/group). *, P < 0.05 vs. initial tumor volume; ǂ, P < 0.05 vs. all other groups. B) Change in tumor volume over a period of 28 days. C) Average body weight before and after treatment (* P < 0.05 vs. initial weight for the treatment group).

For evaluating the gross toxicity, mouse body weight was measured over the course

of the survival studies. As shown in Fig. 47c, oral gavage of Talazoparib was accompanied 143 by a statistically significant (P < 0.05) decrease of ~ 15% in mouse body weight that was not observed in mice treated with 50 µg Talazoparib implants. Talazoparib implants were well-tolerated and produced no significant change in initial vs. final body weight in either treatment group. Mice in the control group lost weight (P < 0.05), which was likely the result of ill health as they became moribund or perhaps a consequence of the relatively short lifespan (~12d) after treatment which did not allow for full recovery from the procedure. The absence of any adverse toxic effects in the Talazoparib implants groups also indicate that the Talazoparib implants have no residual solvents which can cause toxicity.

In order to verify tumor inhibition was a result of successful delivery of localized

Talazoparib and its mechanism of action on tumor cells rather than a result of overt toxicity from the implant, tumor sections were analyzed by immunohistochemistry for assessment of cell proliferation and double stranded DNA breaks, two biomarkers indicative of successful treatment of PARP inhibitors. The results showed that the tumors treated with

Talazoparib implants were not necrotic. Cell proliferation, as measured by PCNA staining

(Fig. 48), was 4-fold lower than in control tumors (13.6 ± 5.3% vs. 53.5% ± 8.3%, P <

0.001). Talazoparib implants also showed significantly increased γ-H2AX expression, a marker of double-stranded DNA breaks. The percentage of γ-H2AX-positive cells in

tumors treated with Talazoparib implants was 12.7 ± 6.6% compared to 2.2 ± 0.6% in

control tumors treated with empty implants (P < 0.001). No overt signs of toxicity were

observed.

144

8.4 Discussion

Improved understanding of the genetic mutations prevalent in breast cancer has led to the development of new therapeutic strategies which target these mutations. Breast cancers with germline BRCA1 and BRCA2 mutations are promising candidates for DNA

repair inhibition using Talazoparib. Talazoparib is a potent PARP1/2 inhibitor for the

treatment of patients with deleterious germline BRCA1 or BRCA2 mutations and locally

advanced and/or metastatic breast cancer. Currently, Talazoparib is administered orally,

which not only limits its bioavailability but also results in treatment-associated toxicities

in a majority of patients. Sustained, localized drug delivery systems that eliminate the need

for daily oral dosing have the potential to improve patient outcomes and reduce systemic

toxicity. Here we show that implant-based delivery of Talazoparib can address the

limitations associated with oral Talazoparib.

Current brachytherapy procedures utilize inert spacers that are delivered

intratumorally via an 18G needle. Our implants, of similar dimensions to brachytherapy 145 spacers, can be easily and readily administered as an outpatient procedure using existing clinical protocols for breast cancer treatment. In this proof-of-concept study, we designed a degradable PLGA matrix containing embedded Talazoparib in order to slowly and continuously release this drug into tumors. Localized delivery of Talazoparib promises to not only reduce the amount of drug required to achieve therapeutic benefit but will also provide a continuous supply of Talazoparib to tumor cells over several weeks. Thus,

Talazoparib implants have the potential to bypass the concerns and challenges associated with oral delivery, metabolic stability, and pharmacokinetics. In the future, we envision that such implant could be used to deliver not just one drug but multiple therapies in order to produce synergistic therapeutic outcomes with reduced toxicity.

PLGA is an FDA-approved, biodegradable polymer with low glass transition temperature provides an excellent matrix for implant fabrication. The hydrophobicity of

Talazoparib ensures a uniform distribution of drug in the PLGA matrix, which was confirmed by scanning electron microscopy. Implant drug loading could be quantified by

HPLC and the total available drug concentration could be tailored by changing implant length or administering multiple implants in the tumor. Talazoparib was selected for in vivo studies as it proved to be the most potent of the PARP inhibitors tested in the cell lines derived from the in vivo mouse model. The in vitro release kinetics showed a slow sustained release of the drug over 28 days, comparable to the 21 day time-frame in which

Talazoparib is administered daily in clinical trials. While we expect these release kinetics to be altered in vivo due to enzymatic degradation of the PLGA polymer, we did observe a therapeutic benefit for as long as 28 days, suggesting that Talazoparib release in vivo is slow and sustained. Typically, PLGA is degraded via the hydrolysis of the ester bonds of 146 the polymer backbone resulting in formation of acidic oligomers and we anticipate a similar degradation mechanism for Talazoparib implants. Since the release of the drug in buffer showed a diffusion based release, it is highly unlikely that there is any interaction between the drug molecules and PLGA matrix. The sustained release of the drug can be attributed to the diffusional constraints and slow degradation of the PLGA matrix. The competing dynamics with the continuous release of Talazoparib and simultaneous inhibitory response on the tumor cells will also alter the release profile of Talazoparib from the implants.

However, the release profile in buffer indicates zero order kinetics suggesting a diffusion based release of Talazoparib. Thus, we predict that inside the tumor matrix, the release of

Talazoparib will be a combination of diffusion as well as enzymatic degradation and surface erosion of the implants. Evidence of increased double-stranded DNA damage and decreased tumor cell proliferation was present at least 2 weeks after implantation in comparison to control tumors treated with empty implants. No gross toxicity was observed in any of the control mice, suggesting the method of fabrication completely absolves the implant of any trace amounts of toxic solvents. During fabrication, we dry the implants above the glass transition temperature of the polymer to induce the temperature dependent phase transition. This heating also allows for high vapor pressure and volatile solvents such as chloroform to completely evaporate during that drying time, which is evident in the lack of toxicity and biomarker changes in tumors treated with the implant. The tumor growth inhibition seen in our Talazoparib implant group weeks after treatment further attests to the long-term benefit of slow releasing Talazoparib implants

For evaluating the therapeutic efficacy of the Talazoparib implants, two different lengths were prepared. The cumulative drug doses used this study are 5-10 times lower 147 than those traditionally administered in preclinical oral dosing studies [141]. The 25 µg dose was well-tolerated and decreased the tumor size by 50%. When the treatment dose was doubled, the higher dose was found to significantly enhance tumor growth inhibition and survival time with no additional body weight loss. Taken together, these data suggest that treatment response is dose-dependent. Given that the implants do not appear to generate overt systemic toxicities, our approach opens new opportunities for safely treating patients that require drug concentrations higher than the conventional dosing scheme.

Localized drug delivery at the target site have several positive attributes which includes reduced systemic toxicity as well as minimal toxicity to surrounding tissues, increased control and decreased variability by eliminating reliance on patient circulation for distribution as in case of intravenous administration of drugs. Localized drug delivery formulations can also impart sustained release behavior of the entrapped drug and thus have the potential to increase the effectiveness of treatment at lower concentrations as opposed to systemic or oral administrations. These formulations also modulate the pharmacokinetics of the drug in systemic circulation with a high concentration of drug delivered directly to the tumor site. While delivering drugs directly at the tumor site can improve the bioavailability in diseased cells while bypassing administration and reducing toxicity to healthy organs, local delivery is a major limitation in the case of metastatic disease.

In this study, mice with spontaneous growing tumors were used to test the implant against the current clinical administration (oral), and in many instances mice grew multiple tumors. There was little difference between the average lifespan of mice treated with

Talazoparib by either the intratumoral implant or by oral gavage. The primary tumor was 148 greatly reduced in size in mice receiving implants, but the drug had no effect on the other spontaneous tumors growing concomitantly in this mouse model, necessitating mouse sacrifice due to tumor burden. Thus mice treated with Talazoparib implants did not show any enhanced survival compared to those treated by gavage. This limitation needs to be carefully considered to determine the appropriate cases for clinical application of

Talazoparib implants (i.e. in unresectable tumors, or in combination with other inhibitors, chemotherapeutics, or radiation treatment). However, it is worth mentioning that these

Talazoparib implants can be used in conjunction with standard of care radiation or chemotherapy. These implants can be administered prior to radiation therapy or chemotherapy to combat the metastatic conditions. Currently, our group have initiated in vivo experiments to evaluate the combination therapy with external beam radiation therapy

(EBRT) and Talazoparib implants in lung cancer animal models as well as a combination therapy with cisplatin and Talazoparib implants in the above mentioned breast cancer animal model. Also, we did observe some evidence of primary tumor regrowth in implant- treated mice that lived beyond 30 days (data not shown). At these later time-points, tumor drug levels were not detectable by HPLC/MS. In future studies, we will evaluate if a second administration of Talazoparib implants, or the placement of implants into concomitantly growing tumors, would further prolong the survival time.

The use of Talazoparib implants as a therapeutic intervention in unresectable tumors either as monotherapy or in combination with other standard of care therapies like radiation or chemotherapy have the potential to provide high therapeutic benefits with minimal systemic toxicities. This approach of localized treatment have strong clinical relevance as the treatment using these implants is based on standard clinical brachytherapy 149 procedure which is routinely followed in clinics and will not require additional procedures or protocols during treatment. The limitation associated with this localized treatment approach is mainly in the metastatic conditions but which can be addressed by careful selection of patient populations as well as combining the localized treatment with the standard of care therapy. Overall, a good therapeutic index, low systemic toxicity and the flexibility of the Talazoparib implants formulation to combine with other therapies provides a promising therapeutic platform which can be applied in treatment of different cancer types.

8.5 Conclusion

In conclusion, we fabricated a PLGA implant for localized and sustained intratumoral release of Talazoparib. This unique route of PARP inhibitor delivery significantly slows tumor growth and prolongs survival time in a spontaneous mouse model of Brca1 -deficient breast cancer. Through optimization of the PLGA formulation, sustained drug release was achieved over a scale of ~28 days both in vitro and in vivo .

PARP inhibition was associated with significantly increased double-stranded DNA damage and decreased tumor cell proliferation for several weeks after implantation, further supporting the long-term effectiveness of these implants. The treatment efficacy was dependent on drug dose of the implant, with 50 µg Talazoparib producing more tumor growth inhibition than a 25 µg Talazoparib dose. No acute or chronic toxicity was observed in mice treated with clinically relevant doses for 28 days, indicating that Talazoparib implants may present an easy and feasible way for controlling treatment outcomes with lower systemic toxicity.

150

Chapter 9. Perspectives and Conclusions

This dissertation has made several important contributions in the field of cancer chemotherapy, including formation optimization and dosing optimization. As tumor biology has become more understood, clinicians are seeking to develop new treatment modalities which can target and treat cancer early on in the disease. Even with billions of dollars in research, translating preclinical findings to relevant clinical treatments remains a great challenge. This dissertation addresses the problems associated with prostate cancer treatment, a leading cause of death in men world-wide and proposes the reformulation of docetaxel for early local treatment. Docetaxel is not recommended or clinically used for treatment in local prostate cancer, because like many other anti-cancer agents, the low bioavailability and toxic side effects outweigh any benefit from the small fraction of drug that accumulates at the disease site. To overcome these challenges, we have designed a biodegradable PLGA implant with embedded docetaxel which can be injected via an 18 gauge applicator needle directly into the prostate. This allows us to achieve high local drug concentration directly to the tumor site while sparing healthy organs. These implants were designed as modified brachytherapy spacers for applications in local radiation therapy.

While inert spacers are used for spatial and temporal guidance during brachytherapy, the spacers themselves offer no clinical value. The modified smart spacers are morphologically and physically similar to brachytherapy spacers and allow the same spatial and temporal guidance with the added benefit of sustained chemotherapy and radiosensitization.

This new platform was designed with versatile applications in monotherapy and combinatorial treatments. The hydrophobicity and degradation of PLGA can be tailored to 151 release a number of agents including nanoparticles, chemo-toxic agents, molecular inhibitors, proteins or siRNA, etc. for local and sustained therapy or theranostics in cancer.

Chapters 2 and 3 discuss the development of INCeRT spacers loaded with silica nanoparticles for a dual release platform. Chapter 4 discusses the optimization of this new platform using optical imaging in vivo to determine the degradation and release kinetics of multiple formulations of spacers. Chapter 5, 6, and 7 are focused on the optimized docetaxel loaded implant for treatment in prostate cancer. First the feasibility of a slow eluding drug depot is explored as a monotherapy, and then synergistic therapeutic benefit is explored in combination with radiation for ultimate applications in brachytherapy. The toxicity of the new implant as a mono and combination therapy is tested. Chapter 8 shows the versatility of the platform with applications in aggressive breast cancer. Here, PARP inhibitor Talazoparib was embedded in PLGA for a four week release in BRCA -deficient mice. As a whole, this dissertation discusses the design challenges and optimization parameters considered for a slow release drug implant. Docetaxel spacers were tested in vitro and in vivo and proved to be significantly more effective in human prostate cancer cell line PC3 subcutaneous xenografts. Low doses of radiation and chemotherapy were able to produce significant therapeutic response with extended survival times in preclinical studies. This work has laid the foundation for moving forward into large animal models and eventually, early clinical trials.

The work presented in this dissertation was designed with great consideration for

clinical translation. PLGA and docetaxel are FDA-approved materials for therapy in

humans and, while developing a new administration modality poses difficulty in

translation, having all components well studied and characterized. The FDA regulates that 152 new formulations of current treatments must not cause any new discomfort or toxicities worse than the current standard of care. To accommodate this, we have developed a slow releasing implant of docetaxel that is to be administered during the current clinical procedure brachytherapy as to not inflict any new procedures to the patient’s treatment.

This one-time intratumoral injection will release drug locally near the implantation site and is envisioned to be placed using similar dose mapping and software as brachytherapy. In order to do so, a great understanding about the degradation of the PLGA implant in human tissue and the diffusivity of docetaxel in the prostate and tumor tissue needs to be studied

It is important to mention that biodistribution and PK studies would typically be

done in the development stages of new chemotherapeutics; however for the scope of this

study, I found more flaws than advantages in the proposition. As mentioned in the

introduction, the prostate is surrounded by a fibromuscular sheath which has been shown

in canine studies to act as a natural barrier and prevent drug diffusion out of the prostate.

The mouse model we are using is effective in demonstrating preliminary therapeutic

efficacy, however the subcutaneous tumor xenograft in no way reflects the physiological

conditions and boundaries we would expect to see in an orthotopic setting of a larger animal

model. Therefore, as resources are limited and to spare unnecessary animal lives the study

is best saved for a canine model, where docetaxel levels in plasma can be determined in

measurable quantities via blood draws rather than complete rodent sacrifice to collect the

amount of blood needed to quantify the docetaxel in blood. It is a safe assumption that we

do not have complete dose dumping of the drug from the implant in vivo, as multiple

animal experiments showed therapeutic response nearly 3 months post treatment. 153

Future studies in large animal models (canines) would provide further insight into the release kinetics and diffusion patterns in vivo . These parameters must be studied to

determine the minimum dose which causes necrosis in healthy tissue (local maximum

tolerated dose). Data in canines would provide researchers with the understanding of the

drug concentration gradient from the implant with respect to time as the implant continues

to degrade. It is imperative to understand these properties to design appropriate dosing

plans for whole tumor treatment while avoiding drug delivery and toxicities to healthy

tissue.

In conclusion, the research in this dissertation provided a safe and translatable

solution for local treatment of early stage prostate cancer. Rather than ‘reinvent the wheel’,

we have reformulated a currently utilized chemo-toxic agent docetaxel for localized disease

using a slow releasing implant. The implant has been designed with specific consideration

for applications in brachytherapy for radiosensitization and synergistic therapy over the

course of 70 days. I am confident that this work can provide the necessary foundation and

proof of the therapeutic benefit resulting from local delivery of docetaxel using a

biodegradable polymer with the hope of clinical translation in the near future.

154

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Figure A1. American Cancer Society prostate cancer staging guidelines

169

Parameters Normal range Tumored Control Radiation Radiation+ DTX Spacer Radiation+IV DTX 24 hours 1 week 3 weeks 24 hours 1 week 3 weeks 24 hours 1 week 3 weeks mean stdev mean stdev mean stdev mean stdev mean stdev mean stdev mean stdev mean stdev ALBUMIN 2.8-3.8 g/dL 2.50 0.16 3.20 2.21 1.16 0.57 1.35 0.52 1.20 0.40 2.45 0.24 2.54 0.55 0.86 0.32 2.36 0.44 2.95 0.60 ALT 28-129 U/L 97.80 76.39 100.80 59.98 75.80 24.81 54.25 22.63 41.00 26.76 85.75 92.82 65.00 32.37 125.80 59.34 56.20 40.01 114.00 73.67 AMYLASE U/L 620.20 124.94 858.80 400.26 793.20 67.32 664.00 120.59 627.50 156.37 721.50 82.26 280.80 255.46 605.40 250.91 579.80 140.67 927.50 275.23 AST 46-392 U/L 1172.50 574.22 1150.80 544.58 1235.20 576.00 676.50 191.55 584.50 516.27 568.33 513.38 269.80 417.07 1368.80 826.06 780.25 449.01 1633.00 1166.47 BILIRUBIN DIRECT MEQ/L 1.28 0.61 0.90 0.00 0.60 0.23 1.03 0.39 0.90 0.00 0.78 0.46 1.14 0.51 0.40 0.16 0.90 0.32 0.95 0.29 BUN 7-28 mg/dL 20.80 1.30 15.20 6.42 19.40 2.19 12.50 1.29 9.50 3.00 18.50 3.42 34.20 18.38 10.00 9.14 16.00 2.83 22.50 5.74 CALCIUM 9.7-12.5 mg/dL 7.62 1.46 7.84 6.15 5.62 2.63 4.23 0.55 3.25 0.68 7.98 1.02 4.74 1.83 4.54 2.73 7.28 1.34 9.90 1.05 CHOLESTEROL 69-169 mg/dL 79.80 10.89 87.20 8.79 90.40 6.31 79.00 8.60 39.00 8.72 80.25 7.41 45.80 31.90 62.20 11.65 80.80 15.59 100.00 2.31 CARBON DIOXIDE MEQ/L 9.02 3.43 25.00 0.28 9.38 2.54 6.00 1.48 <4 12.43 4.51 9.72 7.67 10.64 5.65 11.65 3.63 GGT 0-8 U/L 6.80 2.17 11.20 1.30 7.00 2.45 8.25 2.06 11.75 1.26 9.75 1.50 4.80 3.27 5.40 1.14 7.40 0.89 8.50 3.79 GLUCOSE 172-372 g/dL 258.40 55.97 264.80 28.02 228.40 16.92 209.75 38.87 91.50 24.62 244.25 37.43 44.40 73.84 193.80 81.03 236.60 34.41 279.00 90.34 IRON 108.20 40.07 129.20 54.12 122.20 23.27 115.75 31.93 72.50 15.44 94.75 22.38 69.60 67.84 61.00 39.76 115.40 20.95 142.00 63.46 MAGNESIUM mg/dL 2.68 0.54 9.20 13.92 1.56 0.72 1.45 0.26 1.20 0.69 2.50 0.28 1.74 0.39 1.60 0.79 1.98 0.40 3.15 0.38 PHOSPHORUS 7.9-14.5 mg/dL 5.84 0.98 20.64 33.24 5.36 1.69 3.30 0.45 4.05 1.14 8.58 0.85 11.36 9.60 7.44 2.22 5.80 1.25 7.45 1.24 BILIRUBIN TOTAL 0.0-0.9 mg/dL 1.02 0.43 0.72 0.04 0.62 0.13 0.83 0.45 0.73 0.05 0.53 0.30 0.70 0.27 0.90 0.33 0.74 0.24 0.90 0.38 PROTEIN TOTAL 3.9-6.4 mg/dL 4.60 0.39 4.12 3.77 3.62 1.27 2.83 0.22 2.15 0.62 4.58 0.59 3.96 0.63 3.10 1.84 4.18 0.66 6.15 0.38 TRIGLYCERIDES 67-278 mg/dL 92.80 16.77 106.40 18.78 155.80 29.51 117.00 10.52 61.50 9.00 87.25 34.81 85.40 44.05 97.80 25.31 82.80 8.29 131.50 11.00 URIC ACID 1.2-2.3 mg/dL 4.82 1.29 6.84 1.56 3.58 1.21 4.35 0.31 6.25 0.64 3.30 1.23 3.68 1.84 4.26 0.79 3.94 0.61 7.05 0.60 GLOBULIN g/dL 2.10 0.25 1.05 1.03 2.46 1.41 1.48 0.71 0.48 0.15 2.13 0.43 1.42 0.65 2.24 1.54 1.82 0.23 3.20 0.78 A/G RATIO 1.22 0.13 1.68 0.69 0.64 0.56 1.18 0.75 1.33 0.39 1.18 0.25 2.48 2.16 0.68 0.74 1.30 0.12 0.98 0.39 INDIRECT BILI 0.00 0.00 0.00 0.00 0.10 0.17 0.08 0.15 0.00 0.00 0.00 0.00 0.00 0.00 0.50 0.35 0.00 0.00 0.05 0.10 AST/ALT RATIO 18.88 4.30 19.34 18.55 15.84 2.55 12.93 2.16 13.45 4.19 20.80 16.60 3.74 4.54 10.62 3.97 18.75 2.22 14.13 3.64

Parameters Normal range Tumored Control DTX Spacer IV DTX Blank Spacer 24 hours 1 week 3 weeks 24 hours 1 week 3 weeks 24 hours 1 week 3 weeks mean stdev mean stdev mean stdev mean stdev mean stdev mean stdev mean stdev ALBUMIN 2.8-3.8 g/dL 2.50 0.16 2.46 0.87 1.73 0.80 2.10 0.59 2.50 0.34 1.48 0.43 2.46 0.22 0.88 0.62 3.37 0.75 1.78 0.41 ALT 28-129 U/L 97.80 76.39 118.00 41.44 196.00 65.26 73.00 60.03 92.50 90.02 130.80 68.87 93.60 72.46 112.00 41.21 190.25 78.31 63.00 23.70 AMYLASE U/L 620.20 124.94 583.40 78.14 679.25 48.08 471.20 72.88 693.50 21.30 562.60 150.53 630.80 131.09 446.00 111.00 690.75 194.15 576.60 161.36 AST 46-392 U/L 1172.50 574.22 1458.67 641.63 1386.00 #DIV/0! 966.50 1173.09 816.33 1144.36 1352.33 901.78 1012.60 545.82 1214.33 767.76 2497.33 1161.23 437.00 #DIV/0! BILIRUBIN DIRECT MEQ/L 1.28 0.61 1.42 0.45 1.10 0.47 1.10 0.57 0.88 0.71 0.78 0.08 0.62 0.19 1.10 0.47 1.35 0.64 0.88 0.41 BUN 7-28 mg/dL 20.80 1.30 21.80 1.79 22.75 3.86 26.00 12.92 21.00 2.45 19.20 5.81 21.20 3.35 13.50 7.19 24.25 7.14 11.60 5.41 CALCIUM 9.7-12.5 mg/dL 7.62 1.46 8.00 0.49 7.50 0.53 3.00 1.72 7.93 0.40 5.74 1.93 7.86 1.29 6.28 1.88 31.38 48.43 3.78 0.96 CHOLESTEROL 69-169 mg/dL 79.80 10.89 93.20 5.31 83.25 3.30 55.80 26.78 89.50 9.15 80.60 8.41 73.00 4.30 87.25 21.79 107.67 17.90 103.20 22.35 CARBON DIOXIDE MEQ/L 9.02 3.43 6.75 1.99 10.46 2.57 28.96 43.76 9.10 3.60 11.70 3.75 10.58 3.39 GGT 0-8 U/L 6.80 2.17 4.60 1.95 5.00 1.83 6.80 3.03 4.25 0.50 4.40 1.95 6.40 1.95 9.25 2.22 14.50 4.12 6.60 2.19 GLUCOSE 172-372 g/dL 258.40 55.97 225.60 46.60 212.25 28.58 187.40 121.91 247.25 32.98 195.80 55.44 228.60 69.82 168.75 57.16 246.00 72.72 277.40 60.38 IRON 108.20 40.07 125.40 32.25 165.00 5.29 109.40 18.37 127.50 27.38 70.40 53.70 111.80 8.07 170.00 58.69 177.33 61.10 122.20 22.50 MAGNESIUM mg/dL 2.68 0.54 2.84 0.11 3.00 0.16 1.50 0.32 2.78 0.33 1.88 0.86 2.74 0.55 2.23 0.78 2.83 0.83 1.62 0.61 PHOSPHORUS 7.9-14.5 mg/dL 5.84 0.98 3.82 1.28 5.78 0.35 9.30 3.28 5.68 1.84 5.18 1.55 7.94 1.33 5.43 1.27 6.77 1.95 5.00 1.07 BILIRUBIN TOTAL 0.0-0.9 mg/dL 1.02 0.43 1.32 0.23 1.43 0.26 1.00 0.28 0.85 0.58 0.74 0.43 0.62 0.23 1.00 0.29 1.35 0.54 0.64 0.21 PROTEIN TOTAL 3.9-6.4 mg/dL 4.60 0.39 5.24 0.22 4.93 0.49 3.98 0.68 4.95 0.37 3.98 0.95 4.40 0.82 4.00 1.40 5.20 1.40 3.16 0.55 TRIGLYCERIDES 67-278 mg/dL 92.80 16.77 133.00 29.67 157.50 19.12 233.80 88.71 121.25 35.06 99.00 27.72 76.60 14.77 82.25 12.76 150.75 54.37 123.00 68.82 URIC ACID 1.2-2.3 mg/dL 4.82 1.29 4.54 1.11 5.10 0.22 4.76 1.38 3.88 1.65 5.58 1.74 3.52 1.31 4.95 1.66 8.95 2.73 3.68 0.59 GLOBULIN g/dL 2.10 0.25 2.78 1.08 3.20 1.22 1.88 0.23 2.45 0.42 2.50 0.91 1.94 0.68 3.13 1.65 2.43 0.15 1.38 0.79 A/G RATIO 1.22 0.13 1.06 0.48 0.68 0.50 1.12 0.31 1.05 0.26 0.68 0.33 1.42 0.44 0.48 0.62 1.40 0.26 1.78 1.19 INDIRECT BILI 0.00 0.00 0.18 0.40 0.45 0.54 0.24 0.34 0.13 0.25 0.12 0.27 0.12 0.27 0.20 0.40 0.15 0.30 0.06 0.13 AST/ALT RATIO 18.88 4.30 14.50 0.20 13.10 #DIV/0! 22.35 15.49 10.80 6.44 14.37 0.21 12.12 3.82 11.57 6.70 14.97 0.49 7.30 0.00

Figure A2. Mean Blood Chemistry Results for treatment groups in TRAMP-C1 C57BL/6 tumored toxicity study

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Parameters Normal range Tumored Control Radiation Radiation+ DTX Spacer Radiation+IV DTX 24 hours 1 week 3 weeks 24 hours 1 week 3 weeks 24 hours 1 week 3 weeks mean stdev mean stdev mean stdev mean stdev mean stdev mean stdev mean stdev mean stdev WBC (x1000) 4.45-13.96 2.09 0.45 1.66 1.08 1.67 0.93 5.13 2.12 1.25 1.18 2.83 1.03 3.38 0.55 0.90 0.64 2.51 1.14 3.55 1.04 RBC (x10^5) 7.14-12.20 8.42 1.64 8.73 1.23 8.86 1.01 7.93 2.38 9.29 1.21 8.69 0.42 7.36 3.44 6.77 2.74 8.44 1.29 7.48 2.60 HGB (g/dl) 10.8-19.2 11.74 1.72 12.03 1.59 11.83 1.31 0.00 0.00 12.53 1.38 11.74 0.47 0.00 0.00 9.45 3.46 11.92 1.45 6.88 6.48 HCT % 37.3-62.0 44.12 8.09 44.67 8.18 46.78 5.80 45.63 5.94 46.95 7.52 44.82 3.17 40.35 16.44 34.90 14.28 44.66 7.83 39.36 10.76 MCV (fL) 42.7-56.0 52.50 1.00 51.03 2.40 52.75 0.69 61.03 15.53 50.40 1.69 51.58 1.44 57.90 8.79 51.55 0.21 52.78 1.52 55.34 10.51 MCH (pg) 11.7-16.3 14.08 0.92 13.77 0.15 13.38 0.17 0.10 0.00 13.50 0.24 13.52 0.20 0.13 0.05 14.05 0.64 14.18 0.58 12.30 12.97 MCHC (g/dL) 24.6-34.9 26.76 1.37 27.10 1.55 25.40 0.42 0.10 0.00 26.80 1.52 26.22 1.02 0.15 0.10 27.25 1.20 26.94 1.88 20.34 19.20 CHCM (g/dL) 33.9-35.4 25.60 1.14 24.67 1.25 24.08 0.17 26.73 5.59 24.78 0.99 25.04 0.85 26.80 5.55 25.20 0.71 24.80 0.91 26.48 4.57 CH (pg) 22.5-31.8 13.52 0.94 12.53 0.06 12.68 0.10 16.90 8.40 12.45 0.13 12.88 0.19 15.88 6.02 13.00 0.28 13.06 0.15 15.02 5.99 RDW % 15.9-20.3 14.18 2.71 12.23 0.23 12.50 0.33 12.68 0.59 12.10 0.14 12.60 0.59 13.65 1.84 11.80 0.99 11.86 0.22 13.50 1.45 HDW (g/dL) 2.84-3.35 2.28 1.16 1.74 0.02 1.81 0.02 1.62 0.32 1.70 0.05 1.83 0.07 1.78 0.04 1.78 0.04 1.69 0.08 1.89 0.24 PLT (10^3 cmm) 841-2159 615.00 310.95 390.67 246.51 795.25 338.70 655.00 428.80 663.75 207.14 1026.40 490.97 399.00 415.11 402.00 544.47 484.60 466.15 339.40 552.10 MPV (fL) 4.3-6.1 6.76 1.14 9.90 6.52 6.33 0.44 7.58 1.60 7.18 1.97 7.16 2.43 10.03 5.34 12.25 6.86 12.18 7.96 16.04 6.04 NEUT % 7.36-28.59 13.28 8.81 8.33 0.78 11.35 3.78 19.33 21.24 9.05 2.71 14.50 7.95 14.90 14.44 7.50 3.96 5.48 1.13 23.66 21.11 NEUT COUNT 0.53-3.09 K/UL 0.30 0.26 0.13 0.08 0.16 0.08 1.25 1.83 0.10 0.07 0.42 0.27 0.49 0.44 0.06 0.01 0.13 0.05 0.92 0.93 LYMPH % 61.26-87.18 79.28 11.77 74.40 7.16 57.80 28.94 69.03 28.64 60.90 20.62 71.10 12.99 74.93 23.12 76.25 1.91 88.52 1.58 64.32 26.12 LYMPH COUNT 3.24-11.15 K/UL 1.63 0.27 1.28 0.90 1.12 0.76 3.19 1.30 0.89 1.09 2.02 0.95 2.57 0.97 0.70 0.50 2.22 1.01 2.18 0.95 MONO % 2.18-11.02 1.40 0.76 1.63 1.07 0.58 0.39 0.98 0.28 0.65 0.19 1.34 0.72 1.43 0.70 1.05 0.92 1.30 0.47 1.56 0.66 MONO COUNT 0.15-0.94 K/UL 0.03 0.02 0.03 0.04 0.01 0.01 0.05 0.02 0.01 0.01 0.04 0.03 0.05 0.03 0.01 0.01 0.04 0.02 0.05 0.03 EOS % 0.13-4.42 5.04 3.12 12.67 9.03 29.10 26.69 7.75 9.17 27.38 21.64 11.38 11.69 7.30 9.73 13.75 0.35 2.10 0.34 7.22 7.86 EOS COUNT 0.01-0.42 K/UL 0.11 0.09 0.15 0.01 0.37 0.41 0.50 0.76 0.23 0.12 0.29 0.25 0.24 0.30 0.13 0.09 0.05 0.03 0.30 0.40 BASO % 0.01-1.24 0.40 0.16 1.53 0.67 0.80 0.69 1.93 1.19 1.55 1.16 0.74 0.17 0.68 0.17 1.15 0.49 0.94 0.89 0.98 0.25 BASO COUNT 0.00-0.13 K/UL 0.01 0.00 0.03 0.03 0.01 0.01 0.09 0.07 0.02 0.03 0.02 0.02 0.03 0.01 0.01 0.00 0.02 0.02 0.03 0.01 LUC % % 0.66 0.18 1.43 1.26 0.88 0.64 1.23 0.55 1.78 1.06 0.98 0.79 0.93 0.21 1.00 0.28 1.62 1.06 2.44 2.22 LUC COUNT K/uL 0.01 0.01 0.03 0.04 0.01 0.01 0.06 0.03 0.03 0.03 0.03 0.04 0.03 0.01 0.01 0.01 0.05 0.03 0.08 0.06

Parameters Normal range Tumored Control DTX Spacer IV DTX Blank Spacer 24 hours 1 week 3 weeks 24 hours 1 week 3 weeks 24 hours 1 week 3 weeks mean stdev mean stdev mean stdev mean stdev mean stdev mean stdev mean stdev WBC (x1000) 4.45-13.96 2.09 0.45 6.51 1.58 4.08 0.90 4.36 1.32 2.50 0.83 3.34 0.62 3.08 2.46 2.41 0.62 4.28 2.82 3.34 1.04 RBC (x10^5) 7.14-12.20 8.42 1.64 6.85 2.22 8.19 0.33 8.38 0.86 5.70 4.91 5.75 1.19 8.21 0.71 8.88 1.49 6.76 2.32 8.65 0.57 HGB (g/dl) 10.8-19.2 11.74 1.72 11.70 1.04 11.00 0.38 0.00 0.00 8.90 4.95 7.38 1.83 10.88 0.89 12.07 1.62 10.73 1.07 0.00 0.00 HCT % 37.3-62.0 44.12 8.09 39.70 5.66 42.38 2.36 44.85 5.65 32.20 24.75 28.83 6.47 43.34 4.05 44.13 8.03 38.80 6.22 47.87 5.46 MCV (fL) 42.7-56.0 52.50 1.00 61.00 14.81 44.23 15.43 53.45 1.34 60.30 8.63 49.98 1.17 52.82 2.74 49.60 1.47 60.93 16.59 55.23 2.76 MCH (pg) 11.7-16.3 14.08 0.92 18.77 8.17 13.43 0.17 0.10 0.00 18.85 7.57 12.73 0.56 13.30 0.79 13.67 0.46 17.77 8.52 0.10 0.00 MCHC (g/dL) 24.6-34.9 26.76 1.37 29.87 5.40 25.95 0.70 0.10 0.00 30.70 8.06 25.45 0.64 25.16 0.46 27.53 1.50 28.13 5.44 0.10 0.00 CHCM (g/dL) 33.9-35.4 25.60 1.14 28.13 6.81 24.98 0.57 24.23 1.09 29.50 8.77 24.83 0.56 23.10 0.28 25.83 0.81 28.87 6.29 23.43 1.12 CH (pg) 22.5-31.8 13.52 0.94 17.83 8.89 12.93 0.15 12.90 0.29 18.20 7.92 12.40 0.50 12.14 0.53 12.83 0.06 18.27 9.12 12.90 0.10 RDW % 15.9-20.3 14.18 2.71 12.47 0.91 11.93 0.10 12.10 0.12 13.75 2.90 16.00 1.84 13.82 2.34 12.03 0.71 12.97 0.78 12.13 0.15 HDW (g/dL) 2.84-3.35 2.28 1.16 1.56 0.23 1.79 0.03 1.82 0.05 1.74 0.06 2.02 0.33 1.79 0.12 1.80 0.06 1.62 0.35 1.82 0.02 PLT (10^3 cmm) 841-2159 615.00 310.95 187.67 119.27 126.50 54.21 291.50 247.67 714.00 934.80 251.75 98.02 1034.60 431.02 177.67 155.74 352.00 285.10 333.00 217.46 MPV (fL) 4.3-6.1 6.76 1.14 12.37 4.72 18.70 4.59 12.80 8.72 8.20 2.12 10.95 1.80 7.16 1.18 15.37 7.87 11.70 4.44 9.00 1.71 NEUT % 7.36-28.59 13.28 8.81 18.97 22.77 8.18 2.94 10.15 2.85 24.30 21.35 24.28 13.93 22.40 17.36 9.73 0.97 27.53 26.10 14.30 4.74 NEUT COUNT 0.53-3.09 K/UL 0.30 0.26 1.40 1.81 0.35 0.17 0.47 0.24 0.70 0.74 0.82 0.56 0.98 1.54 0.24 0.08 1.67 2.22 0.50 0.33 LYMPH % 61.26-87.18 79.28 11.77 66.80 34.67 83.18 4.69 82.23 3.92 46.15 10.54 66.65 12.04 59.44 13.07 80.83 3.01 59.83 29.54 77.63 4.24 LYMPH COUNT 3.24-11.15 K/UL 1.63 0.27 4.09 1.95 3.37 0.68 3.55 0.96 1.12 0.12 2.22 0.57 1.66 0.95 1.95 0.55 2.01 0.40 2.57 0.68 MONO % 2.18-11.02 1.40 0.76 1.17 0.58 1.68 0.53 2.10 0.61 0.85 0.21 2.55 0.86 1.88 1.29 2.67 0.67 1.43 0.78 1.70 0.20 MONO COUNT 0.15-0.94 K/UL 0.03 0.02 0.07 0.04 0.07 0.02 0.10 0.05 0.02 0.01 0.08 0.03 0.08 0.12 0.06 0.01 0.05 0.03 0.06 0.02 EOS % 0.13-4.42 5.04 3.12 11.37 13.50 4.00 2.04 3.03 1.59 27.45 9.97 3.33 1.44 13.96 12.43 2.37 1.34 8.60 6.55 3.80 0.78 EOS COUNT 0.01-0.42 K/UL 0.11 0.09 0.85 1.07 0.17 0.09 0.14 0.10 0.65 0.02 0.11 0.05 0.30 0.22 0.05 0.03 0.48 0.53 0.12 0.04 BASO % 0.01-1.24 0.40 0.16 0.43 0.06 0.78 0.31 0.70 0.27 0.60 0.00 1.13 0.54 0.92 0.42 1.17 0.38 1.00 0.78 1.17 1.07 BASO COUNT 0.00-0.13 K/UL 0.01 0.00 0.03 0.01 0.03 0.00 0.03 0.01 0.02 0.01 0.04 0.02 0.02 0.01 0.03 0.01 0.03 0.01 0.04 0.02 LUC % % 0.66 0.18 1.40 0.82 2.20 0.84 1.80 1.14 1.00 0.57 2.10 0.93 1.38 0.83 3.20 1.37 1.80 1.91 1.40 0.69 LUC COUNT K/uL 0.01 0.01 0.09 0.06 0.09 0.03 0.08 0.05 0.03 0.01 0.07 0.04 0.03 0.02 0.07 0.03 0.05 0.03 0.05 0.03

Figure A3. Mean Hematology Results for treatment groups in TRAMP-C1 C57BL/6 tumored toxicity study