Engineering Translational Multi-Therapeutic Targeted Delivery Vehicles
For Disease Management
MASSCHES S9&
MASSACHUSETTS
by
INSTITt)TE
OF fEICHNOLOLQV
Stephen Winford Morton
JUN 2 2 2015
B.S. Chemical Engineering
North Carolina State University, 2010
LIBRARIES
SUBMITTED TO THE DEPARTMENT OF CHEMICAL ENGINEERING IN PARTIAL
FULFILLMENT OF THE REQUIREMENTS FOR THE DEGREE OF
DOCTOR OF SCIENCE IN CHEMICAL ENGINEERING
AT THE
MASSACHUSETTS INSITUTE OF TECHNOLOGY
MAY 21, 2015
w
201Ii
@ Stephen Winford Morton. All rights reserved. 2&1i
The author hereby grants to MIT permission to reproduce
and to distribute publicly paper and electronic
copies of this thesis document in whole or in part
in any medium now known or hereafter created.
Signature of author:
Signature redacted
V
Certified by:
Stephen W. Morton
Department of Chemical Engineering
May 21, 2015
Signature redacted____
Paula T. Hammond
David H. Koch Professor of Chemical Engineering
Thesis Supervisor
Signature redacted
Richard D. Braatz
Professor of Chemical Engineering
Chairman, Committee for Graduate Students
)
Accepted by:
1
Engineering Translational Multi-Therapeutic Targeted Delivery Vehicles
For Disease Management
by
Stephen Winford Morton
Submitted to the Department of Chemical Engineering
on May 21, 2015 in Partial Fulfillment of the
Requirements for the Degree of Doctor of Science in
Chemical Engineering
ABSTRACT
The complexity of growth, survival, and death signaling pathways in cancer continues to
motivate extensive investigations using systems biology approaches to better inform
treatments. Many of these drugs and drug combinations, however, are highly toxic, due to
lack of targeting and poor pharmacokinetics, and are not easily delivered together due to
solubility issues - requiring high levels of patient compliance over the course of treatment
whereby lower, less efficacious doses of therapies are administered during multiple
intravenous infusions. To address this problem, development of delivery systems that are
safe/less toxic, yet capable of housing and delivering higher, more efficacious doses with
sophisticated levels of control over the delivery, timing, and/or sequence of release, in
order to therapeutically re-wire these signaling pathways, is essential to translate these
drug combinations to the clinic. Work detailed herein addresses this challenge by
developing controlled approaches for engineering and manufacturing translational
nanoscale delivery systems for more efficacious, targeted treatment of cancer - specifically
using layer-by-layer nanoscale assembly and liposomal manufacture.
Layer-by-Layer (LbL) assembly of polyelectrolytes on solid substrates, including colloidal
systems, is a well characterized, tunable approach for generating functional thin films for a
variety of applications, including drug and gene delivery, tissue engineering, and bone
regeneration. On the basis of electrostatics, this assembly method allows for incorporation
of a broad range of materials; and due to its water-based synthesis, it allows for
incorporation of a broad range of therapeutics without significant alteration of biological
function. Using this approach, work described in Chapters 2-4 detail the utility of this
approach. In Chapter 2, we demonstrate the ability to coat materials atop drug-loaded
PLGA nanoparticle substrates that control drug release and nanoparticle pharmacokinetics
in a systemic environment. In Chapter 3, we incorporate ligand-functionalized materials
atop a drug-loaded liposomal core to promote tissue-level specificity for targeted
treatment of cancer. In Chapter 4, we further adapt this approach towards addressing the
challenge of scalably and reproducibly manufacturing LbL-functionalized nanoparticle
systems by collaborating with Joseph DeSimone at UNC-Chapel Hill. In this work, we
combine scalable methods, PRINT@ particle fabrication, and spray-LbL deposition, to
2
generate uniform and functional nanotechnologies that are built-to-order with precise
control over composition, size, shape, and surface functionality.
In Chapters 5-7, targeted liposomal technologies were synthesized to deliver drug
combinations with different solubilities. Liposomes, representing the first nanomedicine
systems approved and employed in the clinic, are well-characterized, simple and versatile
platforms for manufacture of functional and tunable drug-carrier systems. Further,
functionalization of these systems for tissue-level specificity is readily achieved by ligand
conjugation of a lipid that is subsequently inserted into the dual drug-loaded liposomal
vesicular membrane. Utilizing the hydrophobic and hydrophilic compartments of
liposomes, we achieved high loadings of both hydrophobic and hydrophilic therapies
within a singular tissue-targeted (e.g. folate) liposomal system. The resultant nature of
compartmentalization inside the vesicular structure promoted the desired sequence of
drug release for synergistic cancer therapy, improving tumor cell killing in vitro and
significant tumor shrinkage in in vivo murine triple-negative breast cancer and non-small
cell lung cancer models. In Chapters 6 and 7, we applied this approach for targeted
combination therapeutic delivery for additional drug therapies (e.g. cisplatin-quisinostat
(HDAC inhibitor) in a folate-targeted liposome; (+)-JQ1-temozolomide in a transferrintargeted liposome) and disease targets (non-small cell lung cancer; brain stem gliomas).
Using both LbL and liposomal manufacture, work detailed herein represents a broad skill
set for engineering and manufacturing translational targeted multi-therapeutic delivery
systems with high levels of control.
Thesis Supervisor: Paula T. Hammond
Title: David H. Koch Professor of Chemical Engineering
3
Thesis Supervisor
Paula T. Hammond, Ph.D.
David H. Koch (1962) Professor in Chemical Engineering
Department of Chemical Engineering
David H. Koch Institute for Integrative Caner Research
Massachusetts Institute of Technology
Thesis Committee
Sangeeta N. Bhatia, M.D., Ph.D.
Harvard-MIT Division of Health Sciences and Technology
Department of Electrical Engineering and Computer Science
David H. Koch Institute for Integrative Cancer Research
Massachusetts Institute of Technology
Robert S. Langer, Sc.D.
David H. Koch (1962) Institute Professor
Department of Chemical Engineering
Harvard-MIT Division of Health Sciences and Technology
David H. Koch Institute for Integrative Cancer Research
Massachusetts Institute of Technology
4
In loving memory of my hero and latefather,
Kenneth Reid Morton;
Formy cherishedgrandparents,
Tim Elmo Morton, Shirley Rose Morton, Giles Winford Clayton, Mary Pegram Clayton;
For my loving chocolate (Labrador)inspirations,
Sable and Sadie Mae,
Formy supportive and kind older brotherand bestfriend,
ChristopherReid Morton;
Andfor the best mom anyone could askfor,
Jill Clayton Morton.
5
Acknowledgements
This section is one of the most difficult to write because it is impossible to include everyone
in the exhaustive list of people that made the work done in my Ph.D. possible.
First of all, I want to thank my lord and savior, Jesus Christ. "I can do all things through Him
who strengthens me." Philippians 4:13
Throughout this process, family has been my rock. My parents have always been my
biggest supporters. Their sacrifice for my growth and success from a boy to a young man
deserves more thanks than I can provide in this lifetime. Believe me... growing up, MIT
would have meant "Mission Impossible: Trouble" for my years as a kid. I aspired to be a
teenage mutant ninja turtle (not an engineer...), so you can imagine what sort of mischief
this inspired. To be where I am at today is a true testament to the loving, caring, brilliant
parents they are. From ball fields to the classroom, they were always there to support me in
any way possible. My dad called me the "Rock" growing up (maybe because I was hardheaded and didn't listen), but it was my parents who were the true Rock's in my life. I lost
my father physically in December 2013, but I know he is and always will be with me
spiritually through my journey in life. To my parents, Jill and the late Reid Morton - I love
(and miss you, pops) you dearly.
I also want to thank my brother, Chris Morton, who has also been my best friend
throughout my life. He is one of the most kind-hearted and giving individuals I know. He is
also, more importantly, as excited to talk and play sports as I am - so special thanks to my
big bro for the passionate sports banter on Tiger Woods and Braves/Sox baseball as well as
many an hour on the links.
I also would like to thank my grandparents: Giles and Mary Clayton, Tim and Shirley
Morton. I was fortunate to have three sets of parents growing up - spoiled by each as much
as the other. I always looked forward to sick days at the grandparents where I was healed
with homemade caramel cake and cinnamon toast while watching the Jungle Book or
Sound of Music. They were so caring and giving - memories I will always cherish.
Of course, I cannot forget to mention the two sweetest, chocolate furry labs, Sable and Sadie
Mae, that always energized me with each visit home.
So many teachers have motivated me through the years, from Mrs. Saunders and Mills in 6th
grade to Kay Sandberg and Lisa Bullard at N.C. State. All of them have been advisers,
friends, colleagues, and mother figures to me - challenging and inspiring me intellectually,
developing me professionally, and shaping me as a young man. They set the standard for
education - motivating so many each year yet maintaining personal relationships with
each, and I am honored to keep them as lifelong friends in my continued journey.
There are too many friends to begin to enumerate them here, but specifically during my
Ph.D. I would like to thank Ben Renner and Kevin Shopsowitz - two of my best and most
6
genuine friends, brothers by all but blood. Lots of Braves baseball talk and good times as
roommates for 5 years with the "Sage" Ben Renner, and on the links and hanging out with
Kevin usually around a "Morty-Shops" feast consisting of Shops' homemade pizza dough
topped with Morty's Carolina smoked pork shoulder.
I also want to thank my fitness family - Elena Byrne, Melanie Smithers, Moses Blumenstiel,
Stephanie Gayle, Vicky Palay, Izzy Weigner-Lodahl, Stacey Hutchins, Adrian Barnett,
Richard Hamilton, Melissa Hyland, and Cindy Carter. Science has no shortage of its
frustrations and taking it out in kickboxing, combat, pump, punishment, and jam with these
awesome friends has been tremendous support to keep me motivated throughout my Ph.D.
I also want to thank my friends and colleagues in the Hammond lab, specifically Zhiyong
Poon (P-dizzle), Nisarg Shah (Nate-dawg), Erik Dreaden (Dr. Dre), Jason Deng (Master
Jedi), Mohi Quadir (Dr. Q), Kevin Shopsowitz (Shops), Liz Welch, Brian Aitken, Peter
DeMuth, Santi Correa (triple-threat), Noemie Dorval (Noeminem), Nasim (the Dream)
Hyder, Chibueze (the Breeze) Amanchukwu, Jouha Min, and Steven Castleberry
(Huckleberry). It has been a great lab to work in and a fun ride these past 5 years. I also feel
obliged to thank the Langer lab as an honorary lab member, attending last year's Langer lab
outing to the Cape as well as helping their team win the KF summer league. Cheers to Mark
Tibbitt, Owen Fenton, Ben Tang, Katy Luly, and so many others for more good times.
Other collaborators include Joseph DeSimone, Kevin Herlihy, Kevin Reuter, Tammy Shen,
Kevin Chu, Charles Bowerman - all from UNC-Chapel Hill and who were gracious hosts for
3 months while I worked in this exciting and brilliant lab; also Michael Yaffe, Michael Lee,
Eilaf Ahmed, Tim Swager, Katharina Ribbeck, and Jeremiah Johnson from MIT - all of which
deserve a heartfelt thanks for their help and support as colleagues and friends in this work.
Ed Lawson, Peter Palmer, and their team at Janssen Pharmaceuticals were also
instrumental to the success of this work as well as developing me professionally.
Above all collaborators at MIT, I want to extend a heartfelt thanks to my advisor and
mentor, Paula Hammond. She is a caring, kind, understanding individual with a great heart
and brilliant mind. I am truly indebted to her sacrifice and support as I grew professionally
in her lab these past 4 years. I am blessed to have worked with a leading expert in the field
of nanotechnology and drug delivery. Paula has been my rock at MIT, supporting me in
many endeavors from traveling to the Chapel Hill for exciting work as well attending many
national and international conferences to share the great work we have accomplished
together. I will always cherish my time in her lab and cannot thank her enough for her
leadership and friendship during my time at MIT.
I also want to thank my brilliant thesis committee members: Sangeeta Bhatia and Robert
Langer, who were generous with their time to advise and support my work. It was difficult
to get my entire committee, including Paula, in the same room with their insanely busy
schedules, but when I was able to catch them all together I was fortunate to gain insight
from 3 inspirational world experts in the field of nanotechnology and drug delivery - who
are equally approachable and honest in their opinions and advice on work and growth as a
professional (or where the cool new spots in town are).
7
Finally, I would like to thank the Koch Institute for Integrative Cancer Research at MIT (and
MIT on the whole) for being such an incredible and stimulating place to work and
collaborate. I could not ask for a better environment to learn, engage, and take action in the
fight against cancer.
Heartfelt thanks,
Stephen Winford Morton
Massachusetts Institute of Technology
May 21, 2015
8
I, Stephen Winford Morton, would like to thank the generous funding sources that made
possible all the work described herein: an NSF graduate research fellowship, Janssen
Pharmaceutical TRANSCEND, P30 CA14051 (NCI), 5 U54-CA151884-02 and U54CA151652 (CCNE), R01-AG029601 (NIA), and R01-EB010246 (NIBIB). U54-CA112967,
R01-ES015339, and R21-ES020466, and a Breast Cancer Alliance Exceptional Project
Grant. I would also like to thank assistance from the Koch Institute Swanson Biotechnology
Center, specifically the Hope Babette Tang (1983) Histology Facility (Kathleen Cormier),
the Applied Therapeutics and Whole Animal Imaging Facility (Scott Malstrom), the
microscopy core (Eliza Vasile) and the Flow Cytometry Facility (Glenn Paradis, Michael
Jennings); also, Bill DiNitale and Institute for Soldier Nanotechnologies, Center for
Materials Science and Engineering (CMSE) at MIT, and Chapel Hill Analytical and
Nanofabrication Laboratory (CHANL) for assistance and facilities that supported part of
this work. I would also like to acknowledge UNC-Chapel Hill and Joseph DeSimone's lab in
Chemistry, as well as Liquidia Technologies for graciously allowing me to work in their labs
for an extended portion of my Ph.D., as well as generously supporting part of this work.
9
Table of Contents
A ckn ow ledgem ents..............................................................................................................................6
T able of Con ten ts ...............................................................................................................................
10
Figu re s ...................................................................................................................................................
11
Chapter 1. Introduction ...............................................................................................................
11
Chapter 2. Architecture, Biological Performance of LbL Nanoparticles......................27
Chapter 3. Osteotropic therapy via targeted LbL nanoparticles....................................46
Chapter 4. Scalable manufacture of built-to-order nanomedicine: spray-LbL on
PRINT@ nanoparticles ................................................................................................................
70
Chapter 5. Liposomal compartmentalization of a hydrophobic-hydrophilic
combination chemotherapy regimen (erlotinib, doxorubicin) for staged, targeted
delivery against triple-negative breast cancer and non-small cell lung cancer...........89
Chapter 6. Liposomal delivery of quisinostat and cisplatin for treatment of non-small
cell lu ng can cer................................................................................................................................
117
Chapter 7. Liposomal delivery of JQ1 and temozolamide systemically across the
blood-brain barrier for treatment of brain-stem gliomas................................................
134
Chapter 8. Framing the impact of this work and future directions...............................
154
Appendix A. Architecture, Biological Performance of LbL Nanoparticles................... 157
Appendix B. Osteotropic therapy via LbL Nanoparticles...................................................
165
Appendix C. Scalable Manufacture of Built-to-order Nanomedicine: Spray-LbL on
PRINT@ Nanoparticles..............................................................................................................
173
Appendix D. Liposomal compartmentalization of a hydrophobic-hydrophilic
combination chemotherapy regimen (erlotinib, doxorubicin) for staged, targeted
delivery against triple-negative breast cancer and non-small cell lung cancer........ 183
10
Figures
Chapter 1
Figure 1.1. Therapeutic window following drug administration......................................................................................
........
Figure 1.2. Passive (EPR-based) and active (receptor-ligandspecificity) targeting strategies.81...
20
21
Chapter 2
Figure 2.1. Synthesis and characterization of a Layer-by-Layer PLGA nanoparticle (LbL NP) system with
39
cardiogreen encapsulatedfor m ulti-color im aging............................................................................................................................
Figure 2.2. In vitro release of uncoated and coated cargo-loaded(cardiogreen,doxorubicin) PLGA LbL NPs..... 40
Figure 2.3. Stability assessments of different LbL PLGA nanoparticlearchitecturesassessed by drug (CG820) and
41
particle (PLL700) tracking via in vivo whole animalfluorescentimaging............................................................................
Figure 2.4. Pharmacokinetic data for the PLGA LbL nanocarriers(PLL 7oo, CG82o), as a function of terminal layer
42
and the encapsulated drug, CCG 82 0. .............................................................................................................................................................
Chapter 3
58
Figure 3.1. Achieving bone tissue level specificity of LbL coated nanoparticles..............................................................
Figure 3.2. in vitro assessment of PAA-Alendronate LbL-targeted QD NPs incubated with 143B osteosarcoma
59
c e lls...........................................................................................................................................................................................................................
60
Figure 3.3. in vivo evaluation of PAA-Alendron ate LbL -targetedQD 8 oo core NPs...........................................................
61
Figure 3.4. in vitro evaluation of PAA-Alendronate LbL-targeted liposomal NPs in 143B cells.............
Figure 3.5. Biological evaluation of PAA-Alendronate LbL -targetedempty liposomal NPs....................................... 63
Figure 3.6. 143B xenograft tumor remediation following treatment with doxorubicin-loaded liposomal
65
f o rm u la tio n s.........................................................................................................................................................................................................
Figure 3.7. Representative histological sections of 143B tumors following treatment with doxorubicin-loaded
66
lip osom alfo rm u latio ns6....................................................................................................................................................................................
Chapter 4
80
Schem e 4.1. Spray-LbL on PRIN T® nanoparticles............................................................................................................................
Figure 4.1. Contactangle and particleiridescence characterizationpre- and post-functionalization................ 81
82
Figure 4.2. AFM characterization(height) of pre- and post-functionalizedNP arrays................................................
83
Figure 4.3. SEM characterizationof particlespre- and post-Spray-LbL functionalization.......................................
84
thickness.........
controlfilm
Figure 4.4. TEM characterizationof particlespre- and post-LbLfunctionalization
uptake
tracking
NP
shape
on
cell
and
PRINT@
LbL
coatings
of
variable
spray
vitro
assessment
Figure 4.5. in
85
w ith PL L cys s.. .........................................................................................................................................................................................................
87
Table 4.1. Nanoparticlephysicochemical characterization..........................................................................................................
Chapter 5
Figure 5.1. Characterizationof combination therapeutic-loadedliposomal system.......................................................106
Figure 5.2. Evaluation of doxorubicin-erlotinibdual drug-loaded liposomal system in vitro.....................................107
Figure 5.3. Folate decoration of combination doxorubicin-erlotinib therapeutic-loaded liposomes for targeted
108
delivery to TNBC (B T-20) and NSCLC (A549) cells. ..........................................................................................................................
Figure 5.4. Biologicalperformanceoffolate-targeted liposomal system in vivo...............................................................109
Figure 5.5. Effect of dual drug-loaded or single drug-loaded, folate-targeted liposomes on A549 and BT-20
1 10
tum o r siz e............................................................................................................................................................................................................
Figure 5.6. Developing the RTK inhibitor-cytotoxicagent combination liposomalsystems as a platform for dual1 11
d rug de liv e ry . .....................................................................................................................................................................................................
Table 5.1. Dynamic light scattering, polydispersity index, and potential measurements for the doxorubicin112
erlotinib dual-drug and single-drug uncoated and coated (FP) liposomal systems. .........................................................
Table 5.2. Dynamic light scattering and polydispersity index measurements for the various dual-drug and
single-drug coated (FP) liposomalformulations- doxorubicin + afatinib/erlotniib/gefitinib/lapatinib...............113
11
Table 5.3. Dynamic light scattering, polydispersity index, and zeta-potential measurements for the various
multi-drug coated (FP) liposomalformulations- cisplatin + afatinib/erlotinib/gefitinib/lapatinib.......................114
Chapter 6
Figure 6.1. Characaterizationof cisplatin-quisinostatdual-drug coated (FP) liposomal system..............................124
Figure 6.2. In vitro release and efficacy of cisplatin-quisinostatdual-drugcoated (FP) liposomal system...........125
Figure 6.3. Pharmacokineticsof FP coated empty liposomal carrier.....................................................................................
126
Figure 6.4. Biodistributionof FP coated liposomal carrierin A549 xenograft bearing ncr nude mice...................127
Figure 6.5. Gross toxicity of free and encapsulated quisinostat formulations following intraperitoneal
administrationin immune-proficient BALB/c mice. ........................................................................................................................
128
Figure 6.6. Gross toxicity evaluated as body weight change in BALB/c mice treated intraperitoneallywith
liposomalformulationsand correspondingfree drug controls...................................................................................................129
Figure 6.7. In vivo efficacy of single drug and multi-drug cisplatin-quisinostatloaded liposomal systems in A549
hind flank xenograft bearing ncr nude m ice1........................................................................................................................................130
Figure 6.8. Cryo-preservationof dual-drug cisplatin-quisinostatFP coated liposomal construct............................131
Chapter 7
Figure 7.1. Liposomal constructfor dual-drug (+)-JQ1-temozolamide targeted carriers.............................................143
Figure 7.2. Intravitalmultiphoton imaging of nanoparticlescrossing the blood-brain barrierin mice................144
Figure 7.3. Representative distributiondata of each fabricated nanoparticlesystem in non-tumor bearing mice.
..................................................................................................................................................................................................................................
1 45
Figure 7.4. Biodistributiqnof targetedliposomal carriersin non-tumor bearing mice.................................................146
Figure 7.5. Histological sections of ncr nude brains 24h post-systemic administration of functionalized
n anop article sy stem s............................................. ...................................................... ...............................................................................
147
Fig 7.6. Mouse m odels of hum an g lioma...............................................................................................................................................148
Human glioma cells are implanted into the brains of mice and visualized using bioluminescent imaging (A) as
well as magnetic resonance imaging (MRI) (B, C)............................................................................................................................148
Figure 7.7. Biodistribution data of transferrin-functionalizedliposomal carriers in M059K brain-implanted
tum o rs...................................................................................................................................................................................................................1
49
Figure 7.8. Kaplan-Meiercurvefollowing treatment of M059K brain tumor implants with single drug and dualdrug temozolamide-(+)-JQ1 transferrin-targeted liposomal carriers......................................................................................150
Figure 7.9. Small animal irradiationplatform for combination radiation-systemicchemotherapy........................151
Appendix A
A.1. Representativescanning electron micrograph (SEM) of LbL PLGA NPs (HA -terminal)....................................157
A.2. CG drug stability - biodistributionof LbL PLGA NPs in ncr nude mice............................................................................158
A.3. NP stability - biodistributionof LbL PLGA NPs.......................................................................................................................159
A.4. Biodistributiondata 30 minutes after systemic administrationin ncr nude mice....................................................160
A.5. Dose-dependent co-injection offree polymer with corresponding terminated 6L LbL PLGA NPformulations.
.................................................................................................................................................................................................................................
1 61
A.6. Blood circulationdatafor LbL PLGA NPs and encapsulatedcardiogreen.....................................................................162
A.7. Variation of core template by modulating L:G ratio in PLGA and effect on biodistribution in ncr nude mice.
.......................... .................................................................
....................................................................................................................................
1 63
A.8. Opsonin (IgG) and subsequent uptake by macrophages in vitro as a function of terminal layer on LbL PS
b e ads......................................................................................................................................................................................................................1
64
Appendix B
B.1. Sy nthesis of PAA -A lendronate............................................................................................................................................................
B.2. OsteotropicPAA-Alendronate LbL QD characterization.......................................................................................................
B.3. Variation of number of layers and impact on biological performance of PAA-Alendronate targetedLbL
.................................................................................................................................
. . ..........................................................................................
B.4. Biodistributionof uncoated and PAA untargetedQD NPs. ............................................................................................
12
165
166
NPs.
167
168
B.5. Biodistribution and tumor targeting of PAA-Alendronate targeted QD NP tracking the core and LbL shell
sim u ltan eou sly . .................................................................................................................................................................................................
1 69
B.6. Physicochemical characterizationfollowing functionalization of empty and doxorubicin-loaded liposomal
carriers.................................................................................................................................................................................................................1
70
B. 7. Biological performance of PAA-Alendronate targetedliposomal carriers...................................................................
171
B.8. microCT scans following treatment of 132B xenografts with PAA-Alendronate targeted liposomal carriers.
.....................................................................................
................................................................................
......................................................
1 72
Appendix C
C1. AFM of PRINT@ NPs pre- and post-sprayLIbL. ..........................................................................................................................
173
C.2. Varying crosslinking conditions, incubation period - characterizingcrosslinking of PVA via contact angle
m easurements of PRINT@ N P arrays1. ....................................................................................................................................................
174
C.3. AFM amplitude data correspondingto Figure 4.3. ..................................................................................................................
175
C.4. Dynamic Light Scattering (DLS) Intensity distribution histograms displayed throughout Spray-LbL
deposition on PRINT@ PLGA 2oox 2oonm nanoparticles.........................................................................................................................176
C.5. Controllingfilm thickness of spray-LbL on PRINT ...............................................................................................................
177
C.6. Tracking deposition with Spray-LbL on PRINT@ using AFM, fluorescence imaging. .............................................. 178
C.7. Varying polycationic component in film - Chitosan - via spray-LbL on PRINT®..................................................... 179
C.8. Spray-LbL on PRIN T@ PLGA8ox3 2onm NPs......................................................................................................................................180
C.9. Spray-LbL on PRIN T@ PLGA2oox2oonm N Ps. ...................................................................................................................................
181
C.10. Uncoated, pre-crosslinked TEMs of PRINT@ NPs. .................................................................................................................
182
Appendix D
D.1. Tumor growth following BT-20 and A549 xenograft seeding at day 0 (n = 5)...........................................................183
D.2. In vitro release experimentation - release of inhibitors (afatinib, erlotinib,gefitinib, lapatinib) and cytotoxic
(doxorubicin)from dual-drug coated (FP) liposomalformulations.........................................................................................
184
D.3. Cleaved Caspase-8, phosphorylated ERK (pERK), and phosphorylated H2AX (pH2AX) western blot
quantificationas function of time and treatmentin both NSCLC and TNBC cancer cells...............................................185
D.4. Characterizationof non-functionalized multi-drug and single-drug loaded liposomes (without folate-PEG).
..... ..........................................................................
.... ......... . . . .......................................................................................................................
1 86
D.5. Biodistribution of Cy5.5 labeled blank FP coated liposomes administered to tumor-bearingncr nude mice.
............................................ ............. ......................................................................................................................................................................
1 87
D.6. Tumor luminescence data temporally resolved following treatment of A549 xenograft bearing ncr nude
mice with single and dual drug doxorubicin-erlotinibFP coated liposomal carriers....................................................... 188
D.7. Tumor luminescence data temporally resolved following treatment of BT-20 xenograft bearing ncr nude
mice with single and dual drug doxorubicin-erlotinib FP coated liposomal carriers....................................................... 189
D.8. Release of doxorubicin visualized by confocal microscopyfrom FP coated liposomal carriers..........................190
D.9. Co-releasing in vivo control - timed administrationof single drug doxorubicin liposome with co-treatment
191
of cy clodextrin-erlotinib com p lex. ............................................................................................................................................................
13
Chapter 1. Introduction
Often in clinical settings, drugs are available that effectively treat a particular
diseased state; however, the issue remains one of delivery. Free drug administration at the
site of action is not always realizable; and in many cases, this form of treatment suffers
from poor drug solubility, limited permeability, acute toxicity, and biological multi-drug
resistance mechanisms that flush the cell of free drug influx via drug-efflux transmembrane
pumps] Further, diseased states are often only accessible via the blood supply, in lieu of
invasive surgery, and may exist as metastases in different compartments in the body.
Treatment, therefore, requires a systemically administrable formulation that can access
these locations; however, introduction of free drug to the systemic environment results in
undesirable pharmacokinetics, biodistribution, and lack of selectivity for particular tissues,
yielding tissue damage or necrosis at off-target locations.
Drug delivery carriers on the nanoscale, ranging in approximate size from 10200nm, are promising candidates for systemic delivery of therapeutic agents.
These
vehicles have been intelligently designed in many different forms, including polymer-drug
conjugates, micelles, dendrimers, hydrogels, vesicles (liposomes, polymerosomes), and
nanoparticles (polymeric, inorganic, magnetic).[2,3] Nanocarriers of these types allow for
incorporation of a variety of drug classes, including both hydrophobic and hydrophilic
small molecules, proteins, and nucleic acids, via non-covalent encapsulation or covalent
modification of the carrier material, and can enhance uptake into the tumor, increase
specificity of targeting to cancer cells, and protect and maintain drug activity while
mitigating off-target toxicity to healthy tissue[4 8 ] In this way, they are able to load high
concentrations of drug cargo and effectively stabilize them for delivery. This allows for
enhanced bioavailability of the drug in circulation, such that the therapeutic window for
access to an efficacious dose of drug is maintained for extended periods of time,
overcoming the issues with multiple bolus injections of free drug formulations oscillating
in and out of this regime, as shown in Figure 1.1.
While nanomaterials offer significant promise in stabilizing drugs for delivery,
several of which have been successfully applied in the clinic[91, there remain a number of
14
challenges to engineering therapeutic nanoparticles for translation. These challenges
include[10. 113:
(1) robust methods for accurate characterization of nanoparticle size, shape, and
composition;
(2) scalable approaches for producing nanomedicines with optimized bioavailability
and excretion profiles;
(3) particle engineering for maintaining low levels of nonspecific cytotoxicity and
sufficient stability during storage;
(4) optimization of surface chemistries for maximum targeted delivery and
minimum nonspecific adsorption;
(5) practical methods for quantifying ligand density and distributions on multivalent
nanocarriers;
(6) the design of multifunctional nanomedicines for novel combination therapies
with supportable levels of bioaccumulation.
In particular for systemic application,
a characteristic bolus drug release upon
administration to physiological conditions in circulation is problematic for non-covalently
encapsulated payloads. Alternatively, polymer-drug conjugates are generally of the size,
<10 nm, that are efficiently cleared via renal filtration, and those that do reach the target
location are often unable to efficiently release their cargo for therapy because of the design
constraints for stability in circulation versus triggered release at the tumor.[1 ] Further,
carriers must avoid rapid immune-mediated clearance via serum proteins which decorate
foreign materials for phagocytosis and clearance by macrophages in the reticuloendothelial
system, specifically Kupffer cells in the liver and red pulp cells in the spleen. Filtering on
the basis of size is also employed by the fenestrated endothelium of the liver and sinusoidal
endothelium of the spleen and is responsible for the clearance of larger particles (generally
>200nm), while renal
filtration clears
smaller particles (generally <20nm) from
circulation.['11 1 Towards this end, particles in the range of 20-200nm decorated with a
hydrophilic, neutral (i.e. poly(ethylene glycol)) or negatively-charged (i.e. hyaluronic acid)
corona have been extensively investigated as delivery vehicles.[2
8
11, 121
These have been
shown to exhibit enhanced systemic stability with long-circulating half-lives and are also
efficient at extravasating into tumor vasculature and accumulating on the basis of the
15
enhanced permeation and retention (EPR) effect. A passive targeting strategy, EPR-based
targeting relies on the leaky, perforated vasculature that results from rapid proliferation of
cancerous cells beyond their ability to generate an adequate blood supply, as shown in
Figure 1.2. Further, the lymphatic drainage is defective such that particles are able to
accumulate in the tumor region without being effectively removed. Active targeting can
enhance delivery via cancer cell ligand-receptor interactions, such that delivery is
increased to target tissue. Targeted delivery allows for lower doses of drugs to be
administered, thereby mitigating adverse side effects while increasing therapeutic
potency1] Furthermore, other nanoparticle parameters such as shape and elasticity are
critical to the performance of a nanoparticle system for systemic delivery applications.[ 14 20 ]
After designing a carrier that is both serum-stable and able to accumulate in the
tumor region, the challenge of cellular entry with the drug-loaded platform is the next
barrier to delivery, particularly for payloads, such as nucleic acids, that require cellular
entry but have no means of facilitating internalization independently and are readily
degraded without a delivery material host.[ 211 Positively-charged materials are extremely
effective at interacting with the negatively-charged glycocalyx on the surface of cell
membranes and trafficking into the cell by fluidizing the bilayer, which often also yields a
cytotoxic effect. However, positively-charged materials are not stable in circulation and are
rapidly cleared and often cytotoxic, therefore not a viable systemic delivery carrier. For
serum-stable particles, which are neutral or negatively-charged and with or without the
complement of a targeting ligand, these particles rely on the process of pinocytosis
("cellular gulps") or receptor-mediated endocytosis, which can be quite effective at
internalizing carriers with size-dependence. The latter introduces the challenge of
endosomal escape for the drug to reach its therapeutic target. It is believed that a protonsponge effect is necessary to mediate endosomal rupture, requiring a positively-charged
material to introduce a proton gradient across the endosome such that it bursts and
releases the drug; however, the acidic pH environment may degrade or inactivate the drug
of interest before it is able to function.[21
22]
The paradoxical nature of serum-stability and cellular entry has generated interest
in engineering "smart" particles that possess serum-stable properties that are unmasked in
the tumor microenvironment. This complementary approach to traditional passively16
accumulating or actively-targeted carriers exploits the nature of tumor progression. For
this purpose, designs hone in on the acidity of the localized tumor region, which results
from the poorly oxygenated
(hypoxic) regions at the tumor due to inadequate
vascularization. Stealth particles that have a cue-sensitive (i.e. pH-sheddable) cloak are an
emerging technology to enhance delivery for this purpose.[2, 2 1, 23-26]
The surface chemistry of nanoparticles plays a major role in determining their
biological properties. However, many of the current methodologies used to modify the
surfaces of nanoparticles are inefficient and labor intensive. Layer-by-layer (LbL)
polyelectrolyte deposition is an attractive process for incorporating functionality onto the
surfaces of nanocarriers to provide improved stability, enhanced cellular uptake, and
targeting capabilities.[ 24 , 27-291 Because small molecule drugs and biologics such as nucleic
acids can be included in the nanolayers, it also affords the ability to incorporate additional
therapeutics that act as synergistic drug combinations[ 3 ,
31]
on a single nanoscopic
platform and in a fashion such that the release of each can be programmed.[ 32,331 Layer-byLayer (LbL) assembly[ 34] of polyelectrolytes on solid substrates is a well-characterized,
tunable approach[ 35 ] for generating functional thin films on the basis of complementary
interactions, such as electrostatics or hydrogen bonding, for a variety of applications,
including drug and gene delivery[ 28,
36, 37]
and tissue engineering.[ 3 8,
39]
For these
applications, this technique is advantageous in that the incorporation of a broad range of
therapeutics (e.g., proteins, nucleic acids, and small-molecule drugs) and is carried out
under physiological conditions, non-covalently, thereby preserving the native properties of
the cargo while stabilizing it for delivery. Furthermore, surface-limited, sequential
adsorption affords nano-scale precision over the composition of each layer; these
advantages, when combined with the ability to incorporate a broad range of therapeutics
and materials with diverse functionalities, greatly facilitate the development of drug
delivery platforms with sophisticated control over the spatial and temporal release of film
constituents, allowing for the construction of a customizable therapeutic delivery
platform [40-43].
While LbL has been successfully implemented on macroscopic surfaces with great
promise for many biomedical applications for many years[ 28,43-531, it has only been more
recently established as a viable tool to coat nanoscale materials systems[ 24, 27,28,43,54-75] to
17
achieve particles that can be sustained for systemic delivery. The ability to generate
conformal LbL coatings on nanoscale templates opened new and exciting opportunities for
nanoparticle technologies with applications in drug and nucleic acid delivery[
76-8 7]-
however, the translation of LbL-based nanocarriers for systemic applications required
further examination of how LbL film architectures can be manipulated to overcome
systemic delivery barriers and to sustain drug delivery in such complex biological settings..
Previously, we investigated the impact of different nanoparticle-bound film architectures
on physiological stability, elucidating key control variables necessary to generate a serumstable particle, as well as the effect of terminal polymer layers on subsequent
biodistribution of the nanocarriers.[ 271 LbL architectures that act as highly effective
stabilizing layers were demonstrated to impart long circulation times and EPR-based
passive targeting to the resulting systems.[ 881 We have also developed dynamic LbL layers
that 'shed' at the lower pH of hypoxic tumor microenvironments and subsequently reveal a
positively charged surface that rapidly mediates uptake by tumor cells.[ 2 4] Similar
demonstrations of dynamic LbL particle design for tumor targeting,[25 , 7 2 , 741 along with the
ability of LbL nanofilms to deliver multiple constituents for combination therapies[
71, 891
and to modulate drug and carrier biodistribution, lay a framework for continued
development of LbL-based nanoparticles towards scalable, built-to-order systems[ 75 ] that
are key for unlocking new biomedical opportunities for these platforms - detailed in the
subsequent Chapters 2 - 4.
Further work detailed herein (Chapters 5 - 7) focuses on the use of liposomal
technologies to deliver combinations of small molecule therapeutics with vastly different
solubilities (hydrophobic-hydrophilic small molecules) that are readily compartmentalized
within the liposomal structure[
90, 911 -
a limitation discovered experimentally to the rapid
and scalable deployment of LbL nanoparticles for this purpose (issues with achieving
therapeutic loadings of both drugs). Combination treatments have been explored for more
intransigent tumor types such as triple negative breast cancer (TNBC), but with little
significant gain thus far reported.[ 9 2 -94 ] Similarly, non-small cell lung cancer (NSCLC) is
often non-responsive to single chemotherapy treatments, and is prone to recurrence.
Although numerous drug combinations have been examined in oncology in the past, in
almost all cases, these drugs were administered simultaneously in simple combinations,
18
without manipulation of drug presentation, scheduling and duration of dose; recent
work[ 95 -97] further supports that many combinations of previously screened existing drugs
may be highly potent if combined in the correct order and relative dose. To capitalize on
the importance of order and timing for multi-drug therapies (Chapter 5) in complex
cellular systems, it is key to devise a means of a) enabling bioavailability of the drugs,
including hydrophobic small molecule inhibitors; b) targeting each of the drugs with equal
or similar efficacy to the tumor; and c) releasing the drug in a manner that enables
temporally-controlled release of each agent.
To achieve this within a single nanoparticle platform, liposomes, which have been
investigated extensively and approved for several cancer treatments,[ 98-10 1] were employed.
Previous research
has illustrated the versatility
of these systems, ranging from
demonstrated pH responsive behavior to the development of theranostic systems[ 1 02
103].
Furthermore, these systems were particularly attractive due to their inherent ability to
compartmentalize vastly disparate drugs to enable encapsulation and delivery of multiple
drug types without the need for two or more independent drug formulations that may
undergo differences in pharmacokinetics and uptake. Recent studies have shown that only
25% of cells contain plasmids delivered from two co-administered nanoparticles[
104],
indicating that uptake of both drugs by the same cell cannot be presumed. Furthermore,
the injection of each nanoparticle drug formulation at different times to achieve staging can
lead to increased cost and more complex dosing schedules that are inconvenient for the
patient. A more critical concern in the use of two separate nanoparticle formulations is the
potential change in vascularity and tumor tissue morphology due to the first drug, which
may impede delivery of the second nanoparticle system to the same tumor cells. Finally,
increased treatments mean higher opportunity for systemic toxicity from the nanocarrier.
For this reason, multi-therapeutic liposomal formulations were investigated as the
nanoparticle basis for work described in Chapters 5 - 7.
19
Therapeutic Window
J
Max tolerated dose
0
C
Min therapeutic dose
Time
Figure 1.1. Therapeutic window following drug administration.
Blue = ideal drug nanocarrierformulation. Red = repeatedbolus injection (arrow) offree drug formulations.
20
*
/1
bpgan
Mrug
I,
-
Nonopulclo
8
Figure1.2. Passive (EPR-based) and active (receptor-ligandspecificity) targetingstrategies. ]
21
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Chapter 2. Architecture,
Nanoparticles.
Biological
Performance
of
LbL
This chapter is in part adapted from:
Stephen W. Morton, Zhiyong Poon, Paula T. Hammond. Biomaterials 34(21): 53285335, 2013.
Introduction
As LbL nanoparticle technology moves towards clinical translation, it is essential to
establish these systems as drug-stabilizing carriers; the focus of this work is on small
molecule therapeutics. The challenges and approaches to small molecule delivery are well
documented 105 ; characteristic bolus release upon injection is one of the most significant
limitations to efficacious treatment for this class of therapeutics.[
1051
Once it is freed from
the carrier, drug is often rapidly cleared, reducing its plasma concentration, and producing
significant off-target cytotoxicity. The ability to control small molecule release and drug
distribution at a cellular and tissue level is, therefore, an active area of investigation.
Little information regarding the real-time fate of small molecule release from
delivery platforms in vivo is understood following systemic administration. The current
work seeks to demonstrate LbL nanoparticle material systems as viable candidates for
small molecule drug delivery, as well as to develop a robust, systematic approach for
screening a library of materials for incorporation in these engineered systems. LbL films
have been previously demonstrated as effective small molecule delivery agents,[3"]
33
with
an enhanced level of control over release. This work examines LbL architectures as dually
functional films on nanoparticle surfaces - acting as membranes that control rate of drug
release from the nanoparticle core and thus impact pharmacokinetics of the drug, as well
as the hydrated, protein-resistive coatings that modulate blood circulation half-life and
biodistribution for both the drug and carrier. The current study probes a series of
nanoparticle architectures assembled on a biodegradable poly(lactic-co-glycolic acid)
(PLGA) drug-loaded nanoparticle core. Using in vivo imaging for simultaneous drug and
particle fluorescence tracking in vivo, this work provides a framework for assessment of
LbL nanoparticles as small molecule delivery agents. This technique for live animal imaging
is
a
convenient
and
robust means
of probing
27
delivery
pharmacokinetics
and
biodistribution. It affords the capability of drug and particle monitoring following
administration to a single animal, allowing for high throughput in vivo screening of delivery
systems. The current study demonstrates this capacity, evaluating LbL nanoparticles using
multiple indicators of stability and performance, in an effort to advance the technology
towards therapeutic settings.
Materials and Methods
All chemicals were purchased from Sigma-Aldrich, except for hyaluronic acid (Lifecore
Biomedical) and doxorubicin-HCl (LC Laboratories). Release dialysis float-a-lyzers were
purchased from Spectrum Laboratories. NCR nude and BALB/C female mice were
furnished by Taconic.
Nanoparticle Synthesis. PLGA particles were prepared under aseptic conditions via an
emulsification-diffusion method with slight modifications to a previously established
protocol.[1 0 61 Briefly, 50 mg PLGA was dissolved in 3 mL acetone with 0.5 mL of 10mg/mL
drug (cardiogreen/doxorubicin) in methanol. This solution was subsequently added to
10mL 10% BSA in PBS and sonicated at RT. The emulsion was stirred overnight at 1000
rpm. The resulting particle suspension was purified via centrifugation (15,000 g, 30 min)
prior to LI. assmh1v
L.L assemh1V
wAs conducted by introducing
~no/,O% of the
recovered drug-loaded nanoparticles in excess polyelectrolyte solution (5 mg/mL for PLL,
DXS; 1mg/mL for HA, Alg) at pH 7.4 under agitation for 30 min. The particle suspension
was purified from the polyelectrolyte solution via centrifugation (15,000 g, 30 min), and
this process was repeated iteratively for deposition of each layer. Final particle
suspensions were characterized by dynamic light scattering (DLS) and zeta potential
analysis (Malvern Instruments, ZS90) and stored at 4C prior to testing.
in vitro Drug Release. Particle formulations were suspended in 1X PBS and incubated in 2
mL dialysis float-a-lyzers (3.5KDa MWCO) while agitated at 370C in 1X PBS under sink
conditions. Small aliquots were collected at various time points, replaced with fresh
28
solution, and analyzed by high-performance liquid chromatography (HPLC, Agilent
Technologies).
in vivo experimentation. For biodistribution experiments, NCR nude female mice
(Taconic) were used. To attenuate gut fluorescence, an alfalfa-free special diet (AIN-93M
Maintenance Purified Diet from TestDiet) was administered to the mice 1 week prior to
and during experimentation. Nanoparticle formulations suspended in IX PBS were
administered via the tail vein. In vivo imaging (IVIS, Caliper Instruments) was performed at
regular time points. Full scale biodistribution harvesting of relevant organs was conducted
at a relevant time point and imaged using the IVIS to quantify recovered fluorescence.
Circulation persistence experiments were performed in BALB/c female mice from Taconic.
Nanoparticle formulations were administered via the tail vein. Blood collection was
obtained from the retro-orbital sinus (-0.2-0.3mL, diluted in -0.1mL 0.5 M EDTA) and
imaged using the IVIS to quantify recovered fluorescence. For cardiogreen recovery
experiments, feces were recovered from three BALB/c mice per experimental group and
fractional recovered fluorescence (IVIS) was quantified across a 48 h window. For coinjection experiments, free polymer corresponding to the terminal layer polymer used for
the NP architecture being investigated were diluted in PBS and co-injected at normalized
concentrations of particles against 10 mg/kg and 20 mg/kg free polymer in NCR mice.
Experiments were performed under the guidance and supervision of the MIT Department
of Comparative Medicine and Committee on Animal Care.
in vitro opsonization experimentation. Texas Red-labeled polymer nanoparticles
(company) were incubated with fluorescently labeled protein (Human IgG-FITC, Sigma) for
1h. Bound protein was quantified using a nanodrop spectrophotometer (GE Nanovue Plus).
These particles were subsequently incubated with the murine macrophage cell line
(J774A.1, ATCC) for 1h and cell-associated fluorescence was quantified via fluorescenceactivated cell sorting (BD LSR II). Geometric mean cell-associated fluorescence was used
for display of particle association with the cells.
29
Results and Discussion
The schematic for generating films using electrostatic LbL assembly is presented in Figure
2.1.A.
The PLGA core particles used in this work are generated from carboxylic acid
terminated PLGA, and present a net negative charge on their surfaces, which provide a
means
for initiating
LbL assembly.
Iterative
adsorption
of alternately
charged
polyelectrolytes is accomplished by incubation of particles in aqueous polycation and
polyanion solutions, with intermediate rinse steps, to generate LbL nanoparticles. Although
this study focuses on the materials tabulated in Appendix A.T, a much larger library of film
components, along with a diverse range of therapeutics, can be generated for incorporation
on this platform. Towards this end, materials with relevant biological functionalities are of
particular interest, such as negatively-charged linear polysaccharides including hyaluronic
acid (HA), dextran sulfate (DXS), and alginate (Alg). Each of these polysaccharides have
demonstrated anti-fouling properties as hydrophilic brush-like hydrated coatings to
minimize protein adsorption and opsonization[ 107 -110 ; as such, they present promising
biodegradable, biomimetic alternatives to the traditional and widely-reported use of
poly(ethylene glycol).[111] Further, HA is a known ligand for CD44 receptors[112], which are
characteristically overexpressed in many aggressive cancer cell types, including triplenegative breast cancer.[ 1131 Alg is known to be an essential protective coating for
pathogenic bacteria to evade host immune responses.[" 4] Each of these materials has been
incorporated in the current study.
Characterization of the resulting LbL particles is shown with regard to particle
hydrodynamic size and charge in Figures 2.1.B and 2.1.C.
Following deposition of six
layers on a 125 nm diameter polymer core, particles range in hydrodynamic size from
180nm for DXS-terminated NPs to 250nm for Alg-terminated NPs, based on a number
distribution.
The dynamic light scattering data taken as a function of the number of
deposited bilayers indicates that the nanoparticles grow at a rate of approximately 18 nm
in diameter with each bilayer pair, corresponding to nanolayer thickness increases of 9 nm
per adsorbed polymer pair, over the first five adsorbed layers.
growth is consistent with traditional LbL processes.[
15 ]
This relatively linear
The three different LbL film
architectures are similar in size until the sixth and final adsorbed polyanion layer.
30
The
DXS-terminated LbL nanoparticles continue to grow linearly; whereas the particles
terminated with HA and Alg both indicate a significant increase in thickness with the final
adsorbed layer, which is 55 and 77 nm, respectively. In LbL assembly, it has been observed
that the thickness and density of the layer deposited is highly dependent on pH for weak
polyelectrolytes due to changes in the degree of ionization along the backbone.[1
16]
Furthermore, certain weakly charged polyelectrolytes, including hyaluronic acid, can
undergo exponential growth behavior due to interdiffusion of the absorbing species above
a critical thickness.[ 117 ,1181
This increased increment may be further impacted by the high
degree of hydration of the high molecular weight HA (500K MW) and Alg (200K MW)
backbones, and to the partially charged nature of the carboxylic acid groups along the
glycopolymer backbone.
The surface charge of the particles is reversed with each
deposition step, confirming that the film is building in a stepwise fashion driven by
alternating electrostatic adsorption. The overall surface potential was (+/-) 50-60mV (in
ultrapure 18 MQ water, 250C), independent of terminal layer.
Previous work established a minimum number of four layers for generation of
serum-stable particles, with diminishing returns beyond eight layers.[ 27] Independent
optimization conducted for the current PLGA-based particle template was found to be
consistent with these findings; therefore, this study reports data from six-layered (6L)
particles with the following nanoparticle architecture: a biodegradable PLGA nanoparticle
core (acid-terminal, hydrodynamic diameter = 125nm; -potential = -5OmV) containing a
drug, which is layered with polycationic/polyanionic pairs deposited iteratively on top of
this substrate, terminated with a final polyanion. The systems were assembled with polyL-lysine (PLL) as the polycation and DXS as the intermediate polyanion, and are indicated
as PLGAdrug/(PLL/DXS) 2 /PLL/Xterm, where Xterm is either HA, Alg, or DXS.
To facilitate our investigation of the in vivo fate of drug cargo after systemic
circulation, we sought a good candidate for a conventional chemotherapeutic, doxorubicin
(logP (pH 7.4) = 1.27), which inhibits topoisomerase II, resulting in DNA damage thus
disabling cellular replication. The model drug chosen for systematic stability assessments
of different nanoparticle architectures is cardiogreen (CG, ?ex = 740nm,
Xen
= 820nm, logP
(pH 7.4) = -0.29), a hydrophilic near-IR dye that absorbs light at 820nm, where there is
31
minimal whole-animal auto-fluorescence.
The fundamental chemical characteristics of
both small molecules are also similar; they are hydrophilic with low molecular weights
(543.5 g/mol for doxorubicin and 775g/mol for CG) and are significantly charged (plcG =
3.3; pIDox = 11) under physiological conditions with water solubilities in the range of
1mg/mL and similar lipophilicities (logPCG = -0.29; logPDox = 1.27; pH 7.4). Additionally, CG
is cleared from circulation readily, with a serum half-life of 2-3 minutes (compared to 10
minutes for doxorubicin),[ 1 91 and in a manner that traffics exclusively via the liver, such
that its clearance from the body can be monitored in the liver and gall bladder via in vivo
imaging, and corroborated with recovered fluorescence from the feces in addition to the
relevant biodistribution and blood sampling.[ 120I These advantageous properties make CG a
particularly good choice to model the in vivo fate of small molecule therapeutics in systemic
delivery carriers, as multiple readouts, including biodistribution, persistence data via blood
and feces collection, are generated from a single animal to guide particle engineering.
To
demonstrate
the
similarity
of
CG
to
conventional
small
molecule
chemotherapeutics, release profiles for both doxorubicin and CG were determined for the
uncoated NP core as well as the layered LbL nanoparticle architectures under physiological
conditions (370C, stirred in large excess of 1X phosphate buffered saline (PBS)) using high
performance liquid chromatography (HPLC) to determine quantitative and fractional
release. The resulting release curves for these two molecules are compared in Figures
2.2.A and 2.2.B. The in vitro release characteristics are similar, with both exhibiting the
same general release profiles and similar fractional release in the coated (60% released
after 48h, following from initial drug loadings of 10ug/mL for CG-loaded NP formulations
and SOug/mL for Dox-loaded NP formulations) and uncoated (100% released after 48h)
versions, proving CG to be an effective model of doxorubicin for the demonstration of LbL
NP technologies as small molecule delivery agents. The observed release kinetics is typical
of biodegradable particle cores and consistent with mathematical models accounting for
small molecule diffusion-based release (initial logarithmic-phase release) and bulk erosion
(inflection after initial diffusion-based release plateaus) of the PLGA matrix.[1 2 1 -123 ] The
characteristic diffusion-erosion based release is observed for both coated and uncoated
systems; however, the LbL coated systems show significant benefit in small molecule
sequestration, minimizing the characteristic bolus release observed initially (from -35%
32
uncoated core to -20% coated after 1h), and sustaining release of the therapeutic over
longer time scales of up to 4-5 days in vitro. Greater control and modulation of bolus
release is important to diminish systemic toxicity while the nanoparticle is en route to the
tumor site, particularly for cytotoxic chemotherapeutics. Extended release of drug from the
nanoparticle once it has reached its target location may also have therapeutic value for
certain types of molecular drugs. The variable terminal layer had little to no effect on
overall release profiles, despite the range of functionalities and molecular weights used for
the end-capping layer. Drug release from the NP is characteristic of the bulk film properties
of the LbL multilayer, and relatively independent of the end-capping layer. It thus appears
that the terminal layer primarily serves as a functional interface directing the
biodistribution and blood circulation behavior of the delivery system, while the bulk film
impacts the pharmacokinetics of the encapsulated therapeutic. These findings suggest the
LbL nanoparticle technology can provide a modular platform for systemic drug delivery.
To understand the impact of these particle architectures on the pharmacokinetics of the
model drug and nanocarrier used for delivery, simultaneous tracking of both the near-IR
labeled polycationic layer 1, PLL 700 (modified with Cy5.5 dye, Xeximax = 673 nm,
Xem,nax
=
707
nm), and encapsulated model drug, CG, was conducted using in vivo imaging (IVIS)
following systemic administration via the tail vein. This approach to systematically screen
libraries of LbL NP architectures as therapeutic vehicles is an exciting, novel means of
generating multiple readouts, including biodistribution and persistence in the animal
(blood, feces sampling), from longitudinal studies of animals. In this way, nanoparticle
stability and its pharmacokinetic behavior is assessed based on persistence in circulation
and biodistribution over time at Xex = 640 nm,
Xem
= 700 nm (PLL 7oo channel), while drug
stability and pharmacokinetics is also evaluated by biodistribution and circulation data, as
well as recovered drug fluorescence in feces as a function of time at ex = 745 nm, Aem = 820
nm (CG 820 channel). Due to the hepatic-specific clearance of CG, the feces serve as a
quantitative measure of drug clearance from the animal to further assess drug
bioavailability as a function of time with each NP system investigated.
The IVIS images achieved for biodistribution studies are shown in Figure 2.3 for
both the drug (3A) and NP (3B) channel. Significant differences in the accumulated signal
intensities were observed for different formulations of the layered architectures in the
33
liver, relative to the controls, within the first 30 minutes post-administration on both the
PLL 7oo and CG820 channels. Normalization of the liver-associated CG fluorescence (circled
and identified with an arrow) with total fluorescence from the amount injected, shows
significant reduction of CG accumulation in the liver when administered as cargo in LbLNPs versus free or uncoated PLGA NPs. The quantified signal for liver accumulation at 5
minutes post-injection for the free drug (70%) and uncoated nanoparticle (55%)
formulations are much higher when compared to the layered architectures, Alg-term
(15%), HA-term (17%), and DXS-term (28%). At the carrier channel for PLL7 oo, 50%
injected dose is associated with the liver after 5 minutes for free PLL 7oo, as compared to
12% for Alg-terminated 6L NPs, 15% for HA500K-terminated, 25% for DXS-terminated.
The full 48h stability panels at both channels are shown in Appendix A.2-3, with
quantification of the total radiant efficiency from the whole animal as a function of time
displayed for the panel of images at each channel to gauge persistence of the individual NP
components (drug, particle). These data reflect the improved biodistribution and overall
enhanced persistence of the coated CG-loaded NP formulations, relative to the uncoated
polymer-encapsulated and free drug formulations.
The results from in vivo whole-animal imaging were further substantiated by
harvesting of the relevant clearance organs. This biodistribution data was collected 30
minutes post-administration (Figures 2.4.A and 2.4.B, Appendix A.4.). Significantly
reduced drug clearance is evident for the coated systems within the first 30 minutes postinjection, with 55% ID/g associated with the liver after 30 minutes for free cardiogreen, as
compared to 43% for the uncoated PLGA CG-loaded core, 33% for DXS-terminated, 18% for
HA-terminated, and 20% Alg-terminated systems. Layered architectures also exhibit
minimal clearance during this time frame as delivery carriers, with 70% ID/g associated
with the liver for control free PLL 7oo, 32% for DXS-terminated, 20% for Alg-terminated
systems, and 22% for HA-terminated systems.
The similarity in percent recovery of
fluorescence between the particle and drug channel from this biodistribution data, along
with the in vivo images in Figure 2.3, illustrate co-localization of the drug with the delivery
platform for the initial period post-administration (up to 4h); however, as shown in
Appendix A.2 and A.3, this does not hold at long time points. At 8h and beyond, due to
34
small molecule out-diffusion ("leakage") from the LbL NP and subsequent bulk-erosion
based release observed in vitro, the particle and drug localize in the liver at different times.
Further, recovered fluorescence from the feces can also serve as readouts for drug
clearance in the animal. As shown in Figure 2.4.C, CG is recovered in the bile following
longer time intervals for the coated systems.
The DXS-terminated system peaks in
fractional CG recovery at only the 4-8h time span post-injection as compared to Alg- and
HA-terminated systems, which peak in the 8-12h time period. This further corroborates the
enhanced persistence and drug bioavailability that the coated delivery systems can
promote.
To try to rationalize the observed differences in feces drug recovery and
biodistribution of DXS-terminated systems as compared to the other two coated systems,
we hypothesized that, as previously implicated,[ 124
125]
the sulfonated polysaccharide
dextran sulfate preferentially binds liver receptors (e.g. liver endothelial cells, Kupffer
cells) that scavenge blood for soluble macromolecules. From the biodistribution and feces
recovery results, it is clear that HA- and Alg-terminated systems delay particle and drug
clearance as well as exhibit improved biodistribution. These results implicate that this
potential receptor-mediated mechanism for clearance impacts drug biodistribution as well
as that for the carrier agent. This is further illustrated by co-injection of free terminal layer
polymer with the corresponding polymer-terminated 6-layered NP. Significant reduction in
CG signal was observed only when free DXS was co-injected with the DXS/sulfonateterminated NPs, as evident in Figure 2.4D. Free polymer competition for receptor-binding
suggests that the sulfonated terminal layer preferentially interacts with the liver
endothelial cells,[ 1 25 , 126] promoting receptor-mediated endocytosis that significantly affects
the clearance of DXS-terminated LbL particles. The differential impact of a co-injection of
free
terminal
layer polymer was quantified by normalizing the liver-associated
fluorescence obtained from co-injection studies with signal intensities for particle-only
studies (30 minute stability panel shown in Appendix A.5.). No effect was observed for
corresponding co-injections with Alg- and HA-terminated systems, in contrast to DXSterminated nanoparticles, with liver-associated uptake, implying receptor-independent
liver uptake.
35
Additionally, in vitro opsonization of the different NP architectures (Appendix A.8.)
showed little dependence on protein adsorption and no dependence on subsequent
macrophage-associated
fluorescence as a function of terminal layer choice, further
suggesting that all coatings have similar protein resistive properties and that the
mononuclear phagocyte system clearance is likely not the primary reason for the different
liver signal intensities (respective to uncoated and PLL controls). Poorer biodistribution of
the DXS-terminated particles results relative to other coatings guided particle and drug
persistence experiments for only the more optimal systems, HA- and Alg-terminated
particles. Results on drug and carrier stability in circulation for these systems are
presented in Figures 2.4.E and 2.4.F (raw data shown in Appendix A.6.). The particles
exhibit improved pharmacokinetic properties in circulation, with Alg-terminated NPs
characterized by half-lives, based on a two-compartment model, of 0.59h and 7.18h and
HA-terminated NPs characterized by half-lives of 0.50h and 7.861h. The lifetime of drug
from these platforms is also, compared to the reported free drug half-life of 2 min and
uncoated PLGACG core half-lives of 0.15h and 1.87h, extended to 0.52h and 4.17h for Algterminated NPs and, for HA-terminated NPs, to 0.48h and 4.54h (based on a 2compartment model).
These findings are consistent with previously reported PEGylated versions of PLGA
nanoparticles that show 30-65% particles remaining (for various MW PEG shells) in serum
3h post-i.v. administration with liver-associated clearance of the particles remaining at or
below 20% up to 5h post-administration.[ 12 71 This is important because, in addition to
imparting similar hydration properties to the particle surface like PEG, LbL can also
provide a surface that is molecularly targeted for cellular engagement and/or cue-sensitive
(e.g. pH-responsive[ 24]) for exploiting the cellular environment in hypoxic tumor
compartments. Alternatively, PEG can be used exclusively as a hydration layer for
minimizing protein adsorption, opsonization, and subsequent clearance by the major
players in the mononuclear phagocyte system (MPS) system - a property that in fact can
also negatively impact interactions with target cells.[12 8 , 1291 These findings clearly highlight
the benefit of LbL-functionalized nanoparticle systems as extended release platforms,
particularly for small molecule therapeutics, allowing for enhanced drug and particle
persistence in circulation and that promote hydrated, protein-resistive properties for more
36
efficacious systemic drug delivery. This provides much promise towards development of
these systems as molecularly targeted entities, in addition to improved biological
performance observed in this investigation.
Small molecule sequestration in delivery platforms for extended periods in
circulation remains a significant challenge to selective delivery of therapeutics to the target
tissue-105 however, the coatings investigated here demonstrate an enhanced level of
control over diffusional bolus release following administration, as well as an ability to
enhance the circulation half-life for longer periods of time in vivo. The findings in this
investigation, namely the improved pharmacokinetic behavior observed with Alg- and HAterminated NPs via biodistribution, feces drug recovery, and persistence in circulation, are
consistent with documented anti-fouling characteristics of HA[ 1 3 0- 1 3 2] and Alg.[
110
, 125,
133]
These systems present exciting new opportunities for enhancing delivery of small molecule
therapeutics; however, due to the accelerated rate at which the uncoated PLGA drugloaded core is hydrolytically degraded and unloaded in circulation (with short half-lives
and poor biodistribution, see Appendix A.7.), the LbL system can be further adapted to
different core materials and encapsulation approaches (e.g. covalent attachment of
therapeutics to core materials or film components), which is currently under active
investigation in our laboratory to probe whether the benefits of enhanced persistence,
improved biodistribution, and controlled drug release provided by LbL-coated systems will
promote more efficacious EPR-based and molecular targeting strategies.
Conclusion
LbL is a facile approach for generating functional thin films for enhanced systemic delivery
of nanoscopic systems. This self-assembly method allows for incorporation of a broad
range of materials; and due to its water-based synthesis, it allows for incorporation of a
diverse set of therapeutics without significant alteration of biological function. To facilitate
continued development of these systems, the current study establishes a two color imaging
methodology that enables efficient, high throughput in vivo screening of a library of
material architectures and establishes each as systemic drug carriers. Using live animal in
vivo imaging, multiple readouts, including biodistribution, persistence, and feces drug
37
recovery can be performed, longitudinally assessing in vivo fate of both particle and drug in
a single animal model. Further, this work demonstrates LbL-drug delivery particles with
enhanced persistence and improved drug/carrier biodistribution using biomimetic
alternatives to poly(ethylene glycol), specifically Alg and HA, as terminal layers for this NP
delivery vehicle. In total, LbL provides a powerful tool to engineer NP systems that control
drug release while incorporating materials that are dynamic, protein-resistive, and
molecularly-targeted for more efficacious systemic treatment of diseased states.
38
A
PL1700
Centrifuge
CG 8 2 0
x layers
Cati
Anion
Centrifuge]
'""I
-.
--*-
240-
C
PLGACG/(PLL/DXS) 3
PLGACG/(PLL/DXS) 2/PLL/HA 500 K
PLGACG/(PLL/DXS) 2/PLL/Aig
50iE
E'
0
180-
25-
U
~
~
~
~
I
~
I
I
I
I
~
I
0.
4-'
-
So
-'
-
0-
I
I
I
I
..---
GJ
I
I
-25-
~
E
I
Ij
I
-50
120-
T
nI
0
U
I
-o
E
-"-
2
Layer
4
-75-
0
6
2
Layer
4
6
Figure 2.1. Synthesis and characterization of a Layer-by-Layer PLGA nanoparticle (LbL NP) system with
cardiogreenencapsulatedformulti-color imaging.
(A) LbL deposition on NP template. (B) Materials library used in current study. (C) Representative mean
hydrodynamic diameter (based on a number distribution) and (D) zeta potential analysis, conducted in
ultrapure water at 250C. All particles in current study are 6 layers (6L) with the following NP architecture:
PLGAs:sodrg/(PLL/DXS) 2 /PLL/Xerm, with drug = doxorubicin (Dox) or cardiogreen (CG) and Xterm = terminal
layer (DXS, Alg, HA500K). (PLGA) poly(lactic-co-clycolic acid), (PLL) poly(L-lysine), (DXS) dextran sulfate, (Alg)
alginate, (HA) hyaluronic acid; n = 3, mean +/- SEM. PDI of uncoated PLGA NPs is 0.1;functionalized NPs ranged
from 0.15 for DXS-terminated NPs to 0.2-0.25 for HA-coated and Alg-coated NPs. Representative scanning
electron micrograph of LbL NPs (HA-coated) shown in Appendix A.1.
39
.
-
-
-
-
-
I it I
__ --t B 1 00-
80-
800)
U)
Cu
0
4
0
U)
(U
6040-
-
a) 6
PLGACG
20 o-k
PLGACGI(PLL/DXS)2/PLUAg
-y -
PLGACG/(PLL/DXS) 2/PLUHA5 00K
0
4
8
20
PLGACG/(PLU-DXS)3
.1.I...4
12
16 20
time (h)
PLGAOX
PLGAD 0
-2/PLLAg
-k- PLGAno0 /(PLUDXS) 3
-V-PLGADox/(PLU/DXS) 2/PLUHA50o
54U
-
-
a)
2440
60
0
.2
0
80
4
8
12
x/(PLUDXS)
1.
1
16
20
2440
60
80
time (h)
Figure 2.2. In vitro release of uncoated and coated cargo-loaded(cardiogreen,doxorubicin) PLGA LbL NPs.
(A) Cardiogreen(CG)-loaded NPformulations and (B) Doxorubicin (Dox)-loaded NPformulationsin PBS, pH 7.4
at 370C under agitation (3.5KDa MWCO) as measured by high-performance liquid chromatography (HPLC).
(PLGA) poly(lactic-co-clycolic acid), (PLL) poly(L-lysine), (DXS) dextran sulfate, (Alg) alginate, (HA) hyaluronic
acid, (CG) cardiogreen, (Dox) doxorubicin; n = 3, mean +/- SEM.
40
A
/-
B
PLL,,
/-PLL700
qWI__CG82
Pre
5min
|15min |30min
opened
Pre
5min
15min
|30min
max
max
min
min
(ii)
(iv)
A
Figure 2.3. Stability assessments of different LbL PLGA nanoparticlearchitecturesassessed by drug (CG8 20) and
particle (PLL 7oo) tracking via in vivo whole animalfluorescentimaging.
(top of each image panel) Schematic detailing the fluorescent trackers used for stability assessment of NP
architectures. CG82 0 is the cargo encapsulatedin the LbL NP architecturewith thefirst polycationic layer of the 3
bilayer structure labeled with Cy5.5. (A) CG82 o stability panel, surveying drug bioavailability up to 30 minutes
post-injection. (i) free CG8 2o; (ii) PLGA5O:5OcG; (iii) HA-terminated 6L NP; (iv) Alg-terminated 6L NP; (v) DXSterminated 6L NP. Representative 48 h study shown in Appendix A.2. IVIS images at CG channel (tex = 745 nm,
Aem = 820nm), surveyed up to 30 min and subsequently opened. Region of interest analysis shows -70% injected
dose associated with the liver (white dashed circle identified with an arrow) after 5 minutes for free CG, -55%
for PLGAso:50 c, -28% for DXS-terminated and -15% for AIg- and HAsOK-terminated 6L NPs; n = 3. (B) Carrier
stability panel, surveying nanoparticlebiodistribution up to 30 minutes post-injection. (i) free PLL 7oo; (ii) DXSterminated 6L NP; (iii) HA-terminated 6L NP; (iv) Alg-terminated 6L NP. Nanoparticle tracking facilitated by
labeling offirst polycationic layer (PLL) in 6L NP architecturewith Cy5.5. Representative 48 h study shown in
Appendix A.3. IVIS images at PLL 7oo carrierchannel (2tex = 640 nm, Aem = 700 nm) were surveyed up to 30 minutes.
Region of interest analysisyields -50% injected dose associated with the liver (white dashed circle identified
with an arrow) after 5 minutes for free PLL 7oo, as compared to -25% for DXS-terminated, -15% for HA500K-
terminated, -12% for Alg-terminated 6L NPs. (PLGA) poly(lactic-co-clycolic acid), (PLL) poly(L-lysine), (DXS)
dextran sulfate, (Alg) alginate, (HA) hyaluronicacid, (CG) cardiogreen;n = 3.
41
A
B
80-
Rn
PLL70 0
60.. o9
CG820
PLGACG/(PLUDXS) 3
PLGACG/(PLUDXS) 2/PLUALg
PLGACG/(PLUDXS) 2/PLUHAoOK
PLGACG
PLGACG/(PLUDXS) 3
PLGACG(PLLIDXS) 2/PLUAIg
PLGACGI(PLUDXS) 2/PLUHA5 00 K
600
40
.40-
0
0
M111
20-
20-
Liver Spleen
&
0-
Liver
Kidneys Lungs Heart
Spleen Kidneys Lungs
Heart
D
C
-
nG
T)
0
0)
0.4-
=
=
=
C
g"
0
PLGAC,/(PLUDXS)2/PLLAg
PLGACG/(PLLJDXS) 2/PLUHAOK
PLGACG/(PLUDXS),
FDCG
150.
~CE
PLGACG
.~0
0~
0O 0 0
6LNP(X-term)
6L NP + Free X (10mg/kg)
6L NP + Free X (20mglkg)
100
2,
0
T
0.2-
~L_
F
111 of
PLGACGI(PLLDXS) 2 /PLUAlg
-10- PLGACG/(PLUDXS)2/PLUHA5 00 K
-0- PLGACG
-0
0)
C
0
0
I.-
0
E 10-1
0
0
20
time (h)
40
- -
E 10-
-
- -
-0
+
-
-0- PLGACG/(PLUDXS) 2/PLUAig
-5- PLGACGI(PLUDXS) 2 /PLUHA50 0K
- PLL700
HA500K
A:g
X-term
DXS
4-8h 8-12h 12-24h 24-48h
time (h)
0-4h
00
-J
-J
0~
1Ii0
.U
U
-
0. -- r
0 pre-inj
50I
60
1
0
10
time (h)
20
30
.
Figure 2.4. Pharmacokinetic data for the PLGA LbL nanocarriers (PLL 7oo, CG82o), as a function of terminal layer
and the encapsulateddrug, CG8 20
(A,B) Biodistributionprofile of carriervia PLL 7oo tracking (Aex = 640nm, em = 700 nm and CG820 (Aex = 745 nm,
Aem = 820 nm) 30 minutes post-injection (n=3). Raw data shown in Appendix A.4. (C) Profile of CG recovery in
murine feces (3 total mice feces/experimental group). Aex = 745 nm, Aem = 820 nm. Data presented as fraction
initial dose over a 48h period. (D) Dose-dependent co-injection of free polymer with corresponding polymerterminated 6L NP. Liver-associatedCG 82 0 fluorescence (Aex = 745 nm, Aem = 820 nm) collected 15 min following iv
administration.Data normalized to average value associated with 6L NP withoutfree polymer co-injection. Raw
data shown in Appendix A.5. (EF) Circulationprofile of CG 82o-loaded NP formulationsfor both the carrier(Aex =
640 nm, tem = 700 nm and encapsulated drug (CG820). Raw data shown in Appendix A.6. (PLGA) poly(lactic-co-
clycolic acid), (PLL) poly(L-lysine), (DXS) dextran sulfate, (Alg) alginate, (HA) hyaluronic acid, (CG) cardiogreen;
n =3, mean +/-SEM.
42
References
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Z. Poon, D. Chang, X. Zhao, et al., ACS Nano 2011, 5, 4284-4292.
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44
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f
I
4
A
Ai
........
..
Chapter 3. Osteotropic therapy via targeted LbL nanoparticles.
This chapter is in part adapted from:
Stephen W. Morton*, Nisarg J. Shah*, Mohiuddin A. Quadir, Zhou J. Deng, Zhiyong
Poon, Paula T. Hammond. Advanced Healthcare Materials 3(6): 867-875, 2014.
Introduction
The development,
maintenance,
and repair of bone require
cell-mediated
remodeling to sustain the structural integrity of the tissue. Disturbances in the
physiological processes of osteoblast-mediated bone deposition and osteoclast-mediated
bone resorption are observed in many bone-related disease states, such as osteosarcoma,
cancer metastasis to bone, osteoporosis, and Paget's disease of the bone.[1 34
Several
therapies have been developed to combat these pathologies;[1 34 -1361 however, the clinical
outcomes for patients with these diseases continue to be very poor. The arsenal of available
agents to treat patients has not made any substantial impact in improving their survival,
and new methods for therapy are critical. Engineering a robust delivery platform with
bone-tissue level specificity to treat these diseases can improve therapeutic efficacy, lower
systemic toxicity, and improve disease management.
Significant fundamental work in the area of bone biology has uncovered the
potential of bone-specific agents, such as bisphosphonates.[1 36 ,1371 These compounds act as
pharmacophores, whereby the pyrophosphate-like structure coordinates calcium ions
within the hydroxyapatite mineral in bone with high affinity and specificity. Accumulation
of bisphosphonates further promotes inhibition of bone resorption by inducing apoptosis
in osteoclasts that are responsible for this action. As such, this class of pharmaceutics is a
potent drug for promoting homeostasis between osteoclast and osteoblast activity in, for
example, osteoporosis, while also promoting tissue-specific binding and localization.
Bone targeting systems[ 138
1421
have been synthesized to utilize the specificity of a
variety of ligands, including bisphosphonates, tetracycline and derivatives, sialic acid, and
bio-inspired materials containing similar functionalities[1 361; however, systems capable of
therapeutic delivery for treatment of bone-related diseases remain limited. Towards this
41 77
end, the modularity of Layer-by-Layer (LbL)-functionalized nanomedicine systems[ , ,143-
46
1491
presents an attractive opportunity to incorporate these bone-specific, highly water-
soluble ligands on the surface of nanotechnology for targeted drug delivery.
In this chapter, the synthesis of tissue-targeted LbL nanoparticles, specifically
manufactured to hone in on bone tissue, is reported. To achieve this, a polyelectrolyte,
poly(acrylic acid) (PAA), was functionalized with a bisphosphonate, alendronate, and
subsequently electrostatically-assembled in a nanoparticle coating. The functionalized
particles accumulated in subcutaneous 143B osteosarcoma xenografts, where they
delivered their payload, doxorubicin, in a mouse model. The targeted particles significantly
attenuated tumor burden and extended animal survival, in some cases even completely
eliminating tumors. The results described herein establish LbL as a modular approach to
develop targeted
drug
carriers
via adsorption
of ligand-functionalized,
aqueous
polyelectrolytes for tissue-specific targeting, further developing these systems towards
clinical application.
Materials and Methods
PAA-Alendronate conjugation. Alendronate (Alfa-Aesar) was conjugated to poly(acrylic
acid) (Sigma) through the primary amine functional group of alendronate. Half of the free
carboxyl end group of the poly(acrylic acid) was targeted for coupling. As both of the
reacting species were only soluble in water, we have used water soluble 4-(4,6-Dimethoxy1,3,5-triazin-2-yl)-4-methylmorpholinium chloride (DMTMM) as the coupling agent
50].
To
perform the coupling chemistry, alendronate (1.0 eq.) was added to a solution of
poly(acrylic acid) (0.5 equivalents considering all carboxylic groups) in 100 mL water. The
solution was stirred for 10 minutes after which DMTMM (2.0 equiv.) was added to the
resulting solution. The reaction was allowed to run for 12 hours at room temperature. After
the reaction period, the solution was dialyzed against water for 36 hours for complete
removal of coupling agent and unreacted alendronate.
The dialyzed solution of
alendronate-conjugated poly(acrylic acid) was lyophilized to yield a white foam-like
product in 65% yield. 1H NMR (in D20): 1.33 - 1.9 ppm [broad signal, poly(acrylic acid)],
2.10-2.40 ppm [broad signal, poly(acrylic acid)], 2.88 ppm (2H, alendronate). The degree of
conjugation was estimated according to the assay protocol reported previously after the
47
hydrolysis of the conjugate at pH 10 with 0.1M NaOH[ 15 11. The estimated degree of
functionalization was found to be 43.2%.
LbL on QDs. Carboxyl-modified quantum dots (QD800, Life Technologies) were suspended
in dilute solution (-lOuL stock in lmL DI water) and injected in excess 10mL
polyelectrolyte (5mg/mL for PLL-Hydrochloride (4-15K, Sigma); 1mg/mL PAA45OKAlendronate) while under agitation at 40C. Each layer was incubated for -30 minutes, prior
to purification via ultracentrifugation (12,500 RPM, 30m minutes - 1st layer required
longer spins, subsequent purifications are much more expedient) and washing with DI
water, repeated twice prior to introduction to the subsequent polyelectrolyte. The final
functionalized particle was re-suspended in 1X PBS after the 2nd wash and filtered in a
0.45pm filter for further experimentation.
Liposome
synthesis.
(DSPC:Cholesterol:POPG
Liposomes
-
were
formulated
at a mass
all purchased from Avanti
Polar
ratio
of 56:39:5
Lipids). These three
components were dissolved in a 2:1 mixture of chloroform:methanol. A thin film of these
materials was generated by rotary evaporation at 400C, 150 mbar for approximately 10
minutes. This film was allowed to dessicate overnight for complete drying. Hydration of the
lipid film was conducted at 650C under sonication in 300 mM citric acid buffer (pH 4) for 1
hour, after which they were filtered through a 0.2 pm PES syringe filter and allowed to cool
to room temperature. The pH of the liposomal suspension was then adjusted to 6.5 by
addition of 300 mM sodium carbonate buffer to create a gradient between the exterior and
interior compartments. Doxorubicin (LC Laboratories) at a feed ratio of 3 mg drug to 50 mg
lipids was then added in a 0.9% sodium chloride solution to load via a pH gradient method.
The final drug-loaded system was subsequently purified out of the high salt buffers and any
excess, unloaded drug via centrifugal filtration (100K MWCO Millipore) and transferred to
DI water for subsequent functionalization via LbL. Blank liposomes were prepared in the
same fashion; however, no drug was added.
LbL on liposomes. Liposomes were diluted in 1 mL DI water (from a 50 mg batch prepared
in a final suspension of 5 mL DI water - use -200 p.L stock to 800 p.L DI water) and injected
48
in excess polyelectrolyte under agitation at 40C as previously described for QDs. Incubate
for -30 minutes, purify via ultracentrifugation at 10,000 RPM for -10 minutes - repeat
twice prior to introduction to subsequent polyelectrolyte. The final, functionalized
liposomal system was filtered through a 0.45 ptm filter and suspended in 1X PBS for further
experimentation.
Physicochemical characterization. Dynamic light scattering and zeta potential analysis
was conducted in 10 mM sodium chloride at 250C using a Malvern ZS90 zeta-sizer. Highperformance liquid chromatography (Agilent technologies, 1:1 acetonitrile to pH 5 water
mobile phase) and nanodrop absorbance measurements (480nm) were performed to
quantify doxorubicin concentrations in the prepared samples. Cryo-TEM was conducted by
imaging a frozen dilute sample of the liposomal suspension at 120 kV.
In vitro Binding/Cytotoxicity.
143B osteosarcoma cells and human mesenchymal stem
cells were seeded in 96-well tissue culture plates at a density of 1x10 3 cells/well and
allowed to proliferate for 48 hours. PLL-Cy5.5 labeled particles were quantified based on
fluorescence, serially diluted in cell culture media and added to the cells. After 24-48 hours,
cells were washed 3X with PBS. Binding was measured by quantifying the amount of
residual fluorescence in the well. Cytotoxicity was measured using a standard MTT assay.
Fluorescence-activated cell sorting analysis. Measurements were performed using a BD
LSRFortessa (BD biosciences, San Jose, CA). PLL-Cy5.5 fluorescence was collected following
excitation at 640 nm and detected at 710/50 nm (AF700 channel). Cell-association data
presented as representative histogram overlays of data collected in triplicate following
143B cell incubation with blank, PLL-Cy5.5 coated, PAA-Alendronate functionalized LbL
liposomes for 2 hours at 370C in 96-well plates. Final samples ran for analysis were washed
3X with optiMEM and trypsinized immediately prior to introduction to the flow cytometer.
Blue is the representative treatment; red is the control sample without treatment.
Confocal Microscopy. Images were taken using a Nikon A1R Ultra-Fast Spectral Scanning
Confocal Microscope (Nikon instruments Inc., Melville, NY). 143B cells were seeded in a
49
CELLview glass bottom dish (Greiner Bio-One GmbH, Germany) at 1 x 105 cells per well and
grown overnight. Cells were then incubated with blank, PLL-Cy5.5 coated, PAAAlendronate functionalized LbL liposomes (or QDs) for 2 hours at 370C. At the end of this
period, cells were washed, fixed via paraformaldehyde, permeabilized via Triton-iX, and
stained with phalloidin-568 for 30 minutes, followed by addition of DAPI for an additional
10 minutes, after which they were washed 3X and imaged.
Xenograft development/targeting/treatment/monitoring - NCR nude. Cells were mixed
in a 1:1 ratio with BD MatrigelTM Basement Membrane Matrix to a final density of 5
x
106
cells/0.1 mL injection. 0.1mL injections of the matrix-cell suspension were performed in
each rear hind flank of NCR nude mice (Taconic). Tumors were allowed to grow until a
visible tumor was established, approximately 25 mm 2 in tumor area prior to treatment
(determined via caliper measurements via the longest length and width dimensions); for
tumor targeting, xenografts were allowed to grow near terminal size (1 cm in diameter) to
reduce effects of EPR. Treatments were injected in 0.1mL 1X PBS at the concentration and
treatment regimen indicated in the figure legend. Imaging was performed using wholeanimal fluorescence imaging (IVIS, Xenogen, Caliper) at the time points indicated pre- and
post-administration at the given fluorescence filters. For targeting experimentation,
functionalized, fluorescent systems were administered at a radiant efficiency of -3x10
9
(700 channel)/-1x10 9 (800 channel) in a 0.lmL injection. This was determined by imaging
the vial prior to injection and diluting as necessary in 1X PBS. Tissue harvest (liver, spleen,
kidneys, heart, lungs, tumors, 'gutted' skeleton) was performed at the terminal point of the
experiment and imaged for fluorescence recovery against an untreated control for autofluorescence determination. Tumor sizes for the remediation study were monitored via
caliper measurements, whereby areas based on the largest visible length and width
dimensions were measured. Depth proved challenging to reliably measure at early time
points due to skin thickness, therefore it was not included for a final volume determination
for real-time tumor size measurements.
Pharmacokinetics (circulation) - immune-proficient BALB/c. Fluorescent systems (PLLCy5.5 coated LbL-targeted empty liposomes) were administered in 0.1mL 1X PBS injections
50
in BALB/c mice (Taconic) at a relative concentration of -3x10
9
(700 channel)/-1x10 9
(800 channel) in a 0.1mL injection, based on fluorescence in the IVIS prior to injection.
Mice
were
imaged pre-
and post-nanoparticle
administration
using whole-animal
fluorescence imaging (IVIS, Xenogen, Caliper) at the time points indicated and at the given
fluorescent filters. Retro-orbital bleeding was performed in a cohort of injected mice to
determine the serum half-life of the particles, presented on the basis of fluorescence
recovery in these isolated blood samples (-0.lmL blood/time point done in 3 different
mice for each time point plotted). Fluorescence recovery was determined by imaging in the
IVIS immediately following isolation and dilution 1:1 in EDTA to prevent coagulation. Final
data is normalized to an untreated control blood sample to remove artifacts of autofluorescence. The data was fit to a two-compartment model (two-phase decay in PRISM@),
from which the fast and slow half-lives characterizing the data were extracted.
Histology. Tumors were resected at the terminal point of experimentation and fixed in 4%
paraformaldehyde (PFA) for 48 hours and transferred to 70% ethanol until processed.
Tissue was embedded in paraffin wax and sections were stained with routine hematoxylin
and eosin (H&E) and Masson's trichrome stain.
Statistical Analysis. Experiments were performed in triplicates, or otherwise indicated.
Data were analyzed using descriptive statistics, single-factor analysis of variance (ANOVA),
and presented as mean values
standard deviation (SD) from three to eight independent
measurements. Statistical comparisons between different treatments were assessed by
two-tailed t tests or one-way ANOVA assuming significance at P < 0.05.
Results and Discussion
LbL-targeted nanoparticle systems were generated by covalent modification of PAA
(Mv ~ 450K, Figure 2.1.A) with alendronate (Figure 2.1.B) at 40% functionalization of the
total carboxylic groups on the PAA molecule (Figure 2.1.C). Alendronate was coupled via
its amino functional handle to PAA through an amide bond. As both of the components
were water soluble, we used methyl morpholine-based water-soluble coupling agents to
51
perform amide coupling (Appendix B.1). The immobilization of the molecule onto the PAA
backbone was confirmed by
31P
and 1 H NMR. The functionalized PAA-Alendronate polymer
was subsequently used as the polyanionic layer in LbL assembly, iteratively adsorbed on a
solid nanoparticle substrate in alternation with polycationic poly-L-lysine (PLL), as
PAA is a well-characterized[15 2 -15 4] weak
schematically illustrated in Figure 3.1.D.
polyanion with a high charge density and a non-erodible backbone that has been listed as
an approved excipient in the FDA's Inactive Ingredient Guide and is used in clinicallyapproved drug formulations as a stabilizer and thickener to tune the rheological properties
of the injectable or topically-applied therapeutic. As such, PAA presented a suitable
material for introducing tissue specificity and targeting via alendronate functionalization.
To establish these systems as osteotropic delivery systems, initial investigations
employed a 25 nm quantum dot (QD) (CdSeTe core) for imaging purposes to understand
how the coating directed particle association with osteosarcoma cells (143B) in vitro and in
vivo, as well as other hard tissue in vivo. These particles were fabricated as shown in Figure
3.1.D, with linear film growth observed for three bilayers of (PLL/PAA-Alendronate) to a
final z-average hydrodynamic diameter of 115 nm, with a PDI of 0.19, and zeta-potential of
-39 mV (10 mM NaCl in DI water, 250C; Appendix B.2).
Incubation of LbL-targeted QD 705 nanoparticles with 143B cells showed significant
binding and cell uptake after 2 hours at 370C (see Figure 3.2.A), where red is
representative of QD 7os nanoparticle fluorescence. Nanoparticle binding, on the basis of
QD 7 os fluorescence, was further characterized and observed to be dose-dependent for a
range of concentrations, over which little cytotoxicity (48 hour incubation) was observed,
as shown in Figure 3.2.B. The number of bilayers (1, 2, and 3, whereby 1, 2, and 3 layers of
the targeted PAA-Alendronate polymer are incorporated on the particle surface) was also
investigated. These results are shown Appendix B.3. It was observed that the in vitro
binding affinities and cytotoxicity profile were similar across different bilayer numbers.
Previous work[ 731 established the improved biological performance of QD nanoparticles
with 3 or more bilayers; therefore, we focused on the 3-bilayer LbL-targeted QD particles
for in vivo assessment of targeting.
To evaluate this system in vivo, LbL-targeted QD8 oo nanoparticles were administered
via the tail vein in NCR nude mice with ectopically induced 143B osteosarcoma xenografts.
52
Particle distribution in live animals was tracked using whole-animal in vivo fluorescence
imaging (Figure 3.3.A). Immediately following administration, the particles rapidly
accumulated in the bone tissue regions, particularly in the parietal region of the cranium,
spinal column, and hind limb regions. At later time points, particles trafficked to the 143B
xenografts, consistent with targeted interactions with the tumor matrix. Controls for
uncoated, untargeted QDs and coated QDs with unconjugated PAA showed no specific
affinity for hard tissue and had little to no accumulation in the xenografts, suggesting the
enhanced permeation and retention effect does not account for much of this tumor
localization (Appendix B.4). Tumor-specific accumulation was significant and observed
over the course of 8 days (Figure 3.3.B). Such a long period of accumulation may be due to
the strong binding affinity between alendronate and the osteosarcoma tissue, leading to
much longer residences times of the nanoparticle in the tumor. The tumors were resected
after 8 days, along with necropsy of other relevant tissue, and analyzed via recovered
fluorescence to investigate post-mortem particle distribution. Significant localization of
particles was observed in xenografts (accounting for -30%
fluorescence recovered),
relative to -40% in the liver and smaller fractions in each of the other organs (spleen,
kidneys, heart, lungs) harvested (Figure 3.3.C, Appendix B.5). Accumulation in the liver is
a common challenge for all nanoparticle delivery platforms, though we have shown in
other work that the use of hyaluronic acid can lower accumulation considerably compared
to other systems. Future work will investigate its use for introducing tissue specificity via
ligand functionalization.
To visualize the biodistribution of both the coating and nanoparticle in real time, the
bisphosphonate-targeted polymer, PAA-Alendronate, was labeled with a near-IR, Cy5.5 dye
and adsorbed onto the surface of QD 8 oo to allow for two-color in vivo imaging of the
targeted, 3 bilayer (PLL/PAAcys.s-Alendronate)
3
nanoparticle. We observed co-localization
of the two components - the PAA-Alendronate outer layer and the quantum dot - in the
xenografts (Figure 3.3.D, Appendix B.5) for up to 9 days, further substantiating these
systems as serum-stable targeted platforms for delivery.
Next, we applied the coating, (PLL/PAA-Alendronate), to drug-loaded particle
systems for targeted delivery and treatment of diseased tissue. For this purpose, the
coatings were adapted to a liposomal carrier for use in drug delivery (Figure 3.4.A). Initial
53
work focused on empty, negatively-charged liposomal carriers, containing DSPC (1,2distearoyl-sn-glycero-3-phosphocholine), cholesterol, and POPG (1-palmitoyl-2-oleoyl-snglycero-3-phospho-(1'-rac-glycerol)), functionalized with one bilayer of PLL/PAA-Al, which
yielded 170 nm nanoparticles with a zeta-potential of -20 mV (Appendix B.6), to validate
these systems for LbL-targeted delivery. Due to size considerations for systemic delivery,
fabrication was truncated after a single bilayer due to the significant increase in size
associated with adsorption of a single bilayer (80-90 nm per bilayer). Differences in bilayer
thickness relative to films on QDs could be due to the relative fluidity of the liposomal
substrate. Consistent with results for the LbL-targeted QDs, binding of the LbLfunctionalized liposomes to 143B cells was observed to be dose-dependent and not
apparently cytotoxic (Figure 3.4.B). Rapid internalization after a 2 hour incubation with
the 143B cells, visualized by confocal microscopy (Figure 3.4.C) and confirmed via flow
cytometry (Figure 3.4.D), was also observed. The high levels of red, diffuse nanoparticle
fluorescence (visualized by incorporation of PLL-Cy5.5 in the LbL coating), along with a
marked shift in cell-associated fluorescence observed by flow cytometry, suggested this
system enabled high levels of nanoparticle binding and uptake, providing a promising
platform for LbL-targeted drug delivery.
Osteosarcomas are known to be responsive to conventional chemotherapeutics,
such as doxorubicin, so we next explored the incorporation of doxorubicin in the liposomal
carrier for subsequent LbL functionalization. We were able to incorporate doxorubicin into
stable DSPC:Cholesterol:POPG liposomes with high drug-loading efficiency (97%) at 5.5
w/w% (drug/lipid). During subsequent coating with a single bilayer of (PLL/PAAAlendronate), no drug loss was observed, and liposomal morphology remained unchanged,
as illustrated in Figure 3.4.A. The LbL-functionalized doxorubicin-loaded liposome was
characterized by a hydrodynamic diameter of approximately 210 nm in size with a zetapotential of -20 mV, measured in 10 mM NaCl (Appendix B.6). The targeted liposomal
carriers were found to be particularly potent against 143B cells, with high levels of toxicity
over a wide-range of doxorubicin-loaded concentrations following incubation for 24 hours
and 48 hours (Figure 3.4.E), whereas the uncoated control showed about 8-fold lower
levels of toxicity (Figure 3.4.F).
54
Translating these systems in vivo, we first investigated the pharmacokinetics of the
coated liposomal system in the absence of drug. Shown in Figure 3.5.A is the circulation
profile of this system on the basis of fluorescence
recovery following systemic
administration of the PLL-Cy5.5 labeled carriers to immune-proficient BALB/c mice. Based
on a two-compartment model, the targeted carriers exhibited half-lives of 0.23 hours (fast)
and 18.7 hours (slow), indicative of a stable, long-circulating system that could promote
enhanced delivery of loaded therapeutics.
In tumor-bearing NCR nude mice, the same PLL-Cy5.5 labeled, targeted system was
observed to accumulate in the target diseased tissue rapidly, with the ability to persist in
the tumor up to the terminal point of the study - 100 hours (Figure 3.5.B, Appendix B.7).
Biodistribution results (Figure 3.5.C, Appendix B.7) at the terminal point of this
investigation further corroborated these systems as stable delivery platforms that
preferentially locate in osteosarcoma xenografts at high levels (-35% on the basis of
fluorescence recovery, relative to -47%
in the liver, ~7% in the kidneys). Elevated
amounts of recovered fluorescence in the liver suggested that hepatic clearance is the
primary means of excretion; however, the large number associated with the liver, in
addition to the values reported for other tissue, is not meant to be wholly quantitative, as
the data is reported as percent of the total fluorescence recovered from only the tissue
collected. The relative ratios of fluorescence between tissue confirmed that we were able to
achieve a high level of tumor specificity relative to the other tissue, which validated this
system for further investigation towards therapeutic delivery.
After observing high levels of tumor localization for the targeted empty liposomal
formulation, the efficacy of the LbL-targeted doxorubicin-loaded liposomes was evaluated
against 143B osteosarcoma xenograft-bearing
NCR nude mice. Initial attempts at
determining the optimal dosing regimen (drug concentration and number/timing of
injections) employed a dose escalation study of serial 1 mg/kg, 2 mg/kg, and 3 mg/kg
(based on doxorubicin loading) injections starting at day 10 post-tumor inoculation with
each treatment separated by one week. As observed in Figure 3.6.A, untreated
osteosarcoma tumors grew beyond terminal size (diameter > 1 cm) after 19 days postinoculation. Treated mice for both the uncoated and coated versions survived repeated
dosing out to 30 days, after which comparisons in terminal tumor size were drawn. As
55
observed visually in Figure 3.6.A and quantified
in Figure 3.6.B,
LbL-targeted
doxorubicin-loaded liposome treated mice showed significantly reduced terminal tumor
sizes relative to the uncoated doxorubicin-loaded liposome control (characterized by
measurements on the longest x-y dimensions). This observation is consistent with
enhancement of the therapeutic potency of the loaded drug due to preferential
accumulation of the targeted system in the 143B xenografts.
Results from the dose escalation study indicated that a higher dose was necessary to
effectively remediate the tumor. For this purpose, LbL-targeted doxorubicin-loaded
liposomes were administered at a doxorubicin concentration of 5 mg/kg, along with an
uncoated, drug-loaded control at the same concentration. Xenograft-bearing mice were
allowed to develop tumors to an area of -25 mm 2 [measured in the longest length and
width dimensions] after which they were treated, with repeated injections one week apart
for 3 total treatments until a terminal point at day 40. We observed enhanced efficacy of
the LbL-targeted doxorubicin-loaded liposome, relative to the uncoated control, as
determined by caliper measurements and visual inspection of the terminal point harvested
xenografts (Figure 3.6.C). Varying levels of therapeutic benefit for the targeted group were
observed from complete remediation (NT - no tumor observed) to tumor maintenance
(tumor reduction of 30% from day 0 on average), whereas the uncoated control group mice
showed significant tumor growth (several beyond the maximum allowable tumor burden
of 1 cm in diameter; growth of 550% from day 0 on average). These results highlight the
capabilities
of
the
LbL-targeted
nanoparticle
approach
for
highly
effective
chemotherapeutic treatment in tumor-bearing mice.
Histological analysis via a Masson's trichrome stain of recovered osteosarcoma
(143B) tumor tissue following treatment with each the targeted and untargeted
doxorubicin-loaded liposomal formulations at 5 mg/kg is displayed in Figure 3.7. Initially,
the tumor demonstrates an abundance of 143B cells (stained red) and connective tissue
(stained blue) in the tumor mass (left column). Particles coated with alendronate for
targeting the tumor tissue resulted in significant cell death (right column, top row). A lower
level of cell death resulted in tumors with uncoated nanoparticles (right column, middle
row). Virtually no change in the tumor infrastructure was observed in untreated animals
(right column, bottom row).
Micro-CT analysis of these tumors immediately prior to
56
resection (Appendix B.8) complements this data, clearly demonstrating the enhanced
efficacy of the targeted system relative to the uncoated control. These observations suggest
the potency of this approach to deliver the payload in an efficacious manner to the 143B
osteosarcoma solid tumor.
Conclusion
LbL is a versatile platform to functionalize nanoparticles in ways that promote
improved biological performance. Prior art has established LbL as a means to impart
protein-resistive, long-circulating properties to nanoparticle systems, with a means to
control biodistribution of both the carrier and drug in a complex systemic environment.
This investigation further demonstrates the modularity of this approach to impart
targeting capabilities to the nanoparticle to diseased tissue to achieve enhanced treatment
outcomes. We capture this potential with the synthesis of osteotropic nanoparticles via
incorporation of alendronate-functionalized PAA as a means of surface modification of
drug-loaded liposomes. Different nanoparticle core substrates may be used for imaging
and treatment for a variety of bone diseases. While surgical resection of a primary tumor
will continue to be the first-line of treatment, subsequent treatment with the osteotropic
nanoparticles has the potential to decrease recurrance rates and increase successful
outcomes.
These functional nanoparticles are also highly promising for future
investigations towards treatment of bone-localized metastases of invasive cancer cell types
such as breast and lung cancer. The potential to further generalize this approach towards
the built-to-order manufacture of different targeted delivery systems continues to provide
much promise for LbL nanoparticles.
57
A
0
OH
B
0
HO 4.OH
OH
H 2N0
C
O
OH
0
HOQI1,,H
P
OH
0
H0O* OH
nn
N
H
HO ~
P'O
OH
D
EIh
Purify
x
layers
M=*
Aneo
Purify
Figure 3.1. Achieving bone tissue level specificity of LbL coated nanoparticles.
(A) Aqueous anionic polyelectrolyte, poly(acrylic acid) [PAA, MW 450K], functionalized with the (B)
bisphosphonate targeting moiety, alendronate [for high specificity to hydroxyapatite in bone], to yield (C) the
aqueous, ligand-functionalizedpolymer at 40% side-chain functionalization (Appendix Bi), which is used for
complementary, iterative adsorption to the polycationic component (poly-L-lysine, PLL) in the film on the NP
substrate, schematically illustratedin (D).
58
A
B
-120
50000- -+- Binding
-.-
40000-
Relative Toxcity
-110
Z
0
0
-100
0
2 -o 20000w 1..
-90
M,
10000-
-80
.2
U)
30000)1%
L.
Cu
Lu.
/
1
-
1~
11.4
1
-
--
0
..
-
100
Incubation Concentration (pM)
10
-
6-
7n
1000
Figure 3.2. in vitro assessment of PAA-Alendronate LbL-targeted QD NPs incubated with 143B osteosarcomacells.
(A) Confocal microscopy of 143B cells incubated with the LbL-targeted QD 7os core NPsfor 2 hours. Blue staining representativeof a Hoescht nuclear stain,
green representative of a phalloidin stain of the actin filamentous cell structure, and red representative of the nanoparticle fluorescence (QD 70s
fluorescence). Scale bar representative of 20pm. (B) Binding and relative cytotoxicity of 3 bilayer [PLL/PAA-Alendronate] LbL-targeted QD8oo core NPs
following incubationfor 2 hours in 143B cells. Fluorescenceemission data correspondsto that of the nanoparticlecore (QDoo).
A
-G
A
2.5
7d
fI
5da
2d
D
8d
Photograph
IAex
C 50
B
9d
3d
= 640nm Aem = 700nm
U
40
I
2.0
$
0
30
$20
1.5
.
0
24h
9hs
5 min
pre
010
0
8
16 24 50
100 150 200
0
Liver
Spleen Kidneys Lungs
hours
Heart
Tumors
1L
Figure 3.3. in vivo evaluationof PAA-Alendronate LbL-targeted QDaoo core NPs.
(A) Representative live-animal imaging of LbL-targeted QDeoo core NPs following systemic administrationto 143B osteosarcoma xenograft-bearing NCR
2
em = 800 nm for up to 8 days. Hind-flank xenografts identified in the pre-injection image; arrows at 5
nude mice. Imaging conducted at Aex 640 nm, A
minutes and 9 hours indicate binding to native bone tissue. (B) Quantification offold tumor-specific accumulation normalized to tissue auto-fluorescence
pre-injectionfor the systemically administeredLbL-targeted QDaoo core NPs, as visualized in (A). n = 3 mice (6 tumors); data presented as mean +/- SEM. (C)
Biodistributiondata corresponding to endpoint of (A) (quantified as percent recovered fluorescencefollowing harvest of relevant tissue after 8 days postadministration). n = 3 mice (6 tumors); data presented as mean +/- SEM. (D) Co-localization of QD8oo NP core with LbL-film containing labeled PAAAlendronate7o; top row [PAA-Alendronate70o channel] - imaging conducted at Aex = 640 nm, Aem 700 nm, bottom row [QDooo channel] - imaging conducted
at A,, 640 nm, 2em 800 nm.
60
B
6x1004rx1 O
C
-120
0
o4x104C
$
-150
4
-
A
1111111 iii
-
3x10 04
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w
-60
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.30
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- - - - ____ I I
0.001
0.1
0.01
1
Concentration (10-2pg/mL)
D
F
-
m40
30
0
o20
A
10
0
2
IIII
I
I
3
II
13
AF700
.0
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s
0.1
120-
..
I
I
10
Dox Concentration (pM)
I
100-
I
I
,
0 102
4
..- 140-
Z
..,
100
80-
'II I
60+-
0.1
1
I
I
I
Ii
10
Dox Concentration (pM)
Figure 3.4. in vitro evaluation of PAA-Alendronate LbL-targeted liposomal NPs in 143B cells.
(A) Cryo-TEM of LbL-targeted doxorubicin-loaded liposomal NPs. Scale bar representative of 200 nm. (B) Binding (2 hour incubation) and relative
cytotoxicity (48 hour incubation) of LbL-targeted empty liposomal NPs over a range of concentrations, based on fluorescence, in 143B cells (Aex = 675 nm,
Aem = 710 nm; tracked using PLL 70 o as cationic component in LbL film). (C) Confocal microscopy of 143B cells incubated with the LbL-targeted empty
liposomal NPs [tracked via PLL 700 polycationic component in LbL film] for 2 hours. Blue staining representative of a Hoescht nuclear stain, green
representativeof a phalloidin stain of the actin filamentous cell structure, and red representativeof the nanoparticlefluorescence (PLLcy 5 .5 polymer shell
fluorescence). Scale bar representativeof 10 pm. (D) Representative cell-associated NP fluorescence following a 2 hour incubation of 143B cells with the
61
100
LbL-targeted empty liposomal NPs. (E) in vitro cytotoxicity of LbL-targeted doxorubicin-loaded liposomal NPs and the corresponding (F) uncoated
doxorubicin-loadedliposomal controlfollowing 48 hour (blue) and 72 hour (red) incubationperiods over a range of doxorubicin concentrations.
62
Ba
PK (2-phase decay)
half-lifefast = 0.23h
half-lifeslo = 18.7h
a-
C 80
60LL
C
)-
EE
E 5
E
1.2
O=
00
0
0oo
Li.
0
20
40
hours
60
80
T
T
)
10
01
4020j
0
20
40 60
hours
80
100
0-
Liver
Spleen Kidneys Lungs
Heart
Tumors
Figure 3.5. Biological evaluation of PAA-Alendronate LbL-targeted empty liposomal NPs.
LbL-targeted empty liposomal NPs tracked via PLL 7oo polycationic component in surface coating. Circulation data in (A) normalized to % remaining
particles recovered, on the basis offluorescence [Aex = 640 nm, Aem = 700 nm], immediately following systemic administrationto immune proficient BALB/c
mice; data presented as mean +/-SEM (n = 3). Two-compartment modelfit to determine half-lives displayed. Data includes background subtraction of blood
auto-fluorescence. (B) Quantification of fold tumor-specific accumulation normalized to tissue auto-fluorescence pre-injection for the systemically
administered LbL-targeted blank liposomal core NPs to 143B xenograft-bearing NCR nude mice; data presented as mean +/- SEM (n = 3). (C)
Biodistribution data corresponding to endpoint of (B) (quantified as percent recovered fluorescence [;tex = 640 nm, Aem = 700 nm] following harvest of
relevant tissue after 100 hours post-systemic administration);data presented as mean +/-SEM (n = 3).
63
A
-a
N B
I
E
E100-
50E
I rI
unpaired t-test, two-tailed
*** p = 0.003
-
U
Untreated
19d
C
N
E
E
150.
100-
03 Untargeted
03 Targeted
Nd
a)
50
L
-r
I-
D 2 4 6 8 101214161820
5mg/kg [Dox]
t
5mg/kg [Dox]
Targeted
30d
Treatment groups - 40d post inoculation
Treatment Regimen: 5mg/kg 22d, 28d, 35d
***-
0
E
Untargeted
days5mg/kgt
[Dox]
64
Figure 3.6.143B xenograft tumor remediationfollowing treatment with doxorubicin-loadedliposomalformulations.
(A) in vivo tumor remediation of 143B xenografts in NCR nude mice for LbL-targeted doxorubicin-loaded liposomal NPs, against untreated and uncoated
doxorubicin-liposomal NPs. Untreated control xenografts were sacrificed 19 days post-inoculation of xenograft due to tumor burden exceeding 1 cm.
Untargeted and targeted Dox-liposomal formulations were sacked at 30 days post-inoculationfollowing a dose-escalation study with the following
treatment regimen: day 10 at 1 mg/kg, day 17 at 2 mg/kg, day 24 at 3 mg/kg. Terminal point shown with final tumor area displayed in (B). Statistics are
from an unpaired t-test, two-tailed to compare the untargeted and targetedformulations at each time point. Data presented as mean +/- SEM; n = 4. (C)
Caliper measurementsfor in vivo tumor remediation with repeateddosing of 5 mg/kg for both targeted and untargetedformulations at 22 days, 28 days,
and 35 days post-inoculation (displayed as day 0, day 6, and day 13 in (C)). Statisticsarefrom an unpairedt-test, two-tailed to compare the untargetedand
targetedformulations; *p < 0.05; **p < 0.01; ***p < 0.001. Data presented as mean +/- SEM; n = 4. Resection of tumorsfrom terminal point of (C), n = 4for
each group, displayedfrom thefinal caliper measurement.
65
Figure 3.7. Representative histologicalsections of 143B tumorsfollowing treatment with doxorubicin-loadedliposomalformulations.
Tumors were harvested at 18 days post-treatment correspondingto the terminal point in Figure 3.6.C. (left column) tumors prior to treatment; (right
column) tumors at the 18 day terminal point post treatment of 3 repeated injections at 5 mg/kg doxorubicin. (top row) tumors treated with PAAAlendronate coated doxorubicin-loaded liposomes, (middle row) uncoated dox-loaded liposomes, and (bottom row) untreated animals. Masson's
trichrome stain - red = 143B osteosarcomacells, blue = connective tissue; scale bars representativeof 100 yMm.
66
References
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68
A
4
r
I
,
"T -T*1QI
Adwo",
Chapter 4. Scalable manufacture of built-to-order
nanomedicine: spray-LbL on PRINT@ nanoparticles
This chapter is in part adapted from:
Stephen W. Morton, Kevin P. Herlihy, Kevin E. Shopsowitz, Zhou J. Deng, Kevin S. Chu,
Charles J. Bowerman, Joseph M. DeSimone*, Paula T. Hammond*. Advanced Materials
25(34): 4706-4713,2013
Introduction
A significant limitation in the design of new nanotechnologies for drug delivery is
the balance between efficacy, safety, and scalability - the three major hurdles to
streamlined approval towards the clinic and, ultimately, adaptation in the pharmaceutical
industry. Oftentimes, these aspects of nanomedicine
are competing, and impose
challenging design requirements on systems synthesized in the laboratory.
Resource-
intensive syntheses or purification procedures, limited material yields, and the difficulty of
precisely controlling particle size, morphology, and composition for increasingly complex
systems are some of the major challenges that continue to prevent the translation of
promising technologies away from the bench-top. For this reason, it is highly desirable to
conceive of simple, scalable, and highly controlled methodologies for the manufacture of
multi-functional nanoparticles that possess the necessary physicochemical characteristics
to be clinically relevant.[7 , 10, 155]
Particle fabrication using the Particle Replication in Non-wetting Templates
(PRINT@) process is a well-established, scalable approach[ 191 for rapidly manufacturing
particles with exquisite control over particle geometry (size, shape) and composition
(cargo, carrier system), important parameters for optimizing cellular engagement and in
vivo pharmacokinetics.[
19 , 156-1591
Through a roll-to-roll process, an elastomeric mold is
implemented in a high-throughput fashion to generate monodisperse particles with welldefined geometries. By varying the properties of the elastomeric mold, particles can be
produced with a dynamic range of shapes and sizes, varying from 10 nm - 200 ltm. Further,
particles prior to recovery are presented in an ideal, ordered spatial arrangement that
presents
an
attractive
opportunity
for high-throughput
surface
functionalization
methodologies, such as LbL. The ability to fabricate precisely controlled nanotechnologies
70
on a large scale is highly attractive towards expedient approval through regulatory
pathways.[ 1 2 ,160]
Purification and surface functionalization of PRINT@ particles, however, endures all
of the challenges faced by other synthesis techniques.
Following particle fabrication,
purification from the solubilized transfer adhesive film is limited by traditional means (e.g.
ultracentrifugation, tangential-flow filtration). Post-purification modification is generally
required to achieve effective surface modification using the more traditional techniques
available to the biomaterials community at-large for yielding functional particle systems.
Oftentimes, these bioconjugation or functionalization
chemistries are difficult and
expensive to scale and result in significant material loss. The ability to control surface
characteristics is essential to improve performance of nanotechnologies regarding
enhanced specific molecular interactions with target cells, as well as avoid non-specific
clearance in vivo via protein-resistive properties.[1 2 8] Surfaces largely mediate interactions
at the interface of biology; therefore, the desire to incorporate a technology with the ability
to include a library of materials on the surface of PRINT@ particles in an analogous high
throughput, scalable fashion is highly desirable, and one that holds much promise for the
future of multi-functional nanotechnology manufacture.
Towards the development of a complementary, scalable approach for the surface
functionalization of roll-to-roll assembled functional nanoparticles, we demonstrate that
spray-assisted Layer-by-Layer (Spray-LbL) deposition may be used to generate highlycontrolled functional coatings in a rapid, reproducible, and facile manner on a generalizable
,
platform of particle systems, as illustrated in Scheme 4.1. Spray-LbL has been shown[1 61
162] to conformally coat materials on the nanoscale in a controllable fashion, with extremely
thin layers deposited on the surface of various charged substrate materials, including 3D
structures such as electrospun mats. The combination of PRINT@ and LbL technologies
offers an expansive toolbox for producing functional particles for a variety of applications,
including catalysis, microelectronics,
nanomedicine.[ 13 ,
1631
photovoltaics,
and
cosmetics, in addition to
Bringing PRINT@ technology and Spray-LbL functionalization
together, this work demonstrates a scalable, reproducible approach for fabrication of
functional carriers with exquisite control over particle composition, geometry, and surface
71
properties, providing an exciting platform for large-scale manufacture of highly-controlled
multi-functional nanocarriers.
Materials and Methods
All materials and equipment necessary for PRINT@ particle fabrication were maintained in
a clean room under controlled conditions at UNC-Chapel Hill. Bulk material, including rolls
of poly(ethylene terephthalate) [PET], poly(vinyl alcohol) [PVA]-coated PET, and pre-made
molds cast from a perfluoropolyether (PFPE) material, were provided by Liquidia
Technologies. All chemicals used were provided by Sigma Aldrich, except for hyaluronic
acid, provided by Lifecore Biomedical, and Cy5.5-NHS from Lumiprobe.
PRINT@ Nanoparticle Fabrication. PRINT® particle fabrication follows a well-established
protocol.[1 5 61 Briefly, the polymeric material, poly(lactide co-glycolic acid) [PLGA], used for
particle molding was dissolved in dichloromethane. A film of PLGA was cast using a mayer
rod (#3) on a high energy PET backing prior to lamination of this film to the pre-cast mold
at 3200F. The filled mold from this step was laminated at 320OF to a transfer adhesive film
(PVA) cast on a PET backing. The particles were preserved in this state via vacuum seal
under N 2 atmosphere until ready for functionalization.
Vapor-phase Glutaraldehyde Crosslinking. Molded nanoparticles were delaminated from
the mold onto the transfer adhesive film. Sections of these harvested particles were
subjected to PVA-crosslinking conditions in an enclosed chamber via vapor-phase
glutaraldehyde crosslinking to further immobilize the particles for spray-LbL deposition.
This was done via incubation of the particle arrays with vials containing 50%
glutaraldehyde (in water) and 10% aqueous acid (HCl) for 15h. Arrays following
(cosslinkling(T WATV-e
~ ~.~.I1Li~11
5
VV%-L% .IkIIJVI.~
iree
Spray Layer-by-Layer
ifor susqn
unct-ialfi
n.
~J1
3.L4LJ.k-JLA%Lt1L ltuflCLiJ1aII11/aLJI.
Deposition.
Materials were deposited onto the surface of
crosslinked nanoparticle arrays by aerosolization of polyelectrolytes for spray times of -3
seconds, with -3 second wash steps between each layer. Polyelectrolytes were sprayed at
72
a concentration of 1mg/mL. A final water rinse was used prior to drying and subsequent
analysis and/or recovery of functionalized particles. Particles recovered following sprayLbL were harvested by sonication of the functionalized particle arrays in water for
-15minutes.
The collected particles were subsequently purified via centrifugation and
filtration through a 0.45 pm syringe filter.
Nanoparticle Characterization. Images to assess iridescence of the nanoparticle arrays
following each processing step were obtained by photography with a 5-megapixel camera
(iPhone
4).
Contact
angle
measurements
and
images
(Rame-Hart
model
500
goniometer with Nikon camera) were obtained immediately following drop-casting a
bubble of water on the NP array. Dynamic light scattering (Malvern ZS90) was used to
determine the hydrodynamic diameter of the nanoparticles. Zeta potential measurements
were also performed using the Malvern ZS90. Measurements were conducted in 10mM
NaCl in millipore water at 250C.
Scanning electron micrographs were collected using a JEOL 6700 high resolution
microscope. Sample preparation included drop-casting on a silicon wafer followed by
sputter coating with gold-palladium (-3nm). Transmission electron micrographs were
collected using a JEOL 2010 Advanced High Performance TEM. Sample preparation
included drop-casting on a carbon/formvar-coated copper grid. Atomic force microscopy
data was collected using a MultiModeTM atomic force microscope with a NSC15/AIBS,
325kHz, 46N/m tip from imasch in tapping mode.
Confocal microscopy (Nikon AiR scanning confocal microscope) was conducted
following incubation of the functionalized nanoparticles (tracked via labeled PLL, PLLcys.5)
at 370C with BT-20 cells. Fixed samples included a DAPI and Phalloidin-488 stain, in
addition to particle tracking Cy5.5 label. Flow cytometry (BD LSRFortessa) was performed
in parallel with confocal microscopy to quantitatively assess particle association (tracking
via PLLcy5.s) with cells following incubation at 370C. Particles administered were
normalized against the fluorescent intensity (Aex=675nm, Xem = 710nm), as measured by a
Tecan Microplate Reader, of the labeled polycationic layer deposited initially on the NP
surface.
73
Results and Discussion
Using the widely-reported top-down approach of PRINT@ particle fabrication,
200x200x200nm PLGA nanoparticles (PLGA 200x2OOnm) were fabricated and collected on a
low molecular weight polyvinyl alcohol (PVA) base coating atop a polyester support film.
While immobilized on the support film, the particles are arranged in a highly ordered 2-D
hexagonal array. The PVA supported particles were visualized by AFM and SEM. The
regular spatial arrangement of PRINT@ particles coupled with the negative surface charge
of the acid-terminated PLGA NPs in water make them ideal candidates for rapid LbL
polyelectrolyte functionalization through spray-assisted LbL deposition (Spray-LbL). Initial
attempts to deposit aqueous solutions of polyelectrolytes onto the supported PRINT
particles resulted in a loss of the particles due to detachment from the substrate.
This
process could be visualized by the loss of diffraction-based iridescence characteristic of the
ordered nanoparticle arrays (Figure 4.1.B, 4.1.C); this characteristic iridescence is
indicative of highly ordered particles on the harvesting layer, allowing for a convenient
qualitative assessment of particle immobilization or detachment. The loss of particles was
confirmed with AFM and SEM, which essentially show a complete absence of particles after
the Spray-LbL process due to the water solubility of the PVA backing.
The reason for
choosing PVA as the support material for PRINT particles is in fact to enable facile particle
harvesting (see Appendix C.1).
To
avoid
PVA
dissolution,
vapor-phase
crosslinking
using
concentrated
glutaraldehyde and acid was employed to selectively crosslink the PVA adhesion layer to
reduce its water solubility, while avoiding any chemical changes to the relatively
impermeable and inert PLGA PRINT@ nanoparticles. Contact angle measurements were
carried out to study the effects of the vapor-phase cross-linking reaction on the wettability
of the PRINT@ films. As shown in the Figure 4.1.A, pre-crosslinked NP arrays show
()nidbl4
w"t tLabilky,
1wIng t
LLthp 1rL
%Jf1ily VVatLeLr-sLUU aniu 11yr1UpIJ11111
PVA polymer. After cross-linking, the contact angle increased from 260 to 780 (Figure
4.1.D), demonstrating that the films become considerably more hydrophobic due to the
formation of a glutaraldehyde cross-linked PVA adhesive. Crosslinking conditions were also
varied with results shown in Appendix C.2. As a result, the cross-linked films can be placed
74
in water without particle detachment. Following crosslinking, it was observed that
particles could be readily functionalized, as evidenced by maintenance of iridescence from
the NP array following each processing step (Figure 4.1.E, 4.1.F).
This step therefore
allowed for the deposition of water-based polyelectrolytes using Spray-LbL.
As shown in Scheme 4.1, Spray-LbL on PRINT@ nanoparticles, following vaporphase cross-linking of the PVA support films, is achieved by spraying an aqueous solution
of cationic polyelectrolyte is sprayed onto the NP containing films for 3 seconds and, after
briefly rinsing with water (3 seconds), an anionic polyelectrolyte is subsequently sprayed
onto the films (3 seconds). This cycle can then be repeated indefinitely to control the
thickness of the polyelectrolyte coating. In the final step, the particles are harvested by
sonicating the films in water, which causes the particles to detach from the PVA substrate.
Using AFM, the shape, persistence and uniformity of the particles were monitored
with each processing step, beginning with the initial NP array, followed by crosslinking,
and subsequent Spray-LbL deposition, as shown in Figure 4.2 (corresponding amplitude
images shown in Appendix C.3). The versatility of this platform was further demonstrated
by coating a second PRINT@ PLGA particle type, 80x80x320nm [PLGA80
x 3 2 0 nm].
AFM
confirms that the particles are coated in a manner that maintains the uniformity of the
PRINT@ particle template while building a film, observed by slight increase in height
(-5nm/bilayer), which avoids bridging of neighboring particles and without causing
significant loss of particles during functionalization.
Particles were subsequently recovered via sonication, purified via filtration, and
concentrated using ultracentrifugation following Spray-LbL. A variety of techniques were
then used to illustrate the uniformity of the particles obtained from this fabrication
method. Dynamic light scattering analysis is shown in Table 4.1, with the corresponding
histograms shown in Appendix C.4. For both particle types, uniformity of the particle
population is maintained as evidenced by the PDI of functionalized particles (see Table
4.1: PRINT@ PLGA20 0 x2Oonm particles - original 0.01, 0.06 following deposition of three
bilayers of PLL/HAsOOK; PRINT@ PLGA80x320nm particles - original 0.05). A slight
decrease in size following crosslinking is observed for both particle types, consistent with
the amount of material exposure to the crosslinking contraction forces on the PVA adhesive
layer. The change in size for functionalized particles is indicative of a very thin LbL film,
75
approximately 15-20nm for 3 bilayers of PLL/HA500K, based on dynamic light scattering
(Table 4.1) and confirmed with electron microscopy (Figures 4.3, 4.4). Histogram
overlays also illustrate the thin coatings applied to the particle surface (Appendix C.4).
Characteristic surface charge reversal of the LbL deposition is also observed, as evidenced
by samples taken at different steps in the functionalization process. c-potential, measured
at 250C in 10mM NaCl, of pre- and post-XL NPs are approximately the same, while
deposition of materials post-XL results in surface charge characteristics consistent with the
material deposited (see Appendix C.4). A range of polyelectrolyte multilayers were
explored, including poly-L-lysine (PLL)/hyaluronic acid (HA) [PLL/HAOOK], PLL/dextran
sulfate (DXS) [PLL/DXS], PLL/poly(acrylic acid) [PLL/PAA], chitosan/HA [Chit/HA5 00 K],
and chitosan/PAA [Chit/PAA] (see Appendix C.4), further illustrating a primary advantage
of LbL as a means of tailoring the surface properties of these NP systems. These films, while
diverse in the range of materials incorporated, behaved similarly in coating the NP systems
without compromising the particle shape or significantly increasing the size beyond a few
tens of nanometers.
Recovered and purified particles visualized by electron microscopy, including SEM
and TEM, are displayed in Figures 4.3 and 4.4. Figures 4.3.A and 4.3.B are representative
of the uncoated, crosslinked PRINT@ PLGA2 00 x 20 0 nm particle array, along with the
corresponding recovered particles observed by both SEM (Figure 4.3.C) and TEM (Figure
4.4.13). The coated array is shown in Figure 4.3.D and 4.3.E for direct comparison. A thin
film coating is observed on the particles in such a fashion that individual particle integrity
is maintained prior to recovery. Subsequent recovery yields the results in Figure 4.3.F,
where a thin coating of approximately 20nm is observed to surround the entire particle
surface by TEM (Figure 4.4.B). Observation of a conformal coating surrounding the entire
nanoparticle is a surprising result due to the inaccessibility of the spray deposition to the
immobilized surface; however; it can be explained by the unique feature of the electrostatic
self-assembly LbL process to "self-heal" defects in the adsorbed film.[1 64 1 The excess film
that corrects for this defect in spray LbL on PRINT@ is observed in Figure 4.3.E, whereby
the film surrounding the individual nanoparticle collapses in a manner that completely and
smoothly seals the nanoparticle with deposited film around the harvested particle
following suspension. This phenomenon is also observed for the
76
PLGA8Onmx320nm
NPs.
Displayed in Figure 4.3.G and 4.3.H are the uncoated, crosslinked PLGA8 Ox320nm NP arrays,
followed by the recovered particles visualized by both SEM (Figure 4.3.J) and TEM (Figure
4.4.E). Compared to the coated NP arrays in Figure 4.3.J and 4.3.K, the particles clearly
grow in size in both dimensions following film deposition. This thin film is also observed
following recovery in using both SEM (Figure 4.3.L) and TEM (Figure 4.4.E), where a
coating of 20nm is observed around the entire particle. From these data, it is clear that we
can conformally coat PRINT@ PLGA NPs with films around the entire particle surface and
in a manner that preserves the monodispersity of the platform.
After demonstrating the feasibility of spray functionalization and observing a
conformal coating surrounding two different PRINT@ NP templates with a variety of
coating materials, it was important to examine the ability to control film growth using
Spray-LbL. Control over film thickness will allow precise tuning of drug release for
therapeutic-containing NP systems as well as functionalization of these material systems in
such a way that does not significantly impact the NP size or shape, which is a primary
advantage of PRINT@ manufacture. From Figures 4.3, 4.4.B, 4.4.E, it was observed that
three bilayers of PLL/HA (3 seconds/spray) generate a film of approximately 15nm. To
observe whether thicker films could be generated, high molecular weight polymers
(Chitosan, 200K; HA, 500K) were sprayed (3 seconds/spray) alternately to yield a 5 bilayer
film. As expected, this film was found to be significantly thicker, approximately 40nm by
TEM (Figure 4.4.C, 4.4.F), while conformally coating the NP surface so as to not
compromise the NP shape. SEMs of the coated backings and subsequent harvested and
purified coated NPs for each PRINT@ template are displayed in Appendix C.5. TEMs of
each corresponding uncoated NP systems are displayed in Figure 4.4.A and 4.4.D for
reference. This effectively demonstrates the capability of Spray-LbL as a means of tuning
film thickness by incorporating higher molecular weight polymers with more bilayer
coatings. Additional parameters to tune film thickness are longer spray times and higher
concentrations of material sprayed.
As an additional test of particle functionalization, the polycationic layer, poly-Llysine (PLL), was labeled with a Cy5.5 dye (Lumiprobe, Cy5.5-NHS ester). Deposition was
tracked using fluorescence imaging and shown to significantly increase following layers 1
and 3, consistent with the subsequent addition of PLLcys.s. This is also demonstrated by
77
directly comparing PLL/DXS films cast both with and without PLLCy5.5 using confocal
microscopy (Appendix C.6).
While particle functionalization is clearly demonstrated and in a reproducible
manner, questions regarding biological functionality remained as this component of the
carrier system is critical to the translational relevance of the technology. Using the labeled
PLLcys.s Spray-LbL functionalized NP systems, particles were incubated with a triple
negative breast cancer cell line, BT-20. These cells, like many aggressive cancer cell types,
characteristically overexpress CD44 receptors[22], which are a convenient target for the
natural ligand, hyaluronic acid.[165 - 1 6 81 Previous work has also shown much promise for HA-
coated
systems
to provide a
serum-stable,
stealth-like
platform
for delivery.[2 7]
Investigations of cell-associated fluorescence in combination with confocal microscopy,
shown in Figure 4.5.A-4.5.D, confirmed that these coatings are functional, as shown with
enhanced levels of cell uptake for the HA-coated systems. Further, increased uptake for HAcoated systems is clear relative to the non-specific uptake of the DXS-terminated system,
suggesting that film construction can be finely tuned towards targeting specific cell
populations. In this way, Spray-LbL on PRINT@ presents exciting opportunities to not only
functionalize particles scalably, reproducibly, and with extremely thin coatings but also in a
manner that molecular engagement of target cells can be achieved.
Particle shape also plays a significant role in cellular internalization, as shown in
Figure 4.5.E-4.5.G. Coated PLGAsox 3 20nm NPs ([PLL/HA] 3 ) exhibited nearly 10-fold higher
levels of cell association relative to identically-coated PLGA 200x2 OOnm NPs within 2h of
treatment, as determined by FACS. This enhanced level of uptake is further complemented
by confocal microscopy, from which significantly more punctate NP fluorescence is
observed inside the cells for the PLGA8Ox32nm NPs. The impact of particle size and shape on
cellular internalization is of particular interest towards designing effective carrier systems
that further mediate endocytosis, in addition to surface chemical functionality. This nroves
highly promising towards combination of optimal PRINT@ technologies with targetspecific Spray-LbL architectures to maximize therapeutic impact of these NP systems.
78
Conclusion
Spray-LbL on PRINT@ is a significant advance in the design of nanotechnology, as it
provides an exciting platform for large-scale production of built-to-order functional
nanoparticle
systems, whereby we have precise control
of the
physicochemical
characteristics of the NP systems developed. This includes the ability to finely tune the NP
shape, size, chemical composition, and surface characteristics towards manufacture of
systems that maximize cellular entry and optimize drug and NP pharmacokinetics in vivo.
Further, the ability to rapidly and scalably manufacture these systems is realizable, as
PRINT@ technology is capable of roll-to-roll NP fabrication while spray can be applied in a
continuous fashion.
This compatibility in continuous manufacture and subsequent
functionalization of NP systems presents a highly attractive opportunity for highly
reproducible manufacture of customizable NP systems for a wide variety of applications.
79
Scheme 4.1. Spray-LbL on PRINT@ nanoparticles.
PRINT@ particleswerefabricatedand used while immobilized in a post-harvestingarrayfor Spray-LbL. Arrays were subsequently crosslinked under vaporphase glutaraldehyde/concentratedacid conditions, followed by LbL application (sequential deposition of polycation/wash/polyanion/washcomprising
one bilayer). Functionalized particles were harvested by sonication of the arrays in water and purified by filtration and ultracentrifugation.
80
U
260 d
780
.
Figure4.1. Contact angle and particleiridescence characterizationpre- and post-functionalization.
(a) Contact angle measurement and image of pre-crosslinked NP array. (b) Particleiridescence observed in the
ordered particle array post-PRINT@ fabrication. (c) Loss of iridescence following spray of water-based
polyelectrolytes, indicative of complete loss of particles. (d) Contact angle measurement and image of postcrosslinked NP array (glutaraldehyde/HCI vapor phase crosslinking). Particle iridescence is maintained
following (d) 15h crosslinking and (f) Spray-LbL deposition of [PLL/HA]3
81
PLGA.
PLGARnv
C
M
51Lm
5RM200nm
In
.j
Ln'
gm
PM
Figure 4.2. AFM characterization(height) of pre- and post-functionalized NP arrays.
(a) Pre-crosslinked PLGA2oox 2oonm NP array; (b) PLGA 20ox 2 oonm NP array following crosslinking; (c)
PLGA 200x2 00nm/(PLL/HASOOK) 3. (d) Pre-crosslinked PLGA8 ox3 2 onm NP array; (e) PLGA8 ox 3 2onm NP arrayfollowing
crosslinking; ) PLGA8 Ox 32 Onm/(PLL/HA5OOK)3. Scale bars representative of 2pm. Color scale representative of
lineargradientfrom 0 [bottom] to 200 nm [top].
82
Figure 4.3. SEM characterizationof particlespre- and post-Spray-LbLfunctionalization.
SEM of (ab) uncoated, crosslinked 200x200nm PLGA NP array; (c) purified uncoated, crosslinked 200x200nm
PLGA NPs; (de) coated 200x200nm PLGA NP array [(PLL/HA)3]; (f) purified coated 200x200nm PLGA NPs
[(PLL/HA)3; (g,h) uncoated, crosslinked 80x320nm PLGA NP array;(i) purified uncoated, crosslinked 80x320nm
PLGA NPs; (jk) coated 80x320nm PLGA NP array [(PLL/HA)3J; (1) purified coated 80x320nm PLGA NPs
[(PLL/HA)3]. Scale bars representativeof 200nm.
83
Figure 4.4. TEM characterizationof particlespre- and post-LbLfunctionalization- con trolfilm thickness.
TEM of (a) uncoated, crosslinked 200x200nm PLGA NPs; (b) purified coated 200x200nm PLGA NPs [(PLL/HA)3 ];
(c) purified coated 200x200nm PLGA NPs [(Chit/HA)s]; (d) uncoated, crosslinked 80x320nm PLGA NPs; (e)
purified coated 80x320nm PLGA NPs [(PLL/HA)3]; (f) purified coated 80x320nm PLGA NPs [(Chit/HA)s]. Scale
bars representativeof 1 00nm. TEMs of uncoated, uncrosslinked NPs are displayed in Appendix C.1 0.
84
d
1000
I
I
I
I
800-
g
I
3fl ',.
-'
I
-*- 200x200nm
I
*- 80x320nm-
-o
S20W0-
600LA
LA
Ln
LA
>ooo
400
2000
0
control
L2 HA
L6 HA
L6 DXS
1
2
3
concentration (X)
Figure 4.5. in vitro assessment of variablespray LbL coatings and PRINT@ NP shape on cell uptake - tracking with PLLcy5s..
Confocal microscopy of (a) PLGA 20 0x 2 0 0nm/PLL/HA5OOK NPs [L2 HA], (b) PLGA 2o0 x2 00n,/(PLL/HASOOK)3 NPs [L6 HA], (c) PLGA 2oox2 oonm/PLL/DXS NPs [L6 DXS]
incubated with BT-20 cellsfor 6h at 370C. NPs labeled with PLLcys.5 asfirstpolycationic component in film and normalized in dose based on the absorbance
of this layer on the function alized NPs. (d) Mean cell-associatedfluorescence [Cy5.5 channel (tex = 640nm, Aem = 700nm)]for each NPformulation displayed
in (a)-(c). Data presented as the average +/- SEM of triplicate measurements. Confocal microscopy of (e) PLGA8Ox32 nm/(PLL/HAooK)3 NPs and (f)
85
PLGA 200 xzoonm/(PLL/HA5OOK) 3 NPs incubated with BT-20 cellsfor 2h at 370C. NPs labeled with PLLcys.5 asfirstpolycationic component in film and normalized
in dose based on the absorbanceof this layer on thefunctionalized NPs. (g) Mean cell-associatedfluorescence[Cy5.5 channel (Aex = 640nm, Aem = 700nm)]for
each NP formulation displayed in (e)-(f). Data presented as average +/- SEM of triplicate measurements. Confocal microscopy images representative of
overlays of individualfluorescencechannels displayed below ifrom left to right under each overlay: DAPI nuclearstain, phalloidin-488stain, NPfluorescence
- Cy5.5].
86
237 +/- 2
229 +/- 3
244 +/- 2
221+/- 1
192 +/- 2
214 +/- 5
0.01
0.03
0.06
0.06
0.1
0.05
-6 +/- 0.1
-8 +/- 2
-28 +/- 1
-10 +/- 2
-13 +/- 0.5
-24+/- 2
Table 4.1. Nanoparticlephysicochemical characterization.
Dynamic light scattering data, based on z-average hydrodynamic diameter, and zeta-potential measurements were conducted at 250C in 10mM
NaCl/deionized water. Data is presented for purified, uncoated
PLGA200xOOnm
and
PLGA8OX32Onm
NPs, purified PLGA200x200nm and
PLGA8Ox32Onm
following
crosslinking, and coated NPs: PLGA 200 x2 onm/(PLL/HA5 OOK) 3 and PLGAx
80 32 0 nm/(PLL/HA500K)3. Data for mean hydrodynamic diameter and zeta-potential are
presented as average +/-standarderrorof the mean for triplicatemeasurements.
87
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88
I
Chapter 5. Liposomal compartmentalization of a hydrophobichydrophilic combination chemotherapy regimen (erlotinib,
doxorubicin) for staged, targeted delivery against triplenegative breast cancer and non-small cell lung cancer.
This chapter is in part adapted from:
Stephen W. Morton*, Michael J. Lee*, Zhou J. Deng, Erik C. Dreaden, Elise Siouve,
Kevin E. Shopsowitz, Nisarg J. Shah, Michael B. Yaffe*, Paula T. Hammond*. Science
Signaling 7(325): ra44, 2014.
Introduction
Cancers represent the end states of accumulated genetic transformations that
disrupt normal cell signaling events, including those involved in DNA repair, cell cycle
regulation, and cell death by apoptosis
[169, 1701,
enabling these mutant cells to proliferate
and metastasize. Paradoxically, although defects in DNA-damage signaling and response
underlie tumor development, they also provide a mechanism for therapeutic tumor cell
killing by conventional anticancer cytotoxic therapies, such as chemotherapy or radiation
therapy. Undesired effects of these DNA-damaging treatments include the development of
highly resistant residual tumors and toxicity to proliferative non-tumor tissues, such as the
bone marrow and epithelial lining of the gastrointestinal tract. Therefore, there is an
important clinical need to identify potent therapeutic strategies that target multiple tumor
cell-specific survival pathways to enhance the extent of tumor cell killing and potentially
reduce total drug exposure during treatment. Most drug-screening efforts have focused on
careful selection of drug cocktails on the basis of the underlying biology of the tumor or the
response to individual agents. Unfortunately, much less is known about the positive and
negative drug-drug interactions for many combination therapies, and the influence of
relative dose, dose duration, or timing of delivery has been much less frequently
explored[ 169 ,170].
Our recent work validated the impact of timing of drug delivery on the efficacy of
multiagent chemotherapy: A pair of drugs that are not particularly beneficial as singular
therapies, are effective when used in combination if a specific time lag between the
administration of each drug is employed
[30,95-97].
90
Thus, systematic study of the adaptive
responses of signaling networks in tumor cells following an initial drug exposure could be
used to design highly effective drug combinations[3 0 ].
We recently demonstrated that triple negative breast cancer (TNBC) cells and nonsmall cell lung cancer (NSCLC) cells are dramatically sensitized to the effects of DNAdamaging chemotherapy by prolonged, but not acute, suppression of epidermal growth
factor receptor (EGFR) signaling
[30].
This sensitization effect resulted from the rewiring of
signaling networks upon persistent EGFR inhibition, which unmasked a caspase-8dependent cell death pathway critical to the ability of doxorubicin and other genotoxins to
more effectively kill tumor cells in culture
[30].
This work demonstrated the importance of
drug order and timing for maximizing synergistic effects of combination chemotherapy for
cancer. However, translating these new findings into the clinic with existing delivery
methods
is
likely
to
prove
challenging
due
to
patient-specific
differences
in
pharmacokinetics, the differing pharmacodynamic parameters for each drug, and the
difficulties in targeting both drugs to the same tumor cells in the proper temporal
sequence.
To overcome these challenges, we created a robust nanoparticle-based delivery
platform capable of producing a precise time-staggered drug release in vivo. Nanoparticles
are colloidal material systems commonly comprised of organic (such as lipids or polymers)
or inorganic (such as silica, iron, or gold) materials and are generally 200 nm or less in size.
These structures are commonly employed as vectors for controlled drug delivery by
containing and protecting the therapeutics from metabolism or destruction. Using this
approach, we designed a single nanoparticle construct to serve as a dual hydrophobichydrophilic depot for timed sequence of both a EGFR inhibitor and a DNA-damaging agent.
This technology facilitated delivery into mice, achieving intracellular co-localization of both
drugs and time-sequenced delivery of the synergistic drug combinations. Such an approach,
if successful in patients, should mitigate off-target side effects and increase the
bioavailability of the drugs, thereby expanding the therapeutic efficacy of selected
chemotherapy combinations identified by ongoing systems pharmacology and signalingbased studies.
91
Materials and Methods
All lipid components were purchased from Avanti Polar Lipids, except for cholesterol from
Sigma. All therapeutics were purchased from LC laboratories, except for cisplatin (cisdiamineplatinum(II) chloride) from Sigma. All other chemicals (citric acid, sodium citrate,
sodium carbonate) and solvents (chloroform, methanol, phosphate buffered saline) were
purchased from Sigma.
Study Design. Erlotinib and doxorubicin as a combination therapeutic treatment have
been previously shown to enable synergistic cell death in certain cell types when
administration of the therapeutics is manually time-staggered[ 31 . We sought to achieve this
temporal control over release in a systemically-administrable nanoparticle formulation.
For efficacy evaluation of these systems, power analysis was performed with G*Power
Analysis, using repeated-measures ANOVA, between-factors test. We assumed an effect size
of 0.5, error probability of 0.05, power of 0.95, and a correlation of 0.2. To achieve
statistical significance, 15 animals bearing dual rear xenografts were used - 5 in each
experimental group. Endpoints were pre-determined for evaluation of the biological
performance of these systems, as well as for those experiments investigating tumor burden
of the animals following treatment. All experiments were randomized and non-blinded.
Liposome
preparation. Liposomes were formulated at a mass ratio of 56:39:5
(DSPC:Cholesterol:POPG). These three components were dissolved, along with the smallmolecule inhibitor (weight ratio to total lipid weight = 3:50), in a 2:1 mixture of
chloroform:methanol. A thin film of these materials was generated by rotary evaporation at
400C at 150 mbar. This film was dessicated overnight until completely dry. Hydration of the
lipid film was conducted at 650C under sonication in 300 mM citric acid buffer (pH 4) for
1h, Functionalization with DSPE-PEG5K-Folate, DSPE-PEG2K, with or without DSPE-PEG 2 KCy5.5 was conducted post-fabrication using a post-insertion technique in which micelles
(in 0.9% sodium chloride solution) of the components desired to incorporate on the
liposomal surface were incubated with the prepared, drug-loaded liposomes under
sonication at 650C for 30 minutes, after which they were filtered through a 0.2 ptm
polyethersulfone (PES) syringe filter and cooled to room temperature. The pH of the
92
liposomal suspension was then adjusted to -6.5 by addition of 300 mM sodium carbonate
buffer to create a gradient between the exterior and interior compartments. The cytotoxic
(doxorubicin - 3 mg, cisplatin - 10 mg) was added in a 0.9% sodium chloride solution (1
mL) to load through a pH gradient method. To facilitate solubilization of the cytotoxic
agents, the dispersed solution was sonicated at 650C for 5 minutes. The final combination
drug-loaded system was subsequently exchanged into PBS (pH 7.4) following centrifugal
filtration (100K MWCO Millipore) to remove the high salt buffers (citric acid, sodium
carbonate) and any unloaded drug. Empty (lacking drug) liposomes (NFP) were formulated
using the same procedure.
Liposome characterization. Dynamic light scattering and z-potential analysis was
conducted in 10 mM sodium chloride at 250C using a Malvern ZS90 zeta-sizer. Highperformance liquid chromatography (Agilent technologies) and nanodrop absorbance
measurements (345 nm for inhibitor; 480 nm for doxorubicin) were used to validate drug
loading of the inhibitor
(Xabs
= 345nm) and doxorubicin
(Xabs
= 480nm). Cisplatin
concentration was quantified by a colorimetric assay using o-phenylenediamine against a
standard curve
[1711.
Cryo-transmission electron microscopy (cryo-TEM) was conducted by
imaging a vitrified dilute sample of the liposomal suspension at 120 kV and 77 K.
Nanoparticle Drug Release In Vitro. Liposomes were incubated under sink conditions [1
L sink for 1 mL liposome suspension] in 1X PBS under agitation in 1 mL 3.5 K MWCO floata-lyzers (Spectrum) at 370C. PBS was replenished each day of the experiment. Samples
were taken of the liposomes to quantify remaining drug concentrations by HPLC (following
dissolution in a 50:50 mixture of acetonitrile:pH 5 water) and absorbance measurements
(nanodrop) [345 nm for erlotinib, lapatinib, /gefitinib,
or /afatinib;
480 nm for
doxorubicin].
Confocal Microscopy. Images were taken using a Nikon A1R Ultra-Fast Spectral Scanning
Confocal Microscope (Nikon instruments Inc., Melville, NY). BT-20 and A549 cells were
seeded in CELLview glass-bottom dishes (Greiner Bio-One GmbH, Germany) at 1 x 10s cells
per well and grown overnight in Opti-MEM (Gibco, Life Technologies). For data displayed
93
in Fig. 2, cells were then incubated with empty (no drug, NP-Cy5.5, NFP-Cy5.5) Cy5.5labeled liposomes for 1 h at 37oC. At the end of this period, cells were washed, fixed with
paraformaldehyde, permeabilized with Triton-1X, and stained with phalloidin-568 for 30
minutes, followed by the addition of DAPI for an additional 10 minutes, after which they
were washed (with Opti-MEM) and imaged (in Opti-MEM) at the DAPI (excitation 360 nm,
emission 460 nm) and Cy5.5 channel (excitation 640 nm, emission 700 nm). For data
displayed in fig. S8, live-cell imaging was performed by exposing A549 cells to DFP-Cy5.5 or
free doxorubicin in Opti-MEM and serially imaging the cells at the DAPI (excitation 360 nm,
emission 460 nm), doxorubicin (excitation 480 nm, emission 560 nm), and Cy5.5
(excitation 640 nm, emission 700 nm) channels.
Flow Cytometry. Measurements were performed using a BD LSR Fortessa-HTS coupled
with a high-throughput system for the 96-well plate format (BD biosciences, San Jose, CA).
Doxorubicin fluorescence was measured following excitation at 488 nm and detected at
530 nm; Cy5.5 fluorescence was measured following excitation at 640 nm and detected at
710 nm. Cell-association data presented as geometric mean fluorescence collected in
triplicate following cell incubation with empty and doxorubicin-loaded Cy5.5-labeled
liposomes for 1h at 370C.
A
2-
Apoptosis.
AX --
VledsUreIIeItILs
_ _--
-
--
__
1
--
J- 1
J
F2-
r
i -1 j
were LU11ULLCU ds previuusIy uesLriued.L-'I Drielly, IUiOuwiing
the treatment time course, cells were washed, trypsinized, fixed in 4% paraformaldehyde
for 15 min at room temperature, resuspended in ice-cold methanol, and incubated
overnight at -20'C. Cells were then washed twice in PBS-Tween and stained with
antibodies against cleaved caspase-3 and poly(ADP-ribose) polymerase (PARP). Secondary
Alexa-conjugated antibodies were used for visualization in a BD FacsCaliber flow
cytometer.
Western Blotting. Experiments were performed as previously described
[301.
Briefly, cell
lysates were prepared in a manner that enabled samples to be used for both Western blot
analysis and reverse phase protein microarray. Cells were washed twice in PBS and lysed
directly on the plate in a buffer containing 50 mM Tris-HCl, 2% SDS, 5% glycerol, 5 mM
94
EDTA,
1 mM
NaF, 10 mM, j-glycero-phosphate,
1 mM PMSF,
1 mM Na 3 VO 4 , and
phosphatase and protease inhibitors (Roche complete protease inhibitor tablets and
PhosSTOP tablets). Crude lysates were filtered using an AcroPrep 96-well 3.0 Im glass
fiber/0.2 prm Biolnert filter plate (Pall), and normalized for protein content using the BCA
protein assay (Pierce). For Western blots, lysates were run on 48-well pre-cast gels and
transferred using a semi-dry fast transfer apparatus onto nitrocellulose membranes (iPAGE, E-BLOT, Invitrogen). Blots were blocked in Odyssey Blocking Buffer (LiCOR
Biosciences),
incubated overnight with primary antibody, stained with secondary
antibodies conjugated to an infrared dye, and visualized using an Odyssey flat bed scanner
(LiCOR Biosciences).
Antibodies for gH2AX, pERK, and cleaved caspase-8 were purchased from Cell Signaling
Technologies. Antibodies against b-actin were purchased from Sigma.
Raw signals for each protein of interest were quantified and background subtracted using
the Li-COR Odyssey software and divided by b-actin signals to normalize for loading
differences, and then each normalized signal was divided by a reference sample contained
on each gel for gel-to-gel normalization.
Pharmacokinetics. All animal experimentation adhered to the NIH Guide for the Care and
Use of Laboratory Animals and was in accordance with institutional guidelines.
BALB/c female mice (Taconic) were systemically-administered (tail vein) empty Cy5.5labeled liposomes (NFP-Cy5.5) at a concentration of 3 mg/mL (corresponding radiant
efficiency, measured by the whole-animal imaging system - IVIS [Xenogen, Caliper
Instruments] of 0.1 mL sample injected in each mouse -1
x 1010). Whole-animal
fluorescence imaging (IVIS - Xenogen, Caliper Instruments) was performed at the indicated
time points for one cohort of mice (n = 3), and a separate cohort was used for retro-orbital
bleeds to determine the circulation half-life of the system. Imaging and circulation data
presented is normalized to auto-fluorescence determination (imaged animals, isolated
blood) obtained prior to injection. Analysis of circulation data based on recovered
95
fluorescence normalized to pre-injection blood auto-fluorescence is displayed with a twocompartment model fit (both slow and fast half-lives presented).
Tumor Targeting and Regression. Tumor targeting data was obtained by imaging a
Cy5.5-labeled DSPE-PEG-Cy5.5 lipid that was inserted into the liposomal membrane. This
dye-labeled lipid was functionalized using a DSPE-PEG-NH2 lipid (Nanocs) and Cy5.5-NHS
(Lumiprobe) at a 1:1 ratio in DMSO at 40C for complete labeling of the lipid. Final dyelabeled lipid product was lyophilized and stored as a powder at -200C. Tumor-targeting
data was collected using a near-IR imager (IVIS, Xenogen, Caliper Instruments) at ?ex = 675
nm, Aem = 720 nm. % injected dose (ID) calculations were performed from fluorescence
intensity of the tumors, following background subtraction
of pre-injection tissue
autofluorescence, and normalized to the amount of fluorescent material injected.
Luminescent xenograft tumors were seeded following stable transfection of BT-20 and
A549 cells with the firefly luciferase plasmid. This enabled assessment of tumor size by a
visual and quantifiable luminescent readout generated by whole-animal imaging (Xenogen,
Caliper Instruments). Cells were mixed in a 1:1 ratio with BD MatrigelTM Basement
Membrane Matrix to a final density of 5 x 107 cells/0.1 mL injection. 0.1 mL injections of the
matrix-cell suspension were performed in each rear hind flank of NCR nude mice (Taconic).
Tumors were allowed to grow until a visible tumor was established and luminescence
LitoringLULIII
WdS
1w1UaidILiVU
fU11U
LUmIIUr
g1VVL1
ee
fig. 1 3, ~7 A I' lraUldilLCe linal
luminescent readout from each xenograft). Xenograft-bearing NCR nude mice were then
systemically-administered therapy (DFP, DEFP) in a 0.1 mL injection at a concentration of 2
mg/kg (based on doxorubicin loading) and monitored for 32 days. Luminescence images
were obtained by 0.1 mL intraperitoneal injections of 30 mg/kg D-luciferin (Caliper) and
imaging (IVIS, Xenogen, Caliper) with an open luminescence filter 15 minutes postinjection. These data are presented, along with region of interest quantification of radiance
corresponding to the xenogratt-specitic luminescence tor each mouse treated. Data is
normalized for each mouse against the tumor luminescence prior to injection and
presented as fold luminescence above this measurement. N = 5 for each treatment
(untreated, DFP, DEFP) and data corresponds to mean +/- SEM for each treatment group.
An unpaired, two-tailed t-test comparing the single drug (DFP) and combination drug
96
(DEFP) systems was performed to determine statistical significance. A control simulating
the simultaneous release of both doxorubicin and erlotinib was performed in vivo by the
systemic co-administration of DFP liposomes (1 mg/kg) and an erlotinib-cyclodextrin
complex (CD-Erl) in PBS (1 mg/kg, hydroxypropyl-beta-cyclodextrin)
xenografts were established.
22 days after
Subsequent booster doses of erlotinib were systemically
administered at 1 mg/kg on days 2 and 4. Live-animal bioluminescence images of hindflank
A549 xenografts shown prior to treatment (pre) and at days 3, 10, and 30 were collected.
Quantification of luminescence (radiance [photons]), normalized to pre-injection tumor
luminescence, is displayed as average +/- standard deviation fold luminescent change
(normalized to pre-injection tumor luminescence).
Statistical Analyses. Prism 5 (GraphPad) was used for all analyses. Results are presented
as mean
SEM, unless otherwise noted. Efficacy data was analyzed by an unpaired, two-
tailed t-test and repeated measures one-way ANOVA comparing all groups to assess
significance in treatment. P < 0.05 was considered significant.
Results
As is typical of most EGFR inhibitors (log P ~ 2 to 6), the EGFR inhibitor erlotinib is highly
hydrophobic (log P 2.7); whereas, as is typical of most DNA-damaging agents (log P
-
1.5 to
-1.5), doxorubicin is relatively hydrophilic (log P 0.9). Therefore, this particular drug
combination presents a challenge in achieving high concentrations of both drugs in a single
nanoparticle. Liposomes, however, provide a unique opportunity to achieve this capability
by virtue of their vesicular structure
[172,173].
By utilizing the lipid envelope for storage of
the hydrophobic drug and the aqueous interior to house the hydrophilic drug, liposomes
enable incorporation of high concentrations of both therapeutics and the potential to
present them in a shell first, core second fashion to produce the desired sequential
staggered release
[30]
- release of the hydrophobic small molecule EGFR inhibitor, erlotinib,
from the shell of the nanoparticle prior to unloading of the cytotoxic, doxorubicin, from the
core.
We fabricated liposomes containing doxorubicin and erlotinib using a lipid filmhydration method
[174]
(Figure 5.1.A). Lipid vesicles were formed following hydration in an
97
acidic citric acid buffer under high heat and sonication in the presence of the hydrophobic
inhibitor, erlotinib. Subsequent pH-driven loading of doxorubicin in the interior of the
vesicles resulted in the final dual-drug liposomal system (see Materials and Methods)
(Figure 5.1.B). Dynamic light scattering showed that the liposomes prepared in this twostep drug-loading process were of uniform size on the basis of their polydispersity index
(PDI) (Table 5.1). We prepared both single-drug (D, doxorubicin) and dual-drug (DE,
doxorubicin + erlotinib) liposomes, containing a mixture of lipids in a 56:39:5 mass ratio of
DSPC (distearylphosphatidylcholine): cholesterol: POPG [1-palmitoyl-2-oleoyl-sn-glycero-3phospho-(1'-rac-glycerol)]
with z-average hydrodynamic sizes (physical size of the
nanoparticle in suspension) relevant for systemic administration and with a (-potential,
measure of surface charge, of -29 mV, indicative of a negatively charged phospholipid
exterior. The drugs effectively accumulated in these, yielding a 2.5:1 mass ratio of
doxorubicin:erlotinib within the liposomes, from an equal mass supply (3 mg of each drug
per 50 mg of lipid used) during fabrication, and the encapsulation efficiency was higher for
doxorubicin than for erlotinib (Table 5.1).
From the compartmentalization of drug in the liposome, we predicted that erlotinib,
which is sequestered in the exterior lipid bilayer membrane compartment, would be
released before doxorubicin, which is concentrated in the hydrophilic core of the
liposomes. To determine the relative release rates of these two drugs, we measured the
amIount 01 Urug reaininiiig -i1
te 1IPjJsUlels aL seal
Id m LIIe pUiti Iuwiuvving IICLdLIUai iiI
phosphate-buffered saline (PBS) at physiological pH (7.4) and temperature (370C). After
24h, 60% of the erlotinib had been released, compared to only 20% of the doxorubicin
(Figure 5.2.A). This differential fractional release suggested that spatial control of
therapeutic loading into liposomes could enable presentation of this drug combination in
the desired fashion for synergistic killing, particularly because the effective local
concentration of erlotinib necessary for EGFR inhibition (-100 nM) is much lower than
that required for doxorubicin-induced cytotoxicity (Z-10 ptM)
[I5,
I/b].
To determine if the differential release rates from the liposome recapitulated the
synergistic cell killing observed previously
[30,
we applied the dual-drug and single-drug
liposomes to BT-20 TNBC and A549 NSCLC cell lines. We observed
an increase in
apoptosis, monitored by the number of cells positive for cleaved caspase-3 and cleaved
98
poly(ADP-ribose) polymerase (PARP), in response to either liposome, but DE liposomes
were more effective at both 24h and 48h compared to the single-drug D liposomes (Figure
2B). Importantly, the enhanced cleavage of caspase-8, which we have shown previously
occurs only in cells treated with time-staggered EGFR-doxorubicin combinations and not in
cells co-treated simultaneously with EGFR inhibitors and doxorubicin
[301,
was detected in
both cell lines upon DE liposome treatment (Figure 5.2.C).
Reduced abundance of phosphorylated extracellular signal-regulated kinase (pERK)
occurs in response to EGFR inhibition. We observed reduced pERK abundance in the DE
liposome-treated cells, confirming that EGFR activity had been effectively suppressed by
release of erlotinib from the dual-loaded nanoparticles (Figure 5.2.D).
Molecular characterization of liposome-treated cells revealed equivalent amounts of
DNA damage, as assessed by yH2AX formation, produced by exposure to either D liposomes
or DE liposomes (Figure 5.2.E). Although both liposomal formulations induced substantial
DNA damage (Figure 5.2.E), the DE formulation produced sustained inhibition of EGFR as
indicated by reduction in pERK for up to 72h (Figure 5.2.D) and were more cytotoxic, as
indicated by the larger apoptotic response (Figure 5.2.B-C).
To further improve the utility of this time-staggered drug delivery for in vivo use,
we prepared liposomes with folate and polyethylene glycol (PEG) as a means of minimizing
protein adsorption and subsequent nonspecific clearance, while promoting tumor targeting
[177-182]. Folate is a commonly used ligand for targeted cancer delivery; because it enhances
delivery of therapeutics to tumor cells on which folate receptor alpha is abundant at the
cell surface. Of particular interest for this investigation, as many as two out of three
patients with NSCLC and TNBC, as well as many other cancer cell types including tumors of
the ovary and prostate, have tumors with abundant folate receptors[18 3]. Thus, to enhance
tumor targeting, we included a folate-functionalized lipid, 1,2-distearoyl-sn-glycero-3phosphoethanolamine-N-polyethylene glycol-5000 (DSPE-PEG5 K), with a 5K PEG linker
(DSPE-PEG5K-Folate) in the lipid formulation at a 0.5 mole% ratio to the total lipid
composition. We also included shorter 2K PEG linker lipid (DSPE-PEG2K) at an equivalent
0.5 mole% ratio to minimize protein adsorption and opsonization of the liposomes while
still enabling access to the folate group on the longer PEG linker (Figure 5.3.A). This type
99
of targeted liposomal design has been used to deliver single-drug therapies
[177, 178, 180-182,
184]
The resulting single drug-loaded liposomes with doxorubicin, folate, and PEG (DFP)
and the dual drug-loaded liposomes with doxorubicin, erlotinib, folate, and PEG (DEFP)
were uniform, with a slight reduction in negative z potential compared to the uncoated
liposomes, due to the addition of PEG to the surface, and a corresponding -10-15%
increase in size from the uncoated liposomes (Table 5.1). Thus, the addition of folate and
PEG on the drug-loaded liposomal exterior only resulted in a moderate increase in particle
size, which is important for systemic administration to minimize physical filtration of the
nanoparticles by the liver and spleen and enhance tumor accumulation.
To demonstrate the cell-targeting capabilities of these functionalized liposomes, we
labeled doxorubicin-loaded, folate-functionalized liposomes with a near-infrared (IR) dye
Cy5.5 (DFP-Cy5.5) by including the Cy5.5-labeled 2K PEG lipid conjugate DSPE-PEG2K-Cy5.5
at 0.1 mole%, which enabled fluorescent tracking
[180, 181]
in vitro and in vivo. We also
created nonfunctionalized PEG liposomes labeled with Cy5.5 (DP-Cy5.5), but lacking the
folate (Appendix Table D.1), so that we could determine the effect of folate targeting on
the cellular uptake of the liposomes. We analyzed nanoparticle uptake by confocal
microscopy (Fig. 3B) and flow cytometry (Figure 5.3.C) of both BT-20 and A549 cells
following 1-hour incubation with either Cy5.5-labeled liposomes containing the folate for
cell LargetLing (LJ-CYJ.J) VI lcinLIlig Lth lfUIlat (LJ-CyJ.Jj. FUinuvvIg IILUUaLIII WILII the
targeted liposomes (DFP-CyS.5), both cell lines exhibited particle fluorescence (red)
throughout the cell cytosol, suggesting uptake and endosomal escape of the liposomal
contents (Figure 5.3.B). Analysis by flow cytometry confirmed that both cells lines
exhibited enhanced uptake of folate-containing liposomes compared with that of the
liposomes without folate (Figure 5.3.C) and showed that saturation occurred at high
concentrations of folate-containing liposomes. Not only did the folate-containing liposomes
result in substantially greater uptake of the nanoparticies by the cells, but also increased
the concentration of doxorubicin associated with each cell line (Figure 5.3.C). For the
untargeted systems, it is likely that most of the doxorubicin that accumulates in the cells
results from nonspecific liposomal uptake or occurs by pinocytosis of the free drug that is
released into the media by the liposomes. We compared the uptake and nuclear
100
accumulation of DFP-Cy5.5 liposomes with that of free doxorubicin by confocal microscopy,
which confirmed that uptake and nuclear accumulation were delayed when delivered by
the nanoparticles (Appendix D.1).
To evaluate these systems for tumor remediation in vivo, we first examined the
pharmacokinetics of the folate-PEG-functionalized empty nanoparticles (lacking drugs)
injected intravenously into healthy BALB/c mice, which are immune proficient. The folatePEG-functionalized, Cy5.5-labeled empty liposomes (NFP-Cy5.5) exhibited relatively low
liver accumulation as a function of time (14%, 23%, 17%, 14%, 12% injected dose
respectively, based on fluorescence recovery, at 30min, 9h, 24h, 48h, 72h, respectively)
(Figure 5.4.A) In the circulation, NFP-Cy5.5 exhibited half-lives of 2.2h and 11.1h,
representing rapid and slow elimination phases calculated on the basis of a twocompartment model (Figure 5.4.B). Furthermore, visualization of firefly luciferaseexpressing tumors revealed NFP-Cy5.5 liposome accumulation in the tumors of xenograftbearing nude mice up to 30 days following a single injection the folate-targeted liposomes;
we observed 2% injected dose at 9h, 5% injected dose at 24h, and 7% injected dose after
30 days based on fluorescence recovery (Figure 5.4.C). We also observed lower intensity
liposome fluorescence in other tissues, such as the liver, kidneys, and brain (Appendix
D.2). Thus, the formulated folate-targeted liposomal systems accumulated appreciably in
the target tissue (xenograft tumors), as well as the liver, kidneys, and brain, for a prolonged
period of time.
Because
tumor size
in BT-20
and
A549
xenograft
models
scaled
with
bioluminescence of the firefly luciferase-expressing tumors measured with the live animal
imaging system (Appendix D.3), we could quantify tumor size in situ during treatment. We
compared the effectiveness of the administration of DEFP liposomes or DFP liposomes, in
limiting or reducing tumor growth in the BT-20 TNBC and A549 NSCLC xenograft models
(Figure 5.5.A, B; Appendix A.4, 5). A single administration of 2 mg/kg of the dual drugloaded (DEFP) liposome 22 days after the tumors had been injected (Appendix D.3)
produced tumor regression, whereas the single drug-loaded DFP liposomes and the
untreated control mice exhibited continued tumor growth (Figure 5.5.A, B) In some of the
untreated control mice injected with A549 cells, tumors grew rapidly, reaching the
101
maximum allowed tumor burden (Appendix D.4, 5). Significant tumor shrinkage was only
observed for combination therapy-treated mice bearing xenografts for each cell line.
We also tested the effectiveness of codelivery of doxorubicin and erlotinib. Because
it was difficult to achieve simultaneous delivery with a single liposomal nanoparticle for
this drug combination, we designed single drug delivery systems for each drug in the
combination therapy and administered them together to mimic a co-release behavior. We
co-injected 2 mg/kg DFP with cyclodextrin-erlotinib (hydroxylpropyl--cyclodextrin) in
phosphate-buffered saline at 0.5 mg/kg in A549 xenograft-bearing mice, followed by
repeated dosing of cyclodextrin-erlotinib at days 2 and 4 to supply additional erlotinib to
match the slow, sustained release of doxorubicin (Appendix D.6). Although some
inhibition of tumor growth was observed at day 3, later the tumors exhibited increased
growth.
Thus, simultaneous, sustained coadminstration was a less effective treatment,
resulting in tumor growth. In contrast, tumor shrinkage and regression was achieved with
staged combination therapeutics packaged in the same nanoparticle delivery vehicle.
So far, we extensively characterized a time-staggered
doxorubicin:erlotinib
combination liposome that effectively and significantly reduced xenografted tumor growth.
We examined if this platform could be adapted for delivery of other therapeutic
combinations. We interchanged
the cytotoxic
drug encapsulated
in the aqueous
compartment [doxorubicin (D) or cisplatin (C)], and incorporated one of the following
initors
nRTK
0f signaling by EGFrIx faily m-nembIIers into the lipid envelope, erlotinib (E),
gefitinib (G), afatinib (A), or lapatinib (L). The physicochemical properties of the
nanoparticles, including size distribution and drug loading efficiency, with the different
drug combinations were similar (Tables 5.2, 3). In vitro drug release kinetics were similar
for each of the combinations, with rapid release of the RTK inhibitor followed by a
sustained release of the cytotoxic agent (Figure 5.6.A, Appendix D.7). We examined the
ability of each drug combination-folate-targeted liposome (DAFP, DEFP, DGFP, DLFP, CAFP,
CEFT, CGFP, (LFP) to mediate cell death in culture by analyzing caspase-3 and PARP
cleavage in BT-20 and A549 cells (Figure 5.6.B, C). Some combinations were more
effective at promoting tumor cell death, whereas others had little effect on the absolute
magnitude of cell death but accelerated the rate at which the cells died. Compared to the
A549 cells, BT-20 cells were typically more sensitive to the cytotoxicity effects of any of the
102
single cytotoxic agents or the dual drug-loaded liposomes, exhibiting >80% cell death in
most conditions. Within each cell line, there were some differences in cytotoxicity. For
example, the afatinib-cisplatin combination was the most effective of the cisplatincontaining combinations in promoting A549 cell death (-20%); whereas -40-50% of these
cells were apoptotic when exposed to any of the doxorubicin-containing combinations. For
the BT-20 cells the cisplatin-erlotinib and cisplatin- afatinib combinations produced the
fastest apoptotic response, which might be an important determinant of efficacy in vivo;
however, the maximum percent of apoptotic BT-20 cells was similar for all dual drugloaded liposome combinations. The cisplatin combinations were generally as effective at
producing cell death as cisplatin alone, with the exception of the cisplatin-afatinib (CA)
combination, which produced a higher maximal cell death than any other combination or
cisplatin alone in A459 cells.
We also examined the different doxorubicin- and RTK-loaded liposomes for
caspase-8 cleavage, as well as their effect on the extent of DNA damage, as indicated by the
abundance of yH2AX, and on downstream signaling, as indicated by phosphorylation ERK
abundance (Appendix D.8).
Discussion
Here, we utilized knowledge of synergistic combination therapies that re-wire signaling
pathways and networks
[301.
However, the synergistic therapeutic response exhibits a
pronounced dependence on timing and sequence of drug release. Therefore, we packaged
the drugs together in a systemically administrable tumor-targeted drug carrier, a liposomebased nanoparticle. Some success towards implementing synergistic cancer therapies using
engineered delivery systems had been achieved previously
[172, 173, 18S-191].
These prior
strategies have largely focused on the simultaneous release of drugs from a single delivery
platform using various techniques, including covalent linkage of drug to the material
comprising the nanoparticle or encapsulation within the nanomaterial. Although these
approaches have demonstrated the enhanced therapeutic efficacies of drug combinations,
there has been no development to date of time-staggered release platforms that are
specifically designed to engage and rewire cancer survival pathways.
103
We found that differential release rates of the drugs occurred in the correct
sequence from a single nanoparticle platform and that the dual drug-loaded liposomes
were cytotoxic to cultures A549 cells and BT-20. Molecular readouts (caspase-8 cleavage,
pERK, yH2AX) further confirmed that the multidrug liposomes disrupted cell function in
the expected sequence with early and sustained suppression of EGFR signaling, thereby
rendering the cells susceptible to apoptosis in response to later exposure to toxic amounts
of DNA-damaging agents. Furthermore, modification of the combination drug-loaded
liposomes to include folate and PEG promoted tumor targeting in vivo for enhanced
efficacy against A549 and BT-20 xenograft tumor models compared to single drug
treatment alone.
The efficacy of these staggered release nanoparticle systems containing a smallmolecule inhibitor with a cytotoxic agent presents a new therapeutic option for cancer
therapy. By exploiting the intrinsic two-compartment property of a liposomal system,
combination therapeutics with different physiochemical properties (hydrophilic and
hydrophobic) were compartmentalized with reasonable efficiencies in a core-shell fashion
that facilitated the staggered release observed. Bevcause liposomal systems provide high
levels of control over fabrication and modularity
[174],
these type of nanoparticle systems
should enable customizable targeting to specific tumors, as well as loading of therapeutic
agents tailored to specific treatment regimens.
The drugs that we tested were naUIrally hydruphobic (eriotinib, afatinib, gefitinib,
lapatinib) or hydrophilic (doxorubicin and cisplatin), and this nanoparticle system should
be adaptable to combinations that have chemically similar properties as these two
combinations (RTK inhibitor and cytotoxic agent). However, staggered release that relies
on the core versus outer layer loading of the two drugs may not always be attainable.
Therefore, other technologies, such as layer-by-layer assembly of polyelectrolytes, can
facilitate coatings on the exterior of drug-loaded nanoparticles to incorporate both
hydrophilic and hydrophobic molecules I'Jl, to introduce compartmentalization of two or
more drugs and enable control over timing and sequence of drug release. Further,
conjugation of one therapeutic to the nanoparticle core, followed by encapsulation or
coating on this core could further promote the staggered release necessary to elicit
synergism between drug combinations. Finally, combinations that require delivery to the
104
same cell for optimal synergistic efficacy further demonstrate the importance of
encapsulating multi-drug combinations within a single nanoparticle formulation.
The complexity of growth, survival, and death signaling pathways in cancers
continues to motivate investigation using systems biology approaches to inform treatments
against cancer[ 95 ,
192].
Furthermore, with the growing appreciation that the pathways
change in response to treatment or drug exposure, development of delivery systems that
are safe, yet capable of sophisticated levels of control over the delivery, timing, or sequence
of release, to therapeutically rewire these signaling pathways are essential to translate
findings to the clinic. Here, we addressed this challenge by developing a simple delivery
platform to target specific cancer cell types that are responsive to sets of synergistic drugs
with predetermined staged treatment regimens. Future tuning and adaptation of these
systems using nanomaterials design approaches should greatly enrich the range of
signaling pathway-based drug combinations and synergistic time-dependent release
profiles that can be achieved for targeted tumor therapy through dynamic network
rewiring.
105
A
B
Erlotinib
Doxorubicin
avg
h
Figure 5.1. Characterizationof combination therapeutic-loadedliposomal system.
(A) Cryogenic transmission electron micrograph (cryo-TEM) of dual drug-loaded liposomes. Scale bar
representative of 100 nm. (B) Schematic of dual loading of a small molecule inhibitor (erlotinib, blue) into the
hydrophobic, vesicular wall compartment and a cytotoxic agent (doxorubicin, green) loading into the aqueous,
hydrophilic interior.
106
A 100
B
D
=11001
w 80.
E
80]
-- ---
.T 60'
o 40
20
2
PBS
D
0 2448
D
8
4
o
0.06
0
8
1
+D
X 0.4
0.2
0
0
24
8
0
48
D
2448 72
48
8
0
72
DE
D
DE
8 244872
8 2448 72
8 244872
1.2
DE
24 48 72
8 24 4872
-+D
+aDE
+ODE
2
.
0.2
72
4
0.3
4
24 48
hours
D
8
0.4
+uDE
-
24
hours
hours
8
+D
0.6
00.02
0
DE
72
Z 0.8
oD
C 0.04
1
0
60
hours
D
aR
7 824
7 48
SM
24 48 72
40
20
0
E
DE
24 48 72
0
60
Z
0 0.0.8
0.1
I
0
0.6
0
8
24
hours
48
72
-0.1
0
8
24
hours
48
729
72
.
5
r.
0.08
-PBS
-- D
+DE
40
ours
2
D
0 2448
U..
6
01
m20
2
0
60
DE
0 2448
SoDE
40-
0
20
40
hours
A549
=t D
0
.
C
60i
.
60j
a40
0
BT-20
D
DE
-
0
8
24
48
72
hours
Figure 5.2. Evaluation of doxorubicin-erlotinibdual drug-loaded liposomalsystem in vitro.
(A) Drug releasefrom dual-drug -loaded liposomes in an excess volume of pH 7.4 PBS at 370C under agitation.
(B) Comparative cytotoxicity of dual drug-loaded liposome relative to the single drug-loaded liposome in BT20
(TNBC) and A549 (NSCLC) cell lines. This was measured by staining against cleaved caspase-3 and poly(ADPribose) polymerase [PARP]. (C) Cleaved caspase-8 in BT-20 cells (upper) and A549 cells (lower) following
addition of the indicated liposomes. In the gel images, red is actin and green is cleaved caspase-8. Quantification
shown below stainedgel images corresponds to relative signal of cleaved caspase-8 to actin. Data are presented
as mean +/- SEM or 3 experiments. (D) The dynamics of pERK in BT-20 cells (upper) and A549 cells (lower)
following addition of the indicated liposomes. The data are representative of 3 experiments and, for
quantification,pERK abundancewas normalized to actin. (E) yH2AXformation in BT-20 cells (upper) and A549
cells (lower) following addition of the indicated liposomes. The data are representativeof 3 experiments and,for
quantification, yH2AX abundance was normalized to actin. D liposomes, doxorubicin only (single drug); DE
liposomes, doxorubicin and erlotinib (dual drug).
107
A
I-
S
E3 A549
-15000
PEGL.
cc-
Erlotinib
,-'
10000
E3
0. .BT-2
.
0
A549
BT-20
BT-20
B
DoxorUbicin
'
5000
PEG,7
z
dh
150nm
0
10 20 30 40
NP (kg/m1)
-
zavg
6000
B
E3 A549
--0
-
-
--
(
E] E3
0
A549
BT-20
BT-20
3 2000
x
DAPN
Ph568
cys's
DAPN
PhS68
(y
'
4--
~I,...
0
.5
0 10 20 30 40
[DOX] of NP (stg/ml)
Figure 5.3. Folate decoration of combination doxorubicin-erlotinibtherapeutic-loadedliposomes for targeted
delivery to TNBC (BT-20) and NSCLC (A549) cells.
(A) Schematic of addition of DSPE-PEG 2K[mole% ratio (0.5%)] to minimize nonspecific protein binding, DSPEPEG 2 KCyss [mole% ratio (0.1%)]for fluorescent tracking, and DSPE-PEG5K-Folate [mole% ratio (0.5%)]for celltargeteddelivery. (B) Cell uptake of thefolate-targetedliposomes in BT-20 and A549 cells, visualized by confocal
microscopy,. Blue = nuclei, labeled with DAPI; green = actin, labeled with phalloidin Ph586; red = DSPEPEG2Kcyss-labeled DFP liposomes. Lower panels represent the fluorescence in each channel, upper panel is a
merged image. (C) Cell-associatedfluorescence measured by flow cytometry as a function of nanoparticle(NP)
concentration in both cell lines after incubation with folate-containingor absent liposomes at 370C for 1h. Data
on left correspondsto Cy5.5 (Aex = 675nm, Xex = 71Onm) correspondingto liposomal associationfor both targeted
(DFP-Cy5.5, squares) and untargeted control (DP-Cy5.5, circles). Data on right corresponds to the amount of
doxorubicin associated with the cells [doxorubicinfluorescence, (Ae = 480nm, Xex = 560nm)] for both targeted
(DFP-Cy5.5, squares) and untargeted control (DP-Cy5.5, circles). Data presented as mean +/- SEM of 3
experiments.
108
Liver-associated Accumulation (%
A
pre
Injected
Dose [IDI)
14% ID
23% ID
17% ID
14% ID
12% ID
30min
9h
24h
48h
76h
C
B 100
Tumor Visualization - Luciferase
NP tracker - NFP-C 5.5
T wo -c mputme nt mod
half-life (fast) = 2.2h
80 1 half-life (slow) = 11.1h
a
60"
40.
20
#
------------
0**
0
20
40
hours
60
80
Figure 5.4. Biologicalperformance offolate-targeted liposomal system in vivo.
(A) Biodistributionpanel offolate-targeted liposomes containing no drug (tracked through Cy5.5 fluorescence)
that were intravenously administered to BALB/c mice. In situ quantification (region identified with circle at 30
min post-injection) of liver-associatednanoparticlefluorescence (normalized to injected dose) presented above
each time point. (B) Circulation data displayed as percent injected dose (ID), on the basis of nanoparticleCy5.5
fluorescence recovered in blood samples. Half-life calculated based on a two-compartment model and presented
as mean +/- SEM. (C) Tumor visualization (left, visualized asfirefly luciferase expressed in the xenografted cells)
and nanoparticlevisualization (right, visualized by Cy5.5 fluorescence) 30 days after injection of single 0.1 mL
administration of folate-targeted empty liposomes (NFP-Cy5.5) to either BT-20- and A549-xenograft bearing
NCR nude mice bearing BT-20 xenografts or A549 xenografts on the hind flanks. The same animals are shown in
the left and right images. Seefig. S2for tissue necropsy.
109
A pre
7d
3d
21d
A549
A549
~11
BT-20
'
10
-A untreated
f
{
G
o
0
32d
20
BT 20 -
10,
DEFP
21d
BT-20 - DEFP
NT
01E
7d
3d
pre
32d
A549 - DEFP
E'.a
DFP
DEFP
10
20
days
30
0.1
unrtd
Q
0.1
0
DFP
- 10
20
days
1
30
Figure 5.5. Effect of dual drug-loaded or single drug-loaded, folate-targeted liposomes on A549 and BT-20
tumorsize.
(A) A549-luciferase expressing, xenograft-bearingNCR nude mice; n = 5, quantification representativeof mean
+/- SEM. (see fig. S5 for all 5 mice) (B) BT-20-luciferase expressing, xenograft-bearing NCR nude mice; n = 5,
quantification representative of mean +/- SEM. (see fig. S4 for all 5 mice). Tumor-imaging data for dual-drug
(DEFP, top), single-drug (DFP, middle), and untreated control, along with luminescence quantification(reported
as fold initial tumor luminescence, presented on a semilog plot) corresponding to tumor size as a function of
time, following a single administration of 1 mg/kg drug-loaded liposomal formulations. Animals with tumor
reaching 1 cm were sacrificed. An unpaired, two-tailed t-test comparing the DFP and DEFP at the terminal 32d
time point shows statisticalsignificance with p-values of 0.0057 and 0.001 for treatedA549 and BT-20 xenograft
bearing mice, respectively. A one-way ANOVA comparing all treatmentsfor the duration of the experiment (all
time points) was also performedfor each xenograft cell line and p-values of 0.0024 and 0.0010 were obtainedfor
each A549 and BT-20 cells, respectively.
110
801
I
U,D
S60
&
S40
B
E
L
V A
G
-
__4F~ ---- --
BT-20
1001
U
60
DEFP
03DFP
DAFP
DGFP
80~
N7
601
0
40
o
0
i
7
N
iu
DFP
DEFP
DAFP
DGFP
7
20
0. 2
4DLFP
0
A549
i
A
7
0
OLFP
20
20
__
i
0
20
40
hours
60
80
C
8 0-
cu 64A02
-V4
C
0' E
-A L
A
1 00
0
4
0
20
40
hours
hours
-
25'
-
15.'
60,
40
0
0
i
20
60
80
- 20-
--
01
0
"
12
24
hours
36
CAFP
jCEFP
0.
o 10'
CGFP
CLFP
5.
48
20
0
60
40
BT-2U
80
0.
- -- -- --
0
20
0
01W
0
hours
60
40
A549
C
CAFP
CEFP
CGFP
CLFP
- -
-Q
I,,
I/
12
-
24
hours
..
-
A
36
48
Figure 5.6. Developing the RTK inhibitor-cytotoxic agent combination liposomal systems as a platform for dual-
drug delivery.
(A) In vitro drug release from dual-drug loaded folate-targeted liposomal formulations. Top represents the
release from folate and PEG-containing liposomes (FP) with doxorubicin and the indicated RTK inhibitor.
Bottom represents the release from folate and PEG-containingliposomes (FP) with cisplatin and the indicated
RTK inhibitor. Liposomes were incubated with pH 7.4 PBS at 370C under agitation in sink conditions. Data
presented as mean of triplicate experiments. (B) Cytotoxicity of dual drug-loaded (RTK and doxorubicin)
liposomes compared with that of single drug-loaded (doxorubicin) liposomes againt BT-20 and A549 cells. Data
presented as mean of triplicate experiments. (C) Cytotoxicity of dual drug-loaded(RTK and cisplatin) liposomes
compared with that of single drug-loaded (cisplatin) liposomes against BT-20 and A549 cells. Data presented as
mean for triplicate experiments. D = doxorubicin; A = afatinib; E = erlotinib; G = gefitinib; L = lapatinib; C =
cisplatin; F =folate; P = PEG.
111
formulation
Mean zavg
dh (nm)
~+/- SEM
PDI
DE
136+/- 10
0.13
-29
D: 97%
D: 5.4%
D
126+/
0.15
-27
D: 97%
D: 5.5%
DFP
156 +/- 7
0.1
-17
D: 97%
D: 5.4%
DEFP
151 +/-
0.16
-15
D: 97%
D: 5.5%
E: 40%
E: 2.2%
______a__n
16
114
Encapsulation
w/w
Efficiency
(drug/lipid)
(drg/___d_
Ef_ cincy
%
Liposomal
potentials
(MV)
E: 40%
E: 2.2%
Table 5.1. Dynamic light scattering, polydispersity index, and potential measurements for the doxorubicinerlotinibdual-drug and single-drug uncoated and coated (FP) liposomal systems.
Measured in 10 mM NaCl at 250C. Encapsulation efficiency (% loadedfrom amount supplied) and mass loading
ratio (g/g drug/lipid) as determined by high-performance liquid chromatography. DE, doxorubicin and
erlotinib;D, doxorubicin; DFP, doxorubicin, folate, PEG; DEFP, doxorubicin and erlotinib,folate, PEG.
112
Liposomal
Formulation
Mean za dh
(nm)
DFP
w/w
(drug/lipid)
156 +/- 7
0.1
D: 97%
D: 5.4%
DAFP
174 +/- 13
0.19
D: 99%
A: 82%
A: 82%
D: 5.5%
A: 4.4%
A: 4.4%
DEFP
151+/- 14
0.16
D: 97%
E: 40%
E: 40%
D: 5.5%
E: 2.2%
E: 2.2
DGFP
160 +/- 11
0.12
DLFP
182 +/- 9
0.13
D:
G:
G:
D:
L:
L:
98%
8%
85%
97%
7%
70%
%
SEM
%
PDI
Encapsulation
Efficiency
D:
D:
G:
D:
L:
L:
5.5%
4.6%
4.6%
5.5%
3.8%
3.8%
Table 5.2. Dynamic light scattering and polydispersity index measurements for the various dual-drug and
single-drug coated (FP) liposomalformulations- doxorubicin + afatinib/erlotniib/gefitinib/lapatinib.
Measured in 10mM NaCl at 250C. Encapsulation efficiency (% loadedfrom amount supplied) and mass loading
ratio (g/g drug/lipid) as determined by high-performance liquid chromatography. D = doxorubicin;A = afatinib;
E = erlotinib; G = gefitinib; L = lapatinib; C = cisplatin; F =folate; P = PEG. Correspondingformulations without
folate-PEG (FP) are shown in table S1.
113
Liposomal
Formulation
Mean Zavg dh
CFP
183+/-12
0.16
C: 62%
CAFP
177 +/- 6
0.06
A: 73%
A: 73%
A: 5.2%
A: 5.2%
CEFP
183 +/- 13
0.16
C: 60%
C: 10.5%
E: 34%
E: 2.4%
CGFP
162 +/- 8
0.09
CLFP
204 +/- 11
0.11
(nm)+/1SEM
_/-____
PDJ
Encapsulatio
Efficiency
C: 58%
E: 64%
C: 55%
w/W%
(drug/lipid)
C: 11%
C: 9.9%
: 2
C: 9.4%
C: 81%
G: 81%
C: 60%
G: 5.8%
C: 10.2%
L: 72%
L: 5.1%
L 2
G: 5.8%
:51
Table 5.3. Dynamic light scattering, polydispersity index, and zeta-potential measurements for the various
multi-drug coated (FP) liposomalformulations- cisplatin + afatinib/erlotinib/gefitinib/lapatinib.
Measured in 10mM NaCl at 250C Encapsulation efficiency (% loaded from amount supplied) and mass loading
ratio (g/g drug/lipid) as determined by high-performance liquid chromatography. D = doxorubicin; A = afatinib;
E = erlotinib; G = gefitinib; L = lapatinib;C = cisplatin; F =folate; P = PEG. Correspondingformulations without
folate-PEG (FP) are shown in table S1.
114
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116
Chapter 6. Liposomal delivery of quisinostat and cisplatin for
treatment of non-small cell lung cancer.
Introduction
Histone deacetylase inhibitors (HDACi) are promising anti-cancer therapeutics that
have demonstrated activity in altering gene expression, inducing apoptosis and cell cycle
arrest in various cancer cells, as well as promoting degradation of oncogenes.[1
9 3 -2 0 1]
One
such HDACi developed by Janssen Pharmaceuticals, Inc. is JNJ-26481585, also known as
quisinostat. It has been shown to have potent anti-tumor activity and favorable oral
bioavailability, biodistribution, and tissue penetration than the current clinically approved
HDACi, vorinostat,[1 93-201] Despite this, free drug administration of HDACi, including
quisinostat, is dose-limiting and shows a potent cardiotoxicity that prevents its immediate
clinical application.[ 20 2]
Lung cancer is the deadliest form of cancer for both men and women in the United
States and worldwide.[ 2 0 3] Over 80% of tumors are non-small cell lung carcinoma (NSCLC),
with most cases diagnosed at a late stage, beyond the point where they can be cured by
surgery alone. Despite chemotherapy and radiation, the mainstays of treatment for
inoperable tumors, the overall prognosis is grim, with a 5-year survival of
-15%.[
2 04
]
Targeted therapies have shown substantial benefit in select patients; EGFR inhibitors
significantly improve survival of patients with adenocarcinomas with EGFR mutations.[ 205 ]
Similarly effective therapies are, however, presently lacking for the majority of patients.
Conventional cytotoxic therapies for advanced NSCLC involving platinates with a second
agent chemotherapeutic agent have shown modest benefits.[2 06]
Towards this end, we sought to address this clinical need by packaging and
delivering quisinostat combination cisplatin in a singular nanoparticle formulation for
more effective treatment of NSCLC.
Materials and Methods
All lipid components were purchased from Avanti Polar Lipids, except for cholesterol from
Sigma and DSPE-PEG 2k/DSPE-PEG 2k-NH2 from Nanocs, Inc. Quisinostat was provided by
Janssen Pharmaceuticals, Inc. Cisplatin (cis-diamineplatinum(II) chloride) was purchased
117
from Sigma. All other chemicals (citric acid, sodium citrate, sodium carbonate, folate,
transferrin)
and solvents (chloroform, methanol, phosphate buffered saline) were
purchased from Sigma.
Liposome
preparation. Liposomes were formulated at a mass ratio of 56:39:5
(DSPC:Cholesterol:POPG). These three components were dissolved, along with the smallmolecule inhibitor (weight ratio to total lipid weight = 3:50), in a 2:1 mixture of
chloroform:methanol. A thin film of these materials was generated by rotary evaporation at
400C at 150 mbar. This film was dessicated overnight until completely dry. Hydration of the
lipid film was conducted at 650C under sonication in 300 mM citric acid buffer (pH 4) for
1h. DSPE-PEG2K-Folate was prepared by reacting DSPE-PEG2k-NH 2 with folate in the
presence of DCC/pyridine as previously described.[ 179 , 2 07, 2 08 1Functionalization with DSPEPEG5K-Folate, DSPE-PEG2K, with or without DSPE-PEG2K-Cy5.5 was conducted postfabrication using a post-insertion technique in which micelles (in 0.9% sodium chloride
solution) of the components desired to incorporate on the liposomal surface were
incubated with the prepared, drug-loaded liposomes under sonication at 650C for 30
minutes, after which they were filtered through a 0.2 ptm polyethersulfone (PES) syringe
filter and cooled to room temperature. The pH of the liposomal suspension was then
adjusted to -6.5 by addition of 300 mM sodium carbonate buffer to create a gradient
between the exterior and interior compartments. The cytotoxic (doxorubicin - 3 mg,
cisplatin - 10 mg) was added in a 0.9% sodium chloride solution (1 mL) to load through a
pH gradient method. To facilitate solubilization of the cytotoxic agents, the dispersed
solution was sonicated at 650C for 5 minutes. The final combination drug-loaded system
was subsequently exchanged into PBS (pH 7.4) following centrifugal filtration (100K
MWCO Millipore) to remove the high salt buffers (citric acid, sodium carbonate) and any
unloaded drug. Empty (lacking drug) liposomes (NFP) were formulated using the same
procedure. Cryo-preservation was done following a previous published protocol.[ 2091
Liposome characterization. Dynamic light scattering and z-potential analysis was
conducted in 10 mM sodium chloride at 250C using a Malvern ZS90 zeta-sizer. Highperformance liquid chromatography (Agilent technologies) and nanodrop absorbance
118
measurements (345 nm for inhibitor; 480 nm for doxorubicin) were used to validate drug
loading of the inhibitor
(Xabs
= 345nm) and doxorubicin
(?abs
= 480nm). Cisplatin
concentration was quantified by a colorimetric assay using o-phenylenediamine against a
standard curve
[171].
Cryo-transmission electron microscopy (cryo-TEM) was conducted by
imaging a vitrified dilute sample of the liposomal suspension at 120 kV and 77 K.
In vitro assessment. Liposomes were incubated under sink conditions [1 L sink for 1 mL
liposome suspension] in 1X PBS under agitation in 1 mL 3.5 K MWCO float-a-lyzers
(Spectrum) at 370C. PBS was replenished each day of the experiment. Samples were taken
of the liposomes to quantify remaining drug concentrations by HPLC (following dissolution
in a 50:50 mixture of acetonitrile:pH 5 water) and absorbance measurements (nanodrop)
[345 nm for erlotinib, lapatinib, /gefitinib, or /afatinib; 480 nm for doxorubicin]. IC50
determined by a dose-escalated MTT cytotoxicity assay with the value empirically
determined by the fit to the data - find value at which 50% cell death occurs corresponding
to fit on a death vs. log(concentration) plot.
Pharmacokinetics. BALB/c female mice (Taconic) were systemically-administered (tail
vein) empty Cy5.5-labeled liposomes (NFP-Cy5.5)
at a concentration of 3 mg/mL
(corresponding radiant efficiency, measured by the whole-animal imaging system - IVIS
[Xenogen, Caliper Instruments] of 0.1 mL sample injected in each mouse -1 x 1010). Wholeanimal fluorescence imaging (IVIS - Xenogen, Caliper Instruments) was performed at the
indicated time points for one cohort of mice (n = 3), and a separate cohort was used for
retro-orbital bleeds to determine the circulation half-life of the system. Imaging and
circulation data presented is normalized to auto-fluorescence determination (imaged
animals, isolated blood) obtained prior to injection. Analysis of circulation data based on
recovered fluorescence normalized to pre-injection blood auto-fluorescence is displayed
with a two-compartment model fit (both slow and fast half-lives presented).
Gross toxicity.
BALB/c mice were dosed intraperitoneally with free and liposomal
encapsulated therapeutics at the given concentrations (dose-escalated). Body weight of
119
cohorts of 5 mce/treatment was recorded during and following repeated doses on the
indicated days. Data is presented as weight versus day from first administration.
Tumor Targeting and Regression. Tumor targeting data was obtained by imaging a
Cy5.5-labeled DSPE-PEG-Cy5.5 lipid that was inserted into the liposomal membrane. This
dye-labeled lipid was functionalized using a DSPE-PEG-NH2 lipid (Nanocs) and Cy5.5-NHS
ester (Lumiprobe) at a 1:1 ratio in DMSO at 40C for complete labeling of the lipid. Final dyelabeled lipid product was lyophilized and stored as a powder at -20 0C. Tumor-targeting
data was collected using a near-IR imager (IVIS, Xenogen, Caliper Instruments) at Ax = 675
nm, Xem = 720 nm. % injected dose (ID) calculations were performed from fluorescence
intensity of the tumors, following background subtraction of pre-injection tissue
autofluorescence, and normalized to the amount of fluorescent material injected. Necropsy
of relevant tissue (liver, kidneys, spleen, heart, lungs, brain, and tumors) was preformed,
with images displaying recovered fluorescence in each tissue via IVIS whole animal
fluorescence imaging shown. Luminescent xenograft tumors were seeded following stable
transfection of BT-20 and A549 cells with the firefly luciferase plasmid. This enabled
assessment of tumor size by a visual and quantifiable luminescent readout generated by
whole-animal imaging (Xenogen, Caliper Instruments). Cells were mixed in a 1:1 ratio with
BD Matrigel TM Basement Membrane Matrix to a final density of 5 x 107 cells/0.1 mL
injection. 0.1 mL injections of the matrix-cell suspension were performed in each rear hind
flank of NCR nude mice (Taconic). Tumors were allowed to grow until a visible tumor was
established and luminescence monitoring was indicative of solid tumor growth (see fig. S3,
-7
x
108 radiance final luminescent readout from each xenograft). Xenograft-bearing NCR
nude mice were then systemically-administered therapy (DFP, DEFP) in a 0.1 mL injection
at a concentration of 2 mg/kg (based on cisplatin loading) and monitored for 50 days.
Luminescence images were obtained by 0.1 mL intraperitoneal injections of 30 mg/kg Dluciferin (Caliper) and imaging (IVIS, Xenogen, Caliper) with an open luminescence filter 15
minutes post-injection.
These data are presented, along with region of interest
quantification of radiance corresponding to the xenograft-specific luminescence for each
mouse treated. Data is normalized for each mouse against the tumor luminescence prior to
injection and presented as fold luminescence above this measurement. N = 5 for each
120
treatment and data corresponds to mean +/- SEM for each treatment group. An unpaired,
two-tailed t-test comparing the single drug (cisplatin/quisinostat) and combination drug
(cisplatin + quisinostat) systems was performed to determine statistical significance.
Quantification of luminescence (radiance [photons]), normalized to pre-injection tumor
luminescence, is displayed as average +/- standard deviation fold luminescent change
(normalized to pre-injection tumor luminescence).
Statistical Analyses. Prism 5 (GraphPad) was used for all analyses. Results are presented
as mean
SEM, unless otherwise noted. Efficacy data was analyzed by an unpaired, two-
tailed t-test and repeated measures one-way ANOVA comparing all groups to assess
significance in treatment. P < 0.05 was considered significant.
Results and Discussion
We fabricated liposomes containing cisplatin (C) and quisinostat (Q) using a lipid
film-hydration method[174 ] (Figure 6.1.A) with the same lipid composition and fabrication
protocol as that described in Chapter 5. The final functionalized liposomal system was
targeted using 0.5mol% surface folate functionalization, which was shown to significantly
enhance NP uptake in NSCLC A549 cells in Chapter 5. Physicochemical characterization of
the fabricated liposomes is shown in Figure 6.1.B. Fabricated cisplatin-containing folatePEG functionalized liposomes (CFP) were observed to be -150nm with a narrow PDI of
0.13 by dynamic light scattering. Dual drug (cisplatin + quisinostat)-containing folate-PEG
functionalized liposomes (CQFP) were observed to be -180nm with a narrow PDI of 0.04.
Loading of quisinostat was conducted analogous to erlotinib in Chapter 5, whereby it was
loaded in the lipid bilayer by incorporating the small molecule in the organic phase during
thin-film formation, while cisplatin was driven in the aqueous center by a pH-gradient
method analogous to that of doxorubicin described in Chapter 5. To further solubilize
cisplatin, sonication and heat was applied to the saline-drug containing solution prior to
loading into the pre-formed liposomes. The final folate-PEG functionalized dual drug
liposomes were visualized by cryogenic-TEM shown in Figure 6.1.C.
121
Drug release properties of this dual drug system was characterized in PBS under
agitation at 370C and is shown in Figure 6.2.A. Quisinostat was released on a faster time
scale than that of cisplatin, likely due to the nature of the core-shell compartmentalization
of the drugs. Approximately 50% of quisinostat was released after -2 days, while that of
cisplatin was achieved after 4 days. The given drug loading and release properties of the
dual drug folate-targeted liposomal system was sufficient to elicit a strong anti-cancer
response in A549 cells in vitro, significantly reducing the IC50 of cisplatin used in the
treatment after 48h and 72h relative to a single drug cisplatin-only folate-targeted
liposomal construct. This data presented in Figure 6.2.B suggests an added or synergistic
benefit of delivering quisinostat and cisplatin together.
To evaluate this system for tumor remediation in vivo, we first examined the
pharmacokinetics of the folate-PEG-functionalized empty nanoparticles (lacking drugs)
injected intravenously into healthy BALB/c mice, which are immune proficient. The folatePEG-functionalized,
Cy5.5-labeled
empty liposomes
exhibited
relatively
low liver
accumulation as a function of time (14%, 23%, 17%, 14%, 12% injected dose respectively,
based on fluorescence recovery, at 30min, 9h, 24h, 48h, 72h, respectively) (Figure 6.3.A).
From blood isolation, the NPs were characterized by serum half-lives of 2.2h and 11.1h,
representing rapid and slow elimination phases calculated on the basis of a twocompartment model (Figure 6.3.B). Furthermore, visualization of firefly luciferaseexpressing A549 tumors revealed the Cy5.5-tracked empty liposomes accumulate in the
tumors of xenograft-bearing nude mice significantly (-8% ID) as early as 48h following a
single injection of the folate-targeted liposomes - shown in Figure 6.4.A. We also observed
little to no fluorescence in the kidneys and heart for these formulations, which was
-
essential given the potent renal toxicity of cisplatin and cardiotoxicity of quisinostat
shown Figure 6.4.B. Gross toxicity of liposomal and free quisinostat as well as liposomal
dual drug cisplatin + quisinostat formulations was conducted and is disDlaved in Figures
6.5, 6. Primary initial toxicity concern centered around the cardiotoxicity of quisinostat,
with previous work suggesting an MTD of free drug quisinostat at -10mg/kg; however, we
observed little gross toxicity as measured by weight change in the mice following 3 repeat
injections at days 0, 2, and 4 out to 24 days. As shown, toxicity requiring euthanasia (body
weight change > 15%) for the given dose range and treatment regimen was not observed,
122
therefore no MTD was observed in this particular experiment. This data suggests the
liposomal carrier does not introduce additional toxicity versus the free drug; we anticipate
that it may also be able to prevent heart toxicity by controlling the release of the drug in
vivo, but more detailed histological studies would be needed to study this effect. Gross
toxicity evaluation of the dual drug (quisinostat + cisplatin) construct showed toxicity
requiring euthanasia was only at 20mg/kg (Figure 6.7), suggesting the MTD for this
formulation lies in the range of 10-20mg/kg for this given treatment regimen. Free
cisplatin was not tested as its MTD is known to be -4mg/kg,[ 2 10] below the tested range.
With
promising
in
vitro
results
and
in
vivo
biological
performance
(pharmacokinetics, tumor targeting, and toxicity) in immune-proficient BALB/c mice and
tumor-bearing ncr nude mice, we evaluated the therapeutic efficacy of these systems
against a hind flank A549 mouse model. Using luminescent signal as a means of tracking
tumor progression (described and validated in Chapter 5), significant differences in the
treatment outcomes were observed, as shown in Figure 6.8, with the combination drug
liposome promoting tumor regression (-30% reduction in tumor luminescence) over the
course of the first three weeks, with overall tumor maintenance observed by the end of 50
days (+/- 5% of initial tumor luminescence) following initial treatment.
In comparison,
each of the single drug controls showed overall tumor growth (-30% for Pt only, ~80% for
HDAC only) at this dose range, with singular benefit in slowing tumor growth relative to
the untreated control (-130% growth by the end of day 50).
For clinical application, a final set of experiments evaluating the ability to preserve
and store these systems was performed. As presented by the light scattering data in Figure
6.8, one measure of nanoparticle stability is a measure of the size of the nanoparticle over
time; typically we anticipate seeing little or no significant change in size, as may occur with
aggregation, during storage. We have shown that these systems are capable of
lyophilization and storage out to 3 months with marginal impact of resuspended size
(-20% increase) and effectively no change on drug loadings, presenting this system as an
attractive approach towards treatment of NSCLC.
123
A
PEG 5 K-Folate
Hydrophobic
Small Molecule
Inhibitor (QUI)
OA
Hydrophilic
Cytotoxic (CPT)
PEG2x/PEG rZK5-5
I
II
I
I
z,,dh= 15--20nm
B
Cp
147.7+4-13.4
0.13
CQFP
183.5+l- 5.05
0.04
C: 5040%
C: 5080%
Q: 25S40%
C: cisplatin; Q: quisinostat; F: folate; P: PEG
C
Figure 6.1. Characaterizationof cisplatin-quisinostatdual-drug coated (FP) liposomalsystem.
(A) Schematic of liposomal system containing cisplatin in the core and quisinostat in the lipid bilayer. (B)
Physicochemicalcharacterization:hydrodynamic diameter, polydispersity index (PDI), and loading efficiency of
each drug. (C) Cryogenic-TEM of cisplatin+quisinostatloaded liposomalsystem.
124
100
-
A
'CC
S 60
--
40
0
B
5O
100 150
hours
200
10
106
10
10
-109
10" i
10
Pt
Pt + HDAC
1
48 h
24 h
72 h
Figure 6.2. In vitro release and efficacy of cisplatin-quisinostatdual-drug coated (FP) liposomal system.
(A) Drug release study of cisplatin+quisinostatloaded liposomal system in 1X PBS at 370C under agitation. (B)
IC50 of cisplatin (Pt) and cisplatin + quisinostat(Pt + HDAC) loaded liposomal systems againstA549 cells.
125
Liver-associated Accumulation (% Injected Dose [ID])
A
pre
14% ID
23% ID
17% ID
14% ID
30min
9h
24h
48h
12% ID
76h
BU
Two-compartment model
half-life (fast) = 2.2h
B0
100
half-life (slow) = 11.1 h
of50
0
0
G"G--------- -------- p
20
40
hours
60
80
Figure 6.3. Pharmacokineticsof FP coated empty liposomal carrier.
(A) Biodistribution of blank folate-functionalized liposomal system following systemic administration in BALB/c mice with the corresponding liverassociatedaccumulation data, calculated as % injected dose, at each time point as quantifiedfrom IVIS whole animalfluorescence imaging. (B) Circulation
persistence data of systemically administeredfolate-functionalizedliposomal system. Quantified as % injected dose with two-phase decay modelfit to data
to determine two half-lives characteristicof the system.
126
A
Tumor Visualization - Luciferase
BLi
S
K H/Lu T
NP tracker - Cv5.5-Lipid
B
Li: liver; S: spleen; K: kidneys; H/Lu:
heart/lungs; T: tumors; B: brain
Figure 6.4. Biodistributionof FP coated liposomal carrierin A549 xenograft bearing ncr nude mice.
(A) Tumor targeting capabilitiesoffolate-functionalized liposomal system. (left) IVIS luminescence of A549 hind flank xenografts; (right) corresponding
nanoparticlefluorescence overlay showing particle distribution in tumor bearing mice in the left image. (B) Biodistribution data of folate-functionalized
liposomal nanoaprticles 48h following systemic administration. Necropsied tissue corresponds to mice displayed in right image of panel A.
127
32"free" quisinostat - 5mg/kg
liposomal quisinostat - 5mg/kg
"free" quisinostat - 10mg/kg
liposomal quisinostat -10ng/kg
"free" quisinostat - 20mg/kg
liposomal quisinostat - 20mg/kg
22
0
injections
30
20
10
days
O, 2, 4d
Figure 6.5. Gross toxicity of free and encapsulated quisinostat formulations following intraperitoneal
administrationin immune-proficientBALB/c mice.
Three repeat injections (at doses: 5, 10, 20 mg/kg) were preformed, each 2d apart and monitored out for 3
weeks beyond the final treatment. Weight reductionfor 'free" drug treatment averaging 6.7%, 9.1%, 9.2% for 5,
10, and 20mg/kg, respectively. Weight reduction for liposomal drug treatment averaging (n=3) 5.4%, 4.7%,
6.7% for 5, 10, and 20mg/kg, respectively.
128
1Omq/kq
04 I
'Iree" quisinostat - 1 Omgkg
U
D
-5
liposomal qui -10mg'kg
liposomal cptiqui -1 0mgkg
-
C
-6
-10
i
-
II
-15
O
5
10
15
20
25
days
20mg/kg
"free" qui - 20mg/qk
liposomal cpb'qui -20mg/kg
-
-5
I
M1 -10
0
liposomal qui - 20rmq1kg
I
-
C
-15
I
-20
219
0
5
10
15
20
25
days
t
t
Dose Od, 7d, 14d
Figure 6.6. Gross toxicity evaluated as body weight change in BALB/c mice treated intraperitoneally with
liposomalformulationsand correspondingfree drug controls.
Repeat administration on days 0, 7, and 14. Mice monitored for one week following final administration.
Cisplatin MTD 7 mg/kg.
129
37
0O
2-
V
V
X
U
A
A
IA
A
V
0
Pt
+
HDACI folate-lipo
Pt folate-lipo
HDAC folate-lipo
untreated
N
0
C-.
I
M
I[1 1x
0
0
10
lt tf
30
20
40
50
days
injections
0, 5, 10d
Figure 6.7. In vivo efficacy of single drug and multi-drug cisplatin-quisinostatloaded liposomal systems in A549
hindflank xenograft bearing ncr nude mice.
Data presented asfold luminescence signal normalized to initial tumor size as quantified by luminescent signal.
Dosage of 5mg/kg for cisplatin only and dual drug-loadedNPs; 4mg/kg for HDACi only liposome - corresponds
to concentration of HDACi in dual drug-loaded liposomalsystem with 5mg/kg cisplatin (5:4 drug loading ratio).
Repeat administrationsperformed on days 0, 5, and 10.
130
nanoparticle + drug I (log P ca. -1.5)
+ drug 2 (lop P ca. +3.0)
+ NP/drug1,2 (150 +/- 10 nm)
NP/drugl,2 - lyopres (700 +/- 100 nm)
-- NP/drugl,2 + tyopres (180 +/- 30 nm)
-+
40-
Drug 1: erlotinib
I1,
A,
-15 N
I
Loadinzs Pre-Ivo:
030-
20-
I.
-45
10-
10
I
-
0'I
Drug 2: doxorubicin
1000
hydrodynamic diameter (nm)
100
460
5.5 wt% doxorubicin
2.2 wt% erlotinib
Loadings post-IVo:
5.2 wt% doxorubicin
2.1 wt% erlotinib
10000
Figure6.8. Cryo-preservationof dual-drugcisplatin-quisinostatFP coated liposomal construct.
Data shown for resuspensionof lyophilized particles both with and without lyopres (the preservation media).
131
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133
Chapter 7. Liposomal delivery of JQ1 and temozolamide
systemically across the blood-brain barrierfor treatment of
brain-stem gliomas
Introduction
Approximately 18,000 new diagnoses of primary malignant central nervous system
(CNS) tumors occur annually in the United States resulting in 13,000 deaths.[ 2111 Gliomas
comprise approximately 80% of primary malignant CNS tumors, and carry one of the
highest mortality rates of all tumors.[ 212 There is currently no cure for gliomas. The most
common and most aggressive form of glioma is glioblastoma multiforme (GBM). Despite an
improved molecular understanding of GBM, there remain no effective treatments to cure
this brain cancer.[ 2 13 -2 1 61 The current clinical standard includes surgery followed by
radiation therapy plus temozolomide resulting in a median survival of only 15 months.[
213]
The prognosis has not changed substantially over 20 years.[ 214 ,217]
Gliomas can arise anywhere in the CNS, however brainstem gliomas (BSG) occur in
the midbrain, pons, medulla oblongata, and upper cervical spinal cord, critical areas of the
brain that control crucial life functions.
BSGs commonly occur in children, comprising
approximately 20% of primary pediatric CNS tumors and are the leading cause of death for
children with brain tumors.[ 218] Given the delicate anatomy in which BSGs are found,
surgery is often not an option for these patients, and radiation therapy is the current
standard of care for these tumors.
Many cancer cells are deficient in their ability to effectively repair DNA damage. This
is apparent in glioma cells, where deficient DNA repair results in genomic instability and a
high rate of DNA mutations that can lead to more aggressive tumors and resistance to
therapy. Radiotherapy is partially effective against glioma by exploiting this deficient
tumor DNA repair and causing extensive DNA damage that can be repaired in normal cells,
but not in tumor cells. Using a novel discovery platform for genes and small molecules that
alter the response to DNA damage, a bromodomain-containing epigenetic reader Brd4 was
identified as a modulator of ionizing-radiation-induced intracellular signaling and cell
survival.[
2191
A range of small molecule inhibitors of Brd4, including JQ1, was found to
further alter DNA damage signaling and demonstrate activity against glioma cells.
134
One factor that limits the utility of many potent chemotherapy agents against
gliomas is the blood-brain barrier (BBB), which restricts entry of many small molecules
into the brain. The BBB is comprised of capillary endothelial cells that act as a highly
selective permeability barrier between the blood in circulation and the extracellular fluid in
the central nervous system. This barrier mediates entry of molecules via a number of
different mechanisms, including a) transcellular lipophilic pathways for lipid soluble
agents, b) paracellular pathways through the endothelial tight junctions for water souble
agents, c) transport proteins for amino acids and glucose, d) receptor-mediated
transcytosis (e.g. transferrin, insulin, folate), and e) adsorptive transcytosis for plasma
proteins such as albumin.[ 2 2 01 The BBB also contains efflux pumps such as P-glycoprotein
and multidrug resistance pump-1 (MRP-1) that can recycle therapeutics back into
circulation.
A number of drug delivery strategies have been employed for treatment of CNSrelated disease, including direct injection via a catheter into the brain or cerebrospinal fluid
driven either by diffusion (natural gradient) or convection (pump), implants of devices
purposed as drug depots, delivery via trans-nasal pathways, and systemic solutions in
combination with barrier disruption mechanisms, such as ultrasound, microwaves,
,
radiation, infection, or injection of various substances such as hyperosmolar mannitol.[ 221
2221 Limitations of localized delivery include limited diffusion of the drug away from the site
of injection and risk of infection following direct disruption of the CNS tissue.[2 231 Systemic
delivery promises to overcome these limitations as a less invasive option for more efficient
delivery throughout the brain where the delivery system would preferentially bind and
deliver cargo to target cells via receptor-mediated interactions; however, issues of tissue
selectivity and biodistribution
continue to hinder these systems towards clinical
application.
Regarding systemic delivery, a number of molecular receptor targets, including
transferrin[ 224 -226], insulin, and low-density lipoprotein, have been investigated towards the
deployment of nanomedicine designed to traverse the BBB.[ 2 2 7 2 2 9 ] Further selectivity
following BBB penetration towards gliomas can be achieved with molecular targets such as
those promoting angiogenesis (e.g. integrins, VEGF, ILs) or invasiveness (e.g. MMP-2). One
such promising strategy employed a chlorotoxin molecular ligand, which is a 36 amino acid
135
peptide isolated from deathstalker scorpion venom that selectively binds isoforms of MMP2 and chloride conductance channels unique to gliomas.[ 2 3 0- 2 3 2] BBB-penetrating decoration
has been investigated on a number of different nanoparticle cores, including liposomes,
solid-lipid nanoparticles, dendrimers, gold nanoparticles, carbon nanotubes, quantum dots,
and magnetic nanoparticles.[ 233 1 Nanoparticles employing this targeted approach show
significant promise towards development of a systemic delivery agent to CNS tissue,
3 2 35
several of which are currently being evaluated in the clinic.[ 2 3 - ]
Materials and Methods
(+)-JQ1 was purchased from Cayman Chemical. All lipid components were purchased from
Avanti Polar Lipids, except for cholesterol from Sigma and DSPE-PEG2/end-functionalized
DSPE-PEG 2k-(NH2/maleimide) from Nanocs, Inc. Chlorotoxin was purchased from Enzo
Life Sciences. Cholesterol, temozolamide, citric acid, sodium citrate, sodium carbonate, and
solvents (chloroform, methanol, phosphate buffered saline) were purchased from Sigma.
Liposome
preparation. Liposomes were formulated at a mass ratio of 56:39:5
(DSPC:Cholesterol:POPG). These three components were dissolved, along with the smallmolecule inhibitor (+)-JQ1 (weight ratio to total lipid weight = 3:50), in a 2:1 mixture of
chloroform:methanol. A thin film of these materials was generated by rotary evaporation at
400C at 150 mbar. This film was dessicated overnight until completely dry. Hydration of the
lipid film was conducted at 650C under sonication in 300 mM citric acid buffer (pH 4) for
1h. DSPE-PEG2K-folate was synthesized as reported in Chapter 5. DSPE-PEG2K-HAlOk was
functionalized by reacting HA with DSPE-PEG2k-NH2 under the same conditions as those
described for folate in Chapter 5. DSPE-PEG2K-Chlorotoxin/Transferrin was prepared by
reacting
a
maleimide-functionalized
DSPE-PEG2k
lipid
with
thiolated-
transferrin/chlorotoxin (obtained by reacting each with Traut's reagent at a 1:40 ratio of
protein:Traut's in pH 8.5 borate-EDTA buffer (0.15M sodium borate, 0.1mM EDTA) and
,
finally purified using amicon ultracentrifugation using previously established protocols.[ 230
236] DSPE-PEG2K, with or without DSPE-PEG2K-Cy5.5, and with or without a ligandfunctionalized lipid (0.5 mol% folate, 2 mol% transferrin, 2 mol% chlorotoxin, 5 mol%
HAOk) was conducted post-fabrication using a post-insertion technique in which micelles
136
(in 0.9% sodium chloride solution) of the components desired to incorporate on the
liposomal surface were incubated with the prepared, drug-loaded liposomes under
sonication at 650C for 30 minutes, after which they were filtered through a 0.2 ptm
polyethersulfone (PES) syringe filter and cooled to room temperature. The pH of the
liposomal suspension was then adjusted to -6.5 by addition of 300 mM sodium carbonate
buffer to create a gradient between the exterior and interior compartments. The cytotoxic,
temozolamide (10mg), was added in a 0.9% sodium chloride solution (1 mL) to load
through a pH gradient method. To facilitate solubilization of the cytotoxic agents, the
dispersed solution was sonicated at 650C for 5 minutes. The final combination drug-loaded
system was subsequently exchanged into PBS (pH 7.4) following centrifugal filtration
(100K MWCO Millipore) to remove the high salt buffers (citric acid, sodium carbonate) and
any unloaded drug. Empty (lacking drug) liposomes (NFP) were formulated using the same
procedure.
Liposome characterization. Dynamic light scattering and zeta-potential analysis was
conducted in 10 mM sodium chloride at 250C using a Malvern ZS90 zeta-sizer. Highperformance liquid chromatography (Agilent technologies) was used to validate drug
loading of (+)-JQ1
(Xabs =
260nm, elution -10min) and temozolamide
(Aabs =
260nm, elution
-2min).
Cranial window technique for intravital multiphoton imaging of intracranial tumors.
Mice were anesthetized using ketamine and xylazine in accordance with institutionapproved animal protocols. A cranial window was fashioned using a handheld micro-drill
(Ideal Micro-Drill TM ) and the skull carefully removed exposing the underlying brain
parenchyma (Fig.). Mice were immediately positioned on the stage of an inverted Olympus
FluoView 1000 multiphoton laser scanning confocal microscope for imaging through the
cranial window. Mice were anesthetized using 2% isofluorane during the imaging sessions.
For visualization of NP transport across the BBB, NPs were labelled with Cy3 and injected
intravenously 24 hours prior to intravital imaging to allow for uptake. For visualization of
intracranial tumors, GFP-expressing U87MG tumors were induced for 10 to 14 days prior
137
to creation of the cranial window and imaging through the multiphoton microscope.
Composite images were compiled using ImageJ software.
Biodistribution. For biodistribution experiments, NCR nude female mice (Taconic) were
used. To attenuate gut fluorescence, an alfalfa-free special diet (AIN-93M Maintenance
Purified Diet from TestDiet) was administered to the mice 1 week prior to and during
experimentation. Nanoparticle formulations suspended in 1X PBS were administered via
the tail vein. In vivo imaging (IVIS, Caliper Instruments) was performed at regular time
points. Necropsy of relevant organs was conducted at a relevant time point and imaged
using the IVIS to quantify recovered fluorescence.
Histology. Brains were resected at the terminal point of experimentation and embedded
and frozen in OCT medium until processed. Tissue was sectioned coronally at 20um thick
with 10 sections performed 100um apart and mounted in DAPI medium on a glass slide
with a coverslip for confocal imaging.
Stereotactic intracranial implantation of U87MG human glioma xenograft tumors.
Human U87MG glioblastoma cells were obtained through ATCC and maintained in 10%
FBS (Gibco) and DMEM. For generation of stable cell lines, cells were transduced with a
mir30-based pMLS vector encoding both GFP and luciferase (gift from M. Hemann). For
xenograft implantation, 6 week old male NCR nude mice (Taconic) were anesthetized using
2% isofluorane, ketamine, and xylazine. The skull was exposed through a midline scalp
incision and a hole was made through the right hemicranium using a 16 gauge needle (2
mm lateral to the midline, 2 mm anterior to the lambdoid, 3 mm below the dura). Mice
were then place in a stereotactic frame (Stoelting Co., IL) and 1 x 10s cells suspended in
PBS were injected slowly into the right cerebrum using a 32 gauge Neuros syringe
(Hamilton). Bone wax was applied to occlude the hole in the skull and the scalp was closed
using interrupted sutures. All procedures were performed under sterile conditions. Tumor
growth was monitored using bioluminescence imaging using an IVIS Spectrumbioluminescent and fluorescent imaging system (Xenogen Corporation). Brain tumors were
also imaged using a Varian 7-Tesla whole-mouse MRI system (Varian/Agilent). Tumors
138
were visualized by bioluminescence and magnetic resonance imaging modalities. Efficacy
of transferrin-targeted
liposomal
drug
carriers
was
evaluated
by
systemically
administering a single dose at 2mg/kg then observing extension of survival for the animal
cohorts in each treatment, generating the Kaplan-Meier curve shown in Figure 7.7.
Results and Discussion
There are currently no chemotherapeutic or biological agents that have proven
additional benefit over radiation therapy for high-grade pediatric BSGs. Even with
radiation therapy, the median survival for children with malignant BSGs is <1 year.[ 2371New
leads toward effective therapy are a significant unmet need in this devastating disease.
Using a unique automated microscopy and image analysis platform to interrogate the DNA
damage response (DDR), Floyd et al probed RNAi libraries for modulators of DDR
signaling,[ 219] which identified Brd4 as a novel and potent regulator of early signaling
events in the DDR. Brd4 contains tandem bromodomains that bind to acetylated lysine. A
small molecule inhibitor of bromodomain binding, JQ1, had potent effects on DDR signaling
in glioma cell lines. Additionally, Floyd et al found that among these glioma cell lines, those
with known defects in DDR repair genes become exquisitely sensitive to ionizing radiation
after pretreatment with JQ1. Delivery of JQ1, however, remains a challenge for delivery due
to stability and solubility issues for clinical application - usually requiring direct
intrathecal or intracranial delivery of the drug.
To overcome these limitations, we have developed nanoparticle technology to
stabilize and solubilize the drug for systemic application. The nanoparticle platform is a
liposomal construct described in Figure 7.1.A. Similar to the construct described in
Chapter 5, we have fabricated dually loaded liposomal carriers by a traditional thin filmhydration method[ 238] with the same lipid composition (56:39:5 DSPC:Cholesterol:POPG),
loading and compartmentalizing the inhibitor (+)-JQ1 and cytotoxic cisplatin as shown. For
targeted transport of these therapeutics, a number of targeting ligands including
transferrin and chlorotoxin, were chosen for transport across the BBB as well as targeting
glioma cells, while other systems decorated with PEG, hyaluronic acid, and folate, which is
overexpressed on glioma cell surfaces,[ 239] were also studied in parallel to determine the
optimal system for delivery and efficacy. Physicochemical characterization of the fabricated
139
targeted liposomes in Figure 7.1.B shows sizes between 130 - 170 nm with PDIs around
0.15. Targeted liposomes (e.g. transferrin, chlorotoxin) have been shown to traverse the
BBB at appreciable levels, preferentially localizing in the brain to deliver their cargo, and
more specifically in areas with direct access to the brainstem. Liposomal technology is
readily adaptable to carry a wide range of small molecules, siRNAs, biological agents, as
well as combinations of cargos analogous to the system developed in Chapter 5, and
subsequently functionalized with a hybrid surface for both targeting and stealth (e.g.
resistant to opsonization and clearance) capabilities.
To investigate the capability of these fabricated systems to cross the BBB, we
developed an intravital imaging technique that allowed us to visualize the BBB and the
engagement of the NPs with this barrier following tail vein systemic administration.
Surgically removing the skull of the mouse, exposing the brain and blood vessels, and
fluorescently tracking the transit of our Cy3-labeled NPs using a multiphoton confocal
imager achieved visualization. Using this technique, we tested each of the different NP
formulations - PEG, HA, folate, transferrin, and chlorotoxin. Intravital imaging results show
rapid uptake of the NPs into the blood vessel walls in the brain (Figures 7.2.A-D, white
arrowheads) but only the chlorotoxin- and transferrin-conjugated particles demonstrate
passage across the BBB into the surrounding brain matter (Figures 7.2.B, C). This is
consistent with published reports of success using chlorotoxin[ 2301 and transferrin[ 230, 236]as
ligands to facilitate delivery of NPs across the BBB. Although the particles tagged with
folate also showed rapid uptake into the blood vessel walls (Figure 7.2.D), they did not
demonstrate diffusion across the BBB, serving as a good negative control that the
penetration was ligand specific.
To further assess the ability of these NPs to penetrate and reside in the brain, we
administered the targeted formulations in both immunodeficient ncr nude and immuneproficient BALB/c mice and monitored the distribution of the NPs throughout their bodies
and organs using live animal fluorescence imaging. Shown here in Figure 7.3, we observed
significant accumulation in the brain, -1.3-1.7%
injected dose based on recovered
fluorescence, for both the chlorotoxin (CTX)- and transferrin (Tf)-functionalized NPs
respectively at 24 hours post-injection, which we expected due to their known capacities
for traversing the BBB and targeting glioma cells. Control PEG and HA-targeted NPs
140
showed no appreciable accumulation, whereas folate showed some, -0.9% injected dose at
24 hours (data not shown), likely owing to the expression of folate receptors on the blood
side in the choroid plexus 3 and possible other avenues of entry yet unknown.
Figure 7.3 informed our selection and further screening of the chlorotoxin- and
transferrin-targeted liposomal systems for efficacy evaluation. In both ncr nude and
BALB/c mice, significant BBB penetration and distribution of NPs to the brain was
observed for both targeted systems, as shown in Figure 7.4, with slightly higher amounts
observed for the transferrin-targeted system (-1.7% ID) relative to the chlorotoxintargeted system (-1.3% ID). Enhanced levels of NP accumulation (shown in red, nuclei
stained with DAPI mount shown in blue) in the brain is further shown in Figure 7.5 from
fluorescent imaging of coronal brain sections of treated non-tumor bearing ncr nude mice.
Similar results were obtained from coronal brain sections of treated non-tumor bearing
BALB/c mice. No apparent accumulation was observed for PEG- and HA-functionalized
liposomal systems, while some accumulation was visualized for folate-targeted liposomal
systems - consistent with data presented in Figures 7.3, 4.
We are currently testing the adult human glioma cell line, U87MG, and have
successfully implanted these cells into the brains (Figure 7.6.A) of ncr nude mice that can
be visualized by MRI (Figures 7.6.B, C). Using the U87MG animal model for initial efficacy
evaluation,
we tested our transferrin-targeted
liposomal
nanoparticles,
observing
significant colocalization of the NPs with the glioma brain xenografts as early as 24h postadministration and out to 5d (Figure 7.7). From necropsy data shown in Figure 7.4, the
amount could be as high as -1.7% injected dose, based on the entry of transferrin-targeted
liposomal manufactured systems in healthy ncr nude mice. Further experimentation
investigating whether the glioma xenograft affects the BBB and entry of our NPs into the
brain is ongoing.
Efficacy evaluation is currently in progress. Preliminary results are shown in Figure
7.8, where transferrin-targeted single drug and dual drug ((+)-JQ1, cisplatin) liposomal
nanoparticles were tested at 2 mg/kg single systemic tail vein administration (each drug,
drug loading ratio -1:1 in dual drug nanoparticle) against the U87MG brain xenograft.
Shown in the Kaplan-Meier data, extension of life was observed for the single drug controls
for 3-4 days longer than that of the control non-treated mice. For the dual drug system,
141
extension of life 12 days beyond the non-treated control and 8-9 days beyond the
corresponding single drug liposomal controls. Testing of free drug controls at the same
drug concentration is currently underway to characterize the impact of drug encapsulation
as well as benefit of multi-drug therapy for this glioma animal model.
Final therapeutic evaluation of this transferrin-targeted dual drug liposomal system
will include combination radiation to investigate any synergism between treatments, as
suggested by previous reports 240] To accomplish this, we have designed a stereotactic
rodent radiation device, which is compatible with the Cesium source irradiator in the Koch
Institute (Figure 7.9.A). The device has interchangeable collimators, which allow us to
administer focused doses of irradiation to the brain or other regions of the body (Figure
7.9.B).
Conclusion
We have successfully adapted the liposomal technology developed in Chapter 5
towards targeting (e.g. transferrin- and chlorotoxin-targeted lipids) and delivering small
molecule chemotherapeutics (e.g. cisplatin) and inhibitors (e.g. JQ1) for treatment of
brainstem gliomas, which are plagued by poor prognoses. Future work will incorporate
radiation therapy with the liposomal drug carriers to screen for synergism to create
effective treatment a1gorithms for this devastating disease.
142
PEGS + transferrin, folate,
A
HA, chlorotoxin
Hydrophobic drug
/
Hydrophilic drug
PEG2K/PEG2KCy5.5
\_7
PEG
Tf-PEG
136
137
0.15
Fol-PEG
HA-PEG
165
160
0.11
CTX-PEG
168
0.18
0.17
0.13
Figure 7.1. Liposomal constructfordual-drug (+)-JQ1 -temozolamide targeted carriers.
(top) Nanoparticle design. (bottom) A list of the different ligands attached to our nanoparticles and the
diameters (zavg dh) of these particles. PEG: poly(ethylene glycol); Tf: transferrin;Fol: folate; HA: hyaluronic acid;
CTX: chlorotoxin.
143
Figure 7.2. Intravitalmultiphoton imaging of nanoparticlescrossing the blood-brainbarrierin mice.
(A) Imaging through a cross section of blood vessels running through the skull (shown in white) shows uptake of
redfluorescentchlorotoxin nanoparticlesinto the walls of blood vessels (white arrows). (B) Imaging deeper into
the brain shows the nanoparticlescoursing through a blood vessel (outlined in white) with diffusion of particles
into the surrounding brain matter. (C) Red fluorescent transferrinnanoparticlesline a blood vessel in the brain
(white arrows) and demonstrate diffusion into the surrounding brain. The skull is shown in blue and outlined in
white. (D) Redfluorescentfolate nanoparticlesshow uptake into the blood vessel wall (white arrows) but do not
show diffusion across the BBB.
144
Nude - 8h
SV:
u Ia..
BALB/c - 8h
Li
Nude - 24h
S K H Lu B
FOL
Li S K H Lu B
PEG
FOL
Tf
PEG
Tf
CTX
CTX
HA
HA
Figure 7.3. Representative distribution data of each fabricatednanoparticlesystem in non-tumor bearing mice.
(left) ncr nude and (middle) BALB/c mice at 8h post-systemic administration. Poly(ethylene glycol) (PEG);
transferrin (Tf);folate (FOL); chlorotoxin (CTX); hyaluronic acid (HA); Li: liver; S: spleen; K: kidneys; H: heart;
Lu: lungs; B: brain. (right) Representative distribution data of each fabricated nanoparticlesystem in nude mice
after 24h.
145
Nude - 8h
Li SKHLuB
Li
UN
C
N-CTX
Tf
UN
B-CTX
C
N-Tf
Tf
B-Tf
SKHLu B
BALB/c - 8h
Figure 7.4. Biodistribution of targeted liposomal carriersin non-tumor bearing mice.
(left) Representative distribution data of chlorotoxin (C) and transferrin (Tfl-targeted nanoparticle
formulations 8 post-systemic administration in ncr nude (top) and BALB/c (bottom) mice relative to an
untreated, autofluorescence control. (right) Representative distribution data of chlorotoxin (CTX)- and
transferrin (TJ-functionalized nanoparticles in mice at 24h post-injection (Li: liver; S: spleen; K: kidneys; H:
heart; Lu: lungs; B: brain).
146
PEG
FOL
CTX
Tf
HA
Figure 7.5. Histological sections of ncr nude brains 24h post-systemic administration of functionalized
nanoparticlesystems.
Polyethylene glycol (PEG); folate (FOL); chlorotoxin (CTX); Tf (transferrin); hyaluronic acid (HA)-targeted
systems. Blue = DAPI stain of nuclei; Red = Cy5.5 tracker lipid in liposomalformulation.
147
Fig 7.6. Mouse models of human glioma.
Human glioma cells, U87MG, are implanted into the brains of mice and visualized using bioluminescentimaging
(A) as well as magnetic resonance imaging (MRI) (B, C).
148
Mouae B - Day I
Fluoresce"ce Lun- esc-en:e
Mouse A - Day 1
Fluorescence LurrjnEsc.2'::e
Transferrin - TMZ + JO1
Mouse A - Day 5
Mouse B - Day 5
Figure 7.7. Biodistribution data of transferrin-functionalizedliposomal carriers in U87MG brain-implanted
tumors.
Whole-animal imaging demonstrating colocalization offluorescently-tracked Cy5.5-labeled transferrin-targeted
liposomal nanoparticleswith luminescently-labeled M059K brain xenografts. Imaging shown 24h and 5d postsystemic administrationof NPs.
149
Vehicle
2 mg/kg TMZ
2 mg/kg JQ1
2 mg/kg TMZ+JQ1
-
100
-ru
50
-
4-,
C)
U
C)
0~
0
0
_ _ ___
5
10
15
20
25
Days
Figure 7.8. Kaplan-Meier curvefollowing treatment of U87MG brain tumor implants with single drug and dualdrug temozolamide-(+)-JQ1 transferrin-targeted liposomal carriers.
(black) vehicle control - transferrin-targetedblank liposome; (blue) temozolamide-loaded transferrin-targeted
liposome; (red) - (+)-JQ1-loaded transferrin-targeted liposome; (grey) ftemozolomade + (+)-JQ1]-loaded
transferrin-targetedliposome. Vehicle dosed at equivalent lipid concentration; drug-loaded systems dosed at
2mg/kg of each drug.
150
e DO"wyGy
sow
Sam, 1aCM
gftcuh
NOW
Figure 7.9. Small animal irradiationplatform for combination radiation-systemicchemotherapy.
(top) The lead irradiatorbox has interchangeablecollimators which will deliverfocused irradiationto rodents.
(bottom) Shown are radiationbeam characteristicsfor a 1 cm circularfield created with custom lead shielding
apparatusdesignedfor use with the Cesium137 irradiatorhoused in the Koch Institute mouse facility.
151
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153
Chapter 8. Framing the impact of this work and future
directions.
Chapters 2 - 4 focus on systems utilizing the capabilities of layer-by-layer (LbL)
customization for design of nanoparticle systems for a variety of disease targets with the
added dimensionality of scalably manufacturing these systems, making them attractive for
clinical adaptation.
-
Future impact of this work
Chapter 2: This work is a fundamental study on LbL films atop nanoparticle platforms and
how that influences drug release and the distribution of the nanoparticle in a systemic
setting. It establishes a framework with which to screen LbL nanoparticle designs and
understand how new biomaterials and/or particle designs compare relative to existing
architectures and NP constructs.
Chapter 3: This work establishes a method for customizing and incorporating biomaterials
atop NP platforms via LbL for tissue targeting. The approach of functionalizing and
incorporating a material via LbL atop a NP platform presents an array of opportunities for
disease management beyond cancer - conceivably any disease target of interest. Regarding
moving this work forward to the clinic, more advanced osteosarcoma models (directly in
bone tissue) are necessary to test, as well as a metastatic model, if possible to establish,
that localizes in bone - that would be the highest impact result.
Chapter 4: This work establishes a scalable method for designing built-to-order medicinal
systems on the nanoscale and that could be easily adapted to the microscale for a variety of
applications. The tunability of this approach for NP engineering opens up a number of
exciting opportunities in the field of biotechnology - treatment, diagnostics. Regarding the
specific technique,
adjusting the hydrophobicity of the sacrificial adhesive layer to
facilitate spray-assisted LbL would be the next technical step towards streamlining the
approach for future scalable manufacturing purposes.
154
Overall, the systemic use of LbL NPs has been shown to exhibit much promise, and
opportunities for clinical development exist in this area with existing technologies
developed in our lab. Next steps would be to consider work on systems for issues where
local
application
(creams/rinses/i.d.
injections/incorporation
in microneedles)
is
necessary, such as vaccines or immunotherapy. Due to the nature of the films to directly
incorporate a wide variety of therapeutics while also serving as surfaces for targeting and
controlled drug release, these systems would be ideal for applications local to the diseased
tissue/site of interest to serve as drug depots for managing the issue.
Chapters 5 - 7 focus on the use of liposomal systems without LbL functionalization towards
disease management. Liposomal formulations became of interest due to their current
clinical
use
and
ease
in
manufacturing
and
manipulation
for
the
desired
compartmentalization and drug release, as well as tissue targeting, necessary for the
applications explored in this work.
Chapter 5: This work takes advantage of the nature of the liposomal structure to load and
compartmentalize hydrophobic and hydrophobic drug combinations together in high
quantities to achieve the desired time-staggered release of the two drugs for synergistic
cancer cell killing - herein lies the novelty, not necessarily in the technology itself. We also
demonstrate this technology to successfully deliver a range of therapeutic combinations for
treatment of TNBC and NSCLC. Future work with this system lies in screening the
nanoparticle formulation against more advanced tumor models, such as the autochthonous
NSCLC model in the Jacks' lab, as well as testing in large animal models such as monkeys. A
current collaborative effort with Merrimack pharmaceuticals is being conducted to further
validate and develops this system towards clinical application.
Chapter 6: This work adapts the technology in Chapter 5 with two different drugs for
treatment of NSCLC, further demonstrating the versatility of this system. Future work with
this system lies in screening the nanoparticle formulation against more advanced tumor
155
models, such as the autochthonous NSCLC model in the Jacks' lab, as well as testing in large
animal
models
such
as
monkeys.
A current
collaborative
effort with
Janssen
pharmaceutical employees is being exploited to investigate the possibilities of VC funding
and development of this technology external to the company towards clinical application.
Again the novelty lies in the delivery of the agents together, not in the technology, so we are
also interested in developing a proprietary LbL formulation designed to achieve the same
efficacy.
Chapter 7: This work adapts the technology in Chapter 5 with two different drugs as well
as incorporating a range of new targeting approaches for traversing the BBB and treating
brain stem gliomas, further demonstrating the versatility of this system. The technology is
not novel in the field, but the capabilities in delivering the combinations of drugs
investigated is exciting, along with the potential to rapidly screen particle designs and
combinations with other treatment modalities, such as radiation therapy. Future
development here lies in developing a more highly controlled and patentable approach in
particle engineering by incorporating LbL and functional materials with the ligands of
interest for BBB penetration and glioma cell targeting ala work in Chapter 3. This would
not only open a number of exciting opportunities for developing a proprietary blend that
could be clinically developed (and IP protected) but also allow for more fundamental
studies on blends using combinations of ligands with high levels of control over the surface
of the nanoparticle for BBB engagement and cancer cell targeting/treatment.
Overall, the liposomal technology is exciting and readily adaptable to the clinic - likely
much moreso than the LbL formulations described herein due to the current use of
liposomal products in the clinic. The issue lies in IP protection. The systems developed are
based on previous systems, with little novelty lying in the technology itself but rather the
action of the technology (staggered release leading to enhanced cancer cell killing,
delivering unique combinations of drugs for various disease targets), hence the interest in
future work focusing on utilizing the knowledge gained with LbL to develop proprietary
blends that can achieve the same level of success with added IP protection and
entrepreneurial prospects.
156
Appendix A. Architecture, Biological Performance of LbL
Nanoparticles
Although this study focuses on a number of biodegradable materials and conventional
therapeutics, a much larger library of film components, along with a diverse range of
therapeutics, can be generated for incorporation on this platform.
PLGA (50:50'; 6 S:3 Sb; 75:25c)
Poly-L-LysIned (PLL)
Dextran Sulfate* (DXS)
Hyaluronic Acid' (HA)
Alginate' (Aig)
a38-54K b24-38K C4-15K
d4-15K
e9-20K '500K '80-120K
A.1. Representativescanning electron micrograph (SEM) of LbL PLGA NPs (HA-terminal).
1pm scale bar.
157
1.5X10I
(i)
C.
1.0x10j1
w
5.0x101
I
20
.,
C
(iii)
min
w
C
(v)
1.0)1011-
5.0X1010
-
(iv)
40
-
1.5X101 1
0
I
U
60
I
-
max
.
(ii)
I
A
0.0
0.1
0.2
0.3
0.4
I
0.5
time (h)
A.2. CG drug stability- biodistributionof LbL PLGA NPs in ncr nude mice.
(left) 48h IVIS imaging panel (Aex = 745 nm, Aem = 820 nm) surveying drug bioavailabilityas a function of nanoparticlearchitecture. (i) free CG82o; (ii)
PLGAso:socG; (iii) DXS-terminated 6L NP; (iii) HA5OK-terminated 6L NP; (iv) Alg-terminated 6L NP. (right) Whole-animal region of interest analysisyielded
from the series of images demonstrates enhanced persistence of drug in the animal.
(i)
max
1.5x10
U
(ii)
1.0x10
wc
5.0X10 0-
(iii)
0
0
min
20
40
60
time (h)
(iv)
A.3. NP stability - biodistributionof LbL PLGA NPs.
(left) 48h IVIS imaging panel (Aex = 640 nm, Aem = 700 nm) surveying carrierbioavailabilityas a function of nanoparticlearchitecture.(i) free PLL 7 oo; (ii)
DXS-terminated 6L NP; (iii) HAsoK-terminated 6L NP; (iv) Alg-terminated 6L NP. (right) Whole animal region of interest analysisyielded from the series of
images. Whole-animal region of interest analysisyielded from the series of images demonstrates enhanced persistence of the layered nanoparticlesin the
animal.
159
0i)
max
P)
max
(()
(wi)
(iv)
I
(iv)
min
min
A.4. Biodistributiondata 30 minutes after systemic administrationin ncr nude mice.
(left) PLL 7oo tracking - Aex = 640 nm, Aem = 700 nm; (i) free PLL 7oo; (ii) Aig-terminated 6L NP; (iii) HA SOOK
terminated 6L NP; (iv) DXS-terminated 6L NP; (right) CG9 20 tracking - Aex= 745 nm, A)em = 820 nm; (i)free CG82o;
(ii) PLGA50:50CG; (iii) HA50oK-terminated 6L NP; (iii) Alg-terminated 6L NP; (iv) DXS-terminated 6L NP.
160
max
(iii)
(ii')
(Iv)
min
(v)
150-
6L NP (X-term)
6L NP + Free X (10mg/kg)
6L NP + Free X (20mg/kg)
E
U
S
t E
0
o100
0
50
0
DXS
Aig
HASOOK
X-term
A.5. Dose-dependent co-injection offree polymer with correspondingterminated 6L LbL PLGA NPformulations.
(left: 10 mg/kg, right: 20 mg/kg) co-injection of 6L nanoparticlearchitectures with corresponding terminal
layerfree polymer. IVIS imaging surveys circulation of drug (tex = 745 nm, A)em = 820nm). (i) Free HA500K + HA500K
6L NP; (ii) HA500K-terminated 6L NP; (iii) Free Alg + Aig-terminated 6L NP; (iv) Alg-terminated 6L NP; (v) Free
DXS + DXS-terminated 6L NP; (vi) DXS-terminated 6L NP. Region of interestanalysis was conducted 15 min postiv injection to capture liver-associateddrug fluorescence (Aex = 745 nm, )tem = 820 nm) shown in bottom graph.
Data is normalized to maximal radiantefficiency in data setfor each NP architecture.
161
6L Alginate
6L HASOOK
PLL7 00
o.5h
Epi-fluorescence
4.0
lh
35
3.0
4h
25 A0
7
2.0
8h
1.5
1.0
Radiant Efficiency
p/seccm q/sr
Pw/cm I
24h
Color 5cale
Min =
1.00e7
Max
4.00e7
=
48h
O.5h
Epi-f
rescence
1h
x120
10
4h
05
Radant Efficiency
opsecacmnsr)
pW/cm2
8h
(Min- 5.00e6
24h
A.6. Blood circulationdatafor LbL PLGA NPs and encapsulatedcardiogreen.
(top) IVIS quantificationof recoveredfluorescence at carrierchannel, Aex = 640 nm, Aem = 700 nm; (bottom)
quantificationof recoveredfluorescence at CG820 channel, Aex = 745 nm, Aem = 820 nm.
162
IVIS
(I)
(I)
max
(iV)
(i)l
Ail
min
A.7. Variation of core template by modulating L:G ratio in PLGA and effect on biodistribution in ncr nude mice.
7 5
: CG core; (top and bottom, right) PLGA :2 CG core - (top, left and right) Ax = 640
(top and bottom, left) PLGA 65 35
nm, Aem = 700 nm; (i) free PLL 7oo, (ii) DXS-terminated 6L NP, (iii) HASOOK-terminated 6L NP, (iv) Alg-terminated
6L NP; (bottom, right and left) Aex = 745 nm, Aem = 820 nm; (i) free CG82o, (ii) PLGA 65 :35 CG core, (iii) HA50OKterminated 6L NP, (iv) Alg-terminated 6L NP, (v) DXS-terminated 6L NP. Similar in vivo stability profiles to that
of PLGASO:SOcc(Figure 2.3); however, significantly higher CG loadings could be achieved by tuning the L:G ratio.
6 35
Typical loadings (based on UV-Vis absorption via HPLC) are: PLGAo:soCG - 10 ug/mL; PLGA S: cG - 20 ug/mL;
7 25
PLGA s: c - 50 ug/mL. 30 min stability assessment representative of 48 h panel of images (data not shown).
Data deemed consistent with findings in main text and not shown to impact the stability of the drug in the core.
163
140001
0.4
E
110000
C.,'
0.3-
U_
E
4000-
0.2C
0
0.
0
U)
0.1-w
}
-I-
3000-
2000-
Il,
E
0
U.
8000
5000-
10000-
PSTXR
PLL DXS 4-1 5 K
Aig
core
HA5 OOK
PLL DXS 4.1SK
AIg
HA500K
A.A. Opsonin (IgG) and subsequent uptake by macrophages in vitro as a function of terminal layer on LbL PS
beads.
(left) Opsonin binding offluorescently labeled IgG (Human IgG-FITC) (blue = absorption pre-opsonin incubation;
red = absorption post-opsonin incubation), and subsequent (right) opsonization by a murine macrophage cell
line (774A.1)
of LbL nanoparticles as a function of terminal layer. Particles used for this investigation were
polystyrene Texas red-labeled latex beads (PSTXR). Experimentalgroups include: blank PSTXR core, PSTXR/PLL, and
PSTXR/PLL/Xerm, where Xterm was variedfor DXS-, Alg-, and HA50K-terminated systems; n = 3, mean +/- SEM.
164
Appendix B. Osteotropic therapy via LbL Nanoparticles.
0
HO iiOH
OH
n
OH
P.-o
HO- '-OH
0.50
0.50
H 3CO
N
/
N
N
0
N+
0
H 3CO
DMT MM
r.t., 12h, H 2 0
B.1. Synthesis of PAA-Alendronate.
Approximately 40% functionalization achieved.
165
N
0
HO 1i OH
OH
H
HO-
P
'
O
O
OH
H 2N
-OH
- - - - - --
I QDSW
I L6
B
S30-
.5E
r
*iI110100
size (nm)
-20
-
Z 1010
60-
40- 0 20-
A20E
nI
C
600M
-
A 40 -
-60-
500
0
I
I
5um
1
2
3 4
Layer
56
I
B.2. OsteotropicPAA-Alendronate LbL QD characterization.
(A) Dynamic light scattering histogram overlay of pre- and post-LbL functionalized QD8oo NP core. (B) Atomic Force Micrograph of QD 8oo/(PLL/PAA-AI)3
NPs. (C) (-potential datafollowing each LbL deposition step.
166
A4
B
QD 8 0 0/PLL/PAA-Alendronate
-1.2
40000-
+
1
-.-
30000-
Binding
Relative Toxicity
-1.1
-1.0
0
20000-
-
-0.9
W1-
Je
S10000-
-0.8
500O40000-
. . .....
. . .......
,
e . .......
0
0.001
0.01
0.1
Incubation Concentration (nM)
QD/8 0 0 (PLL/PAA-Alendronate)
-+
-+-
Binding
Relatie Toxicity
0.7
1
2
.
-12
C
M
C
-1.0
30000-
1
I-
-0.8
0
-
E
0
-0.9
-
---
20000-
0
)
CO
.2
10000-
It 1!
.A
.
W M W0,18"POP""M
JP
-1 ------
10
100
Incubation Concentration (pM)
1000
.7
D
2.0*j ~
00
Ir,
1.5(
Representative
Q
'I
UM
U.
0
50
100
150
200
hours
B.3. Variation of number of layers and impact on biologicalperformance of PAA-Alendronate targeted LbL NPs.
(A) in vitro binding of 1- and 2-bilayer targeted LbL NPs [QD8oo/(PLL/PAA-Al)x] (B) in vivo fluorescence imaging (640nm, 800nm) following systemic
administration of 4L targeted LbL NPs. (C) quantification (n = 3 mice, 2 hind flank tumors/mice) of tumor-specific accumulation above pre-injection
autofluorescence. (D) representative biodistributiondata at terminal pointfrom imaging (7d) following systemic administrationof 2-bilayer LbL-targeted
QD8oo NPs. Li = liver; S = Spleen; K = Kidneys; H = Heart; Lu = lungs; T = Tumors.
167
QDago/(PLL/PAA 4wK) 3
B.4. Biodistributionof uncoatedand PAA untargeted QD NPs.
(top) uncoated carboxylate-modified QD8oo NPs. (bottom) untargeted, 3-bilayer (PLL/PAA) LbL-functionalized NPs.
168
SA
Q PAA-AIY5 -5
f3QD O
1.6-
00
CU
4-,
80 0
rl
1.4-
0..
Representative bioD
B
C
CU
1.20<
0
0
-3
U,
U,
1.0i
0
50
100
hours
150
3II
200
C
CU
3
0I
0)
a)
L-
B.5. Biodistribution and tumor targetingof PAA-Alendronate targeted QD NP tracking the core and LbL shell simultaneously.
(A) Fold tumor specific accumulation as a function of time of the Cy5.5-labeled PAA-Al terminal layer polymer on the QDqoo surface. (B) Corresponding
representativebiodistribution (n = 3) of the mice at the terminal point (7d). Tissue harvest (liver, spleen, kidneys, heart, lungs, tumors, and gutted skeleton
[ventral and dorsal]) displayed for same animal at terminal point for both channels [PAA-AlCY55; QD8oo]. (C) Representative biodistribution data
correspondingto quantificationin Figure 3.4.D at terminal point. Li = liver; S = Spleen; K = Kidneys; H = Heart; Lu = lungs; T = Tumors.
169
B 30
-~~---"
1,111,-po0en
-7
*Nw
30-
dh =,
1-potential
-t- 1 nm
1 +1-
I
z-avg
dh= '73 -- in,-pote ntial = -2: +- 2f'v
I
I
Dox lipo
z-avg dj = 12:. *4 '-potential - -_14 4-
r
t-"P
!
L PAM-Ai
d.<- 211
tial
13 1r!.
- -2 '_+" - D.;-I W-0
20
20-
E
z 10-
01-10
L2 PAA-Al
.
A 40 I. I blank
z-avg
-
-Q
.
z= 10
E
n
100
I
10
1000
size (nm)
El
ID
100
I.
1000
size (nm)
B.6. Physicochemical characterizationfollowing functionalizationof empty and doxorubicin-loaded liposomal carriers.
Dynamic light scatteringoverlays offormulated liposomes and correspondingfunctionalization via LbL for both (A) empty liposome core (for PK data in
BALB/c) and (B) doxorubicin-loaded liposomes in efficacy studies.
170
A
B
I a
-
100h bioD
T
60-
I
-
40
20-
=
-
0'
e~
'v%'~
C pre
30min
8h
24h
W""""
0
\~
Cs
72h
B.7. Biologicalperformance of PAA-Alendronate targetedliposomal carriers.
(A) in vivo imaging of NCR nude mice at 100h following systemic administration of LbL-targeted empty
liposomal NPs. (B) Corresponding biodistribution data of harvested tissue and skeleton (dorsal), along with
quantification presented as percent recovered fluorescence for tissue (sans skeleton). (C) Representative
biodistribution (ventral in vivo fluorescence imaging) following systemic administration of LbL targeted NPs
(empty lipo/PLLcY5s/PAA-Al) in BALB/c mice for PK studies (presented in Figure 3.5). Li = liver; S = Spleen; K =
Kidneys; H = Heart; Lu = lungs; T = Tumors.
171
B.8. microCT scansfollowing treatment of 132B xenografts with PAA-Alendronate targeted liposomal carriers.
Initial and final representative MCT images of tumor burden (red arrows) in live animals treated with LbLtargeted doxorubicin-loaded liposomes.
172
Appendix C. Scalable Manufacture of Built-to-order
Nanomedicine: Spray-LbL on PRINT@ Nanoparticles
PLGA 200x 2oonm
No XL - PLGA200X2 00/PLL/DXS
750mV
5pgm
5pm
200nm
C.1. AFM of PRINT@ NPs pre- and post-spray LbL.
Particles washed offfollowing deposition (right) on non-crosslinked harvestedparticle array (left).
173
1h
2h
4h
11h
23h
50% GA,
1%HCI
50%
1h
2h
4h
11h
50% GA,
5%HCI
23h
50%
50%
50%
50%
50%
50%
1h
50% GA,
12%HCI
GAconc
50%
1.5h
2h
4h
50%
50%
50%
50%
50%
50%
50%
HCI conc time (h) CA (deg)
37
1
1%
1%
2
56.2
70
4
1%
1%
11
71.8
1%
23
71
1
72.2
5%
73.1
5%
2
74
4
5%
11
76.6
5%
5%
23
75.6
12%
1
76.3
12%
1.5
73.1
2
74.9
12%
12%
4
74.8
37%
0.5
70.3
C.2. Varying crosslinking conditions, incubationperiod - characterizingcrosslinking of PVA via contact angle measurements of PRINT@ NP arrays.
Crosslinking at 37% HCI results in significant loss of particles beyond O.Sh (which was also found not to be long enough to maintain particles during
functionalization),presumably due to acid-catalyzed hydrolysis of PLGA, therefore no data shownfor this sub-optimal crosslinking condition.
-ir
5Pm
5Pm
750mV
I
LA
V4I
S
5&pm OZ5pJm
LO
C.3. AFM amplitude data correspondingto Figure4.3.
175
0.63mV
20-
1
01II
1
1000
100
size (nm)
30
-20-
1 0-
size (nm)
10
1000
I
I
LL
- 10
0
DXL
ID L
0
10
size (nm)
100
1000
1000
1000
*12HOK
oL2
XL
size (nm)
100
30
'20
*core
L6 HA500K
HASOOK
0 -----------------
L4 DXS
20-
-E
S20-
_potential
-27.1 / 3.47mV
/ 1.16mV
-20-
-M20-
0
size (nm)
-283
2.02mV
-388+/
-20-
0-
. qn
r potential
potential
-6.2 +/- 1.04mV
.1
100
lo .
I in-
r potential
-
AU I
S_potential
3.4
-20-
0-
0_
100
PLGAaMoz,,d(Chit/HAsow),
.
- -~
0
1000
size (nm)
PLGAm2Wwn;J(PLL/DXS)
E -4-
0
2
PLGA
1000
size (nm)
nJPLL/HAWW
I
iJIL 15-IL
______________________
75OiiV
5potential
-
25potential
-13.4
20-
0.5mV
1.47mV
10-
10-
0
100
5pM
-23.7
20-
.
&P
II..,6
10-
0
10-
size (nm)
1000
0
200mmn
100
size (nm)
1000
320s XL
20-
L6 HASOO,
15-
10
0o
size (nm)
C.4. Dynamic Light Scattering (DLS) Intensity distributionhistograms displayed throughoutSpray-LbL deposition on PRINT@ PLGA2oox2 oonm nanoparticles.
(-potential analysis ran concurrently with DLS characterization.
176
25
20No.
M
I1
"thin"
"thick"
zavg dh = 221nm
PDI = 0.05
Zavg dh = 238nm
PDI = 0.07
15-
IIIN
105-
-II
0
100
25
20-
I
size (nm)
*
"thin"
1000
"thick"
Zavg dh = 229nm
zavg dh = 214nm
PDI = 0.05
PDI = 0.03
1510C
5
0
II
100
I
I
I I
size (nm)
II
I
1000
C5. Controllingfilm thickness of spray-LbL on PRINT®.
(top, left) Overlay of coated PRINT@ PLGA 200x200nm particles with "thin"[(PLL/HA)3] and "thick"[(Chit/HA)s] coatings. (bottom, left) Overlay of coated
PRINT® PLGA 80x320nm particles with "thin" [(PLL/HA)3] and "thick" [(Chit/HA)s] coatings. (right) Coated (Chit/HA)s 200x200nm and 80x320nm PLGA
particle arrays and subsequent harvested and purified coated NPs.
177
2h XL
PLGA-
/P LL
P LGA,,,,.,,,,-/PL L,.. ./DXS
7somv
I
,I
I
2 )On-
C.6. Tracking deposition with Spray-LbL on PRINT® using AFM,fluorescence imaging.
(top) AFM of each layer on PRINT@ from crosslinked particles (2h) to L4 (PLL/DXS) 2. (bottom) fluorescence imaging using LiCor (left) and confocal
microscopy (right) to track PLLcys.5 deposition.
178
fiChit/HAI.
PLGA.--'sxt
e
etk
*re
9si
bbN
0
0
as
s4
sm
spsp
sm
pm
spm
oobe .ipN6
.unctionOliziston
179
T
loop
C.8. Spray-LbL on PRINT® PLGA8 ox 32 onm NPs.
(top) uncoated, pre-crosslinked 80x320 PLGA particlearrays. (bottom) coated (PLL/HA) 3 80x320 PLGA particle arrays.
180
C.9. Spray-LbL on PRINT@ PLGA 2 00x 2 00nn NPs.
(left) Uncoated, pre-crosslinkedPRINT@ PLGA 200x200nm particle arrays. (right) Coated (PLL/HA) 3 200x200 PLGA particlearrays.
181
C.10. Uncoated, pre-crosslinked TEMs of PRINT® NPs.
(A) PRINT@ PLGA 2oox 2 oonm NPs and (B) PLGA8 ox3 2 onm systems as visualized under TEM.
182
Appendix D. Liposomal compartmentalization of a
hydrophobic-hydrophilic combination chemotherapy regimen
(erlotinib, doxorubicin) for staged, targeted delivery against
triple-negative breast cancer and non-small cell lung cancer
BT-20
-
7.0)108
0
0.
..
-
6.0x108.
0.-
5.0x10 8
M
-o
4.0x108 E
3.0x108
..--
I
S5
1.0) 109.0
0
*a 8.0x108.
15
days
A549
10
20
25
-
6.0X108
M
1+.4. .4.
4.0x108.
2.0x108.
.4.
-
...
(
-G
) am
0
15
days
10
20
25
D.1. Tumor growth following BT-20 and A549xenograft seeding at day 0 (n = 5).
Treatment demonstrated in Figure5.5 began on Day 22 (terminal pointJbrdata above).
183
1 00
Cu
i
Y
pH 5.5
80-
pH 7
SGef
- Afa
0 Er
a) 60
z
40
-*
0
A Dox
A Lap
20'
0
200
400
hours
600
D.2. In vitro release experimentation- release of inhibitors(afatinib, erlotinib,gefitinib, lapatinib)and cytotoxic
(doxorubicin)from dual-drug coated (FP) liposomalformulations.
Release study conducted under sink conditions and agitation at 370C with variation on pH of PBS buffer
(adjusted by titration of citric acid buffer, pH = 4) to 7 and 5.5, respectively, as denoted on the graph.
184
B
A549
1.2
BT-20
1.6
-
A
-+-D
-.-DE
DA
DG
+DE
1.2
-
DA
-- DG
(U)
0.8
-
0.8.
0.6
4
0
24
8
48
24
8
0
72
0.25
+wDE
0.15
-DL
- DA
-- D
-
0.07
+mDE
- DA
-- DG
DL
0
-0.05
8
48
24
0.03
-
0.01
-
-
0.05
0.05
72
48
0.35
-0.01
0
8
24
48
72
8
24
48
72
1
-+D
+oDE
- DA
0.6
-DA
-- DG
DL
0.
2
-DL
0.4
-
-
4.
0.8
-
+aDE
-
0.2
0
0
0
8
24
48
72
0
HOURS
HOURS
D.3. Cleaved Caspase-8, phosphorylated ERK (pERK), and phosphorylated H2AX (pH2AX) western blot
quantificationasfunction of time and treatment in both NSCLC and TNBC cancercells.
(A) A549 and (B) BT-20 cells. Ordinate representative as relative signal intensity with time in hours plotted on
the abscissa. D = doxorubicin liposome; DA = doxorubicin + afatinib liposome; DE = doxorubicin + erlotinib
liposome; DG = doxorubicin + gefitinib liposome; DL = doxorubicin + lapatinibliposome.
185
D
97
0.13
DA
139
0.14
DE
136
0.13
DG
131
0.10
DL
204
0.12
C
99
0.14
CA
112
0.11
CE
134
0.17
CG
127
0.10
CL
204
0.16
D: 97%
D: 5.4%
D: 99%
D: 5.5%
A: 82%
A: 4.4%
D: 97%
D: 5.5%
E: 40%
E: 2.2%
D: 98%
D: 5.5%
G: 85%
G: 4.6%
D: 97%
D: 5.5%
L: 70%
L: 3.8%
C: 97%
C: 11%
C: 99%
C: 9.9%
A: 82%
A: 5.2%
C: 97%
C: 10.5%
E: 40%
E: 2.4%
C: 98%
C: 9.4%
G: 85%
G: 5.8%
C: 97%
C: 10.2%
70%
L: 5.1%
D4. Characterizationof non-functionalized multi-drug and single-drug loaded liposomes (with out folate-PEG).
__L:
Mean hydrodynamic diameter, polydispersity index (PDI), and the drug loading characteristicsare presented.
186
N~il
D.5. Biodistributionof CyS.5 labeled blank FP coated liposomes administeredto tumor-bearingncr nude mice.
Time point from end of experiment (BT20 - 48d post-injection; A549/MDA-MB-468 - 55d post-injection). Li
liver; S = spleen; K = kidneys; H/Lu = heart/lungs; T = tumors; B = brain. Single systemic injection of liposomal
formulation.
187
D.6. Tumor luminescence data temporally resolved following treatment of A549 xenograft bearing ncr nude
mice with single and dual drug doxorubicin-erlotinibFP coated liposomal carriers.
Remediation panels for (right to left) untreated, dox-only targeted liposome (DFP), and combination therapy
liposome (doxorubicin:erlotinib targeted liposome; DEFP) in A549 (NSCLC) xenografts. Panels display tumor
luminescence detected via bioluminescence imaging post-treatment A single administration was conducted
immediately after the "pre"image. Mice were imaged up to 35d as displayed.
188
D.7. Tumor luminescence data temporally resolved following treatment of BT-20 xenograft bearing ncr nude
mice with single and dual drug doxorubicin-erlotinibFP coated liposomal carriers.
Remediation panels for (right to left) untreated, dox-only targeted liposome (DFP), and combination therapy
liposome (doxorubicin:erlotinib targeted liposome; DEFP) in BT20 (TNBC) xenografts. Panels display tumor
luminescence detected via bioluminescence imaging post-treatment. A single administration was conducted
immediately after the "pre"image. Mice were imaged up to 35d as displayed.
189
1 min
10 min
15 min
40 min
1 min
3 min
10 min
20 min
w
0
z
a
CL
LU
x
0
z
D.8. Release of doxorubicin visualized by confocal microscopyfrom FP coated liposomal carriers.
Time course comparing uptake and release of doxorubicin for the liposomal doxorubicin [DFP] (A) and free
doxorubicin (B). Nuclear stain via Hoescht (blue), doxorubicinfluorescence observed in green, nanoparticle(NP)
fluorescence observed in red (Cy5.5).
190
4d
2d
Od
II
I
Dox lipo
(2mg/kg)
-
I
i
CD-Erl
(1mg/kg)
CD-Erl
(1mg/kg)
CD-Erl
(1mg/kg)
)-o
0.
a>) 0
4a2
-
/
/
-
-
-
-
3a>) -WM
-J
0
L.
E 1E
0
L.
0
I4
FE
1
10
20
30
days
Co-releasing in vivo control - timed administrationof single drug doxorubicin liposome with co-treatment
of cyclodextrin-erlotinib complex.
Treatment displayed at top: co-injection of doxorubicin-loaded liposome (DFP, folate-PEGylated) + erlotinibD.9.
cyclodextrin complex in PBS (0.5mg/kg, hydroxypropyl-beta-cyclodextrin) co-injection at beginning of
treatment, followed by subsequent boost doses of erlotinib at 0.5mg/kg on day 2 and 4. Live-animal
bioluminescence images of hind-flank A549 xenografts shown prior to treatment and at days 3, 10, and 30.
Quantification of luminescence (radiance[photons]), normalized to pre-injection tumor luminescence, displayed
at bottom. Average +/- standard deviation fold luminescent change (normalized to pre-injection tumor
luminescence) at day 3 (0.83 +/- 0.61), day 10 (1.74 +/- 0.47), and day 30 (2.37 +/- 1.2) is plotted.
191