Degradable polymeric nano-films and particles as delivery
platforms for vaccines and immunotherapeutics
by
XINGFANG SU
Bachelor of Arts, Master of Science, Materials Science and Metallurgy
University of Cambridge, UK, 2004
Submitted to the Department of Materials Science and Engineering
in partial fulfillment of the requirements for the degree of
DOCTOR OF PHILOSOPHY IN MATERIALS SCIENCE AND ENGINEERING
at the
MASSACHUSETTS INSTITUTE OF TECHNOLOGY
ARCHNES
EC
MASS
OLOGY
September 2012
© Massachusetts Institute of Technology, 2012. All rights reserved.
LIBRARIES
Signature of
Author:
Depa~nintf Materials Science and Engineering
August 2, 2012
Certified
Darrell J. Irvine
Associate Professor of Materials Science and Engineering & Biological Engineering
Advisor
Accepted
by:
Gerbrand Ceder
R.P. Simmons Professor of Materials Science & Engineering
Chair, Departmental Committee on Graduate Students
I
E
Degradable polymeric nano-films and particles as delivery
platforms for vaccines and immunotherapeutics
by
XINGFANG SU
Submitted to the Department of Materials Science and Engineering
On August 2" , 2012 in partial fulfillment of the requirements for the degree of
Doctor of Philosophy in Materials Science and Engineering
ABSTRACT
Degradable polymeric materials provide opportunities for the development of
improved vaccines and immunotherapies by acting as platforms that facilitate the
delivery of molecules to appropriate tissue and cellular locations to achieve therapeutic
outcomes. To this end, we have designed and characterized nano-films and particles
employing a hydrolytically degradable polymer for the delivery of vaccine antigens and
immunotherapeutics. We first describe protein- and oligonucleotide-loaded layer-by-layer
(LbL)-assembled multilayer thin films constructed based on electrostatic interactions
between a cationic poly(p-amino ester) (PBAE, denoted Poly-1) with a model protein
antigen, ovalbumin (OVA), and/or immunostimulatory CpG oligonucleotides for
transcutaneous delivery. Linear growth of nanoscale Poly-1/OVA bilayers was observed.
Dried OVA protein-loaded films rapidly deconstructed when rehydrated in saline
solutions, releasing OVA as non-aggregated/non-degraded protein, suggesting that the
structure of biomolecules integrated into these multilayer films are preserved during
release. Using confocal fluorescence microscopy and an in vivo murine ear skin model,
we demonstrated delivery of OVA from LbL films into barrier-disrupted skin, uptake of
the protein by skin-resident antigen-presenting cells (Langerhans cells), and transport of
the antigen to the skin-draining lymph nodes. Dual incorporation of OVA and CpG
oligonucleotides into the nanolayers of LbL films enabled dual release of the antigen and
adjuvant with distinct kinetics for each component; OVA was rapidly released while CpG
was released in a relatively sustained manner. Applied as skin patches, these films
delivered OVA and CpG to Langerhans Cells in the skin. To our knowledge, this is the
first demonstration of LbL films applied for the delivery of biomolecules into skin. This
approach provides a new route for storage of vaccines and other immunotherapeutics in a
solid-state thin film for subsequent delivery into the immunologically-rich milieu of the
skin.
In parallel, we also developed biodegradable core-shell nanoparticles with a
PBAE core enveloped by a phospholipid bilayer shell for cytosolic delivery, with a view
towards delivery of messenger RNA (mRNA)-based vaccines. The pH-responsive PBAE
component was chosen to promote endosome disruption, while the lipid surface layer was
selected to minimize toxicity of the polycation core. mRNA was efficiently adsorbed via
electrostatic interactions onto the surface of these net positively charged nanoparticles. In
vitro, mRNA-loaded particle uptake by dendritic cells led to mRNA delivery into the
2
cytosol with low cytotoxicity, followed by translation of the encoded protein in these
difficult-to-transfect cells at a frequency of -30%. Particles also promoted cytosolic
uptake of co-delivered anti-tumor toxins in tumor cells resulting in synergistic killing,
demonstrating potential for cancer therapy. In vivo, particles loaded with mRNA
administered intranasally or intratracheally in mice led to the enhanced expression of the
reporter protein luciferase compared to naked mRNA. This system may thus be
promising for noninvasive delivery of mRNA-based vaccines.
Thesis Supervisor: Darrell J. Irvine
Associate Professor of Materials Science and Engineering & Biological Engineering
3
Table of Contents
ACKNOW LEDGEM ENTS.......................................................................................................................
10
1.
12
BACKGROUND AND SCOPE OF THESIS.............................................................................
1.1.
DELIVERY OF VACCINES AND IMMUNOTHERAPEUTICS...........................................................
1.1.1.
1.1.2.
1.2.
1.3.
1.4.
Transutaneousand mucosal delivery.............................................................................
Cytosolic delivery............................................................................................................................18
SCOPE AND OUTLINE OF THESIS...................................................................................................
2.
SYNTHESIS AND CHARACTERIZATION OF POLYELECTROLYTE MULTILAYER
FILM S......................................................_...............................................................................................
2.1.
2.2.
INTRODUCTION.....................................................................................................................................
MATERIALS AND M ETHODS.........................................................................................................
2.2.1.
2.2.2.
2.2.3.
2.3.
12
13
NON-INVASIVE TRANSCUTANEOUS AND MUCOSAL ADMINISTRATION ROUTES ................. 14
CYTOSOLIC DELIVERY FOR VACCINES, IMMUNO-MODULATION AND THERAPY..................15
DEGRADABLE POLYMERIC MATERIALS AS A DELIVERY PLATFORM......................................16
1.4.1.
1.4.2.
1.5.
Vaccines..............................................................................................................................................12
Imm unomodulatoryagents................................................................................................
Materials............................................................................................................................................22
Film preparation............................................................................................................................22
Film characterization..................................................................................................................
RESULTS AND DISCUSSION.................................................................................................................
17
18
21
21
22
23
23
2.3.1.
2.3.2.
2.3.3.
2.3.4.
2.3.5.
2.4.
Ovalbumin loading in PBAEfilms....................................................................................
23
Rapid release of ovalbuminfrom deconstructedfilms upon rehydration...........25
Film assembly on flexible substrates.............................................................................
26
Characterizationof releasedovalbumin via SDS- and native PAGE................27
Dual loading and release of ovalbumin and CpG .....................................................
29
CONCLUSIONS......................................................................................................................................
31
3.
SYNTHESIS AND CHARACTERIZATION OF LIPID-ENVELOPED POLYMER
NANOPARTICLES....................................................................................................................................
3.1.
3.2.
INTRODUCTION.....................................................................................................................................
MATERIALS AND METHODS .........................................................................................................
3.2.1.
3.2.2.
3.2.3.
3.2.4.
3.2.5.
3.2.6.
3.3.
3.4.
33
34
Materials............................................................................................................................................34
Synthesis and Characterizationof Lipid-coatedPBAE Nanoparticles............34
pH-sensitivity assay.......................................................................................................................35
mRNA preparation........................................................................................................................
35
Loading and release of RNA ...............................................................................................
36
RNA DegradationProtectionAssay .................................................................................
36
RESULTS AND DISCUSSION................................................................................................................
3.3.1.
3.3.2.
3.3.3.
3.3.4.
33
36
Design, Synthesis, and Characterizationof Lipid-coatedPBAE particles............37
pH-responsivenessof lipid-coatedPBAE nanoparticles.......................................... 40
Efficient RNA loading on particlesurface....................................................................
42
Protectionof surface-loadedRNA from nuclease degradation..........................
45
CONCLUSIONS......................................................................................................................................
46
4.
TRANSCUTANEOUS DELIVERY MEDIATED BY POLYELECTROLYTE MULTILAYER
FILM S IN M URINE EAR SKIN.........................................................................................................
47
4
4.1.
INTRODUCTION.....................................................................................................................................47
4.2.
M ATERIALS AND METHODS ...............................................................................................................
4.2.1.
4.2.2.
4.2.3.
4.3.
M aterials............................................................................................................................................48
In vivo m urine earskin penetration................................................................................
Analysis of cells in auricularlym ph nodes....................................................................
48
48
49
RESULTS AND DISCUSSION.................................................................................................................49
49
4.3.1.
Disruptionof stratum corneum barrierby tape-stripping...................................
4.3.1.
Penetrationof ovalbumin releasedfrom films into murine ear skin and uptake
by Langerhan cells...............................................................................................................................................50
Dualpenetration of ovalbumin and CpG releasedfrom films into murine ear
4.3.2.
53
skin and uptake by Langerhan cells............................................................................................................
55
4.3.3.
Transportof antigen to draining lymph nodes...........................................................
4.4.
CONCLUSIONS........................................................................................................................................56
5.
CYTOSOLIC DELIVERY MEDIATED BY LIPID-ENVELOPED POLYMER
NANOPARTICLES IN VITRO...........................................................................................................
5.1.
INTRODUCTION.....................................................................................................................................58
5.2.
M ATERIALS AND METHODS ...............................................................................................................
5.2.1.
5.2.2.
5.2.3.
5.2.4.
5.2.5.
5.2.6.
5.3.
59
M aterials............................................................................................................................................59
60
Analysis of endosomal disruption......................................................................................
Cytotoxicity assay...........................................................................................................................60
60
In vitro transfection of DCs..................................................................................................
Cytosolic delivery of antitumor toxin and in vitro tumor cell killing................61
Statisticalanalysis.........................................................................................................................62
RESULTS AND DISCUSSION.................................................................................................................62
5.3.1.
5.3.2.
5.3.3.
5.3.4.
5.4.
58
Endosomal escape mediated by lipid-coatedPBAE nanoparticles....................62
65
Cytotoxicity of lipid-coatedPBAE nanoparticles......................................................
Transfection of dendritic cells with lipid-coated PBAE nanoparticlesin vitro.66
Synergistic tumor cell killing with lipid-coatedPBAE nanoparticlesin vitro ...70
CONCLUSIONS........................................................................................................................................74
MUCOSAL DELIVERY OF MRNA BY LIPID-ENVELOPED POLYMER NANOPARTICLES
6.
76
IN VIVO......................................................................................................................................................
6.1.
6.2.
6.2.1.
6.2.2.
6.2.3.
6.3.
INTRODUCTION.....................................................................................................................................76
M ATERIALS AND METHODS ...............................................................................................................
M aterials............................................................................................................................................77
Anim al studies .................................................................................................................................
Statistical analysis.........................................................................................................................78
77
77
RESULTS AND DISCUSSION.................................................................................................................78
In vivo transfection with lipid-enveloped nanoparticlesadministered
6.3.1.
intranasally.............................................................................................................................................................78
In vivo transfection with lipid-enveloped nanoparticlesadministered
6.3.2.
intratracheally.......................................................................................................................................................82
6.4.
7.
CONCLUSIONS AND FUTURE W ORK..................................................................................
7.1.
8.
CONCLUSIONS........................................................................................................................................85
86
DEGRADABLE PBAE-BASED NANO-FILMS AND PARTICLES AS A PLATFORM FOR
DELIVERING VACCINES AND THERAPEUTICS .............................................................................................
7.2.
ISSUES FOR FUTURE W ORK ................................................................................................................
86
BIOGRAPHICAL NOTE...........................................................................................................
90
5
87
9. APPENDIX A: LIPIDS COMPRISING THE LIPID SHELL ON PBAE
NANOPARTICLES .....................................................................................................................
91
10.
APPENDIX B: IDENTIFICATION AND DISSECTION OF AURICULAR LYMPH
N O D ES .......................................................................................................................................
92
11.
APPENDIX C: PROTOCOL FOR INTRATRACHEAL ADMINISTRATION IN MICE.. 94
12.
APPENDIX D: INTRAPERITONEAL IMMUNIZATION WITH MRNA.....................
13.
APPENDIX E: CYTOSOLIC DELIVERY OF SOLUBLE MOLECULES IN TUMOR
CELLS .......................................................................................................................................
14.
REFERENCES...........................................................................................................................111
6
98
104
List of Figures
Figure 1-1. Schematic illustrating the processes involved during antigen presentation by
den dritic cells............................................................................................................
16
Figure 2-1. Chemical structures of Poly- 1 and Poly-2 used in this study. ................... 22
Figure 2-2. Loading of OVA into multilayers prepared by sequential deposition of 15
bilayers of Poly-l/O VA . ........................................................................................
24
Figure 2-3. Growth curve of (Poly-1/OVA) multilayers for two different deposition
con d ition s..................................................................................................................
25
Figure 2-4. Cumulative normalized kinetics of OVA release and representative thickness
changes with time from rehydrated LBL films......................................................
26
Figure 2-5. Photograph of (Poly- 1/OVA-Texas Red) 40 film assembled on a flexible
PD M S substrate. ...................................................................................................
27
Figure 2-6. Non-reducing SDS-PAGE of supernatants from (Poly-1/OVA) 40 films
rehydrated for 30 m in in PBS. ..............................................................................
28
Figure 2-7. Native PAGE analysis of OVA released from (Poly-1/OVA)40 films after a
30 m in rehydration in PB S....................................................................................
28
Figure 2-8. Dual loading and release of OVA and CpG...............................................
31
Figure 3-1. Schematic of structure and composition of lipid-coated PBAE particles and
m RN A cargo association. .....................................................................................
37
Figure 3-2. Schematic illustrating the double emulsion and nanoprecipitation process
used to synthesize lipid-coated PBAE nanoparticles............................................
38
Figure 3-3. Size histograms of synthesized lipid-coated PBAE nanoparticles............. 39
Figure 3-4. Representative cryoEM image of lipid-enveloped PBAE particles....... 40
Figure 3-5. Proposed mechanisms of endosomal disruption by PBAE particles. ......... 41
Figure 3-6. pH-dependent solubility profile of lipid and PVA stabilized particles..... 42
Figure 3-7. RNA adsorption to lipid-coated PBAE particles in water. ........................
43
Figure 3-8. Loading of poly I:C and mRNA on lipid-coated PBAE nanoparticles.......... 44
Figure 3-9. Size histograms of PBAE particles before and after mRNA adsorption. ...... 44
Figure 3-10. Release profile of poly I:C from PBAE particles at physiological conditions.
...................................................................................................................................
45
Figure 3-11. Surface loading of mRNA on PBAE particles protects mRNA against
nuclease degradation..............................................................................................
46
Figure 4-1. Schematic illustration of the events taking place following the application of
PE M patch onto skin.................................................................................................
50
Figure 4-2. Histological sections of murine ear skin before and after tape-stripping....... 50
Figure 4-3. Penetration of ovalbumin in tape-stripped murine ear skin. ....................... 52
Figure 4-4. Co-penetration of ovalbumin and CpG in tape-stripped murine ear skin...... 54
Figure 4-5. Transport of ovalbumin to draining lymph nodes.......................................
56
Figure 5-1. Endosomal escape by lipid-coated PBAE particles. ...................................
64
Figure 5-2. Cytotoxicity profiles of PBAE-based particles..........................................
66
Figure 5-3. mRNA adsorption does not hinder endosomal rupture by particles.......... 67
Figure 5-4. Lipid-coated PBAE particles mediate cytosolic mRNA delivery and
transfection in dendritic cells in vitro. ..................................................................
70
7
Figure 5-5. Schematic illustration of the events taking place following the co-delivery of
71
im m unotoxin and particles.....................................................................................
Figure 5-6. Cytosolic delivery of soluble phalloidin by lipid-coated PBAE particles. .... 72
Figure 5-7. Killing of A431 target cells by particle-chaperoned immunotoxin. ........... 74
Figure 6-1. In vivo transfection with intranasally administered lipid-coated PBAE
82
particles.....................................................................................................................
Figure 6-2. In vivo transfection with intratracheally administered lipid-coated PBAE
83
p articles.....................................................................................................................
Figure 6-3. Transfected tissues in dissected mice imaged post-mortem. ..................... 84
Figure 6-4. Particle uptake by MHC-Class II-GFP-expressing cells in lungs.............. 85
Figure 10-1. Schematic illustrating the location of the auricular lymph nodes in mice... 93
95
Figure 11-1. Intratracheal instillation procedure. .........................................................
Figure 11-2. Preparation of the Exel Safelet IV catheter for intratracheal instillation..... 96
Figure 11-3. Recovery following intratracheal administration..................................... 96
Figure 12-1. ELISPOT assay on mice immunized with mirus-mRNA complexes relative
100
to particle-adsorbed mRN A . ...................................................................................
101
of
lipid-coated
PBAE
particles..................
administration
Figure 12-2. Intraperitoneal
Figure 12-3. Transfected tissues in dissected mice imaged post-mortem. ..................... 102
Figure 12-4. Distribution of bioluminescence at various tissue sites. ............................ 103
Figure 13-1. Cytosolic delivery of calcein by lipid-coated PBAE particles in tumor cell
10 5
lin es.........................................................................................................................
Figure 13-2. Cargo molecules must be localized in the same endolysosome as particles
106
for increased cytosolic uptake.................................................................................
Figure 13-3. Effect of molecular weight and charge of cargo molecules on cytosolic
110
uptake w ith lipid-coated PBAE particles................................................................
8
List of Tables
Table 3-1. Size and zeta potentials of PBAE nanoparticles determined by dynamic light
scatterin g ...................................................................................................................
40
Table 9-1. Lipid molecules and fluorescent lipid analog comprising the lipid shell of
PB A E nanoparticles ..............................................................................................
91
9
Acknowledgements
I would sincerely like to thank the following people:
My thesis advisor, Prof. Darrell Irvine for his tutelage, patience and encouragement in
numerous aspects and being an inspiration throughout the course of my PhD;
My thesis committee, Prof. Paula Hammond and Prof. Michael Rubner for their
enthusiasm, guidance and insights at different stages of this work;
My collaborators, without whom much of this work would not have been possible:
Hammond Lab, MIT:
Dr. Byeong-Su Kim for his expertise in layer-by-layer assembly and imparting his
experience in managing projects and preparing manuscripts in my early years.
Kavanagh Lab, Ragon Institute of MGH, MIT and Harvard:
Jennifer Fricke and Dr Daniel Kavanagh for their expertise in mRNA and providing
invaluable help and support with supply of reagents and assistance in running
experiments.
Past and current members of the Irvine Lab:
Dr. Andrew Miller for introducing me to the lab and helping me get things started. Sheree
Beane and Erin Morehouse for being responsible lab managers and keeping the lab
together. Postdocs, fellow graduate students and technicians who provided discussions,
advice and help on multiple occasions.
The veterinarians and staff at the MIT Division of Comparative Medicine, especially Liz
Horrigan, who provided animal husbandry and worked with me to learn techniques,
establish protocols, and address the technical challenges of working with animal models.
Staff at MIT's core facilities whose diligent services facilitated various aspects of this
work. Thanks also to the students from other labs who kindly agreed to help when
approached: Michel Dupage for help with intratracheal delivery. Recent collaborations
with Dr. Giuseppe Battaglia and his student Carla Pergorano, Prof. Dane Wittrup and his
student Nicole Yang, also provided unique opportunities and insights.
Agency of Science, Technology and Research for providing me the opportunity and
funding to pursue a PhD in the United States. This research was also supported by the
Institute for Soldier Nanotechnolgy, Howard Hughes Medical Institute and Ragon
Institute of MGH, MIT, and Harvard.
10
Friends in MIT who have generously shared their experiences, given timely advice on
various fronts and made a pleasant difference in my MIT experience: Shyue Ping,
Zhiyong, Henry, Shireen, Wui Siew, Trina, Vincent and Huili.
My family:
Siak Joo Soh, Ee Lee Yeo, Songliang Chua for their understanding, support and concern.
This thesis is dedicated to them.
11
1. Background and scope of thesis
1.1. Delivery of vaccines and immunotherapeutics
In recent years there has been a surge in the perceived potential for applying
vaccination and immunotherapy approaches towards the prevention and treatment of a
diverse range of diseases. In addition to the traditional role of vaccination in prophylactic
prevention against infectious diseases, strategies targeting the immune system for
protection and therapy against cancer, allergies, autoimmune and inflammatory diseases
in both therapeutic as well as prophylactic mode have been increasingly pursued.' In the
area of infectious diseases, vaccines against HIV and malaria are the subject of intensive
research. 2 For immunotherapy against cancer, ways to promote more effective and
targeted killing of tumor cells are being sought after. Beside allergic and autoimmune
diseases such as asthma and multiple sclerosis, researchers also hope to modify the
immune system functions to treat conditions with an inflammatory component such as
Alzheimer's and heart diseases. 3 In all cases, the key lies in the effective delivery of
vaccines and immunodulatory agents via suitable administration routes to access target
cells and subcellular compartments to initiate processes leading to a protective or
therapeutic response.
1.1.1.
Vaccines
The first generation of vaccines have been developed based on the use of live
attenuated or inactivated naturally immunogenic pathogens, such as the highly successful
smallpox, polio, and diphtheria vaccines. 8 A large proportion of currently licensed
vaccines are based on live or inactivated organisms, primarily attributed to their high
potency.9 However, due to their complex nature, such vaccines can vary widely in quality
from batch to batch. Inactivation of pathogens can also lead to changes in antigenicity
and possible reversion to virulent forms. Moreover, live systems have associated adverse
reactions which can range from simple headache to encephalitis (MMR), intussception
(rotavirus), vaccine associated disease (polio) and even death (smallpox).10 Whilst rare,
inactivated vaccines can also cause serious adverse effects varying from nausea to
anaphylactic reactions and neurological complications.1 0 Storage and effective delivery
of these vaccines can also be a major issue in developing countries with limited health
service infrastructure. In order to address the above deficiencies, a second generation of
vaccines based on defined recombinant protein-, peptide-, or nucleic acid-based antigens,
termed subunit vaccines has emerged, following recent rapid advances in genomics and
proteomics. Although improving the safety profile, their effective implementation is
limited since they usually require adjuvants to induce the desired immunological
responses.",
12
1.1.2.
Immunomodulatory agents
When subunit vaccines lacking endogenous danger signals present in pathogens
are delivered alone, a lack of responsiveness may occur. NaYve antigen-specific T cells
may recognize the antigen but are not sufficiently activated and can even become
tolerized. This can be remedied by delivering antigens in the presence of appropriate
immunomodulatory agents as adjuvants that can significantly enhance the immune
response to the antigen. Adjuvants modulate humoral, cellular, and/or mucosal immunity
to alter the immunogenicity of a vaccine antigen, competition between multiple antigens,
the duration of immunity, and/or the type of immune response (such as the avidity,
specificity, or isotype of antibodies produced).' 3 This is achieved by influencing cytokine
production through the activation of major histocompability complex (MHC) molecules,
costimulatory signals, or through related intracellular signaling pathways.
Antigen presenting cells known as dendritic cells (DCs) are known to respond to
pathogen-derived biomolecules referred to as pathogen-associated molecular patterns
(PAMPs) through germline-encoded pattern recognition receptors (PRRs). One of the
major PRR classes is the Toll-like receptor (TLR) family, members of which recognize a
large number of pathogen-derived ligands. Although TLRs constitute the most
extensively studied class of PRRs, other classes include virus-sensing (RIG-I)-like
receptors (RLRs) such as RIG-I and MDA-5, and bacteria-sensing nucleotide-binding
oligomerization domain (NOD)-like receptors such as NOD 1/2 and NALP1/3.14
A number of TLR agonists are known. TLR1 and TLR2 form dimers that
recognize triacyl lipopeptides such as Pam3CysK4. TLR2 and TLR6 dimerize to
recognize diacyl lipopeptides (Pam2CSK4), lipoteichoic acid, zymosan, porins, bacterial
peptidoglycan, and lipoarabinomannan. TLR3 recognizes double-stranded RNA,
including the synthetic agonist Poly I:C, which can also trigger RIG-1 and MDA5
cytosolic RNA sensors. TLR4 recognizes a broad range of molecules, the best-studied of
which is bacterial lipopolysaccharide (LPS) and its safer analog, Monophosphoryl Lipid
A (MPLA). TLR5 recognizes flagellin, which, being a protein, can also serve as an
antigen and source of CD4' helper peptide epitopes. TLR7 and TLR8 recognize singlestranded viral RNA, and can be triggered with small-molecule drug agonists such as
imiquimod and resiquimod. TLR9 recognizes bacterial DNA, and can be triggered
through synthetic unmethylated CpG motifs. These compounds constitute an extensive
repertoire of candidate immunostimulatory agents that can be co-delivered with subunit
vaccines to drive immunity. 15-19
In addition to acting as adjuvants to enhance immune response to subunit vaccines,
the delivery of immunomodulatory agents themselves as therapeutics is also of
considerable interest. For example, imiquimod, a TLR7 agonist, is the active ingredient in
the topical cream Aldera for treatment of skin lesions caused by the papilloma virus
infection. 20 In addition to activating innate immunity, synthetic CpG oligonucleotides, a
TLR9 agonist, is also known to induce cytokines that promote Thi immunity and can be
used to treat or prevent undesired Th2-dominated immune responses, such as allergy.2 1
Furthermore CpG oligonucleotides have also be used in tumor immunotherapy where it
exerts its effect by improving stimulation of DCs, thereby enhancing tumor antigen
presentation. 22, 23 Therefore, delivery of immunomodulatory agents can facilitate the
development of prophylactic and therapeutic vaccines as well as immunotherapy.
13
1.2. Non-invasive transcutaneous and mucosal administration routes
To achieve a desired immune response, vaccine antigens and immunomodulatory
agents must be effectively delivered to target cells in the immune system. The immune
cells most frequently targeted include B cells, macrophages and dendritic cells (DCs),
which are known as professional antigen-presenting cells (APCs).24 Among APCs, DCs
are of particular interest because they present antigen to their cognate naive T-cell
partners and direct the resultant immune response - induce anergy, tolerance or
immunity. 25
APCs reside in almost every tissue in the body, sampling antigen in their
environment. Epithelial and mucosal tissues have strong immunosurveillance activity
especially since they form barriers to the outside world through which most pathogenic
entry occurs. In these tissues, dendritic cells and macrophages detect danger signals and
antigens, signal other immune cells to the site by means of chemokine secretion and
migrate to the nearest draining lymph node to activate an immune response. The skin and
mucosal tissues are therefore common target tissues for vaccines and immunotherapy and
different immune responses can be achieved in different target tissues. 26
Administration of vaccines and immunotherapeutics through the skin and mucosal
surfaces is both appealing and challenging. Since most pathogens invade the body
through these surfaces, establishment of an immune response in these tissues would be
especially beneficial for providing protection. However, administration via parenteral
routes is typically only effective in eliciting systemic immunity while mucosal and
transcutaneous administration can elicit both systemic and mucosal immune responses.
Moreover, vaccines administered at one mucosal site have been shown to elicit an
immune response in other mucosal sites mediated by trafficking of effector immune cells
between mucosal tissue compartments. For example, nasal immunization has been found
to give rise to substantial IgA and IgG antibody responses in the human cervicovaginal
mucosae. 27-29 Similarly in mice, transcutaneous immunization may induce a mucosal
immune response in the female genital tract.30 This is of special interest for vaccination
against HIV where mucosal immune response at the vaginal tissue is critical for
protection.
From a drug delivery point of view, the non-invasive nature of these
administration routes also confers practical advantages Current vaccine and therapeutic
delivery is largely needle-based, 3 1 but a number of inherent risks and disadvantages to
needle-based delivery have been recognized, such as the need for cold storage of liquid
formulations, 3
the requirement of trained personnel for administration, and reduced
safety due to needle re-use and needle-based injuries. 33 To address these limitations,
vaccination and therapeutics administration through the skin and mucosae represents a
promising alternative administration route.34-36
Despite the multi-fold benefits associated with transcutaneous and mucosal
administration, currently there exist only a few examples of such vaccine and
immunotherapy formulations such as NasalFluMist, an intransal vaccine against
influenza, and Aldera, a topical cream containing imiquimod for treating skin
carcinomas. A large part of this is due to the difficulties encountered in delivering
molecules across the barriers put forth by these tissues, which are also responsible for
keeping pathogens out. As such, it is increasingly appreciated that delivery of vaccines
14
and immunotherapeutics via these routes will require delivery platforms that can help
deliver molecules more efficiently to the immune system in these tissues.
1.3. Cytosolic delivery for vaccines, immuno-modulation and therapy
In addition to understanding the tissue and cellular targets for vaccines and
immunotherapeutics, it is important to also understand the subcellular targets. APCs
process antigen from the cytosol or following endocytosis for display in major
histocompability complex (MHC) class I molecules and/or MHC class II molecules,
respectively.3 7 Antigens presented in the context of MHC class I molecules are
recognized only by CD8' T cells giving rise to cellular immunity based on cytotoxic
CD8' T cells, whereas those bound to MHC class II molecules are recognized by CD4+ T
cells which can prime B cells to produce antibodies thereby promoting humoral
immunity. The detection of both antigens and immunomodulatory agents is complex and
takes place in different compartments of the cell. For example, whereas the endosomal
compartment is the interesting target for MHC class II loading, MHC class I presentation
requires the antigen payload to be present in the cytosol. (Figure 1-1). With regards to
danger signals for APC activation, in the TLR family, receptors for some ligands (such as
bacterial cell-wall components, which bind to TLR4, or bacterial flagellae components,
which bind to TLR5) are present on the plasma membrane, whereas receptors for others
(double-stranded RNA, which binds to TLR3, single-stranded RNA, which binds to
TLR7, or unmethylated bacterial CpG DNA, which binds to TLR9) are present and active
within the endolysosome. 38, 39 The spatial details of antigen and danger-signal delivery
are therefore important. In particular, cytosolic delivery has significant impact on the
effectiveness of vaccine antigens and immunotherapeutics.
For antigen delivery to drive cellular immune responses (especially critical for
diseases such as HIV and cancer) antigens must be available within the cytosol in order
for presentation by MHC class I molecules to occur. This can be also achieved in some
degree by a mechanism known as cross-presentation in non-infected DCs that take up
exogenous antigen.3'40 However, the presentation of antigens to cytotoxic T cells can be
greatly amplified by delivery of protein antigen to the cytosol, where the DC intracellular
machinery can load them efficiently onto MHC class I molecules for presentation to
CD8+ T cells.41, 42 In the case of gene-based vaccine antigens such as antigen-encoding
DNA or mRNA, they must be delivered to the cytosol and subsequently the nucleus (for
DNA) in order to be expressed to produce the encoded antigen.
Likewise, for the delivery of immunotherapeutics, certain immunostimulatory
molecules, such as viral RNA and DNA mimics that trigger anti-viral immune responses,
operate by binding to RNA and DNA sensors present in the cytosol of DCs.43 In addition
to triggering the production of cytokines that can induce the synthesis of cell proteins
with antiviral activity and also shape the adaptive immune response, they have also been
shown to have anti-tumor effect. Recently, cytosolic delivery of poly I:C, a synthetic
dsRNA analog, using cationic liposomes or polyethyleneimine (PEI), has been shown to
enhance tumor cell killing in vitro and reduce tumor growth in vivo in
44 several different
tumor cell lines and models, demonstrating promise for tumor therapy. -49
Immunotoxins represents another class of immunotherapeutics effective against
cancer and are a promising approach to the targeted delivery of highly potent, cancerspecific, cytotoxic agent. 50' 51 They are composed of a targeting moiety specific for cancer
15
cell (derived from antibodies or other cell-binding proteins) either chemically conjugated
or genetically fused to highly cytotoxic plant or bacterial protein toxins. The potency of
an immunotoxin is dependent on the ability to deliver the toxin to the cytoplasm, which is
commonly considered to be the rate-limiting step. While the inclusion of toxins with
domain facilitating cytoplasmic access can address this, they can also lead to increased
nonspecific toxicity in vivo. 52, 53 The use of endosomal disrupting materials to facilitate
delivery of targeted toxin may be useful to overcome this limitation.
Finally, efficient cytosolic delivery in DCs can also be used to deliver plasmid
DNA or siRNA to amplify or suppress adaptive immune responses for vaccines or
immunotherapy. For example dual delivery of interleukin 10 (IL-10)-targeted small
interfering RNA (siRNA) and DNA vaccines to DCs has been shown to enhance immune
response and modulate it toward a stronger Thi phenotype.5 4 However, transfection of
DCs is notoriously inefficient.ss-57 As DCs would readily take up particles through
endocytosis or phagocytosis, particulate delivery systems that can trigger efficient
endosomal escape could potentially be used to mediate intracellular delivery to DCs.
Pahgn Macropinocytosts
SiA .,&%
Endocylosis
I
__
.N.
Nucleus
Figure 1-1.
Schematic illustrating the processes
involved
during antigen
presentation by dendritic cells.
(From Jeffrey A. Hubbell, Susan N. Thomas & Melody A. Swartz, Nature 462, 449-46,
2009.58)
1.4. Degradable polymeric materials as a delivery platform
The need for suitable delivery systems for vaccine and immunotherapeutics has
remained an enormous challenge for the pharmaceutical and vaccine industries. Indeed,
16
drug delivery was recently cited as one of the top 10 technologies that will significantly
impact global health.59
Improving delivery platforms is critical for the development of new generation
vaccines and immunotherapies that are increasingly designed to elicit cellular immune
responses, paramount for targeting chronic infectious diseases that may have an
intracellular stage (e.g. HIV, herpes, hepatitis C, malaria and tuberculosis) as well as for
therapeutic vaccines against cancer, autoimmune diseases or allergies. 60 New vaccines
are also being developed to elicit mucosal immune responses, for example to protect
against pathogens such as influenza virus, HIV, HSV or human oncogenic or wartassociated papilloma viruses. These efforts require the recruitment of cellular or mucosal
immune effector mechanisms and necessitate delivery platforms designed for cytosolic
delivery, alternative routes of administration, new formulations, and new adjuvant
systems.6 0
For clinical translation, biodegradable systems are also desired to ensure
biocompatibility. The use of degradable polymeric materials in vaccine delivery is largely
based on the use of micro- and nanoparticles, motivated by the effective uptake of many
particulate systems by APCs and the desire to control the duration of exposure to antigen
via controlled release. 61-64 Such systems improve immune responses not only due to their
ability to control the release rate of their components, but also due to the inherent potency
of degradable particles as materials for vaccine delivery. 65 -67 Additionally, particles can
co-deliver immunostimulatory molecules on the same particle, targeting multiple classes
of molecules to the same intracellular compartment. 68-7
1.4.1.
Transutaneous and mucosal delivery
As mentioned in Chapter 1.2, transcutaneous and mucosal delivery is
predominantly limited by the need to overcome barriers associated with the transport of
molecules across these tissues. In the skin, the stratum corneum constitutes the principal
barrier that blocks the access of molecules to underlying viable cells. In the nasal
passages and airways, the mucus layer overlying the epithelium not only acts as a
penetration barrier and an active site for enzymatic digestion, the associated mucociliary
clearance mechanisms also actively clear away any deposited molecules.
As such, the development of trancutaneous and mucosal vaccines and
immunotherapy requires efficient antigen and adjuvant delivery systems. Ideally, such
systems should (i) facilitate antigen/adjuvant transfer across membranes by protecting
molecules from physical elimination and enzymatic degradation, (ii) promote uptake by
relevant cells including APCs and specialized epithelial cells that contribute to the
induction of an immune response.
For transcutaneous delivery of vaccine antigens and immunotherapeutics, solidstate delivery systems that can stabilize sensitive biomolecules and facilitate storage in a
dry state are desired. In addition, the ability to incorporate and separately regulate the
release kinetics of multiple components, that may include both skin penetration enhancers
as well as therapeutic cargoes, will allow the effective orchestration of transport across
barrier and subsequent immunological processes necessary to achieve the therapeutic
goals. Finally, the delivery system should be compatible with a variety of transcutaneous
17
delivery settings ranging from simple skin-adhesive patches to microneedle arrays for
maximal flexibility.
In the case of mucosal delivery, particulate delivery systems may provide various
advantages over the delivery of soluble molecules. First and foremost, there is a need to
protect the biomolecules against enzymatic degradation, particularly in the mucus layer
that has been evolved to inactivate any foreign entities. Furthermore, particulate antigen
can be taken up by specialized epithelial cells known as M-cells, a member of the
mucosal-associated lymphoid tissue (MALT), where they are processed and
preferentially directed to the APCs, in contrast to soluble antigen. 72-76 It has also been
demonstrated that particulate delivery systems, particularly those bearing cationic charges
such as chitosan particles and cationic liposomes, can enhance muco-adhesiveness,
reducing the clearance rate and thereby 73
increasing
the contact time of the delivery system
75 77
with the mucosae, facilitating transport. ' '
1.4.2.
Cytosolic delivery
To enable delivery of membrane-impermeable molecules into the cytosol of cells,
much research has been directed at the development of synthetic chaperones that can
facilitate transport of hydrophilic and macromolecules to the cytosol. 78 Approaches
include the use of membrane-penetrating peptides, 79 pathogen-derived pore-forming
proteins, 80, 81 and "endosome escaping" polymers or lipids that disrupt the endosomal
membrane in response to reduced pH, which occurs in these compartments.82-88
While many of these approaches show promise, strategies that can promote highly
efficient delivery of molecules into the cytosol while avoiding unacceptable cytotoxicity
are still sought. In addition, many of the chaperone molecules that efficiently aid
transport of macromolecules into the cytosol are formulated with drug cargos by physical
complexation of the chaperone and drug (e.g., polyplexes or lipoplexes of cationic
polymers/lipids with DNA), forming nanoparticles whose size, stability, and properties
are highly dependent on formulation parameters including the properties of the drug
cargo, the drug-to-chaperone weight ratio, and the characteristics of the surrounding
environment (pH, ionic strength, and presence/absence of serum proteins).82' 89 This
implies that the choice of cationic residues will simultaneously influence both drug
binding and endosomal escape properties of these materials. Furthermore, the lack of
control over chaperone/drug particle size and stability is a concern since particle size is a
critical determinant of cellular uptake in vitro and biodistribution and toxicity in vivo. A
modular design that would allow for the separate formulation of the cargo and delivery
agent is therefore also desirable.
1.5. Scope and outline of thesis
This thesis explores nanofilms and lipid-coated nanoparticles based on a pHsensitive and hydrolytically degradable polymer, poly(p-amino ester) (PBAE), 9 as
versatile platforms for combinatorial vaccines and immunotherapy, with a focus on
administration via noninvasive routes. The goals of the thesis are to explore the ways in
which delivery of vaccine and immunotherapeutics can be improved with these
degradable polymeric delivery platforms.
18
PBAEs have been previously shown to be biocompatible and degradable, to build
polyelectrolyte multilayers (PEM) with DNA that transfect cells in vitro for potential
gene delivery applications. 91-95 It has also been used to fabricate PEM films with
controlled erosion and tunable drug release. 96 98 The same polymer has also been used to
construct microparticles to deliver DNA vaccines effective against cancer, 84 nanoparticle
capable of facilitating tumor targeted delivery of chemotherapeutics 99-101 and intracellular
delivery of siRNA.10 ~104 Finally, while beyond the scope of this thesis, it is also possible
to envisage embedding the particles within films to construct a modular delivery platform
that can combine functionalities, segregate materials assembly/release properties from the
characteristics of individual therapeutic molecule, thereby creating a platform truly
generic for delivering a wide range of antigens and immunotherapeutics.
We contemplate the targeting and penetration of barriers as objectives for material
design. One way to target APCs is to administer therapeutics to a tissue that contains a
relative high frequency of such cells, for example, the epithelial and mucosal tissues that
guard the gateways to the body. To gain access to APCs residing in these tissues, the
delivery system must help mediate the transport of intact molecules past the barrier layers
after depositing them onto the skin and mucosae such as nasal passages and the airways.
Within the cell, the barrier of entry to the intracellular space presented by the endosomal
membrane also needs to be overcome for therapeutics whose functioning is contingent
upon access to cytosolic cellular machinery. We also explore the potential of using these
delivery materials to control the way that these therapeutics are delivered, for example
via co-delivery of multiple moieties to the same target cell as well as controlled release
mechanisms which affect the temporal sequence or kinetics of release.
Considerations of clinical translation was also taken into account in our design
where we assemble our system based on biocompatible and biodegradable components
such as a hydrolytically degradable polymer, PBAE and a biomolecule, lipids. Stability
of therapeutic molecules during incorporation into materials and subsequent storage was
also considered since many of these novel vaccine antigens and adjuvants are based on
relatively sensitive biomolecules that must be protected to preserve their function. The
ability to store therapeutics in a dry state without the need for refridgeration will also
have significant impact on the cost of delivery, especially crucial for delivery in less
developed countries. Finally, improving vaccine and immunotherapy administration
generally for both the physician as well as patient, towards pain-free and safe needle-less
delivery will help ease constraints on medical personnel and improve patient compliance.
Chapter 2 describes protein- and oligonucleotide-loaded layer-by-layer (LbL)assembled multiplayer films incorporating PBAE to demonstrate the potential for simple
and versatile materials encapsulation into conformal thin films, providing robust control
over materials release, solid-state stabilization of environmentally-sensitive encapsulated
materials, and nanometer-scale control over film structure and composition. Notably, the
versatility of multilayer assembly would allow this concept to be implemented in a
variety of transcutaneous delivery settings, including coatings of simple skin-adhesive
patches, woven-fiber adhesive dressings, or microneedle arrays, providing a foundation
for transcutaneous delivery that is tested in chapter 4.
Chapter 3 examines methods of synthesizing biodegradable core-shell structured
nanoparticles with a poly(p-amino ester) (PBAE) core enveloped by a phospholipid
bilayer shell, with a view towards delivery of mRNA-based vaccines. The core-shell
19
particle structure enabled the physical and compositional segregation of the functions for
the particle into an endosome-disrupting pH-responsive core and a shell whose
composition could be separately tuned to facilitate particle targeting, cell binding, and/or
drug binding. Messenger RNA was efficiently adsorbed via electrostatic interactions onto
the surface of these net positively charged nanoparticles post-synthesis, avoiding
denaturation typically encountered with encapsulation into particles and was also
protected subsequently from nuclease degradation.
Chapter 4 focuses on the challenges associated with transcutaneous delivery of
antigens to APCs to initiate an immune response. Using confocal fluorescence
microscopy and an in vivo murine ear skin model, we demonstrated delivery of a model
protein, ovalbumin and synthetic adjuvant, CpG oligonucleotide from LbL films into
barrier-disrupted skin, uptake of the protein by skin-resident APCs (Langerhans cells),
and transport of the antigen to the skin-draining lymph nodes. To our knowledge, this is
the first demonstration of LbL films applied for the delivery of biomolecules into skin
Chapter 5 explores the capability of lipid-coated PBAE particles to mediate
endosomal disruption and enhance cytosolic delivery of molecules in DCs while avoiding
excessive cytoxicity. The study begins with delivery of a model small molecule, calcein
and culminated in the successful transfection of DCs in vitro in serum-containing
medium with mRNA. The use of these particles to enhance delivery of anti-cancer agents
in tumor cells to achieve synergistic tumor cell killing was also examined.
Chapter 6 investigates the ability of lipid-enveloped PBAE particles to deliver
intact molecules across mucosal tissues by evaluating the administration of particles
loaded with mRNA via the nasal passages and airways in mice. Particle-treated mice
showed enhanced expression of a reporter protein, luciferase compared to naked mRNA
following intranasal and intratracheal administration. This system may thus be promising
for noninvasive delivery of mRNA-based vaccines.
20
2. Synthesis and characterization of polyelectrolyte multilayer
films
2.1. Introduction
Polyelectrolyte multilayers have attracted much interest for their versatility, ease
of preparation, and ability to conformally coat virtually any substrate. 93, 105-109 Pioneering
studies demonstrated that proteins could be assembled into such multilayers" 0' 1" and
retain their functionality," 2 stimulating a broad interest in potential biomedical
applications of these materials.107' 113 Further, the ability of multilayers to be built from
biocompatible and bioresorbable polyelectrolytes has led to additional applications,
including modification of cell adhesion on surfaces,"4 tissue- and skin-bonding films,115
, or even
coatings directly applied to epithelial or endothelial cell layers in situ
18-120
cells.
single
living
on
coatings
In concert with a growing interest in applying multilayers to biomedical
applications, erodible multilayers that deconstruct in aqueous conditions via disassembly
and/or breakdown of the constituent polymers have begun to be explored as potential
Drug-loaded degradable multilayers have
controlled release drug delivery films.
been explored for the sustained release of small-molecule antibiotics, protein
therapeutics, plasmid DNA and siRNA. 94, 98, 107, 110, 121, 123-127 The mild aqueous
conditions for encapsulating molecules into multilayer films preserves the bioactivity of
fragile biomolecules such as proteins and nucleic acids. 94, 123, 128 By employing
degradable polyelectrolytes as building blocks, the ability to tune the degradation kinetics
of multilayer assemblies has been demonstrated and used to control the release kinetics of
compounds embedded in these films. 122, 129 Applications envisioned for such drug-loaded
films include antimicrobial- or anti-inflammatory coatings on implants and drugreleasing coatings for stents. 126,130,131
In this chapter, we report on the assembly of polyelectrolyte multilayer films
based on a degradable polymer paired with protein and DNA with a view towards
transcutaneous immunization. We first demonstrate that biodegradable multilayers can be
loaded with a model protein antigen at doses physiologically relevant to vaccination, that
dried multilayers release the embedded protein in a monomeric, unaggregated form on
rehydration and erosion of the films. Film assembly was successfully translated to
flexible substrates suitable for application onto skin. We further demonstrate the
incorporation of multiple drug cargos in films, illustrated for the case of vaccine design
by embedding a protein antigen together with an adjuvant molecule, single-stranded CpG
DNA oligonucleotides. We show that antigen and adjuvant can be loaded together in
decomposable multilayer coatings and released with distinct kinetics, as may be desirable
for temporally controlling the induction of an immune response (or other therapeutic
response).
21
2.2. Materials and Methods
2.2.1.
Materials
Poly(#-amino esters) (PBAE), Poly-1 and Poly-2, were synthesized via the
Michael-type addition of bifunctional amines, N,N'-dimethylethylenediamine and
piperazine respectively, to 1,4-butanediol diacrylate according to previous literature. 90
N,N'-Dimethylethylenediamine, piperazine, and 4,4'-trimethylenedipiperidine were
purchased from Aldrich Chemical Co. (Milwaukee, WI). 1,4-Butanediol diacrylate was
purchased from Alfa Aesar Organics (Ward Hill, MA). Polymerization proceeded
exclusively via the conjugate addition of the secondary amines to the bis(acrylate ester)
to form poly(p-aminoesters) containing tertiary amines in their backbones and readily
degradable linkages. In a typical synthesis reaction, 1,4-butanediol diacrylate (0.750 g,
0.714 mL, 3.78 mmol) and diamine (3.78 mmol) were weighed into two separate vials
and dissolved in THF (5 mL). The solution containing the diamine was added to the
diacrylate solution via pipet. A Teflon-coated stirbar was added, the vial was sealed with
a Teflon-lined screw-cap, and the reaction was heated at 50 'C. After 48 h, the reaction
was cooled to room temperature and dripped slowly into vigorously stirring diethyl ether
or hexanes. Polymer was collected and dried under vacuum. Both Poly-1 and Poly-2 were
specifically designed to degrade by hydrolysis of the ester bonds in the polymer
backbones at physiological conditions. However, Poly-2 is designed to have a slower rate
of degradation relative to Poly-1 due to the presence of more hydrophobic blocks in the
backbone owing to the choice of diamine (Figure 2-1).
o
0
Poly 1
o
0
Poly 2
Figure 2-1. Chemical structures of Poly-1 and Poly-2 used in this study.
Their molecular weights are 8,000 - 10,000 g/mol and undergo different rates of
hydrolytic degradation.
Alexa Fluor 488, Alexa Fluor 647 and Texas Red-conjugated ovalbumin,
NuPAGE 10% Bis-Tris gels, 10% Tris-Glycine gels, and all PAGE reagents were
purchased from Invitrogen (Eugene, OR). TAMRA-labeled and non-labeled CpG
oligonucleotides were purchased from Integrated DNA Technologies (Coralville, IA).
2.2.2.
Film preparation
22
All LbL films were assembled with a modified programmable Carl Zeiss HMS
DS50 slide stainer. Typically, films were constructed on a glass slide with approximate
size of 1 x 2 in2, which was treated in a plasma cleaner (Harrick Scientific Corp.) with 02
plasma for 5 min prior to use. The substrate was then dipped into Poly-l solution (2.0
mg/mL in 100 mM NaOAc buffer) for 10 min and followed by three sequential rinsing
steps with pH-adjusted water for 1 min each. Then the substrate is dipped into OVA
solution (0.10 mg/mL in 100 mM NaOAc buffer) for 10 min and exposed to the same
rinsing steps as described above. The PDMS substrate was initially coated with
poly(allylamine hydrochloride) (50 mM, pH 7.5), followed by the same protocol to build
(Poly-1/OVA)n multilayer as described above. Co-delivery film of OVA with CpG was
constructed in a similar manner with CpG solution (0.10 mg/mL in 100 mM NaOAc
buffer, pH 6, mixture of 10% TAMRA-labeled DNA
2.2.3.
Film characterization
Protein loading on the film was quantified by measuring the absorbance of protein
within the film using a UV/Vis spectrophotometer (Varian Cary 600). OVA and CpG
release from the film was followed by measuring the fluorescence spectra of released
OVA-AF 488, OVA-AF 647, and TAMRA-CpG in PBS (Quantamaster Fluorimeter,
PTI). Film thickness was measured with a Tencor P-10 Surface Profilometer.
To assess the structure and form of ovalbumin released from eroded films upon
rehydration, non-reducing SDS PAGE and native PAGE were performed using NuPAGE
10% Bis-Tris gel and 10% Tris-Glycine gels respectively. Samples were diluted in LDS
sample buffer for SDS PAGE according to the manufacturer's instructions, heated to 90
'C for 2 min and ran using MES SDS running buffer at 200 V for 35 min. For native
PAGE, samples were diluted in native sample buffer without heating and ran under native
running buffer at 75 V for 2.5 h. Finally, gels were stained with a silver staining kit
(Pierce Biotechnology, Rockford, IL) according to the manufacturer's instructions.
2.3. Results and Discussion
2.3.1.
Ovalbumin loading in PBAE films
To fabricate a PEM film capable of releasing macromolecular cargos into skin, we
utilized polyelectrolytes belonging to the family of poly(p-amino esters), known to
degrade hydrolytically under physiological conditions, focusing on two members of this
family designated Poly-1 and Poly-2 (Figure 2-1).90 The biocompatibility of Poly-1 has
been established in earlier studies, and we have previously employed this polymer to
fabricate LbL films with controlled erosion and tunable drug release profiles.9 1 ,9, 12 1 As
a model protein cargo, we examined the assembly of multilayers containin
ovalbumin
(OVA), a 45 kDa globular protein routinely used as a model vaccine antigen.
We first prepared PEM films by alternating adsorption of Poly-1 and OVA from
aqueous buffers onto glass slides, utilizing electrostatic interactions between the protein
and polymer to mediate assembly and monitoring protein incorporation via the
absorbance of fluorophore-conjugated OVA. Based on the pI of OVA (-4.6) and the pKa
23
of Poly-1 (between 4.5 - 8), we tested film assembly at pH 5 - 7, varying the pH of both
Poly-1 and OVA adsorption solutions (Figure 2-2).'0 As summarized in Figure 2-3, film
growth was approximately linear for deposition buffer pH values of 5/6 and 6/6 for Poly1/OVA bilayers. Maximal incorporation of OVA was achieved by adsorbing both Poly-1
and OVA at pH 6. For OVA adsorption buffers at pH > 6, loading decreased, possibly
due to decreasing charge density of Poly-1, which contains tertiary amine groups, at
higher pH.121 Measurement of the total amount of OVA released from completely
dissolved films showed that for films assembled at pH 6, 40-bilayer films carried 5
gg/cm 2 OVA protein total (40-bilayer films are ca. 150 nm). In the context of
vaccination, this order of dose/area is well within a meaningful range as antibody
m
responses have been reported for transcutaneous antigen doses as low as 3 gg. 35-137
0.08
0.06
.'
0.04
'U
gO 0.02
0.00
5.0
5.5
6.0
6.5
pH of Poly 1
Figure 2-2. Loading of OVA into multilayers prepared by sequential deposition of
15 bilayers of Poly-1/OVA.
For each condition, OVA was deposited from pH 6 buffered solutions and Poly-1 was
deposited from buffered solutions of varying pH (5.0-6.5).
24
0.15 -
u
pH 5/6
MpH 6/6
0.10.
0
0.05.
0.000
5
10
15
20
25
30
35
40
Number of Bilayers (n)
Figure 2-3. Growth curve of (Poly-1/OVA) multilayers for two different deposition
conditions.
OVA loading at different pHs of OVA and Poly-1 solutions was assessed by measuring
absorbance from fluorophore-conjugated OVA incorporated in films.
Rapid release of ovalbumin from deconstructed films upon
2.3.2.
rehydration
When dried as-assembled (Poly-1/OVA), PEM films on glass were rehydrated in
phosphate-buffered saline (PBS, pH 7.4), films prepared across all deposition conditions
exhibited rapid protein release, as illustrated in Figure 2-4. Characterization of (Poly1/OVA) 40 film thickness vs. time following rehydration revealed rapid dissolution of the
films coinciding with protein release. These films fall apart significantly more rapidly
23, and also greatly exceed
than others assembled with other proteins or biopolymers
the rate of polyaminoester hydrolysis anticipated for Poly-1. For this reason, we believe
that OVA, which is a relatively small and globular protein with low surface charge
density, is able to readily diffuse out of the Poly 1 films upon contact with aqueous
solution, leading to a charge destabilization of the electrostatically assembled thin film
and subsequent rapid dissolution of the film. Though the pH/ionic strength conditions
used for PEM assembly here are near the buffer conditions used for analysis of
rehydration and disassembly, our previous studies based on multilayer complexes of
Poly-1 and anionic polysaccharides illustrated that subtle changes in solution pH can
have dramatic effects on film disassembly kinetics.1 21 The overall release kinetics were
modestly altered by either using the more hydrophobic poly(p8-amino ester) (Poly-2) for
film assembly, or by decreasing the ionic strength of the dipping solutions from 100 to 10
mM to reduce charge screening and increase electrostatic interactions between the
polyions during assembly, resulting in more tightly-associated PEM films (Figure 2-4).
Although there are undoubtedly additional strategies that could be used to further regulate
protein release from these films, for our initial studies we hypothesized that rapid film
25
dissociation on rehydration could be potentially advantageous for allowing maximal
delivery of antigen/adjuvant molecules into barrier-disrupted skin prior to sealing of the
epidermis (discussed in subsequent chapter).
-
1.0
1..
0
CF 0.80.8
0
a Poly 1 (1OOmMNaOAc)
. 0.6
*
0.6
Poly 1 (10 mM NaOAc)
A Poly 2 (100 mM NaOAc)
0.4
0.4
~, 0.2.
.
0
0.0
0.0
0
1
2
3
Time (h)
4
5
Figure 2-4. Cumulative normalized kinetics of OVA release and representative
thickness changes with time from rehydrated LBL films.
Films prepared with Poly-1 vs. Poly-2 and with Poly-1 deposited at low or high ionic
strength were rehydrated in PBS (pH 7.4 at 25 *C) to determine release kinetics of OVA.
Representative thickness changes with time for a Poly-1/OVA film assembled at high
ionic strength were also measured.
2.3.3.
Film assembly on flexible substrates
The above basic characterization measurements were made for films on glass
substrates, which may not be ideal for application on skin. However, as there is virtually
no restriction in the choice of the substrate for LbL assembly, we next tested assembly of
Poly-1/OVA films on substrates that might be useful in skin patch formulation. We tested
poly(dimethylsiloxane) (PDMS) as an elastomeric substrate that could form the basis of
easily-handled macroscopic patches, exploiting its well-known utility in microcontact
printing to achieve intimate contact between drug-loaded multilayers and the
inhomogeneous surface structure of skin. As illustrated in Figure 2-5, film assembly was
readily achieved on flexible PDMS substrates, and incorporation of fluorescent OVA
(OVA-Texas Red) was readily observed from the blue hue of multilayer films on the
transparent rubber. We were also able to deposit films onto fibrous substrates such as
electrospun polymeric fibers that may be suitable materials for wound dressing (image
not shown).
26
Figure 2-5. Photograph of (Poly-1/OVA-Texas Red) 40 film assembled on a flexible
PDMS substrate.
Characterization of released ovalbumin via SDS- and native
2.3.4.
PAGE
We next tested whether OVA protein assembled into Poly-1 LbL films, dried, and
then rehydrated and released into buffer remained intact and unaggregated by running
both non-reducing SDS-PAGE (Figure 2-6) and native PAGE (Figure 2-7) on
supernatants taken from (Poly-1/OVA) 40 films assembled at different pHs, dried, and
then rehydrated in PBS for 30 min. As seen in Figure 2-6 and Figure 2-7, almost all of
the OVA released from poly-1 films existed as monomeric protein running at the same
position as unmanipulated control OVA solutions in PAGE gels, indicating that the
structural integrity of OVA was preserved during both integration into and subsequent
release out of the films. OVA released from films did not show significant signs of
aggregation either with itself or with released Poly-1, as indicated by the lack of protein
detected at positions above the main band of free OVA in the PAGE gels. Lack of
aggregation or degradation in the released OVA suggests that proteins can be
incorporated into poly-1 LbL films, dried, and released in a native monomeric state on
rehydration without irreversible complexation to the polycation component of the films,
an important outcome for the bioactivity of therapeutics or for the correct elicitation of
antibody responses for vaccine antigens.
27
98k
38k
Figure 2-6. Non-reducing SDS-PAGE of supernatants from (Poly-1/OVA) 40 films
rehydrated for 30 min in PBS.
LbL films were prepared by depositing OVA at fixed pH 6 and Poly-1 at pH 5.5, 6 or 6.5,
respectively. Released protein is compared to control solutions of OVA in the dipping
buffer or PBS (OVA cntl). The majority of OVA is released from the LbL films as free
monomer (46 kDa). The molecular weight of released protein is determined by
comparison with a molecular ladder comprising of standard proteins at known molecular
weights.
Film release
pH of Poly
aS
5.5 6.0 6.5
5
o
0 W
Figure 2-7. Native PAGE analysis of OVA released from (Poly-1/OVA)40 films after
a 30 min rehydration in PBS.
Films were prepared by depositing OVA at a fixed pH 6 and Poly-l at pH 5.5, 6 or 6.5,
respectively. Protein released from films is compared to controls of OVA solutions in
PBS pH 7.4 or in the multilayer deposition solution (pH 6,100 mM NaOAc buffer).
28
2.3.5.
Dual loading and release of ovalbumin and CpG
An attractive property of multilayers for controlled drug release is the ability to
load multiple components into multilayer films and potentially tune the release kinetics of
these cargos separately via the architecture of the films. 12' To explore multi-component
drug delivery in the context of a transcutaneous vaccine, we next assembled LbL films
incorporating both a protein antigen and an adjuvant cargo. Adjuvants are compounds
typically co-injected with vaccines to increase their potency, reduce the required antigen
dosage for a protective response, and/or to elicit immune memory with an appropriate
protective profile. 138 We tested the co-incorporation of OVA antigen in LbL films with
CpG DNAs, short single-stranded oligonucleotides that contain unmethylated CpG
(Cytosine--phosphate diester--Guanine) sequences characteristic of bacterial DNA. CpG
DNA oligos are ligands for Toll-like receptor-9 on dendritic cells and other innate
immune cells, and are a potent adjuvant currently in clinical trials for a variety of vaccine
and cancer immunotherapy applications.139 Immunologically, adjuvant signals such as
CpG trigger dendritic cell activation, which is accompanied by a downregulation of
antigen internalization by these cells.140 Thus, we hypothesized that for vaccination, films
capable of releasing antigen coincident with or prior to adjuvant release would be
attractive. Further, sustained adjuvant delivery (mediated in basic immunological studies
by repeated injections) have been shown to help break tolerance to tumor antigens in the
setting of cancer vaccines, and so the ability to release adjuvant molecules for a period of
several days could be a second important opportunity for LbL vaccine films.141
To regulate the release of antigen and adjuvant in composite films containing
OVA and CpG, we explored different sequences of layering in order to determine the
impact of film architecture on the loading and release kinetics for the two components.
We were also interested in examining potential interactions between OVA and CpG
within the film structure. Using Poly-2 as a complementary polycation for the two
polyanionic cargos, co-delivery films containing OVA and CpG were constructed with
several different architectures (Figure 2-8): control films of OVA (A) or CpG (E) alone
with Poly-2, CpG layered on top of OVA (B), alternating OVA and CpG (C), or OVA
layered over CpG (D). In each case 20 bilayers of Poly-2/OVA and Poly-2/CpG were
included. AFM imaging of dried films (Figure 2-8B) showed that each film composition
exhibits a distinct surface morphology: films containing OVA were generally rougher
than CpG-loaded films (mean rms roughness 31±3 nm for (Poly-2/OVA) 20 films vs.
8.5±2 nm for (Poly-2/CpG) 20 films). As described earlier, the release characteristics of
the OVA films suggest a high degree of interdiffusion of OVA within the film, and high
concentrations of OVA at the top surface, typical of interdiffusion growth behavior.
The roughness of the surfaces of the OVA films may be a result of this interdiffusion, and
accumulation of OVA protein in the top layers. Film structures A and D, that were both
terminated with multiple Poly-2/OVA bilayers, were quite similar in their rough surface
topography, whereas film structures B and C were rougher than E (pure Poly-2/CpG)
films even though each film terminated with a Poly-2/CpG bilayer. These results suggest
the presence of OVA near the surface in some substantial concentration in each of these
two film architectures as well, and thus indicates the potential for interdiffusion of the
adsorbing species during assembly. We postulate that OVA is a more diffusive species
within the LbL layers compared to the more densely charged CpG oligonucleotides,
29
despite its larger molecular weight. It has been shown that charge density is a key
controlling element in the ability of macromolecular species to diffuse within multilayer
films.
122, 143, 144
The CpG-only films illustrate a more linear release over much more
extended time frames that correspond well with surface hydrolysis of the multilayer
films. This hypothesis is further supported by the release profiles collected in Figure
2-8C, showing rapid release of OVA from films rehydrated in PBS for all film
architectures, similar to the results observed for pure Poly-1/OVA films. In contrast, a
sustained release of CpG adjuvant over 3 days was observed in all cases. Although each
film architecture was prepared with the same number of bilayers of OVA or CpG, the
total amount of each component within the final films varied significantly with film
architecture. In earlier work we found that if an underlying multilayer system exhibits a
high degree of interdiffusion during its construction, then a second set of multilayers
assembled atop the first will often also exhibit interdiffusion into the pre-existing film12 1;
this behavior provides opportunity for high loadings of both components via diffusion
and trapping of excess molecules within the film. Notably, architecture "B" with OVA
deposited first under bilayers containing CpG, enabled the loading of large quantities of
both components in the film, and exhibited attractive staggered release kinetics, with the
majority of antigen rapidly released within 24 h while the CpG adjuvant exhibited
sustained release over a period of -3 days. Although further tuning of component
loading and individual release kinetics can be readily envisioned by manipulation of film
assembly conditions and the choice of polycation components, the staggered release of
antigen and adjuvant presented here demonstrates the potential of this approach.
30
a
lOva) 2 0
b
IOva) 2 0(
(
ICpG) 20
(
ICpG) 20
lOva
(
ICpG) 20(
C
IOva) 20
CpG) 20
C
-08
10.
C
.
1.0.
0.8
e6
000.6.
4
0.4
0
2
0.2
2
0
0
0
0
1
2
3
4
5
6
Time (day)
7
0
1
2
3
4
5
Time (day)
6
7
0.0
0
1
2
3
5
4
6
7
Time (day)
Figure 2-8. Dual loading and release of OVA and CpG.
(A) Schematic architectures of antigen (OVA) and adjuvant (CpG DNA) co-delivery
films tested. (B) AFM surface morphology of the corresponding dried films (10
x
10 ptm2
images, z-scale of 200 nm). (C) Kinetics of OVA and CpG release from dried LbL films
rehydrated in PBS (pH 7.4, 25 *C). Released OVA and CpG were measured by collecting
non-overlapping fluorescence signals (kmax,OVA = 665 nm, kmaxCpG = 560 nm),
respectively. Overlay of normalized OVA and CpG release curves from film architecture
B.
2.4. Conclusions
In this chapter, we demonstrated protein- and oligonucleotide-loaded LbLassembled films as platform materials for delivering vaccine antigens and
immunostimulatory adjuvants. Rapid release of protein antigen was achieved, potentially
suitable for transcutaneous delivery following disruption of the stratum corneum. When
an adjuvant was also incorporated, staggered and sustained release of the adjuvant over a
few days was observed, which may be appropriate for modulating the desired immune
response. The ability of multilayer films to incorporate protein that can be stored in a dry
state prior to application and then released from rehydrated films in a non-
aggregated/non-degraded
state may allow extended storage/shelf life of vaccine
formulations and avoid the "cold chain" of distribution-the materials, equipment and
procedures required to maintain vaccines within a temperature range for preserving their
potency, thereby addressing challenges in vaccine distribution in developing countries.
In addition, the ability of multilayers to conform-ally coat substrates should allow
this approach to be used in tandem with substrates suitable for mediating transcutaneous
31
delivery such as flexible elastomer for skin patches, electrospun fibers for wound
dressing, and microneedles that can breach the stratum corneum. This sets the stage for
applying LbL films towards transcutaneous immunization in chapter 4. The generality of
LbL assembly also suggests that a variety of antigen and adjuvant molecules could be
readily incorporated into these films. However, to make this platform truly generic for
delivering a wide range of synthetic antigens and adjuvants, biodegradable nano-carriers
may be embedded as delivery vehicles within LbL films, to segregate film
assembly/release properties from the characteristics of individual antigens/adjuvant
molecules. Furthermore, nano-carriers can act as depots to increase the amount of antigen
loaded in a single step and reduce the time needed to deposit multiple layers to achieve
sufficient dosage. This led us to explore the development of nanoparticles based on the
same polymer, PBAE, the topic of discussion in the next chapter.
32
3. Synthesis and characterization of lipid-enveloped polymer
nanoparticles
3.1. Introduction
In Chapter 2, we showed that a hydrolytically degradable poly(p-amino-esters)
(PBAEs) can be used to assemble multilayer films capable of loading multiple antigen
and adjuvant components and releasing intact molecules with some degree of temporal
control. In addition to LbL films, PBAEs have also been formulated as micro- and nanoparticles capable of mediating intracellular delivery of plasmid DNA, RNA, as well as
tumor-targeted chemotherapeutics.84'99-104 This is motivated by the pH-sensitive nature of
the polymer due to the presence of tertiary amine groups in the backbone.
To achieve cytosolic delivery, our lab has previously developed a pH-responsive
core-shell nanoparticle system prepared by sequential emulsion polymerization of a
secondary-amine-containing monomer (diethlyaminoethyl methylacrylate) forming a pHresponsive core, followed by a second monomer (aminoethyl methacrylate) forming a
hydrophilic corona. 145, 146 The core-shell particle structure enabled the physical and
compositional segregation of the functions for the particle into an endosome-disrupting
pH-responsive core and a shell whose composition could be separately tuned to facilitate
particle targeting, cell binding, and/or drug binding. These particles efficiently delivered
associated protein or oligonucleotide cargos to the cytosol of dendritic cells through
endosomal disruption mediated by the core polymer via the proton-sponge effect, while
maintaining low cytotoxicity by sequestering the cationic charge and hydrophobicity of
the polymer core within a more hydrophilic polymeric shell. However, a limitation of
this proof-of-concept system is its lack of biodegradability, which hinders clinical
translation
In this chapter, we adapted this core-shell design approach to a fully degradable
system comprised of a pH-responsive PBAE core and a phospholipid shell, endowing the
particles with properties relevant for enhanced vaccine and immunotherapeutic delivery.
In addition to pH responsiveness for enhanced intracellular delivery, strategies based on
the binding of therapeutics to the surfaces of particles post-synthesis that can address the
challenges of low antigen encapsulation efficiency and denaturation during the
encapsulation process are also of interest, 90' 145 especially for the delivery of novel
sensitive subunit antigens such as proteins, DNA and RNA. Examples of this approach
include adsorption of antigens to charged PLGA articles 147' 148 and covalent coupling of
protein to reactive groups on particle surfaces. 1 '9
150
As a first step toward enhanced
messenger RNA (mRNA) vaccines, we show that negatively-charged RNA can be
adsorbed efficiently via electrostatic interactions onto the surface of these cationic
nanoparticles, protecting the nucleic acids from degradation in serum. The protocols
established in this section will provide a basis for in vitro and in vivo testing of this nanocarrier in the upcoming chapters.
33
3.2. Materials and methods
3.2.1.
Materials
The poly(p-amino ester) poly-I with a number averaged molecular weight of -10
kDa was synthesized as previously reported.90 Poly(lactide-co-glycolide) (PLGA) with a
50:50 lactide:glycolide ratio and molecular weight of -46 kDa was purchased from
Lakeshore Biomaterials (Birmingham, AL). The lipids 1,2-dioleoyl-sn-glycero-3phosphocholine (DOPC), 1,2-dioleoyl-3-trimethylammonium-propane (chloride salt)
(DOTAP),
1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-[methoxy(polyethyleneglycol) -2000] (ammonium salt) (DSPE-PEG), 1,2-distearoyl-sn-glycero-3phosphoethanolamine-N-[methoxy-(polyethyleneglycol)-2000-N'-carboxyfluorescein]
(ammonium salt) (DSPE-PEG-CF) and 1,2-dioleoyl-sn-glycero-3-phosphoethanolamineN-(lissamine rhodamine B sulfonyl) (ammonium salt) (DOPE-rhodamine) were
purchased from Avanti Polar Lipids (Alabaster, AL). Poly(vinyl-alcohol) (PVA) with a
molecular weight of -78 kDa was purchased from Polysciences Inc. (Warrington, PA).
Sodium acetate buffer were purchased from Sigma Chemical Co. (St. Louis, MO). Triton
X-100 (molecular biology grade) was from Promega Corp. (Madison, WI). Poly (Laspartic acid) (sodium salt) was from Alamanda Polymers (Huntsville, AL). 1,12dioctadecyl-3,3,32,32-tetramethylindodicarbocyanine,
4-chlorobenzenesulfonate
salt
(DiD), ribogreen and SYBR Green II RNA gel stain were purchased from Invitrogen
(Eugene, OR). Poly I:C was purchased from InvivoGen (San Diego, CA). RNAse
inhibitor was purchased from Roche (Indianapolis, IN). Cy3 Label IT@ nucleic acid
labeling kit was purchased from Mirus Bio LLC (Madison, WI). All materials were used
as received unless otherwise noted.
3.2.2.
Synthesis and Characterization of Lipid-coated PBAE
Nanoparticles
Lipid-coated nanoparticles with a poly-1 core were synthesized via two different
processes:
a double emulsion/solvent evaporation approach or a solvent
diffusion/nanoprecipitation strategy.
For double emulsion synthesis, we adapted a previous approach we used for
preparing lipid-enveloped PLGA particles 5 1 : 30 mg of poly-1 (or PLGA for pHinsensitive control particles) and 2 mg of the phospholipids DOPC, DOTAP, and DSPEPEG in a 7:2:1 molar ratio were co-dissolved in 1 mL of dicholoromethane (DCM). PBS
(200 pL) was then added to the mixture on ice during a 1 min sonication step at 7 W
using a probe tip sonicator (Misonix XL2000, Farmingdale, NY) to form a first emulsion.
The primary emulsion was then dispersed into 6 mL of distilled, deionized nuclease-free
water with sonication at 12 W for 5 min, before leaving on a shaker for 18 h at 25 'C to
evaporate the organic solvent.
For nanoprecipitation synthesis, the same quantities of DOPC, DOTAP, and
polymer were co-dissolved in 4 mL of ethanol and added drop-wise to 40 mL of distilled,
deionized nuclease-free water, followed by gentle stirring for 5 h to evaporate ethanol.
For particles prepared by nanoprecipitation, we found that adding DSPE-PEG in the
34
organic phase together with DOPC, DOTAP and polymer reduced particle yield
significantly, so DSPE-PEG was introduced into the lipid coating via a post-insertion
process: DSPE-PEG lipid was added at 1 mM to 0.5 mg/mL particles in distilled,
deionized nuclease-free water and the suspension was stirred for 16 h at 25 'C. The
particles were collected and washed once via centrifugation, resuspended in fresh water
and stored at 4 'C until use.
Lipid-free PVA-stabilized particles were prepared by similar processes except the
organic emulsion or solution containing polymer only was dispersed into a 2% w/vol
PVA aqueous solution.
A fraction of each particle batch was dried in a vacuum oven to determine the
particle concentration (mg/mL) by measuring the dry mass. Dynamic light scattering
(DLS) and zeta potential measurements were used to determine the particle size and
surface charge using a ZetaPALS dynamic light scattering detector (Brookhaven
Instruments).
To estimate the percentage of PEG-lipid incorporated and the resultant mol% of
PEG-lipid present in the lipid surface coating, lipid-enveloped particles were prepared by
nanoprecipitation as described above with 1 mol% DOPE-rhodamine added as a tracer
for measuring the total amount of lipid incorporated into the particles, before the postinsertion of a DSPE-PEG lipid labeled with carboxyfluorescein (DSPE-PEG-CF). The
lipids were then stripped from the particle surface by treatment with 2% triton X-100 for
15 min and the supernatant was measured for both rhodamine and fluorescein signals to
quantitate the amount of PEG lipid incorporated.
To investigate the surface structure of lipid-coated PBAE particles by
cryoelectron microscopy (cryoEM), particles were embedded in ice by blotting a particle
suspension (3 uL) on a 1.2/1.3 gm holey carbon-coated copper grid (Electron Microscopy
Sciences) and immediately freezing the sample in liquid ethane using a Leica plungefreezing machine. Samples were transferred to a cryogenic holder and imaged using a
JEOL 2200FS transmission electron microscope at 185 p.A emission current and 40 000x
magnification.
3.2.3.
pH-sensitivity assay
To verify the pH-responsiveness of lipid-coated PBAE particles, lipid-coated
particles as well as particles lacking lipids and stabilized by PVA were suspended in 100
mM phosphate buffers of different pHs and the absorbance of the particle suspensions
was measured at 350 nm as a surrogate measure of particle light scattering.
3.2.4.
mRNA preparation
Synthetic mRNA was prepared as described previously.152, 153 Briefly, linearized
plasmid DNA bearing the T7 promoter, an open reading frame (GFP, 811 nucleotides or
firefly luciferase, 1714 nucleotides), and poly-A tail was used as a template for in vitro
transcription using the Message Machine T7 Ultra Transcription kit (Ambion). The
mRNA product was precipitated with LiCl, resuspended in nuclease-free water, and
quantitated with a NanoDrop spectrophotometer. RNA size, purity and integrity were
ascertained with an Agilient Bioanalyzer 2100. An RMA cell line electroporated with
35
these mRNA transcripts was also used to confirm the functionality of these mRNA in
vitro (data not shown). mRNA was labeled with a fluorescent Cy3 tag using a Label IT@
nucleic acid labeling kit (Mirus Bio LLC) according to the manufacturer's protocol.
3.2.5.
Loading and release of RNA
Poly I:C or mRNA were loaded onto the surface of lipid-coated PBAE particles
by first diluting the particle suspension to 1.5 mg/mL in nuclease-free water, then adding
200 pL of the diluted suspension dropwise to 100 pL containing 4 gg of RNA under
gentle vortexing. The vial was then incubated at 4 'C for 2 h on a rotator to allow RNA
adsorption. Bound RNA was determined indirectly by measuring fluorescence of either
Cy3-labeled or ribogreen-stained RNA remaining in the supernatant after centrifugation
of the particles using a fluorescence plate reader (SPECTRAmax, Molecular Devices
Corp.). RNA bound to the particles was determined by subtracting the quantity of RNA
detected in the supernatant from the quantity measured in identically-treated control vials
containing RNA solutions but no particles.
To assess the binding kinetics and capacity of the particles, the particle
concentration was fixed at 1 mg/mL while the binding time and RNA amount was varied
as described in the text. Binding efficiency was defined as the percentage of RNA
initially present in solution that bound to the particles. Loading was defined as the
amount of RNA bound ([tg) per mg of particles.
To determine the release kinetics of bound RNA from particles in RPMI 1640
culture medium containing 10% FBS, aliquots of particle-adsorbed poly L:C (70 Rg
particles containing 1.87 ig poly I:C) were resuspended in 140 pL of serum-containingmedia and incubated at 37 'C under gentle mixing. At each designated time-point, an
aliquot was removed and particles were washed once with nuclease free water before
resuspending in a digestion buffer (100 mM sodium acetate, 2% triton X-100 and 1
mg/mL Poly (L-aspartic acid)) to dissolve the particles and disrupt any lipid-RNA or
poly-1-RNA complexes. The amount of RNA remaining on the particles was then
determined by measuring the fluorescence following staining with ribogreen as above,
and the amount of RNA released was calculated by subtracting the amount of remaining
RNA from the original amount.
3.2.6.
RNA Degradation Protection Assay
1 pg of mRNA (encoding for GFP) either naked or loaded onto PEGylated PBAE
nanoparticles was incubated in the presence of 2 or 10% FBS in nuclease free water for 5
min or 1 h at 37 'C. After the incubation period, 100U of RNAse inhibitor was added to
quench degradation and the samples were analyzed by electrophoresis on a 1% agarose
gel at 75 V for 1 h. The gel was visualized following staining with SYBR Green II RNA
gel stain according to manufacturer's protocol.
3.3. Results and Discussion
36
3.3.1.
Design, Synthesis, and Characterization of Lipid-coated PBAE
particles
Our laboratory previously demonstrated that core-shell structured hydrogel
nanoparticles composed of a pH-responsive proton sponge core and a hydrophilic
crosslinked shell could be loaded with anionic proteins or oligonucleotides by
electrostatic adsorption of these cargos to the particle surfaces. 146 Uptake of these coreshell particles by cells led to disruption of acidifying endolysosomes and efficient
145
delivery of the cargo molecules into the cytosol of cells with minimal toxicity. , 146
However, these particles were prepared from vinyl monomers and are effectively nonbiodegradable. To translate this proof of concept to a resorbable materials system for in
vivo delivery, we sought to prepare biodegradable core-shell particles with the structure
schematically outlined in Figure 3-1: The particle core is composed of a pH-sensitive
biodegradable poly(p-amino ester) (PBAE), chosen to promote endolysosomal disruption
on particle uptake by cells. We sought to surround the poly-1 core by a phospholipid
shell, with the aims of limiting contact of the hydrophobic, cationic polymer core with
cellular constituents prior to degradation and providing a tunable surface to mediate
binding of mRNA to the particle surfaces.
surfa c-bound mRNA
DOPC
pH < -7
DOTAP
0
DDPE-PEG
Figure 3-1. Schematic of structure and composition of lipid-coated PBAE particles
and mRNA cargo association.
Apropos of this goal, we recently carried out a detailed study of the structure of
nanoparticles formed by an emulsion/solvent-evaporation process, where an organic
phase containing poly(lactide-co-glycolide) (PLGA) polymer co-dissolved with
phospholipids was emulsified in water, followed by evaporation of the organic solvent to
form solid polymer particles.' In this process, the lipids act as surfactants and selfassemble to form a tightly-apposed lipid bilayer envelope around each particle. 151 In
addition to providing a separate tunable surface for surface binding of molecules by
tuning the chemistry of the lipid headgroups, such lipid coatings can also enable facile
attachment of targeting ligands and mimic cell or pathogen surfaces.15 4-1s1 We tested
whether variations on this synthesis strategy could be applied to form lipid-enveloped
poly-1 particles (Figure 3-2): First, we prepared particles by co-dissolving poly-1 with a
mixture of lipids (DOPC:DOTAP 70:20 mol:mol) in dichloromethane (DCM), followed
by emulsification of this organic phase in water and evaporation of the organic solvent.
37
We also tested an aproach designed to avoid the need for toxic DCM, based on
nanoprecipitation:101'
148 PBAE and lipids were co-dissolved in ethanol, and then
1'
dispersed into excess water. In this process, solid PBAE nanoparticles form as the
solvent rapidly diffuses out of the polymer droplets into the surrounding aqueous phase,
leading to the precipitation of the polymer core with lipids stabilizing the particle surface.
Double Emulsion
Aqueous
buffer
I
I
so
lipid and
polymer
organic
solutionf;
(Ok.
sonication
n
solvent evaporation/
lipid self-assembly
Nanoprecipitation
lipid and
polymer
organic
solution
PEG-lipid
aqueous
solution
PEG-lipid
post-insertion
stirrisu
Particle
solution
solvent displacement
lipid self-assembly
Figure 3-2. Schematic illustrating the double emulsion and nanoprecipitation
process used to synthesize lipid-coated PBAE nanoparticles.
For both synthesis strategies, particles were separated from free liposomes by
centrifugation. To promote colloidal stability of particles prepared by either method,
poly(ethylene glycol)-phospholipid conjugates (DSPE-PEG) were introduced into the
bilayer surfaces of the particles during lipid self-assembly (for the double emulsion
process) or by post-insertion 14 9' 150, 157 (for nanoprecipitation), resulting in 6-10 mol%
PEG-lipid in the lipid shell. Notably, as discussed below, PEGylation did not affect
electrostatic adsorption of polynucleotides onto the highly charged particle surfaces; but
PEGylation did make the particles resistant to aggregation and easier to resuspend during
centrifugation/washing steps.
38
Particles prepared by the emulsion/solvent evaporation or nanoprecipitation
processes had similar size distributions as measured by dynamic light scattering (Figure
3-3), and similar net positive surface charges indicated by their zeta potentials of 40 ± 9
mV and 42 ± 8 mV in deionized water, respectively. As a control non-pH-responsive
core polymer, we also prepared lipid-enveloped PLGA nanoparticles by the
emulsion/solvent evaporation process, which had mean diameters of 445 ± 67 nm.
Functionally, PBAE-core particles prepared by either synthesis method exhibited
indistinguishable endosomal disruption capacities, RNA loading, and in vitro and in vivo
transfection in subsequent assays, and thus we report on data collected with particles
prepared by both methods for this thesis.
A
B
100-
100-
80-
80-
S60-
60-
20<
20
A
0
0
0
100
200
300
Particle diameter (nm)
0
100
200
300
Particle diameter (nm)
Figure 3-3. Size histograms of synthesized lipid-coated PBAE nanoparticles.
Particles were prepared via emulsion/solvent evaporation (A) or nanoprecipitation (B)
methods.
To provide direct evidence for the assembly of the lipid coating at the surface of
the particles, we also imaged samples by cryoelectron microscopy. Similar to our prior
findings with PLGA particles prepared using lipids as a surface modifier, we saw that
PBAE particles prepared by the emulsion/solvent evaporation process had electron dense
lipid bands coating the surfaces of the particles (Figure 3-4).
39
(scaie
Dar - ou nm)
Figure 3-4. Representative cryoEM image of lipid-enveloped PBAE particles.
CryoEM image of lipid-coated PBAE nanoparticles and zoomed-in section showing the
presence of the lipid coating at the particle surfaces (scale bar = 50 nm).
Table 3-1 summarizes the size, polydispersity indices (PDI) and zeta potentials of
"naked" PBAE particles (PVA-stabilized particles lacking lipid coating), lipid-coated
PBAE particles with and without PEGylation, and PEGylated particles following mRNA
adsorption (~12-14 ptg mRNA/mg particles loaded, discussed below), all measured in
deionized water, the medium used for particle synthesis and mRNA loading. A positive
zeta potential was measured in all cases and is likely attributable to some degree of
ionization of the PBAE core in deionized water (pH ~ 5.5). In contrast with the pHdependent solubility profile measured in subsequent section in high ionic strength buffers
(100 mM phosphate buffers), particles remained stable in deionized water despite its
acidic pH. We hypothesize this reflects the lack of sufficient ionic strength to fully ionize
the core, which would otherwise lead to particle dissolution. At very low ionic strength,
polyelectrolyte ionization can be dramatically suppressed relative to the expected
1
equilibrium at high ionic strength. 158 ~16
However, when particle zeta potentials were
measured in high ionic strength basic pH buffers where no ionization of the core is
expected, the zeta potentials of the particles were reduced accordingly (10 ± 4 mV and -7
± 3 mV for non-PEGylated and PEGylated particles at pH 8, respectively).
Table 3-1. Size and zeta potentials of PBAE nanoparticles determined by dynamic
light scattering.
Zeta Potential in
Effective
deionized water
Sample
Diameter (nm)
PDI
(mV)
PVA NP
Non-PEGylated lipid-coated NP
PEGylated lipid-coated NP
mRNA-loaded PEGylated lipidcoated NP
3.3.2.
300 ± 50
250 ± 45
230 ± 40
0.141 ± 0.08
0.120 ± 0.06
0.100 ± 0.05
32 8
41 ± 10
42 ± 8
280 ± 70
0.182 ± 0.10
40
pH-responsiveness of lipid-coated PBAE nanoparticles
40
7
Poly-1 is a weak polyelectrolyte that is water insoluble at elevated pH but
ionizes and dissolves in aqueous solutions below pH -7. This selective solubility has
been exploited to promote cytosolic delivery of drug cargos following uptake by
cells. 6 2 Theoretically, once internalized by cells, particles containing this polymer
can trigger disruption of the endosomal membrane by at least two (non-exclusive)
mechanisms (Figure 3-5): (i) via a strong osmotic pressure gradient generated across
the endosomal membrane as the solid PBAE particles dissolve into individual
polymer chains;' 63 (ii) via an osmotic pressure buildup due to buffering of
acidification by the charged polymer backbone and subsequent counterion buildup in
the endosome.164, 16s
Dissociation-Induced osmotic pressure
Proton Sponae Effect
VANu
pH-7A.
Figure 3-5. Proposed mechanisms of endosomal disruption by PBAE particles.
We first verified the pH-responsiveness of "naked" PBAE particles (lacking lipids
and stabilized only by PVA) by measuring the 350 nm absorbance of particle suspensions
in 100 mM phosphate buffers of different pHs, using this absorbance as a surrogate
measure of particle light scattering. As shown in Figure 3-6, intact particles gave highly
scattering milky suspensions at pH > 7.3, but became transparent solutions at lower pHs
modeling the acidic environment found within endosomes. Consistent with earlier studies
of poly-1 PBAE microparticles,1 62 particle dissolution was rapid (within 5 min) upon
exposure to acidic buffers. The same trend was observed for particles prepared with lipid
as surfactant, showing that the lipid coating did not alter the dissolution properties of the
particles. The presence of intact particles at pH > 7 and disappearance of particles at
acidic pH was confirmed by confocal microscopy (data not shown). We thus expect that
following acidification of endolysosomal compartments, mRNA on the particles would
be quickly released into the cytosol upon rapid build up of osmotic pressure triggered by
particle dissolution, thereby minimizing contact with acidic conditions inside
endolysosomes.
41
Endolysosomal pH
E
Extracellular pH
1.0
C
.0
0.5-
0.0
pH
Lipid NP Absorbance 0 PVA NP Absorbance
Figure 3-6. pH-dependent solubility profile of lipid and PVA stabilized particles.
Measurement of absorbance of particle suspensions as a function of solution pH for both
lipid and PVA-coated PBAE nanoparticles.
3.3.3.
Efficient RNA loading on particle surface
Loading of RNA cargos on these lipid-enveloped particles by adsorption post
synthesis provided a means to avoid exposing RNA to harsh processing conditions.
Although encapsulation of nucleic acids is often pursued in order to protect these
compounds from premature enzymatic degradation in vivo, prior studies where plasmid
DNA was electrostatically adsorbed to charged PLGA microparticles showed that
particle-adsorbed DNA had a greatly increased half-life in vivo compared to naked
plasmids 166, suggesting that simple adsorption to particle surfaces may sterically protect
polynucleic acids from nucleases. We hypothesized that loading of RNA onto the
particle surfaces would increase bioavailability of the RNA by allowing faster release in
cells, while avoiding exposure of the polynucleotides to the acidic/hydrophobic
environment inside degrading particles that would be expected if we encapsulated RNA
within the particle cores.
To assess the efficiency of electrostatic adsorption for loading of RNA on
surfaces of lipid-coated PBAE particles, we first quantified RNA binding to the particles
as a function of time and RNA concentration using poly I:C, a synthetic double-stranded
RNA (dsRNA) with an average size of 0.2-1 Kb, comparable in size to mRNA transcripts
of interest. Poly I:C is an immunostimulatory analog of viral dsRNA and activates innate
immune cells through Toll-like receptor 3167, 168 and the cytosolic RNA sensor MDA-5 169'
170, which further motivated these experiments.
Poly I:C was mixed with lipid-coated PBAE particles in nuclease-free water to
allow RNA binding, followed by washing to remove remaining unbound RNA. We found
that RNA binding equilibrated rapidly, reaching a plateau by 2 h when 4 ptg of RNA was
incubated with 300 ptg particles in 300 piL nuclease-free water (Figure 3-7A).
Interestingly, PEGylated particles also bound poly I:C to similar levels, suggesting that
the PEG layer on the particle surface did not prevent electrostatic attraction of the RNA
to the charged lipid surfaces (Figure 3-7A). We next measured binding of varying
42
amounts of poly I:C added to a fixed concentration of 1 mg/mL particles, and saw that 1
mg of particles was capable of absorbing up to 133 ptg of RNA before binding was
saturated (Figure 3-7B). However, we observed that the particles exhibited increasing
aggregation when higher RNA concentrations were used (not shown), likely due to the
bridging of particles by RNA.
A 6-
0
B
Pylated NP
6~0*
0
0 NP
0
PEGylated NP
NPyatdN
00
.0
$150-
4'
0
250
0150.0
0
01
0
6
8
10
12
0
Incubation time (h)
100
200
30
40
Poly I:C in solution
(pg RNAJ mg NP)
Figure 3-7. RNA adsorption to lipid-coated PBAE particles in water.
(A) Kinetics of poly I:C adsorption to PEGylated or non-PEGylated PBAE particles for
300 jig particles incubated with 4 pg RNA in 300 ptL of water. (B) Binding of Poly I:C at
varying concentrations to 300 jig particles in 300 jiL of water for 2 h.
To compare poly I:C and mRNA loading, we used single-stranded mRNA
transcripts encoding GFP or luciferase. Fixing the mass of particles at 300 jig and RNA
at 4 pig in 300 piL of nuclease free water, both poly I:C and mRNA bound to particles
with over 95% efficiency (Figure 3-8A) without inducing substantial aggregation (Figure
3-9), resulting in ~12-14 jig of RNA loading per mg of particles(Figure 3-8B), a total
were73
loading comparable to other biodegradable particle systems where polynucleotides 171
-1
micro/nano-particles.
in
encapsulation
or
loading
surface
either
by
incorporated
This non-saturating level of mRNA loading on the particles did not lead to a significant
change in the zeta potential of the particles in deionized water (Table 3-1); we believe
this reflects the dominant role of the charge density present in the partially-ionized PBAE
core on the zeta potential of the particles.
43
A.
B
PEGy~NedN
M NP
p100
CD50-
,-
ZM
NP
-'1s
w
510
C
13
so.
mRNA
Poly I:C
mRNA
Poly I:C
Figure 3-8. Loading of poly I:C and mRNA on lipid-coated PBAE nanoparticles.
Binding efficiency (A) and RNA loading (B) of mRNA or poly I:C assessed for RNA (4
tg) added to 300 pg PEGylated or non-PEGylated particles for 2 h in 300 pL of water.
*
*
PEGylated NP
PEGylated NP + mRNA
Z.100-
0.~0
0
0
100
200
300
400
500
Particle diameter (nm)
Figure 3-9. Size histograms of PBAE particles before and after mRNA adsorption.
Size of PEGylated particles measured by dynamic light scattering before and after mRNA
adsorption at optimized loading conditions (4 ptg mRNA added to 300 tg particles in 300
pL of water).
To assess how well adsorbed RNA would remain bound to particles upon
exposure to physiological conditions, we measured the kinetics of poly I:C release from
particles in RPMI containing 10% serum (FBS) at 37 *C. A burst release of
approximately 30% of the poly I:C desorbed within the first hour while the remaining
RNA was gradually released over several days (Figure 3-10), suggesting that a significant
amount of RNA would likely remained bound onto the particles long enough for cellular
uptake.
44
PEGylated NP
0
100
50
IL
0
0
10
20
30
40
140
160
time (h)
Figure 3-10. Release profile of poly I:C from PBAE particles at physiological
conditions.
Release of poly I:C from nanoparticles was measured by incubating particles in RPMI
medium containing 10% FBS at 37 'C to stimulate physiological conditions.
3.3.4.
Protection of surface-loaded RNA from nuclease degradation
A major issue for RNA-based therapeutics is their susceptibility towards
nuclease degradation. To determine if mRNA loaded onto the surface of lipidenveloped PBAE particles would be protected from nuclease degradation, we
analyzed free or particle-bound mRNA by gel electrophoresis following exposure to
serum containing RNA nucleases. As shown in Figure 3-11, naked mRNA began to
undergo degradation following as little as 5 min exposure to buffer containing 2%
FBS, and was almost entirely degraded within 1 h when incubated with buffer
containing 10% FBS (bright band detected at shorter fragment size after 5 min
incubation and no bands after 1 h incubation) compared to untreated mRNA (cntl). In
contrast, minimal degraded fragments were detected for particle-loaded mRNA,
where the majority of the nucleic acid remained intact on the particles following
nuclease treatment (trapped in the loading wells, top band of the gel), suggesting that
mRNA adsorbed onto the particles was protected from immediate nuclease activity.
45
NP
Naked
n
Cnd
1
,
2% 10% 2% 10%
A
2% 10% 2% 10%
Figure 3-11. Surface loading of mRNA on PBAE particles protects mRNA against
nuclease degradation.
1 ig mRNA (either free mRNA (Naked) or bound to PEGylated PBAE nanoparticles
(NP)) was incubated with 2 or 10% FBS for 5 min or 1 h at 37 'C (or left untreated, Cntl)
and analyzed by electrophoresis on a 1% agarose gel. Upper panel shows mRNA detected
in the loading wells when mRNA was loaded onto particles and lower panel shows intact
free mRNA (Cntl) and degraded mRNA fragments obtained when samples were exposed
to serum nucleases. Naked mRNA is completely destroyed by 1 h when exposed to 10%
FBS.
3.4. Conclusions
In this chapter, we described protocols for creating biodegradable core-shell
particles comprising of a functional polymer enveloped within a lipid bilayer. The core
containing a pH-sensitive polymer, PBAE enable these particles to sense changes in the
pH of the surrounding environment and potentially disrupt endosomes in cells, thereby
99 0
mediating targeted and intracellular delivery as demonstrated by several studies. -1 4, 1 2
The lipid shell can be composed of a variety of lipids, providing a versatile platform for
functionalization of and loading of molecules onto particle surfaces.
The chapter closes with a detailed examination of the latter with RNA as a novel
model antigen. We demonstrated RNA can be simply mixed with fully-formed particles
to allow for binding via electrostatics to particle surfaces. A key advantage of this
strategy is the ability to associate potentially fragile antigens and immunotherapeutics
under mild aqueous conditions and avoid exposure of to harsh processing conditions
commonly employed for encapsulation strategies. Taken altogether, this core-shell design
may be useful for enhancing cytosolic delivery of vaccine antigens and cancer
immunotherapeutics, as we will see in chapter 5.
46
4. Transcutaneous delivery mediated by polyelectrolyte
multilayer films in murine ear skin
4.1. Introduction
Beginning with this chapter, we turn to examine the application of PBAE-based
degradable delivery platforms, as described in the earlier sections, for addressing some of
the key issues currently limiting the delivery of vaccines and immunotherapeutics,
starting with the challenges associated with transcutaneous delivery.
Given the ability of multilayers to be conformally coated on a broad range of
substrates, to load both small-molecule and macromolecular drugs, and to regulate the
release of drugs over a period of hours to days, we became interested in the potential of
these polyelectrolyte films in the area of transcutaneous drug and vaccine delivery. Skin
is an attractive site for non-invasive delivery of therapeutics, due to the ease of access to
this organ and the ability of transcutaneous delivery to limit first-pass drug metabolism
and alter the pharmacokinetics and pharmacodynamics of drugs. 174, 175 Epicutaneous
immunization, the topical delivery of vaccine antigens into the epidermis or dermis of
skin, offers the advantages of needle-free
delivery while targeting an immune sentry-rich
76 78
natural portal of pathogen entry.1 -1
The skin is replete with APCs, namely Langerhan cells (LCs) that populate the
epidermis just above the basal layer and deeper-lying dermal DCs. There, they exert a
sentinel role where they sample antigen and increase their rate of migration out of skin, in
response to activating stimuli, to LNs where they present antigen and costimulatory
molecules to T and B cells and thereby initiate an immune response. In particular, LCs
underlies 25% of the skin area 179 and are the principal target in strategies targeting the
skin immune system. Current immunizations are typically performed by either
subcutaneous or intramuscular injections and would therefore bypass the LCs that could
be of most interest for vaccination. Furthermore, the muscle is an inefficient site for
antigen presentation due to the lack of suitable quantities of APCs.180 Myocytes lack
expression of MHC-Class II and costimulatory molecules and thus cannot directly prime
T cells.' 8 1 They also do not have ready access to T cells in lymphoid tissues, as is the case
for DCs.1 82, 183 Many studies have found that intradermal immunization requires smaller
doses of antigen compared to intramuscular immunization, 184-191 probably due to better
availability of APCs at the site of inoculation. Moreover, cutaneous immunization can
elicit mucosal immunity in addition to systemic immunity, generating a more potent
immune response for enhanced protection.
To better access and mobilize the skin APCs, we hypothesized that drug-loaded
multilayers coated on the surface of a skin-adhesive patch or wound dressing could
provide several attractive features for transcutaneous delivery: (i) Interactions of
embedded proteins with the multilayer components might stabilize proteins against
aggregation on drying, facilitating storage of protein therapeutics in a dry state; (ii) The
concentration of a substantial cargo of drug into an ultrathin coating directly apposed to
47
the surface of the skin would provide a strong diffusive driving force promoting
penetration of drug cargos into the skin; (iii) multilayers could be designed to release
multiple components and separately regulate the kinetics of release of individual drug
components to optimize a therapeutic response; and (iv) the versatility of multilayer
assembly would allow this concept to be implemented in a variety of transcutaneous
delivery settings, including coatings of simple skin-adhesive patches, woven-fiber
adhesive dressings, or microneedle arrays. Advances are also being made to fabricate
free-standing films. 192
In this chapter, model elastomeric substrates coated with multilayers were applied
to tape-stripped skin (a simple procedure often used to permeabilize the outer layers of
the stratum corneum in transcutaneous vaccination studies193 - 95 ), rehydrate and dissolve
in situ, releasing protein into the epidermis. Using confocal fluorescence microscopy and
a transgenic mouse model where epidermal APCs including LCs express green
fluorescent protein (GFP)-tagged Major Histocompatibility Complex (MHC) molecules,
we demonstrate that protein antigen, as well as a synthetic adjuvant, CpG oligonucleotide
released from multilayer patches is taken up by LCs within hours of application of the
film and eventually transported to the skin-draining lymph nodes. To our knowledge, this
is the first demonstration of LbL films applied for the delivery of biomolecules into skin.
4.2. Materials and methods
4.2.1.
Materials
Reagents involved in assembly of ovalbumin- and CpG-loaded PEM films are the
same as those described in Chapter 2.
All antibodies, anti-mouse PE-I-Ab, anti-mouse APC-CD1lc, anti-mouse PECD1lb, anti-mouse PE-CD40, anti-mouse PE-CD86 were purchased from BD
Biosciences (San Jose, CA).
4.2.2.
In vivo murine ear skin penetration
Animals were cared for in the USDA-inspected MIT Animal Facility under
federal, state, local and NIH guidelines for animal care.
Tape-stripping was performed on wild type C5BL/6J mice (Jackson Laboratories,
Bar Harbor, ME) and MHC-Class II GFP mice (a gift from Professor Hidde Ploegh)
immobilized under isofluorane anesthesia. Ears were hydrated briefly with PBS before
being subjected to tape-stripping 10 times using cellophane adhesive tape (Office Depot
invisible tape). Tape-stripped ears were subsequently swiped with PBS and pat dry before
the application of LbL films (on their supporting substrates) onto the dorsal side. The
application site was then sealed using Tegaderm and Nexcare medical tape (3M,
Minneapolis, MN). Patches were kept on site for 2 days, following which mice were
sacrificed and their ears excised.
After removing the patches and rinsing with PBS, ears were split into dorsal and
ventral halves. For wild type mice, the cartilage-free dorsal halves were immunostained
overnight with anti-mouse PE-I-Ab (MHC-Class II) antibody (2 [tg/mL) to mark the LCs
present. The stained ear sheet was then mounted on a glass slide and 3D Z-section images
48
were obtained on a Zeiss LSM 510 confocal microscope using a 63x objective to
visualize the penetration into skin.
Relative fluorescence was determined by using Metamorph imaging software to
compute the integrated intensity of images taken at each depth. Values obtained were
then normalized by the maximum value measured in the entire stack of images; for the
case of LC fluorescence, background signal was also subtracted prior to normalization.
Histological sections of intact and tape-stripped mouse ears were prepared by
fixing ears with formalin for 20 h and embedding in paraffin. Sections were sliced and
stained with haemotoxylin and eosin.
4.2.3.
Analysis of cells in auricular lymph nodes
The lymph nodes (LNs) draining the ear skin, the auricular LNs recovered from
mice were digested in a cocktail of Liberase Blendzyme 2 (Roche Applied Sciences,
Indianapolis, IN) (0.28 WU/mL) and DNase I (Sigma, St Louis, MO) (50 ng/mL)
prepared in serum free RPMI with gentle shaking for 1 h at ambient temperature. EDTA
(5 mM) was added to quench enzymatic activity and the suspension was allowed to settle
for 10 min, following which the freed cells in suspension were aspirated from the
remaining undigested tissue matrix. Recovered cells were rinsed once in PBS containing
1%BSA and EDTA (5 mM), then twice in PBS containing 1%BSA before resuspending
them in flow cytometry buffer (1% BSA, 0.1% NaN 3 in Hank's balanced salt solution,
pH 7.4) at 4 'C. The cells were blocked with anti CD 16/CD 32 antibody (10 gg/mL) for
10 min and then stained with fluorescent antibodies against surface markers for 20 min
on ice, followed by three washes with flow cytometry buffer and addition of propidium
iodide (PI) (1.25 pg/mL) for viability assessment. Stained cells were analyzed on a BD
FACSCalibur flow cytometer.
4.3. Results and Discussion
4.3.1.
Disruption of stratum corneum barrier by tape-stripping
To assess antigen penetration into skin from LbL multilayers, we employed a
murine ear skin model often used for analysis of antigen delivery for epicutaneous
vaccination. 196 , 197 This site provides a clearly defined draining lymph node (the auricular
node) for immunological analysis, and Langerhans cells (LCs) within the ear are readily
visualized by microscopy, allowing direct observation of protein delivery to these
dendritic cells in intact tissue. Importantly, despite differences between murine and
human skin structure, murine studies have been remarkably predictive for human clinical
trials.1 77
In order for vaccine agents to reach viable immune cells in the epidermis or
dermis, the stratum corneum (SC), the thin, nonliving, exterior lipid-rich barrier layer of
the skin must be bypassed. 198 One approach to transcutaneous vaccination is to employ
adjuvant compounds that disrupt the integrity of the SC, such as cholera toxin derivatives
or heat labile-enterotoxin from bacteria.1 5 , 136 A simple and safe alternative is to
physically disrupt the SC. This can be achieved using an approach such as glue or tapestripping, based on the repeated application and removal of adhesives on skin, a process
49
shown to remove the outer 10 - 20 pm of the SC. This barrier disruption process is
painless, highly effective in augmenting the transfer of macromolecules such as proteins
or DNA into skin, and has been successfully utilized in human clinical trials. 19 7 , 1 9-201 An
additional advantage is that physical barrier disruption itself has been shown to trigger
activation of keratinocytes and LCs in skin, providing some level of 'built-in' adjuvant
response.137,
196, 197, 200-203
We thus tested a model where the dorsal ears of anesthetized mice were tapestripped followed by application of as-assembled, air-dried (Poly-1/OVA) 4o films on
either glass or PDMS substrates, fixed in place by surgical tape against the exposed ear
skin, mimicking a common strategy used in skin vaccination studies' 96 , 197 (Figure 4-1).
Tape-stripping reproducibly removed significant portions of the SC but did not disrupt
the viable epidermis, as assessed by histology (Figure 4-2).
Poly 1
Ear Skin
Ovalbumin protein antigen
(1)Apply PEM patch to tape-stripped skin
(2) Internalization of antigen
released by LCs
Figure 4-1. Schematic illustration of the events taking place following the
application of PEM patch onto skin.
(1) PEM patch on a substrate is applied onto tape-stripped ear skin. (2) OVA is released
from the patch as the multilayer disassembles, and OVA then penetrates to reach the LCs
located at the basal layer of the epidermis.
Figure 4-2. Histological sections of murine ear skin before and after tape-stripping.
Histological sections of intact (left) and tape-stripped (right) mouse ears stained with
haemotoxylin and eosin. Most of the stratum corneum is removed following tapestripping.
4.3.1. Penetration of ovalbumin released from films into murine ear
skin and uptake by Langerhan cells
50
To assess protein release from LbL films into tape-stripped skin, films carrying
fluorescent OVA were assembled on PDMS patch substrates and applied to tape-stripped
ear skin of C57Bl/6 mice or transgenic MHC II-GFP mice, whose Langerhans cells
express GFP-tagged MHC molecules and are thus intrinsically labeled in the tissue. For
comparison, we applied gauze pads loaded with an equivalent amount of fluorescent
OVA in saline to tape-stripped skin, mimicking the approach typically taken in vaccine
studies. Mice were euthanized 48 h later, the PDMS patch or gauze was removed, and
the skin sites were examined by confocal microscopy.
Optical sections revealed the penetration of OVA released from LbL films into the
epidermis down to depths where LCs reside, 10 - 15 pm below the surface of tapestripped skin (Figure 4-3a-b; LCs are readily distinguished by their branched/dendritic
morphology). In separate experiments using C57Bl/6 mice, LCs deep in the epidermis
below the main front of skin-penetrating OVA were observed with pockets of
internalized OVA (Figure 4-3c), indicating that the cells were able to take up the
delivered protein antigen from their surroundings. At the 48 h time point, LCs were
observed with less pronounced dendritic morphology and at deeper depths than seen in
fresh skin explants, suggesting that the cells were activated and in the process of
beginning their slow migration to lymph nodes - a process known occur over a few
5
In contrast, OVA solutions applied to tape-stripped skin showed much less
days.
protein accumulation in the skin, and LCs were present at depths and densities in the
epidermis more reminiscent of untreated skin (Figure 4-3b).
51
a
.J
0
-J
bOva film
Ova gauze pad
z
X6x10
ox10
lOx10
*
3x10
4g*
Ux108
7
-
x10
2x103
0
lxlO
0
0
0
0
10 20 30 40 50 60 70
Depth (pm)
Figure 4-3. Penetration of ovalbumin in tape-stripped murine ear skin.
(a) Overlay confocal micrographs (upper panel) taken at 5 gm intervals showing the
penetration of Texas Red-conjugated OVA (red) released from LbL films into tapestripped ear skin of a MHC-Class II-GFP mouse following application of a Poly-1/OVA
PDMS patch for 48 h. Corresponding GFP fluorescence confocal micrographs (lower
panel) showing Langerhans cells (LCs, green) appearing at z-depths of 10-15 ptm below
the tape-stripped skin surface, just above the epidermis-dermis boundary. (b)
Representative vertical cross-sections and quantitative plots of fluorescence intensity
normalized to maximum values for OVA (red) and LC fluorescence (green) in skin
treated with OVA-loaded LbL films (filled symbols) or gauze pads with OVA solution
(open symbols). (c) Micrograph from an independent experiment, demonstrating
colocalization of Alexa Fluor 488-conjugated OVA (red) released from a Poly-1 LbL
film and immunostained LCs (green) in the skin of a C57bl/6 mouse observed at a deeper
depth where free OVA in the surroundings is no longer detected, suggesting that
activated LCs have taken up penetrated antigen and begun migration to lymph nodes. All
scale bars are 20 prm.
52
Dual penetration of ovalbumin and CpG released from
4.3.2.
films into murine ear skin and uptake by Langerhan cells
We next tested the ability of OVA/CpG composite films coated on PDMS
"patches" to deliver both antigen and adjuvant into tape-stripped murine ear skin. Similar
to the experiments with Poly-1/OVA films, multilayer-coated PDMS patches loaded with
fluorescently-tagged OVA and CpG (architecture "B", see Figure 2-8a) were applied to
ear skin of MHC-Class 1I-GFP mice for 48 h, then patches were removed and the skin
sites were analyzed by confocal microscopy.
Optical sections tracked the penetration of both labeled OVA and CpG released
from the film into the epidermis of tape-stripped skin down to 15 - 20 Rm below the tapestripped surface (Figure 4-4a). Colocalization of GFP* LCs with OVA and CpG
fluorescence (Figure 4-4b) were likewise observed, suggesting LCs internalized both
penetrated antigen and adjuvant from their surroundings.
Interestingly, at the 48 hr time-point LCs were detected at greater depths of 15 20 ptm compared with films containing only OVA, possibly reflecting the effect of CpG
activating LCs and enhancing their migration toward lymphatics located deeper in the
skin en route to regional lymph nodes. Similar results were observed for films with
architecture "C" (see Figure 2-8a).
Thus, erodible ultrathin films loaded with antigen and adjuvant compounds can
deliver both components into skin, providing delivery to dendritic cells in the skin as
required for the initiation of immune responses.
53
aJ
5
20
0
0
4
_J
CL
4 r
5
0
0
.
5
10
20
Z
Figure 4-4. Co-penetration of ovalbumin and CpG in tape-stripped murine ear skin.
(a) Overlay confocal micrographs (top and middle panels) taken at 5 pim z-intervals (zdepth indicated in upper left of each image) showing the penetration of fluorophoreconjugated OVA (red) and CpG (blue) released from LbL films (architecture "B" from
Figure 2-8a) into tape-stripped ear skin of MHC-Class II GFP mouse following patch
application for 48 h. Corresponding GFP fluorescence confocal micrographs (bottom
panel) showing LCs (green). (b) Representative horizontal (left) and vertical (lower right)
cross sections showing colocalization of fluorescent OVA and CpG with GFP+
Langerhans cells (indicated by white arrows; inset figure shows individual fluorescent
channel images of boxed region), suggesting uptake of penetrated antigen and adjuvant
by the dendritic cells. Similar results were also observed for film architecture C (Figure
54
2-8a, data not shown). Depths (below tape-stripped skin surface) of x-y confocal optical
slices are noted (in im) in upper left of micrographs. All scale bars 20 pm
4.3.3.
Transport of antigen to draining lymph nodes
Immune responses are initiated in lymph nodes, where antigen-bearing dendritic
cells directly interact with T- and B-lymphocytes, and thus antigen delivery to the nodes
is critical for primary immune responses. Transport of antigens from the skin to the local
lymph nodes can occur through direct draining of antigen deposited into the skin via
lymphatic vessels or by cell-mediated transport, where activated LCs or other dendritic
cells internalizing antigen in the skin migrate to the lymph nodes. To determine whether
protein antigen released from LbL films applied to skin was transported to draining
lymph nodes, murine ears were tape-stripped and LbL films on glass substrates carrying
Alexa-fluor-conjugated OVA were applied to the skin using surgical tape.
As before, we compared LbL films to skin treated with gauze loaded with an
equivalent amount of OVA in saline. After 48 h, the skin-draining auricular lymph nodes
were excised, digested, and recovered leukocytes were stained with antibodies for flow
cytometry analysis. At this time-point, OVA fluorescence could be directly visualized
within intact lymph node tissue at low magnification (Figure 4-5a).
CD 11 c is generally regarded as a marker for dendritic cells while CD 11 b is used
to mark cells of myleoid lineage including monocytes and macrophages. Murine LCs and
other dendritic cells migrating from skin have been shown to express both CD 1l c and
CDl lb surface molecules. 206 207 As illustrated in Figure 4-5b-c, protein transport to the
lymph nodes was detected following treatment with a (Poly- 1/OVA) 5 o LbL patch, when
examining both the total lymph node cell population as well as individual cell subsets.
Though CDl lcCDl lb+ dendritic cells make up only a few percent of the total lymph
node cells (e.g., as shown in the CDI Ic vs. CD 11b flow cytometry plot for OVA- cells),
these cells were greatly preferentially enriched in the OVA+ cell population, with ~64%
of OVA+ cells being of the CDl lcCD1 lb+ subset (Figure 4-5b). As shown in Figure
4-5c, -14% of all CDl lcCDl lb+ dendritic cells acquired fluorescent OVA following
LbL film release of OVA into the skin. Compared to tape-stripped skin treated with
OVA solution, -1.8-fold more DCs and 2-fold more macrophages were positive for OVA
in the draining lymph nodes following LbL film delivery of the protein.
OVA*CD 11 c+ dendritic cells exhibited up-regulation of the cell-surface activation
markers CD40, CD86, and MHC II compared to OVA~ dendritic cells (Figure 4-5d),
suggesting that antigen-bearing DCs originating from skin acquired a mature phenotype
during migration to the draining lymph nodes, consistent with prior studies. 208,09 Mature
dendritic cells are poised to present antigen and provide the appropriate stimulation
required to trigger T- and B-cell activation. Thus, protein antigen released into barrierdisrupted skin from LbL films is delivered to skin-draining lymph nodes, where it
accumulates in mature dendritic cells.
55
a
b
O*
......
C
0
.$63
1
0.01
200 400 800
CU
12.
M
000
1&
10
10
10
Ova-
PBs
CDllc*/CDllb'
14
Ova
Film
0.00
CD11c/CD11b1
0
0
10
CD11cI/CD11b
>
0
10x
10
6
1U
410
2L
2
0.0
0#
00
0 L
(Polyl/Ova)
200
X 60
0
160
le
I
CD40
41
400
600 .00
1c1
1000
Forward scatter -
Ova gauze pad
0
100
0
9.37
100
60
60
i
1
lf
CD86
10
1
1.
lo1p
101
10?
CD11b
100
0
id)
0
id)4
1
-
4
MHC Class 11
Figure 4-5. Transport of ovalbumin to draining lymph nodes.
(a) Confocal micrograph of intact auricular lymph node draining the ear site where OVA-
loaded LbL film was applied, showing OVA fluorescence (green) accumulated in the
tissue. (scale bar = 20 pm) (b) Left panels: Flow cytometry plots showing the proportion
of total lymph node cells carrying OVA (percentages quoted quadrant corners) as
detected in the draining lymph node following skin treatment with OVA-loaded LbL
films or PBS as a control for 48 h. Right panels: Plots of CDl lb vs. CD l c for OVA+
(green quadrant) and OVA~ (red quadrant) cells from the lymph nodes of mice treated
with OVA-loaded LbL films, showing that the majority of OVA-bearing cells were
CDllb*CDllc* Langerhans cells and dendritic cells. (c) Percentages of OVA+ cells
detected in macrophage and dendritic cell population subsets of skin-draining lymph
nodes determined by staining recovered leukocytes for CDlic and CDllb surface
markers. (d) Flow cytometry histograms of surface expression of costimulatory
molecules CD40, CD86 and MHC-Class II by OVA*CD 11 c dendritic cells (green line)
and OVA~CD 11 c dendritic cells (red line) recovered from lymph nodes of mice treated
with OVA-loaded LbL films.
4.4. Conclusions
In this chapter, we demonstrated a proof-of-principle for the use of dry LbL films
as skin vaccine patches in a murine skin model, whereby films are rehydrated in situ on
56
barrier-disrupted skin. The antigen and adjuvant are thereby released into the epidermis
where they are taken up by skin-resident dendritic cells. The antigen is eventually
transported to draining lymph nodes and the activation of antigen-bearing dendritic cells
is also observed.
To access skin APCs, a physical method, tape-stripping was successfully
employed to disrupt the barrier imposed by the SC and has been demonstrated to be
feasible clinically. 97, 199-201 Notably, such physical disruption of the SC is not only
effective at removing the stratum corneum, thereby enhancing the permeability of
molecules across epithelial tissue; they can also activate keratinocytes and skin resident
APCs and act as an adjuvant, enhancing the immune response. 137, 196, 197,200-203
While not explored in this thesis, it is worth mentioning that it is also possible to
build in a skin-penetrating capability as part of the LbL patches themselves. One way
would be to incorporate skin penetration enhancers as a component in a multi-component
LbL film to help overcome the SC barrier and enhance the penetration of immunogens.
Alternatively, the ability of multilayers to conformally coat substrates should allow films
to assembled on microneedles, a strategy based on an array of micron-sized needles that
has been shown to enable painless skin delivery. 1 , 21 Such a strategy is the subject of
on-going studies in our lab.
We envisage that these PEM-based skin delivery platforms will be especially
attractive for vaccine and drug delivery in the developing world, where a transcutaneous
route of administration has the potential to increase ease and speed of delivery without
the need for medical training, decrease costs, as well as offer improved safety and
compliance by reducing the pain and risks of injuries and disease transmission associated
with the use of needles, the existing mode of delivery for most therapeutics.
57
5. Cytosolic delivery mediated by lipid-enveloped polymer
nanoparticles in vitro
5.1. Introduction
Many strategies for the prevention or treatment of disease require the delivery of
therapeutic agents into the cytosolic or nuclear compartments of cells. Examples include
0
gene therapy mediated by plasmid DNA, 104, 212, 213 gene silencing via siRNA,1 3, 104, 214
antitumor toxin delivery, 9,2s,216 and therapeutic protein delivery. 17,218 However the cell
membrane is largely impermeable to a majority of these molecules by virtue of their large
size or hydrophilic nature relative to the hydrophobic lipid bilayer constituting the cell
membrane. Although cells may take up macromolecules via endocytosis,
macropinocytosis, or phagocytosis, these processes confine the internalized compounds
to vesicles such as endosomes or phagosomes, where the pH is progressively lowered to
5.5-6.5. 212 The fusion of these vesicles with lysosomes, intracellular compartments
carrying the degradation machinery of the cell at a pH as low as 4.5,219220 often leads to
rapid destruction of therapeutic molecules with little or no release into the cytosol. The
limited cytoplasmic delivery of enzyme susceptible biomolecules therefore remains a
major challenge limiting the development of effective intracellular therapies. In
particular, we are interested in the cytosolic delivery of therapeutics for immunization
and immuno-modulation/therapy.
One prominent example is in the area of gene-based vaccination where the
utilization of DNA or mRNA as a source of antigen requires, as a first step, delivery into
the cytosol, in order for processes leading up to the translation of nucleic acids into
protein antigen to take place. Gene-based vaccination was born in 1990 after Wolff et al.
demonstrated the local uptake and expression of exogenously injected plasmid DNA and
in vitro transcribed messenger RNA (mRNA),m followed shortly thereafter by the
demonstration that injected nucleic acids could promote immune responses to encoded
antigens.222, 223 Both DNA and mRNA-based vaccines share many advantages over
alternative protein, peptide, or live vector-based vaccine strategies in terms of safety,
manufacturability and suitability for long-term storage, and the ability to promote broad
cytotoxic T-cell responses.224,22
Another area of research being actively pursued is the delivery of immunotoxins as
cell-targeted therapeutics for cancer therapy. This is achieved by attaching a targeting
moiety (derived from antibodies or other cell-binding proteins) selective for the tumor
cells to cytotoxic toxin molecules. However, the potency of such constructs is still
ultimately dependent on their ability to reach intracellular targets present in the cytoplasm
following internalization, which is commonly considered to be the rate-limiting step.
While the inclusion of toxins with domain facilitating cytoplasmic access can address
this, they can also lead to increased nonspecific toxicity in vivo.5' 5 Materials designed
to enhance endosomal escape can facilitate cytosolic delivery of antitumor toxins and
improve their efficacy.
58
To deliver macromolecules intracellularly, synthetic vectors are preferred over
approaches based on viral vectors for their low cost, ease of large-scale production and
potential for improved safety. 89, 226-228 A variety of polymer and/or lipid-based drug
delivery systems with the capability to disrupt endosomes have been developed,82 -ss 22
230 but systems that achieve efficient cytosolic delivery or transfection in vivo with
minimal cytoxicity are still sought. In the case of antitumor toxins, the target cell of
interest is the tumor cell. For vaccine applications, cytosolic delivery of mRNA into
dendritic cells (DCs), immune cells that play a key role in the initiation of adaptive
immune responses,2 3 1 is desired but the transfection of these cells poses a significant
challenge as synthetic agents generally achieve only 10-35% transfection of these cells in
vitro.,5556,232
In addition, many of the chaperone molecules that efficiently aid transport of
macromolecules into the cytosol are formulated with therapeutic cargos by physical
complexation of the chaperone and therapeutic (e.g., polyplexes or lipoplexes of cationic
polymers/lipids with nucleic acids), forming particles whose size, stability, and properties
are highly dependent on formulation parameters including the properties of the cargo
molecule, the cargo-to-chaperone weight ratio, and the characteristics of the surrounding
environment (pH, ionic strength, and presence/absence of serum proteins). 82, 89 Lack of
control over particle size and stability is of concern because particle size is a critical
determinant of cellular uptake in vitro and biodistribution and toxicity in vivo. 85
Therefore, a modular design that would allow for the separate formulation of cargo and
delivery agent is also desirable.
In this chapter, we evaluate such a system, the lipid-enveloped PBAE
nanoparticles described in chapter 3 in a DC cell line and various tumor cell lines in vitro
for mRNA vaccine antigen delivery and antitumor toxin delivery respectively. We first
demonstrated the capability of particles to disrupt endo-lysosomes and mediate efficient
cytosolic delivery of a model therapeutic molecule, calcein, following co-delivery with
particles. We further verified the effect of the lipid shell in mitigating undesirable
cytotoxicity by sequestering the cationic charge in the polymeric core. We were then able
to go on and show that these particles delivered mRNA, adsorbed onto the particle
surface post-synthesis, into the cytosol of DCs with minimal cytotoxicity, leading to in
vitro transfection of a DC clone at levels comparable to the best reports for DC
transfection from the literature. Finally, we also explored the application of these
particles for enhancing intracellular delivery of antitumor toxins to achieve synergistic
killing of tumor cells in vitro.
5.2. Materials and methods
5.2.1.
Materials
Reagents involved in the synthesis of lipid-coated PBAE nanoparticles and the
preparation of mRNA are the same as those described in Chapter 3.
Calcein and DAPI were purchased from Sigma Chemical Co. (St. Louis, MO)
Annexin V Alexa 350 conjugate, LysoTracker Green, unlabeled Phalloidin and
Phalloidin Alexa 488 conjugate were purchased from Invitrogen (Eugene, OR). WST-1
reagent was purchased from Roche Applied Science (Indianapolis, IN).
59
The immunotoxin E4rGel, a fusion protein based on a plant-derived toxin gelonin
linked to an engineered fibronectin domain targeting the epidermal growth factor receptor
(EGFR) was constructed and produced according to previous literature. 233 Briefly,
expression plasmid containing the genes encoding the recombinant form of the gelonin
toxin and an engineered fibronectin fragment (based on the tenth human fibronectin type
III domain (Fn3)) binding EGFR was constructed and expressed. Immunotoxin protein
was subsequently extracted and purified via column chromatography.
5.2.2.
Analysis of endosomal disruption
The dendritic cell clone DC2.4 (gift from Prof Kenneth Rock) was plated at 1.2 x
10s cells/well in Lab-Tek chambers (Nunc) for 18h, and then calcein (150 pg/mL, 0.24
mM) was added to the cells with or without 75 gg/mL of lipid-coated nanoparticles
containing either a poly-1 or PLGA core in RPMI 1640 complete medium (10% fetal
bovine serum (FBS), 5mM L-glutamine, 10mM HEPES, and penicillin/streptomycin) for
1 h at 37 'C. After washing with medium to remove extracellular calcein/particles, the
cells were imaged live under a confocal microscope (Zeiss LSM 510) at 63x. To quantify
the percentage of cells displaying cytosolic/nuclear distribution of calcein, cells were
detached with Trypsin/EDTA, resuspended in flow cytometry buffer (1% BSA, 0.1%
NaN 3 in Hank's balanced salt solution, pH 7.4) and analyzed by flow cytometry, gating
on the cell population expressing high mean fluorescence intensity beyond that exhibited
by cells treated with calcein alone.
5.2.3.
Cytotoxicity assay
Cytotoxicity of the particles was assessed by incubating DC2.4 cells (6 x 10 5
cells/well in 12 well plates seeded 18 h prior to experiments) with 50, 75 or 100 pg/mL of
lipid-coated PBAE particles with or without PEG-lipid incorporation, or PBAE particles
stabilized with poly-vinyl-alcohol (PVA) as surfactant, in complete medium for 1 or 12 h
at 37 'C. After washing with medium to remove extracellular particles, the cells were
detached with Trypsin/EDTA, resuspended in complete medium and rested for 1 h at 37
'C. The cells were then stained with DAPI and annexin V according to manufacturer's
recommendation to identify cells undergoing apoptosis and necrosis. The percentage of
live cells was quantified via flow cytometry (BD LSR II) by counting cells that were
negative for both stains.
5.2.4.
In vitro transfection of DCs
To test whether mRNA loaded onto particles was able to escape from endosomes
into the cytosol, DC2.4 cells were incubated in the presence of a pH-sensitive fluorescent
indicator LysoTracker Green (1 1 M) and 1 gg of Cy3-labeled mRNA, either naked
mRNA in serum-free Opti-mem media or the same quantity of labeled mRNA loaded
onto PBAE particles (75 Rg particles/mL) in RPMI medium containing 10% FBS. After
1 h, the cells were washed and imaged live by confocal microscopy.
In vitro transfection of DC2.4 cells was also visualized via fluorescent confocal
microscopy using Cy3-labeled mRNA encoding for green fluorescent protein (GFP).
60
Because mRNA loading on particles was limited by particle aggregation at high
mRNA:particle ratios, conditions for transfection studies were determined by identifying
the maximum concentration of particles that could be incubated with DCs in vitro
without giving rise to overt cytotoxicity (75 pg particles/mL), and maximizing the
amount of the mRNA that could be loaded onto this quantity of particles without
inducing particle aggregation. This was determined to be -4-8 ptg of mRNA per 300 pig
particles. mRNA-coated lipid-coated PBAE particles with or without PEGylation (75
pg/ml), loaded with 1 pg of mRNA, were diluted into RPMI complete medium
containing 10% FBS and incubated for 12 h with 6 x 105 cells seeded in 12-well plates 18
h prior to experiments. mRNA uptake and transfection efficiency were further assessed
by flow cytometry, following rinsing with fresh medium and trypsinization, to quantify
the percentage of cells positive for Cy3 and/or GFP, compared with control cells
incubated with naked mRNA. Particle uptake efficiency was quantified simultaneously
by labeling particles with a lipophilic tracer DiD and measuring the fraction of cells
positive for the tracer.
Cytosolic delivery of antitumor toxin and in vitro tumor cell
5.2.5.
killing
Cytosolic delivery of antitumor toxin was visualized via fluorescent confocal
microscopy using a model toxin, phalloidin, and its alexa 488 conjugate. B16F 10 cells, a
murine melanoma cell line, were plated at 1.2 x 105 cells/well in Lab-Tek chambers
(Nunc) for 18 h, and then phalloidin (10-25 piM, 5 mol% alexa 488 conjugated) was
added to the cells with or without 75 pg/mL of lipid-coated PBAE nanoparticles in
DMEM complete medium (10% FBS, 4mM L-glutamine, 4500mg/mL glucose, sodium
pyruvate and penicillin/streptomycin) for 3 h at 37 *C. After washing with medium to
remove extracellular phalloidin/particles, the cells were imaged live by confocal
microscopy.
In vitro tumor cell killing was tested using the immunotoxin, E4rGel on antigenexpressing cells A431, a human epidermoid carcinoma cell line positive for EGFR.
Tumor cells were seeded on 96-well plates at 2500 cells/well. Cells were allowed to
adhere overnight, after which fresh growth medium (DMEM supplemented with 10%
FBS, 4mM L-glutamine, 4500mg/mL glucose and sodium pyruvate) containing varying
concentrations of immunotoxin and/or particles was added to triplicate wells. Toxins
and/or particles were incubated with the cells for up to 24 h before the treatmentcontaining medium was removed and replaced with fresh medium. At 72 h, medium was
replaced with fresh medium containing the WST-1 reagent according to manufacturer's
recommendation. The assay was allowed to develop for 1-3 h under normal culture
conditions, after which plates were measured for absorbance at 450 nm. Untreated cells
and cells lysed with a 1% Triton X-100 solution were used as positive and negative
controls, respectively. Measurements were compared with the base line and normalized
to control treatments, triplicates were averaged, and standard errors were calculated.
Cytotoxicity measurements were conducted at E4rGel concentrations between 3 x 10-9
and 3 x 10-8 M and particle concentrations between 12.5 and 37.5 ptg/mL (particle
concentration was reduced accordingly due to the lowered cell density when particle
61
treatment was initiated compared to the earlier assay setups with a shorter time period
before readout).
To better assess the enhancement in tumor cell killing by co-delivering particles
as chaperones for soluble toxin, we analyzed the cytotoxicity data via a combination
index that has been reported in literature for comparing the efficacy associated with
combination therapy (in this case, toxin and particle co-treatment) relative to monotherapy (toxin or particle treatment alone).2 4 This is computed as follows: we first
calculate the observed percentage of viable cells (% viable cells,
observed)
in each treatment
group normalized to untreated control. We then compute the expected percentage of
viable cells in combination treatment groups (% viable cells, combination treatment, expected) by
multiplying the percentage of viable cells measured in groups given only particles or
toxin at the corresponding concentrations:
/ viable cells, combination treatment, expected
=
/ viable cells, particles, observed X
0
viable cells, toxin, observed
The combination index is defined as:
Combination index
= % viable cells, combination treatment, expected/% viable cells, combination treatment, observed
A ratio of greater than 1 indicates a synergistic effect while a ratio of less than 1 indicates
a less than additive effect.
5.2.6.
Statistical analysis
One-way ANOVA followed by Bonferroni's Multiple Comparison Test was
applied to determine the significance of the difference between the percentage of cells
displaying a cytosolic/nuclear distribution of calcein with different treatments.
5.3. Results and Discussion
5.3.1.
Endosomal escape mediated by lipid-coated PBAE nanoparticles
To test the ability of lipid-enveloped PBAE particles to disrupt endosomes,
calcein, a membrane-impermeable fluorophore, was used as a tracer to monitor the
stability of endosomes 235 following particle uptake by dendritic cells (DCs), a key
cellular target of interest for vaccine delivery. 45, 236 As shown in Figure 5-1A, DCs
incubated with calcein alone showed a punctate distribution of fluorescence indicative of
endolysosomal compartmentalization of internalized dye. In contrast, cells co-incubated
with calcein and fluorescently-tagged lipid-enveloped PBAE nanoparticles exhibited
calcein fluorescence throughout the cytosol and nucleus (at uniform high levels in both
the cytosol and nucleus due to free diffusion of cytosolic dye throughout intracellular
compartments), suggesting escape of calcein from intracellular vesicles following cointernalization of extracellular fluid containing both dye and particles (Figure 5-1B, C).
Cytosolic delivery of calcein triggered by lipid-coated PBAE nanoparticles was not
limited to DCs, as we observed similar results in tumor cells lines. (See appendix E)
Calcein entry into the cytosol triggered by the presence of nanoparticles required the
62
PBAE core, as calcein remained in an endosomal distribution in cells co-incubated with
calcein and lipid-enveloped PLGA particles (Figure 5-1D).
To quantitatively compare the endosomal disruption capability of PBAE-core
particles vs. PLGA-core particles vs. calcein alone in the entire cell population, cells
given various treatments were collected following trypsinization and their calcein
fluorescence intensity was measured via flow cytometry. Cells incubated with calcein
alone showed a homogeneous shift to increased green fluorescence (Figure 5-lE, F). In
contrast, cells co-incubated with calcein and lipid-enveloped PBAE NPs showed two
populations of calcein* cells: a calceinlow and calceinhigh population (Figure 5-1E, F).
Calcein, when trapped within endosomes, can be quenched via 2 mechanisms: selfquenching due to concentration of the dye in endolysosomes 237 , and quenching due to
calcein's sensitivity to pH in the endolysomal pH range. 238 Thus, the net fluorescence of
a given quantity of internalized calcein and other pH-sensitive dyes is seen to increase on
release from endolysosomes, a phenomenon reported by several groups. 239, 240 We used
this phenomenon to quantitate the fraction of DCs showing endosome disruption on the
population level, by scoring the fraction of cells exhibiting bright calcein fluorescence
beyond that observed for control cells incubated with calcein alone (Figure 5-1F). As
shown in Figure 5-1G, following a 1 h incubation with PBAE particles, -50% of the cell
population showed an enhanced calcein intensity in accordance with the cytosolic/nuclear
distribution of calcein, compared to -0% of cells incubated with calcein alone or calcein
+ PLGA particles.
To confirm that endocytosis of the nanoparticles/calcein was required for calcein
delivery to the cytosol, we incubated DCs with calcein and nanoparticles at 4 'C to block
endocytosis and found that neither calcein nor nanoparticles were internalized by cells up
to 3 h at 4 'C (data not shown). This suggests that calcein/nanoparticle uptake and calcein
entry into the cytosol of DCs required the active process of endocytosis and excluded the
possibility of calcein de-quenching due to nanoparticle fusion with the plasma membrane.
63
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cells +
calcein
cells
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cells + calcein
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Cytosoidulear calcein
cells + PBAE NPs
cells + calceln
+ PBAE NPs
u4020.
0
10
2
3
10
4
10
calceln
-af
6
10
Figure 5-1. Endosomal escape by lipid-coated PBAE particles.
pH-responsive lipid-enveloped PBAE particles chaperone the delivery of the membraneimpermeable dye molecule calcein into the cytosol of dendritic cells. DC2.4 cells were
incubated for 1 h with calcein alone or calcein and lipid-coated poly-1 or PLGA particles,
washed to remove unbound particles, then imaged live by confocal microscopy (A-D) or
64
analyzed by flow cytometry (E-G). Confocal images of DC2.4 cells incubated with
calcein alone (A) or calcein and lipid-coated PBAE particles (B) (LHS = fluorescence
overlays: red, nanoparticles; green, calcein. RHS = bright-field images). (C, D) Cells coincubated with calcein and either lipid-coated PBAE particles (C) or lipid-coated PLGA
particles (D) (LHS = Green, calcein. RHS = bright-field images). Representative images
of PEGylated particles are shown here, but a similar trend was observed with nonPEGylated particles. Flow cytometry scatter plots of particle fluorescence vs. calcein
fluorescence (E) and histograms (F) for cells treated with calcein or calcein + PBAE
particles compared with untreated control cells or cells incubated with particles alone.
(G) Average percentage of cells exhibiting cytosolic/nuclear calcein distribution after 1 h
incubation with lipid-coated PBAE particles or lipid-coated PLGA particles compared to
calcein-only control. Shown are the mean ± SD from triplicate samples.
5.3.2.
Cytotoxicity of lipid-coated PBAE nanoparticles
241 2 43
Polycations used for intracellular delivery are notorious for their toxicity;
thus we next analyzed the toxicity of this lipid-enveloped particle delivery system. We
incubated DC2.4 cells with lipid-enveloped poly-1 particles or "naked" PVA-stabilized
poly-1 particles prepared by emulsion/solvent evaporation synthesis for 1 h or 12 h, and
subsequently stained with annexin V and DAPI to identify cells undergoing apoptosis or
necrosis, respectively. PVA-stabilized Poly-1 NPs appeared less toxic than some other
polycations (e.g., polyethyleneimine, which is often reported to induce 50% cytotoxicity
at concentrations of -20-30 ptg/mL 244), but still led to -50% cell death after only an hour
incubation of cells with 100 ptg/mL of the particles, and showed an LD50 between 75 and
100 ptg/mL for longer incubations of 12 h. Lipid-enveloped PBAE NPs, irrespective of
PEGylation, showed lower toxicity at both time-points across all tested particle
concentrations (Figure 5-2); notably, particle concentrations that promoted robust
endosome disruption by the calcein assay (75 ptg/mL) showed low toxicity at both time
points. Thus, the lipid shell achieved the desired effect of mitigating toxicity of the poly- 1
polymer core.
65
U 100 pg/ml NP
PEGylated Lipid N
*75
g/mi NP
1 h incubation
Lipid NP
Z 50 jig/ml NP
PVA
PEGylated Lipid NP
12 h incubation
Lipid MIP
PVA NP
0
-J-4
-NOW-
50
100
% Live (dapi- annexin v-) relative to untreated
150
Figure 5-2. Cytotoxicity profiles of PBAE-based particles.
DC2.4 cells were incubated with increasing doses of PEGylated or non-PEGylated lipidcoated poly-i particles, or PVA-stabilized poly-i particles for 1 h or 12 h at 37 'C,
washed, detached and then stained with DAPI and annexin V to detect apoptotic and
necrotic cells by flow cytometry. Shown are the percentages of live (DAPF annexin V-)
cells relative to untreated controls. Error bars represent standard deviation of triplicate
samples.
5.3.3.
Transfection of dendritic cells with lipid-coated PBAE
nanoparticles in vitro
Efficient loading of mRNA onto the surface of particles via electrostatic
interactions was demonstrated for lipid-coated PBAE nanoparticles in chapter 3 (see
section 3.2.5) We confirmed here that mRNA coating did not block the ability of lipidenveloped poly-1 particles to disrupt endosomes, by quantifying the fraction of DC2.4
cells exhibiting cytosolic/nuclear calcein following co-incubation of cells with the dye
and mRNA-loaded or unloaded particles for 1 h via flow cytometry as discussed in the
previous section. (see section 5.3.1) mRNA-loaded lipid-enveloped PBAE particles
trended toward slightly reduced endosome disruption compared to unloaded particles, but
this did not reach statistical significance (p >0.05) (Figure 5-3).
Thus, simple
electrostatic adsorption allowed for rapid loading of the particles with substantial
quantities of RNA cargo, which could be retained on the lipid-coated poly-1 surface
without blocking their endosome disruption activity.
66
80
60
Cytosolic/nuclear calcein
--
20
20
Figure 5-3. mRNA adsorption does not hinder endosomal rupture by particles.
Average percentage of cells exhibiting cytosolic/nuclear calcein distribution after 1 h
incubation with mRNA-loaded vs. bare particles (0.2 pg mRNA adsorbed to 15 pg NPs
added in 200 gL cell culture medium). Shown are the mean ± SD from triplicate samples.
For an mRNA vaccine, these lipid-coated PBAE particles need to deliver intact
mRNA into the cytosol of antigen presenting cells such as DCs, where it can then be
translated into protein antigens for processing and presentation. Dendritic cells are
agents generally achieve only
notoriously difficult to transfect and nonviral transfection
55 56 232
10-35% average transfection of these cells in vitro. , ,
To visualize directly that mRNA loaded onto lipid-enveloped poly-1 particles was
capable of escaping endosomes to access the cytosol in dendritic cells, we imaged DCs
incubated with labeled mRNA and LysoTracker tracer to stain the endolysosomal
compartments, and examined the intracellular trafficking of mRNA loaded onto particles
compared to naked mRNA. When DCs were incubated with naked mRNA in serum-free
medium, we observed colocalization of labeled mRNA with endolysosomes (Figure
5-4A). In contrast, particle-delivered mRNA fluorescence was detected in regions of the
cell outside endolysosomes, providing direct evidence of the escape of particles and
mRNA from endosomes into the cytosol (Figure 5-4B). Notably, when naked mRNA was
incubated with DCs in complete medium containing 10% FBS, no mRNA fluorescence
could be detected in the cells (data not shown), indicating the labile mRNA is readily
degraded in the presence of serum nucleases without protection from the particles.
We then tested transfection of DCs in vitro by lipid-enveloped poly-1 particles in
the presence of serum, as a precursor to in vivo transfection studies.
DC2.4 cells were
incubated with Cy3-labeled mRNA encoding for green fluorescence protein (GFP);
mRNA was added to the cells in soluble form or adsorbed to fluorescently-tagged lipid-
enveloped NPs. Using this set of fluorescence markers, mRNA, NPs, and translated
protein (GFP) were simultaneously traced in live cells. Naked mRNA is extremely labile
67
in serum-containing medium, and in fact essentially no intact mRNA was seen to be
taken up by DCs in this study when added to the cells in serum-containing medium
(Figure 5-4C), though some uptake was detected in serum-free medium, data not shown).
In contrast, mRNA loaded on poly-1 NPs was efficiently delivered into the cells (Figure
5-4C), supporting the contention that particle adsorption protected the RNA from serum
nucleases. Further, a fraction of the cells clearly expressed GFP at the 12 h timepoint
(Figure 5-4C) . Quantifying the results via flow cytometry analysis, we found that
particles were taken up by 80% of the cells and a similar fraction of the cells internalized
mRNA (Figure 5-4D-F). Looking at protein expression, ~30% of the total cells
expressed GFP (Figure 5-4G, H), a frequency comparable to many prior best reports of
DC transfection in the literature.55 , 56, 232 Irvine et. al compared transfection efficiencies in
human monocyte-derived DCs using a cationic peptide-plasmid DNA complex with a
series of commercial nonviral agents at optimized conditions including lipofectin,
lipofectamine and DOTAP lipid where they demonstrated a superior average transfection
efficiency of 17% using a reporter GFP protein compared to 1-3% by commercial agents.
Aswathi et. al also reported similar transfection efficiency in JAWS II dendritic cells
(derived from the bone marrow of C57/BL6 mice) using a nonviral non-liposomal lipid
polymer, TransIT-TKO reagent. Strobel et. al reported transfection efficacy of up to 20%
when GFP RNA was delivered to monocyte-derived human DCs using commercial
liposomal formulations. In contrast to our experiments where transfection was tested in
the presence of serum, dendritic cells were incubated with nonviral transfection agents in
serum-free medium in each of these prior studies, limiting translation of the results to
conditions in vivo.
Particle uptake and GFP expression was similar for both PEGylated and nonPEGylated particles (Figure 5-4D, H). Notably, under these conditions providing 30%
transfection, the particles were non-toxic and >95% of the cells were negative for DAPI
and annexin V (data not shown).
68
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20 pm
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mRNA
GFP
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labeled mRNA (red) in serum-free Opti-mem (A) or Cy3-labeled mRNA loaded on
particles in complete medium (10% FBS) (B). To study transfection, DC2.4 cells were
incubated for 12 h in complete medium (10% FBS) with either naked mRNA (1 ptg/ml)
or an equivalent dose of mRNA adsorbed to PEGylated or non-PEGylated lipid-coated
PBAE particles (75 pig particles/mL), then washed and imaged live by confocal
microscopy at 37 *C or analyzed by flow cytometry. (C) Confocal images of DiD-labeled
particles (blue), Cy3-labeled mRNA (red), and GFP expression (green) with
fluorescence/bright-field overlays in DC2.4 cells. (D-H) Flow cytometry analysis of
DC2.4 cells showing mean particle uptake (D), representative histograms of fluorescent
mRNA uptake (E) and mean frequency of mRNA* cells (F), representative cytometry
histograms of GFP expression (G), and mean frequency of GFP* cells (H). Shown are
means +SD from duplicates in 2 independent experiments.
5.3.4.
Synergistic tumor cell killing with lipid-coated PBAE
nanoparticles
in vitro
To achieve selective killing of tumor cells, immunotoxins should be allowed to
first seek out and bind to intended target cells, before internalization and subsequent
delivery into the cytosol in order for the toxin to exert its cytotoxic effect. As such, we
envisage a scenario where the particles act as a chaperone to facilitate uptake and
70
cytoplasmic access of soluble toxin molecules co-administered with particles to target
cells.(Figure 5-5) The toxin and particles should be taken up in the same endosomes so
that the particles can subsequently lyse the vesicle and release the toxin into cytosol.
Therefore, we wanted to first evaluate the ability of lipid-coated PBAE particles to
enhance the uptake of co-administered free cytotoxin (compared to mRNA which was
explicitly bound to the particles in a separate step prior to delivery) into the cytosols of
tumor cells. We have in a way demonstrated the feasibility of such a strategy in section
5.3.1 where we made use of the cytosolic/nuclear distribution of calcein observed in a
tumor cell line, B16F10 co-treated with particles as evidence for endosomal rupture by
lipid-coated PBAE particles and now seek to confirm this with a model toxin, phalloidin
using a similar setup.
Tumor cell
Particles
+p
Immunotoxin
Figure 5-5. Schematic illustration of the events taking place following the codelivery of immunotoxin and particles.
Phalloidin (788 Da) is a cytotoxin isolated from the Death Cap mushroom
Amanita phalloides. It is a polar, cell-impermeable, cyclic heptapeptide and binds tightly
to actin filaments, preventing its depolymerization thereby poisoning the cell.24 s247 When
a sufficient amount of phalloidin is microinjected into the cytoplasm, cell proliferation is
inhibited. 248 The actin-binding properties of phalloidin make it a useful tool for
investigating the cytoskeletal organization in cells by labeling phalloidin with fluorescent
analogs and using them to stain actin filaments for light microscopy. As such,
commercial fluorescent conjugates are readily available that would allow us to evaluate
the potential of lipid-coated PBAE particles for enhancing uptake and cytosolic delivery
of soluble membrane-impermeable toxins, as well as augmenting tumor cell killing, if
phalloidin is sufficiently toxic to tumor cells when delivered into cytosol as reported
recently: 2 49 In this study, a pH-responsive cell-penetrating peptide was shown to facilitate
translocation of phalloidin into the cytosol and inhibit the proliferation of cancer cells in a
pH-dependent fashion.249
As we would expect from previous experiments with calcein delivery, we were
able to observe increased cytosolic uptake of soluble phalloidin in B 16F 10 cells that were
co-administered with particles compared to cells given phalloidin only (Figure 5-6).
Green fluorescence was observed predominantly in the cell periphery, corresponding to
71
the cortical actin cytoskeleton contained within the cell's cytoplasm, validating that
particle-chaperoned phalloidin was able to escape from endosomes and enter the cytosol
to stain actin filaments.
A
Figure 5-6. Cytosolic delivery of soluble phalloidin by lipid-coated PBAE particles.
Confocal images of B16F10 cells incubated with (A) 10 RM or (B) 25 gM phalloidin
alone (left) or phalloidin and PEGylated lipid-coated PBAE particles (right). (red,
nanoparticles; green, phalloidin alexa 488 conjugate). Cells co-treated with phalloidin
and particles displayed staining of cytoskeletal structures detected at the cell periphery,
consistent with the toxin accessing intracellular targets following particle co-delivery.
We also attempted to measure the cytotoxicity in cells treated with phalloidin to
see if the improvement in intracellular delivery afforded by the chaperoning particles can
be translated to an enhancement in the potency of the membrane-impermeable toxin.
However, we were unable to detect a significant improvement in killing efficacy relative
to cells treated with particles alone (data not shown). We suspect this may be in part due
to the relative modest potency of phalloidin as a cytotoxic therapeutic. We expect that the
use of a more potent toxin candidate such as an immunotoxin designed specifically for
targeted cancer therapy will be better suited for demonstrating such an enhancement in
therapeutic potency concomitant with the enhancement in cytosolic access, which was
2 33
identified as a rate-limiting step in therapies employing immunotoxins previously.
Such a study was conducted with E4rGel (-40 kDa), an immunotoxin based on
gelonin, targeting the epidermal growth factor receptor (EGFR). Gelonin (-30 kDa) is a
plant toxin and classified as a type I ribosome-inactivating protein because it lacks any
cell-binding or cytoplasmic delivery domains. The use of recombinant gelonin in tumor72
targeted cytotoxic agents has been well studied.2 so, 251 The targeting moiety is based on
the tenth human fibronectin type III domain (Fn3) which has been designed using various
directed evolution approaches for specific affinity toward numerous different targets.2254 The resultant immunotoxin with engineered fibronectin fragments binding EGFR has
an IC 50 of ~ 30 nM when incubated with antigen-expressing A431 cells for 72 h.m Note
that this is a significantly lower concentration relative to the pM concentration range
required by phalloidin to achieve the antiproliferative effect in tumor cell lines reported
recently 249 , reflecting the higher potency of targeted gelonin as a cytotoxic agent. We see
that treatment with E4rGel alone at concentrations less than 30 nM resulted in low levels
of cytotoxicity especially at short treatment times. In contrast, co-delivery with particles
enhanced the potency of the toxin treatment even at 10-fold lower concentrations, and
also with a reduced treatment time of 6 h (Figure 5-7 top). This is in agreement with the
inefficiency in endosomal escape, limiting the ability of immunotoxin to reach
intracellular targets, reported previously.2 3 3 The therapeutic benefit can be elucidated
more clearly by using the cytotoxicity data to calculate a combination index to determine
if the enhancement in killing associated with particle co-delivery was synergistic in
nature. We thus confirmed that co-delivery of low concentrations of E4rGel and particles
resulted in a modest synergistic effect for 6 h treatment which can be increased severalfold at higher toxin and particle concentrations, and up to 10 to 60-fold for 24 h treatment
(Figure 5-7 bottom). In an ideal setting, for targeted therapy, immunotoxin would bind
only to target cells and upon subsequent co-internalization with particles, then be released
into cytosol following particle-mediated endosomal lysis. In future work, it will be of
interest to demonstrate that cell killing remains tumor cell specific and that particle codelivery is not mediating intracellular toxin delivery indiscriminately in all cells that may
result in off-target toxicity in vivo.
73
6 h treatme nt
24 h treatmentd
0 ugimI NP
*125 ug/mI NP
2ugmI NP
T0
37.2
G/mI NP
00 ugmI NP
112.5 ugimI NP
26ug/mI NP
ugimi W
T0
4
j37.5
E4rGel
E4rGel
6 h treatment
24 h tratient
112.6 ugVmI NP
S26ug/mI NP
37.5 ugimI NP
*12.5 ugmi NP
126 uWmi NP
03.6 ugjmI NP
Figure 5-7. Killing of A431 target cells by particle-chaperoned immunotoxin.
A431 cells were incubated with various concentrations of immunotoxin (3, 10, 30 nM
E4rGel) and/or PEGylated lipid-coated PBAE particles (12.5, 25, 37.5 pg/mL NP) for 6
or 24 h. Viability was measured via a WST-1 assay 72 h following treatment and used to
compute the combination index where a value > I indicate synergistic tumor cell killing
in combination treatment (particles + E4rGel) groups.
5.4. Conclusions
In this chapter, we studied the application of lipid-enveloped PBAE particles to
facilitate cytosolic delivery of therapeutics for immunization and immunomodulation/therapy purposes. Building on the pH-responsivity of the PBAE-based
particles first introduced in chapter 3, we demonstrated conclusively here in vitro that the
particles were indeed capable of lysing endosomes in target cells via mechanisms
hypothesized earlier (Figure 3-5). The particles were synthesized with a pH-responsive
core and hydrophilic charged shell designed to decouple the endosomal escape and drug
binding functionalities respectively, thereby affording greater. control over the
formulation of the particles being delivered. Moreover, by sequestering the cationic
polymeric component within a more hydrophilic pH-insensitive shell, effective delivery
of small molecules and mRNA to the cytosol was achieved in dendritic cells with
minimal cytotoxicity, a key cell type of interest in the context of vaccination, culminating
indin vitro transfection of a DC clone at levels comparable to the best reports for DC
transfection from the literature. The chapter closes with promising preliminary data on
the synergistic killing of tumor cells in vitro that can be achieved by using particles as a
chaperone to enhance intracellular delivery of antitumor toxins. In the next chapter, we
shall extend the in vitro results obtained with mRNA in this chapter to a mouse model in
74
vivo and examine the non-invasive mucosal delivery of mRNA as a novel vaccine
antigen.
75
6. Mucosal delivery of mRNA by lipid-enveloped polymer
nanoparticles in vivo
6.1. Introduction
Gene-based vaccine technology first emerged from observations that
intramuscularly injected naked DNA resulted in transfection of cells in vivo and
expression of the encoded protein, leading to the induction of immune responses against
that protein.222 Subsequently, DNA vaccines encoding antigens from a variety of viral,
bacterial and parasitic pathogens have been found to be effective at eliciting potent
systemic immune responses and protection from infection in preclinical models. 255
The favored route of administration of naked plasmid DNA, intramuscular
injection, generates systemic antibody and T-cell responses255 but has been reported not
to enhance mucosal antibody production or to protect against entero-invasive pathogens
at the mucosal surface. 256 Therefore, considerable interest has been generated in targeting
gene-based vaccines to mucosal tissue sites such as the nasal passages and airways to
enhance the protective immune responses at mucosal surfaces. Moreover, vaccine
administered at one mucosal site has been shown to elicit an immune response in other
mucosal sites mediated by trafficking of effector immune cells between mucosal tissue
compartments. For example, it has been reported that mucosal antibody responses were
detected in lung and vaginal washes following intranasal instillation with a DNA vaccine
encoding HIV antigens. This is of particular interest for vaccination against HIV where
a mucosal immune response at the vaginal tissue may be critical for protection. Although
naked DNA can induce immune responses at mucosal surfaces, the immunogenicity does
not appear to equal that observed by the intramuscular route. Consequently, novel
formulations have been used to enhance the efficacy of DNA vaccines delivered by
mucosal routes.
A variety of systems, including biodegradable microparticles, liposomes and
bioadhesive polymers that are effective for mucosal delivery of protein vaccines, have
been successfully translated to enhance the immunogenicity of mucosally-delivered DNA
vaccines.2ss Formulation of DNA into liposomes,2s8 lipid complexes 259, 260 or
macroaggregated albumin conjugates261 is efficient at targeting plasmid DNA to mucosal
surfaces, especially the respiratory tract. Intranasal immunization of a DNA vaccine in
liposomes with plasmids encoding IL-12 and granulocyte-macrophage colonystimulating factor induces strong mucosal and cell-mediated responses against HIV-1
antigens.2 This approach is also being used to treat Thl-mediated inflammatory
diseases: mucosal administration of plasmid DNA encoding IL-10 inhibits Th 1 cytokines
in mice infected with the herpes simplex virus. 262
The strategies above using DNA as antigen or immunodulatory agent are likely to
hold true when mRNA is used as an alternative source of genetic material. Since the first
demonstrations, DNA-based vaccines have undergone extensive preclinical and clinical
testing, while mRNA-based therapies remained less thoroughly investigated. However,
DNA vaccines have failed to show potency in large-animal models and humans, in
76
discord from results in small-animal studies, 263 which may reflect in part the difficulty of
overcoming not only the barrier posed by the plasma membrane of cells but also the need
to transport DNA through the nuclear membrane of non-dividing cells. Vaccines based
on mRNA, by contrast, require nucleic acid delivery only to the cytosol, thereby allowing
the transfection of quiescent and post-mitotic cells that comprise the majority of target
cells in vivo.264, 265 To date, strategies reported for non-viral delivery of mRNA for
vaccines and gene therapy applications include injection of naked mRNA,2, n4
polyplexes, 266, 7 lipoplexes or liposome-entrapped mRNA,
27 3 via
272 268, 269 bolistic delivery
, 274
270
gene gun,
,
271
particulate carrier-mediated delivery
and electroporation.
Recently, clinical trials in which human patients were vaccinated with naked or
protamine-complexed mRNA against tumor antigens via intradermal injections were
completed, demonstrating feasibility, lack of toxicity and promising responses based on
clinical and immunological read-outs. 275' 276 However, more efficient transfection in vivo,
and the ability to deliver these nucleic acids non-invasively and/or to mucosal sites,
would be expected to enhance the prospects of this strategy for vaccination.
In this chapter, we build upon the data described in the previous chapter where
lipid-coated PBAE particles were successful in mediating cytosolic delivery of intact
mRNA leading to the expression of a reporter protein antigen in vitro and moved forward
with testing the same system in mice in vivo. In addition, to demonstrate the usefulness of
particles for facilitating non-invasive mucosal delivery, we chose to carry out this study
by administering the particles via both the intranasal and intratracheal routes, targeting
the mucosal tissue in the nasal passages and respiratory tract respectively, and were thus
able to confirm the general applicability of our system in vivo.
6.2. Materials and methods
6.2.1.
Materials
Reagents involved in the synthesis of lipid-coated PBAE nanoparticles and the
preparation of mRNA are the same as those described in Chapter 3.
2-2-2 Tribromoethanol (Avertin) and 2-methyl-2-butanol was purchased from
Sigma Chemical Co. (St. Louis, MO). D-Luciferin (potassium salt) was from Caliper Life
Sciences (Hopkinton, MA). Exel Safelet IV catheters were from Fisher Scientific
(Pittsburgh, PA).
6.2.2.
Animal studies
Female C57Bl/6 mice 6-8 weeks of age were purchased from Jackson
Laboratories and cared for in the USDA-inspected MIT Animal Facility under federal,
state, local and NIH guidelines for animal care. For both intranasal and intratracheal
delivery, mice were anesthetized by intra-peritoneal (i.p.) injection of 0.4 mg/g body
weight avertin prior to administration.
For intranasal administration, the anesthetized mouse was held gently ventral side
up, in a supine position in one hand, tilted at a slight angle so that the head is positioned
above its feet. The other hand is used to dispense the particle suspension dropwise using a
micropipette by placing the pipette tip gently over the opening of one nostril and
77
expelling the suspension slowly until the droplets are completely inhaled. 5 pL of particle
suspension was administered into each nostril alternately with a 5 minute interval
between each aspiration during which the mice were kept in a supine position. Mice were
administered 4 pg of luciferase-encoding mRNA adsorbed to 150 pg particles in 20 pl
serum-free RPMI medium this way.
For intratracheal administration, we followed a procedure using an intubation
platform and a fiber optic light source reported previously. 277 (See appendix C for
detailed protocol and pictures illustrating the procedure.) The anesthetized mouse is
placed on the platform so that it is hanging from its top front teeth on a bar and the light
source is then directed at the mouse's chest. A catheter (Exel Safelet IV catheter) with the
needle withdrawn was used to deliver the formulation into the trachea. The mouth of the
mouse is pushed open using the catheter and its tongue gently pulled out with a flat
forceps. The opening of the trachea was located via the white light emitted from the
trachea and the catheter inserted vertically into the trachea. The needle was then removed
and the formulation pipetted directly into the opening of the catheter. The mouse was
then allowed to completely inhale the formulation before removing the catheter. 4 ttg of
luciferase-encoding mRNA adsorbed to 150 ptg particles in 75 pL serum-free RPMI
medium per mice was administered this way.
After administration was completed, mice were kept warm by placing them either
on heat pads or on a latex glove filled with warm water to recover from the anesthesia.
At various time points following administration, particle fluorescence and
luciferase expression were tracked via bioluminescence imaging following injection of
300 pl of 15 mg/mL luciferin i.p. in anesthetized mice (Xenogen IVIS Spectrum Imager).
For post-mortem bioluminescence imaging of transfected tissues, mice were
euthanized 5 minutes after i.p luciferin injection and dissected immediately to expose the
tissues of interest. The exposed tissues, together with the entire mouse carcass were then
imaged as before, within 15 minutes of death.
6.2.3.
Statistical analysis
One-way ANOVA followed by Bonferroni's Multiple Comparison Test was
applied to determine the significance of the difference between the fluorescence and
bioluminescence radiance detected in mice of different treatment groups.
6.3. Results and Discussion
6.3.1.
In vivo transfection with lipid-enveloped nanoparticles
administered intranasally
We tested lipid-coated PBAE particles loaded with mRNA encoding for luciferase
protein in vivo, since many systems that transfect cells in vitro fail to function or have
serious toxicity in vivo.
Intranasal delivery is a convenient and effective route of mucosal vaccination,
where the development of antigen particulate carrier systems is of considerable interest. 73'
75, 77, 278 For vaccine applications, administration via mucosal surfaces is
an attractive
route for inducing a protective immune response since most pathogens invade the body
78
through mucosal surfaces. Mucosal immunization elicits both systemic and mucosal
immunity279 , 280, while the latter is usually not achieved via parenteral administration. It
has also been established that mucosal vaccines administered at one site can elicit an
immune response in mucosal tissues remote from the site of initial antigen exposure,
mediated by trafficking of effector immune cells between mucosal tissue
compartments.281, 282 Although parenteral injections of naked mRNA in vivo have been
reported in both preclinical models and early clinical trials to elicit protein expression and
immune responses 221,224,27s,276,283, it was not clear if naked mRNA administration would
be effective when administered via a non-invasive route such as intranasal delivery,
where the mRNA may need to traverse both the epithelial layer and the overlying mucus
barrier layer to access target cells. In nasal inoculation, particulate antigens can be taken
up by M-cells of nasal-associated lymphoid tissue (NALT), where they are processed and
preferentially directed to the antigen-presenting cells, in contrast to soluble antigen. 72-76 it
has been demonstrated that particulate delivery systems, particularly those bearing
cationic charges, can enhance muco-adhesiveness, reducing the clearance rate from the
the contact time of the delivery system with the nasal
nasal cavity, and thereby increasing
73 5
mucosa, facilitating transport. ,7 , 7
We chose to test PEGylated particles since the grafting of polymers such as PEG
to nanoparticles has been explored as a way of blocking the adsorption of particles to
components of the mucus. It has been observed that even very large particles, 200-500
nm in diameter, can penetrate mucus when appropriately grafted with surface PEG
chains.284 Moreover, we have shown in the previous chapters that the presence of PEG in
the lipid shell did not interfere with mRNA loading and the later processes of cellular
uptake and endosomal destabilization. In light of this, we decided to incorporate PEGcontaining lipids in the hope that it may help enhance particle penetration. For fluorescent
imaging, the particles were labeled with Rhodamine-DOPE lipid (1 mol%) as opposed to
DiD which was used previously in vitro, as we found the former to be a more faithful
tracer for particles in vivo relative to DiD which has been observed to dissociate from
particles and label tissues/cells in vivo.
Luciferase-encoding mRNA (4 ptg), in soluble form or adsorbed to 150 [ig labeled
NPs, was administered intranasally to anesthetized C57BL/6J mice, and the presence of
particles and luciferase expression were tracked longitudinally via fluorescence and
bioluminescence imaging in live animals (Figure 6-1). Immediately following intranasal
administration, localized particle fluorescence was detected in the nasal region, but
fluorescence above background could not be detected by 6 h (Figure 6-1A, B), suggesting
that the particles are rapidly cleared from the nasal passage. However, at 6 h after
administration, significant bioluminescence (p <0.01 vs. untreated animals) could be
detected at the inoculation site, demonstrating successful transfection by mRNA-loaded
particles, while mice treated with naked mRNA showed no signal above background
(Figure 6-1D). Imaging at later time points revealed that the luciferase activity peaked at
12 h before declining to background levels 24 h after administration, consistent with the
expected short half-life of transfected mRNA in vivo. 22 1, 283 At 12 h, although
bioluminescence was detected in some mice receiving naked mRNA, expression was
observed in all mice treated with particles (statistically significant above naked mRNA
treatment group, p <0.05), demonstrating that lipid-enveloped poly-1 NP delivery was
more effective and consistent in eliciting protein expression in vivo (Figure 6-iC, D).
79
Significant variations in levels of transfection were consistently observed for naked
mRNA-treated mice in multiple independent experiments, reflected in the increased
variance in bioluminescence signal detected at the peak of protein expression at 12 h
(st.dev.naked = 2500 vs. st.dev-NP = 1200). Generally, 1-2 mice/group always failed to
respond to naked mRNA administration, possibly reflecting a fundamental limitation in
the consistency of transfection efficiency of naked mRNA by this route, compared to
nanoparticle administration where all mice were transfected.
Though more detailed safety analyses are a topic for future study, PBAE
nanoparticles did not elicit overt signs of toxicity in mice: no signs of granuloma
formation or chronic inflammation were observed at i.n. or parenteral administration
sites, nor were changes in the body mass or behavior of mice detected up to 3 months
following particle administration (data not shown).
80
A
Control
Lipid-coated PBAE NP
1.5x108 -
,,
1.OxI0S -
8
fluorescnce
(pleclcmIsr)
Blank
Lipid-coated PBAE NP
0
0
BY
5x10*
0
0
5.0x107.
-. +-
-'4-
-tO-
na
Time after administration
C
Control
Firefly luciferase-encoding mRNA
1x10'
Lipid-coated
PBAE NP
Naked
6x10*
doluminescence
(plseclcmIsr)
D
r
ir--,*
*
0
0
6000.
e*
4000-
@
0
e0
2000-
g,
e@0
Lipid-coated PBAE NP
naked
0
8
.2
0
Blank
1-1
0
0%e,
0
Time after administration
81
0
.0
@0
Figure 6-1. In vivo transfection with intranasally administered lipid-coated PBAE
particles.
In vivo transfection in C5BL/6J mice following intranasal administration of firefly
luciferase-encoding mRNA either in soluble form or adsorbed on fluorescent lipidenveloped PBAE particles. (A) Fluorescence image of mice immediately following
intranasal particle administration. (B) Total fluorescence signal at the inoculation site at 0
and 6 h quantified from groups of particle-treated or control mice. (C) Bioluminescence
images of mice 12 h following mRNA administration. (D) Bioluminescence signal at the
inoculation site at 6, 12 and 24 h quantified from groups of treated and control mice. ***
,p <0.001, **, p <0.01 and *, p <0.05. Data shown from one representative of two
independent experiments (n = 4 mice/group).
6.3.2.
In vivo transfection with lipid-enveloped nanoparticles
administered intratracheally
As additional proof that lipid-coated PBAE particles can indeed play a role in
protecting and delivering intact mRNA across mucosal tissue barrier, resulting in the
expression of protein antigen in vivo, we examined an intratracheal administration route
for pulmonary immunization, a model of vaccine delivery in humans via inhalation, in
parallel with intranasal delivery described above.
Pulmonary immunization is a promising way to administer vaccines since the
lungs contain a highly responsive immune system where professional antigen-presenting
cells (APCs) such as pulmonary macrophages and dendritic cells (DCs) play important
roles in both innate and adapted immunity. Alveolar macrophages are abundant in the
periphery and interstitium of the lungs. 285 DCs are found in epithelial linings of the
conducting airways, submucosa below the airway epithelium, within alveolar septal walls
and on the alveolar surfaces. 286 These APCs are able to phagocytose, process and present
antigens to stimulate T cells. Primary stimulation of T-cell clones within the pulmonary
lymphoid tissue is induced when macrophages and DCs migrate to bronchial lymph
nodes and home to the T-cell paracortical area.287, 288 In mammals, bronchus-associated
lymphoid tissue (BALT) - the respiratory equivalent of NALT - is located mostly at
bifurcations of the bronchus. There is evidence that mice genetically lacking spleen,
lymph nodes and Peyer's patches can generate strong primary B- and T-cell responses to
inhaled influenza. These responses appear to be initiated at sites of the induced BALT,
which functions as an inducible secondary lymphoid tissue for respiratory immune
responses.289 Exposure of the lungs to various aerosol formulations designed to protect
against influenza virus was more effective than either intranasal administration or
parenteral injections, indicating that a local response was generated in the respiratory
tract.290 The abundance of APCs as well as the physiological features of lymphoid tissues
in the respiratory tract makes the lung mucosa an attractive target for exploring mucosal
vaccine delivery with our particles.
From Figure 6-2, we see that transfected cells were readily detected in the
pulmonary region for mRNA delivered on particles consistently while little or no
transfection is seen for naked mRNA, as expected with a non-parenteral administration
route where mRNA is likely degraded before it is able to access any cells. We were thus
able to reproduce the trend we observed when administering mRNA intranasally and
82
demonstrate the general applicability of using particles to enhance mucosal delivery at
multiple sites. Note that we were unable to detect fluorescence due to particles when
imaging live mice, which we attribute to greater attenuation of fluorescent signal since
particles were likely delivered to the lungs which are well-enclosed by the ribcage and
body tissue.
A
Firefly luciferase-encoding mRNA
2.5x10'
Lipid-coated
PBAE NP
Naked
1.5x10*
bloluminescence
(p/seclcm2lsr)
B
e
0 15000'
Blank
* Lipid-coated PBAE NP
" naked
E
0
10000'i
C
5000,
iE
M
U
4'
0
-0
0
*
'" '
o
S
0e
p
e0
**0
e@,0
a
Time after administration
Figure 6-2. In vivo transfection with intratracheally administered lipid-coated PBAE
particles.
In vivo transfection in C5BL/6J mice following intratracheal administration of firefly
luciferase-encoding mRNA either in soluble form or adsorbed on fluorescent lipid-
enveloped PBAE particles. (A) Bioluminescence images of mice 6 h following mRNA
administration. (B) Bioluminescence signal at the inoculation site at 6, 12 and 24 h
quantified from groups of treated and control mice. *, p <0.05. Data shown from one
representative of two independent experiments (n = 4 mice/group).
83
To more clearly elucidate the anatomical location where transfection with
particles took place, we also imaged treated mice dissected post-mortem to isolate and
reveal tissues of interest. This can also help us obtain a more sensitive readout of the
fluorescence and bioluminescence signal since particle uptake and antigen expression is
likely to occur in deeper tissues following intratracheal instillation relative to intranasal
administration, and any signal will therefore be attenuated by the intervening layers of
body mass it has to traverse through for detection. In doing so, we were able to show that
a majority of transfected cells that also take up fluorescent particles were present in the
lungs (Figure 6-3), a potential APC-rich site for delivering antigen to APCs that can then
go on and prime immune responses in the draining lymph nodes. We were also able to
resolve a more dramatic difference between the transfection levels measured in the lungs
of mice given particle-adsorbed mRNA relative to naked mRNA (Figure 6-3), likely a
result of better protection of mRNA against enzymatic degradation provided by the
particles. When the isolated lung tissue of MHC-Class II-GFP mice treated with
fluorescently labeled particles carrying mRNA was examined via confocal microscopy, a
fraction of the particles were found to colocalize with MHC-Class 11-expressing cells that
may represent target APCs such as DCs and macrophages (Figure 6-4).
Lipid-coated PBAE NP
Naked
1x10'
Trachae
Lungs
b
1x103
umincnce
(Pls"ccm2lsr)
Trachae
Lungs
3x10'
fluorescence
(pIecIcmIdsr)
Figure 6-3. Transfected tissues in dissected mice imaged post-mortem.
Representative fluorescent and bioluminescence images of dissected mice 6 h following
mRNA administration. Particle fluorescence and luciferase protein expression was
detected mostly in the lungs of mice only in particle-treated group. (Note the unlabeled
bright circular spots are due to autofluorescence and autoluminescence from the pins used
to hold tissues in place. Also autofluorescene is detected in the gut region due to the
presence of alfalfa in mouse feed and should be disregarded.)
84
Figure 6-4. Particle uptake by MHC-Class II-GFP-expressing cells in lungs.
Representative confocal image of lung tissue in MHC-Class II-GFP mice administered
fluorescently labeled particles showing particle (red) uptake (indicated by arrows) by a
subset of MHCII-expressing cells (green), likely to be APCs.
6.4. Conclusions
Altogether, the data presented in this chapter show that mucosal delivery of intact
functional mRNA is achieved by this lipid-enveloped PBAE nanoparticle delivery system
in vivo with minimal toxicity, via the electrostatic adsorption of cargo molecules to the
particle surfaces. mRNA bound to these pH-sensitive particles was sufficiently protected
from nucleases in the mucus barrier layer and facilitated in vivo transfection of potential
target cells relative to naked mRNA. This approach provides a simple strategy to
enhance mRNA delivery of interest for mRNA vaccine design, and may be useful for
other instances where local gene delivery may be of therapeutic value.
85
7. Conclusions and future work
7.1. Degradable PBAE-based nano-films and particles as a platform for
delivering vaccines and therapeutics
We have designed and characterized two different systems for the delivery of
vaccines and therapeutics that utilizes a biodegradable and pH-responsive polymer,
poly(#8-amino ester) (PBAE). Using these materials as delivery platforms, we were able to
demonstrate ways in which the delivery of antigens, adjuvants and other immune- as well
as non-immune-based therapeutics can be improved to both subcellular as well as tissue
targets. Although similar approaches to drug delivery have been reported 84 , 90, 92-94, 99-101,
103, 104, 121, 123, 127, 162, the applications presented in
this thesis is unique since to our
knowledge, similar technologies have not been tested for non-invasive delivery of
biologics targeting the immune system when we first embarked on this work.
We were initially motivated by the prospects of designing delivery systems
capable of facilitating transcutaneous delivery to better target the APCs in the skin for
triggering protective immune responses. Using protein- and oligonucleotide-loaded layerby-layer (LbL)-assembled multiplayer films incorporating PBAE, we were able to
demonstrate the facile and versatile encapsulation of biologics into conformal thin films,
coupled with robust control over materials release and nanometer-scale control over film
structure and composition. Applying these films in vivo to tackle the challenges
associated with transcutaneous immunization, we demonstrated effective delivery of a
model protein, ovalbumin and synthetic adjuvant, CpG oligonucleotide from LbL films
into barrier-disrupted skin, uptake of the protein by skin-resident APCs (Langerhans
cells), and transport of the antigen to the skin-draining lymph nodes. To our knowledge,
this is the first demonstration of LbL films applied towards the delivery of biomolecules
into skin.
We then zoomed in to examine at a cellular level, how materials can be used to
mediate endolysosomal escape and achieve cytosolic delivery of biologics. To do so, we
explored the possibility of taking advantage of the pH-sensitivity of PBAE as a delivery
system in a reducing pH environment. By double emulsion and nanoprecipitation
methods, we successfully synthesized fully degradable, pH-sensitive PBAE-core lipidshell nanoparticles whose size was around 200 nm. We demonstrated that these core-shell
nanoparticles were capable of efficient cytosolic delivery of membrane-impermeable
molecules (such as calcein, mRNA, antitumor toxins) to dendritic cells and tumor cells in
vitro via a 'proton sponge effect' or dissolution-induced osmotic pressure. These
preliminary applications indicated that the nanoparticles may be of utility for delivery of
membrane-impermeable vaccine antigens and therapeutics to the cytosol of antigenpresenting cells and tumor cells for immunization and tumor therapy respectively. These
nanoparticles were then applied towards in vivo delivery of mRNA where we showed
significant improvement in the expression of a reporter protein by protecting and
86
delivering a functional biomolecule across distinct mucosal tissues, a promising result for
the use of mRNA as a novel gene-based vaccine antigen or for local gene therapy.
Design of both systems was guided foremost by the goal of biocompatibility.
Presently, all the components of both the PBAE-based films and particulate delivery
platforms are biodegradable and did not show acute toxicity in vitro or in vivo. Other
functional and practical advantages are also incorporated as part of the materials design
to aid future clinical translation.
For the LbL films, the fully aqueous-based processing during film assembly and
solid-state stabilization of incorporated molecules upon drying the films can help protect
environmentally-sensitive biologics and preserve their function. The ability to keep
therapeutics in a dry state and avoid the "cold chain" associated with its storage and
distribution will likely have a significant impact on the cost of delivery, especially
important when considering delivery in less developed countries.
For the core-shell nanoparticles, the synthesis of nanoparticles, particularly the
process of nanoprecipitation that avoids the use of harsh organic solvents and sonication
steps, can have applications in delivery of fragile biologics. Moreover, the strategy of
loading molecules onto particle surface post-synthesis can further ensure that biologics
such as proteins and RNA may be delivered without risking loss of activity frequently
associated with encapsulation within particles.
The pH-sensitivity of polymers for endosome escape has been widely used to
design intracellular drug delivery system. However, distinct from many such systems
typically formed by physical complexation of the polymer and therapeutic cargo, our
strategy is to use a pre-formulated nanoparticle of known size and structure to assemble
formulations with well-defined properties and improved stability, critical for in vivo
translation. We further introduced a core-shell structure to the particle design. By
sequestering the cationic, pH-buffering polymeric component within the core surrounded
by a biocompatible lipid shell, these nanoparticles effectively disrupted endolysosomes
and delivered molecules to the cytosol of cells without overt cytotoxicity, which is often
observed with polymeric carriers that disrupt endolysosomes. This design exemplifies an
approach to lower cytotoxicity by shielding toxic components from direct contact with
cells. In addition, the segregation of cargo/cell binding function of the shell from the pHresponsive and endolysosomal escape function of the core provides flexibility to tune the
surface chemistry to bind molecules of interest and to introduce targeting ligands.
Finally, facilitating alternative non-invasive delivery routes, a feature
demonstrated with both systems for the case of transcutaneous and mucosal delivery, will
help transition current needle-based vaccine and therapeutic administration towards
needle-free delivery. This will help improve safety, patient compliance and ease
requirements on medical expertise, thereby promoting wider access to vaccines and
(immuno)therapy in the effort towards better prevention and treatment of various
diseases.
7.2. Issues for future work
Based on our studies thus far, we conclude that the PBAE-based nano-films and
particles proposed here are highly applicable and show promise as delivery platforms for
a range of biological applications harnessing the immune system as well as beyond the
realm of immunology. However, these delivery platforms may be further improved for
87
better efficacy by making the system more generic and versatile, as well as by
incorporating modalities to explicitly overcome tissue barriers and achieve direct cell
targeting.
The LbL films used in the current studies were assembled on glass or elastomeric
substrates and rely on a physical barrier disrupting method, tape-stripping, to bypass the
barrier imposed by the stratum corneum. However, there exist at least 2 ways in which a
skin-penetrating modality can be assimilated with LbL-based technology to create a selfcontained system capable of overcoming the SC barrier. One way would be to include
penetration enhancers as a component in a multi-component LbL film to perturb the SC
and enhance penetration. A myriad of molecules has been explored for facilitating skin
penetration and combinations of several such moieties has been shown to be particularly
effective. 175, 291, 292 This can play well into the multi-component nature of LbL films
where multiple agents can be incorporated and release either simultaneously or with
distinct kinetics to achieve a greater synergistic effect. Alternatively, the ability of
multilayers to conformally coat substrates should allow films to be assembled on
microneedles where an array of micron-sized needles (short enough to avoid the nerve
endings thereby avoiding pain) is used to pierce through the superficial skin layer to
permit access into the viable tissue underneath. Recently, such a strategy has been
pursued by multiple groups and is the subject of on-going studies in our lab.293 294In
addition, to make this platform truly generic for delivering a wide range of therapeutics,
biodegradable nano-carriers may be embedded as delivery vehicles within LbL films, to
segregate film assembly/release properties from the characteristics of individual
molecules. The examination of such a strategy has begun and is a topic of ongoing work
in our lab.293 We are hopeful that a combination of both aspects will translate the existing
system into one of great utility as a practical delivery system for transcutaneous delivery
applications.
For the application of lipid-enveloped PBAE nanoparticles as vaccine carriers for
immunization, based on the data presented in chapter 6, we see that particles are effective
at mediating delivery of mRNA antigen to multiple APC-rich mucosal sites. In the case
of intratracheal administration where it is relatively straightforward to isolate and image
the transfected lung tissue, we were able to verify the uptake of particles by a subset of
MHC-Class-II expressing APCs. However the same image also show particle uptake by
taken by non MHC-Class-II-expressing non-APCs (Figure 6-4). Moreover, in the lungs,
the alveolar macrophages greatly outnumber DCs and there are studies sugesting that
antigens delivered to macrophages drive tolerance rather than immunity. 295 9 Dendritic
cells are the only cells capable of priming naive T cells and are generally considered to be
superior to other APCs in antigen processing and presentation. As such, we expect that
vaccines delivery in these tissues can be made more efficient by directly targeting DCs of
interest for inducing a stronger immune response. This is likely to play a critical role in
eliciting effective immune response with our system since a preliminary intraperitoneal
immunization study carried out with the existing particles failed to generate a satisfactory
immune response and live animal imaging experiments conducted in parallel indicate that
the anatomical location and cells being transfected is an important correlate for the
generation of a good immune response. (See appendix D)
In the case of mRNA-based antigen, one way to increase the chance of DC
encountering antigen is to design mRNA that is secreted so that mRNA delivered to other
88
cell types can eventually reach DCs even when they are not directly transfected or able to
phagocytose dying transfected cells. Preferential delivery of vaccine particles to DCs can
also be achieved by modifying particles with monoclonal antibodies directed against
several C-type lectin receptors expressed by DCs, including DEC-205 (CD205), the
mannose receptor (CD206), and DC-SIGN (CD209), which have been implicated as
promising targets due to their physiological roles in antigen uptake.300 The first
demonstration of using anti-DEC 205 to target vaccine particles to DCs resulted in -3fold increases in DC uptake of particles and -2-fold increases in the magnitude of T-cell
responses in vivo.301 Particle vaccines can also be targeted to CD206 via its natural
carbohydrate ligand, mannose. Surface-display of mannose increased particle uptake by
DCs both in vitro and in vivo and enhanced expression levels of co-stimulatory markers,
inducing antigen-specific CD4* and CD8' T-cell responses.302-305 Incorporating such
targeting ligands is achievable with our lipid shell structure which not only provides a
wide range of surface chemistry for conjugation but also offers a convenient means of
inserting lipid-conjugated ligands. Again, the core-shell structure makes this modification
very flexible, without affecting the endolysosomal escaping function. Selective and
enhanced antigen delivery to DCs is expected to enhance the immune responses to
antigen of interest compared to the passive delivery of antigen.
For facilitating cancer therapy, particles are shown to act as chaperone to enhance
uptake and cytosolic access of soluble membrane-impermeable toxins, resulting in the
synergistic killing of tumor cells. However for truly effective therapy, selective targeted
killing of tumor cells which spare healthy bystander cells is desired. This can be made
possible by delivering tumor specific immunotoxins that bind to tumor cells. Particles can
then be delivered to help promote cytosolic uptake of these bounds molecules and
facilitate killing. Such a strategy can be explored in future studies with the targeted
gelonin construct described in chapter 5. Furthermore, particles can also be explored for
enabling cytosolic delivery of immunostimulatory molecules such as poly I:C where it
has been shown in a number of studies in various tumor cell lines that cytosolic delivery
can stimulate greater tumor killing by triggering cytosolic RNA sensors. 44 ~49 Finally,
there may also be a role for particles as a cancer therapeutic themselves, stemming from
recent literature implicating lysosomal rupture and autophagy in inducing tumor cell
death.306 3- 08 The use of particles themselves as a therapeutic with a physical basis for its
cell killing mechanism is especially attractive for combating the development of chemical
drug resistance in tumor cells.
Altogether, the design of our existing platforms guided by relatively simple
rationales have produced delivery systems capable of facilitating delivery of various
vaccine antigens, adjuvants and therapeutics both in vitro and in vivo. Further progress in
the assembly, functionalities and applications of these biodegradable delivery platforms
will provide momentum to translate discoveries from basic immunology into novel
vaccines and therapies for numerous diseases, including cancer, infectious diseases, and
autoimmunity.
89
8. Biographical note
EDUCATION
2006-2012
Massachusetts Institute of Technology (Cambridge, Massachusetts,
USA)
Ph.D. Candidate, Materials Science and Engineering
Graduate fellowship:
- National Science Scholarship, A*STAR -- 2006-2012
2003 - 2004
Massachusetts Institute of Technology (Cambridge, Massachusetts,
USA)
Exchange Student under the Cambridge-MIT Institute
2001 -2005
University of Cambridge (Cambridge, UK)
B.A and M.Sci. in Materials Sciences and Metallurgy
PUBLICATIONS
Darrell J. Irvine, Xingfang Su and Brandon Kwong "Routes of
Delivery for Biological Drugs" In Biological Drug Products:
Development and Strategies, edited by Wei Wang and Manmohan
Singh, Wiley-Blackwell, 2012, in press
Xingfang Su, Jennifer Fricke, Daniel G. Kavanagh and Darrell J.
Irvine "In Vitro and in Vivo mRNA Delivery Using Lipid-Enveloped
pH-Responsive Polymer Nanoparticles" Molecular Pharmaceutics
8(3), 774-787, 2011
Peter C. DeMuth, Xingfang Su, Raymond E. Samuel, Paula T.
Hammond and Darrell J. Irvine, "Nano-Layered Microneedles for
Transcutaneous Delivery of Polymer Nanoparticles and Plasmid
DNA" Advanced Materials 22(43), 4851-4856, 2010
Xingfang Su, Byeong-Su Kim, Sara R. Kim, Paula T. Hammond
and Darrell J. Irvine, "Layer-by-layer Assembled Multilayer Films
for Transcutaneous Drug and Vaccine Delivery" ACS Nano 3,
3719-3729, 2009.
SELECTED
PRESENTATIONS
American Chemical Society 2010 (poster)
Controlled Release Society 2010 (oral)
Biomedical Engineering Society 2008 (oral)
90
9. Appendix A: Lipids comprising the lipid shell on PBAE
nanoparticles
Table 9-1. Lipid molecules and fluorescent lipid analog comprising the lipid shell of
PBAE nanoparticles
Full name
1,2-Dioleoyl-sn-Glycero-3PhosphocholineI
Abbreviation
DOPC
1,2-Dioleoyl-3trimethylammonium-propane
DOTAP
1,2-Distearoyl-sn-Glycero-3Phosphoethanolamine-N[methoxy(Polyethylene
DSPE-PEG
Structure
N
N+
0
Glycol)2000]
Rhodamine-labeled 1,2-Dioleoylsn-Glycero-3Phosphoethanolamine
RhodamineDOPE
1,1 '-dioctadecyl-3,3,3',3'tetramethylindodicarbocyanine
perchlorate 1
DiD
N
MC 3
H3
CH3
i
(CH=Cs-C,
I
CM3
C10
91
CH3
10.
Appendix B: Identification and dissection of auricular
lymph nodes
The auricular lymph nodes are the lymph nodes draining the ear region. The
identification and dissection of these lymph nodes is therefore important for studying
immune responses following transcutaneous delivery of vaccine antigens and
immunostimulatory agents in the ear skin. For initial identification and practice, mice
may be injected with a dye to ensure proper isolation of the appropriate node. This can be
done by injecting approximately 0.1 mL of 2% Evan's Blue Dye (prepared in sterile
PBS) intradermally into the pinnae of the mouse ear. Euthanize the mouse after several
minutes and continue with the dissection procedure below.
The most commonly used dissection approach is from the ventral surface of the
mouse. This approach allows both right and left draining nodes to be obtained without
repositioning the mouse. With the mouse ventrally exposed, the neck and abdomen area
is wetted with 70% ethanol. Using scissors and forceps, carefully make the first incision
across the chest and between the arms. Make a second incision up the midline, midline,
perpendicular to the initial cut, and then cut up to the chin area. Reflect the skin to expose
the external jugular veins in the neck area. Care should be used to avoid salivary tissue at
the midline and nodes associated with this tissue. The auricular nodes draining the ear are
located distal to the masseter muscle, away from the midline, and near the bifurcation of
the jugular veins. (Figure 10-1) The nodes can be distinguished from glandular and
connective tissue in the area by the uniformity of the nodal surface and a shiny
translucent appearance.
92
Figure 10-1. Schematic illustrating the location of the auricular lymph nodes in
mice.
(Figure adapted from:
http://iccvam.niehs.nih.gov/docs/immunotox docs/llna/LLNAProt.pdf)
93
11.
Appendix C: Protocol for intratracheal administration in
mice
As part of our goal to explore the use of particulate carriers to mediate noninvasive mucosal delivery of vaccines and immunotherapeutic agents, we sought to
deliver these formulations via an intratracheal route in mice, a model of immunization in
humans via inhalation. To do this, we followed a protocol published by the Jacks Lab in
MIT and the procedure is as follows2 77 :
1) Anesthetize mice by intra-peritoneal injection of room temperature 20 mg/mL avertin
(use 0.4 mg/g body weight for females and 0.45 mg/g body weight for males). Confirm
the mice are fully anesthetized by ensuring that they lack a toe reflex.
2) Place the mouse on an intubation platform (purchased from Steve Boukedes,
labinventions@gmail.com) so that it is hanging from its top front teeth on the bar (Figure
11-la-c).
3) Push the mouse towards the bar so that the chest is vertical underneath the bar
(perpendicular to the platform) (Figure 11-Ib).
4) Direct the Fiber-Lite Illuminator (Model 3100-1, Dolan-Jenner, 660000051001), a
fiber optic light source, to shine on the mouse's chest, in between the front legs (Figure
11-Ib,c).
5) Prepare the Exel Safelet IV catheter (22 gauge, 1 inch, Fisher, cat. no. 14-841-20) for
the instillation procedure. To ensure that the needle does not become exposed and impale
the mouse, hold the square part of the needle with the thumb and the index finger, and
using the middle finger, push the catheter over the end of the needle completely and
continue to hold the catheter in place during the administration protocol (Figure 11 -2a,b).
6) Using the Exel Safelet IV catheter, open the mouth and gently pull out the tongue with
the flat forceps (Figure 11-Id).
7) Locate the opening of the trachea by peering into the mouth and looking for the white
light emitted from the trachea (Figure 11-1 e).
8) While holding the Exel Safelet IV catheter vertically, position the catheter over the
white light emitted from the opening of the trachea, and allow the catheter to slide into
the trachea until the top of the catheter reaches the mouse's front teeth (Figure 11-1 f).
There should be no resistance while inserting the catheter into the trachea.
9) While stabilizing the Exel Safelet IV catheter with one hand, remove the needle from
the mouth (Figure 1l-Ig).
94
10) The proper placement of the catheter in the trachea can be confirmed by visualizing
the white light shining through the opening of the catheter in the mouth (Figure 11-1h).
11) Pipette the formulation directly into the opening of the catheter to ensure the entire
volume is inhaled (Figure 11-li).
12) If the catheter is correctly inserted into the trachea, the mouse will begin inhaling the
formulation immediately. Once the formulation is no longer visible in the opening of the
catheter, wait a few seconds for the entire volume to travel down the catheter before
removing the catheter from the trachea and disposing of it.
13) Place the mouse under a heat lamp (Figure 11-3a) or on a latex glove filled with
warm water (Figure 11-3b) or heat pad to recover from anesthesia.
Figure 11-1. Intratrachealinstillation procedure.
Anesthetized mice are placed on the platform by their front teeth so that their chest hangs
vertically beneath them (a,b). The light is directed on the mouse's upper chest (b,c), on
the spot marked by the 'X' (c). The mouth is opened using the Exel Safelet IV catheter
(d), and the tongue is gently pulled out using the flat forceps. After locating the white
light emitted from the trachea (e), the Exel Safelet IV catheter is slid into the trachea (f),
and the needle is removed (g). The mouse with the inserted catheter (h) on the platform is
95
moved into a biosafety hood, where the virus is dispensed into the opening of the catheter
(i).
Figure 11-2. Preparation of the Exel Safelet IV catheter for intratracheal
instillation.
Upon opening the Exel Safelet IV catheter, the needle is exposed (a). Slide the catheter
over the end of the needle to completely cover the tip (b) and the Exel Safelet IV catheter
is now ready to use.
Figure 11-3. Recovery following intratracheal administration.
Mice can be placed under a heat lamp (a) or on a glove filled with warm water (b) to
recover following anesthesia
Text and figures adapted from:
Michel DuPage, Alison L Dooley and Tyler Jacks; Nature Protocols 4, p. 1 064 - 1072
(2009).
Notes:
Inadequate anesthesia increases the probability that the mouse will swallow the
formulation, therefore the amount of avertin administered to the mice is crucial to the
success of the procedure. Mice administered with too much avertin are more likely to
stop breathing during the procedure and recover poorly from the anesthesia. Conversely,
mice administered with too little avertin may struggle to inhale the formulation and
96
should be given more avertin before continuing. Therefore, it is recommended to start
with the smallest volume of avertin needed to keep mice anesthetized during the
procedure and if necessary, administer more avertin, 50 -100 p.L at a time. If mouse is
over-anesthetized and breathing very slowly, wait until breathing becomes more regular
but ensure the mouse still lacks a reflex response before attempting administration.
Following the procedure, mice will recover better if they are kept warm to maintain their
normal body temperature after anesthesia.
To locate the white light marking the opening of the trachea, it is recommended to
look at the ventral surface of the throat. This can be done by leaning further over the
mouth and pushing the tongue with the catheter towards the ventral surface of the throat.
Sometimes, saliva may be covering the opening of the trachea. It is recommended that
one gently probe at the back of the throat with the Exel Safelet IV catheter to expose the
trachea. If the catheter is correctly inserted into the trachea, the mouse will begin inhaling
the formulation immediately. If it does not inhale the formulation (formulation stays in
the catheter), it is likely that the catheter is inserted into the esophagus. In this case, one
should pipette the formulation out of the catheter for reuse and begin the catheter
insertion procedure again.
97
12.
Appendix D: Intraperitoneal immunization with mRNA
Having demonstrated the successful transfection of a protein antigen in vivo (see
Chapter 6), we proceeded with an immunization trial in mice where we wanted to include
a positive control with mRNA to verify that the mRNA being delivered is capable of
inducing an immune response. Previously, our collaborators at the Ragon Institute have
obtained positive but inconsistent immune responses by immunizing mice
intraperitoneally using a commercial mRNA transfecting reagent, TransIT@-mRNA
Transfection Kit from Mirus (herein referred to as mirus). However, we were unable to
administer mirus-formulated mRNA intranasally or intratracheally since the mirusmRNA complex needs to be formulated in a relatively large volume and we are limited
by the small volume of the murine nasal and lung cavity available for delivery. As such,
we decided to carry out the immunization using an intraperitoneal route.
For this immunization, we sought to deliver mRNA encoding the SIINFEKL and
3S peptide sequence (3 gg per mRNA per mouse) which represent the MHC Class I H2Kb restricted peptide epitope from ovalbumin and a T helper epitope from HIV gp41
protein respectively, where the former constitutes the antigen of interest and the later is
expected to provide stimulation for helper T cells that can aid in the initiation of a CD8 T
cell response against SIINFEKL epitope. We also included in the vaccine formulation 2
adjuvants, flagellin and CL097 (both purchased from Invivogen, 0.5 and 5 jig per mouse
respectively), to stimulate innate immunity and boost any resulting immune response. In
this experiment, we compared mice immunized with mRNA adsorbed onto lipid-coated
PBAE particles (150 pg per mouse) with mRNA formulated with Mirus (n = 5 mice per
group). Mirus-mRNA complex was formulated according to manufacturer's instruction
for delivering the same amount of mRNA. Both mirus and particle associated mRNA was
mixed with the adjuvants immediately prior to administration in a 260 jiL volume
intraperitoneally in mice.
To measure the resultant immune response, at day 40 after the 1 st immunization,
we assayed for antigen-specific immune response generated following a prime and 2
boost immunizations, each immunization spaced 2 weeks apart, using an ELISPOT assay
to detect IFNy-secreting splenocytes upon restimulation with SIINFEKL and 3S peptide
antigens. We also included pneumocytes isolated from lungs to test for mucosal immune
response, in additional to systemic response in spleens.
The ELISPOT assay was performed according to manufacturer's instruction using
the mouse IFNy ELISPOT kit from R&D systems. Figure 12-1 shows the images of the
wells seeded with 1 million splenoctyes and pneumocytes respectively. Sample wells
were stimulated with SIINFEKL and 3S peptides at 1 ptg/mL and compared to negative
control wells that were not given any stimulation. Cells were incubated at 37 *C for 48 h
before developing the plate with the detecting antibody and reagents. Positive control
wells were also set up where cells were stimulated with ConA to produce cytokines nonspecifically, which was shown to be the case here by the high density of spots indicating
the presence of IFNy secreting cells, validating that the assay is working properly.
98
For the group immunized with mirus-mRNA complexes, we observed a
significant response in 3 out of 5 mice marked in red boxes in Figure 12-1 (top) and the
remaining 2 mice displaying little or no response. This result agrees well with our
prior experiments immunizing mice with mirus complexes
collaborators'
intraperitoneally where a positive but inconsistent immune response was obtained. A
stronger response was also observed in the spleen compared to the lungs for all 3
responding mice, consistent with a parenteral immunization route which predominantly
induces a systemic immune response. For the group immunized with particle adsorbed
mRNA, we were only able to detect a relatively weak response in the spleen in 1 out of 5
mice in the group and no response was seen in the lung (Figure 12-1 bottom).
99
Mirus
Mirus I
Spleen Lung
O
-0
2
. Mrus 2
Spleen Lung
0
13
0
13
Mirus 3
Spleen Lung
5
3
.0.0,--
-
3
-
0
" -
Minus 5
Spleen Lung
Mirus 4
Spleen Lung
0
0
I
0
0
0
4-
3
*
4,
7
ounted by: ImmunoSpt 5.0.3
rrpid -coated PBAE particles
o
NM
SpSen Lung
-
-NP2
oASpen Lung
o
-
1
t
0 w
e
NP4
Spleen Lung
NP3
Splen Lung
Sp
NPS
Lung
then
2
m
U.
0
0
4r
Si
10
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ountd byI
SINELppie
0
0
0
z
mmunoSpot 50
next~0 2 row
rw
0
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2
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nex
0
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oswregtv
2
oto
Figure 12-1. ELISPOT assay on mice immunized with mirus-mRNA complexes
relative to particle-adsorbed mRNA.
Splenocytes and pneumocytes were isolated on day 40 following the I"t immunization
and processed to obtain cells, where was then seeded at 1 million cell/well. Cells from
each mouse were seeded in columns. The 1st 2 rows of cells were stimulated with
SLINFEKL peptide, next 2 rows with 3S peptide. The next 3 rows were negative control
wells that were not given any stimulation and the last row is for positive control wells
stimulated with ConA that induce cells to produce cytokines non-specifically. Cells were
incubated at 37 'C for 48 h before developing the plate with the detecting antibody and
reagents. The plate was allowed to dry before imaging and analysis with the
ImmunoSpot®g Series 5 Analyzer.
100
In parallel with the immunization experiment above, we ran separate imaging
experiments on mice given mirus-, particle-associated or naked mRNA encoding for
luciferase reporter protein, to compare the distribution of transfected cells and correlate
any observations made with our immunization data. Also, we wanted to verify that the
addition of adjuvants in the formulation did not affect antigen expression negatively since
we previously observed that certain adjuvants co-delivered with mRNA can interfere
with mRNA transfection and abolish protein expression (data not shown).
To do this, we carried out the same imaging experiments as described in chapter
6. Fluorescence imaging was used to track the fluorescent labeled particles and we see
from Figure 12-2 that particles mostly remained at the injection site with fluorescent
intensity gradually decreasing with time as particles are taken up and degraded. We then
dissected the mice and isolated tissues to determine where the transfected cells were
found (Figure 12-3). Interestingly, no signal was detected in the mice given naked mRNA
despite having seen a strong signal when imaging the mice alive (data not shown). We
postulate that this may indicate the injection with naked mRNA is transfecting free
peritoneal macrophages which are washed away when opening up the animal (in
accordance with what we may expect with parenteral injection of mRNA forcing the
mRNA directly into freely-circulating cells, a likely consequence of the pressure
associated with the act of injection itself. More on this in the paragraph below.) In the
mirus group we see that the bioluminescence signal is distributed between the site of
injection, the spleen, the gut-draining mesenteric lymph nodes and mediastinal lymph
node which drains the lungs. In contrast, transfected cells were concentrated near site of
administration for the particle treated group. Overall, the results correlates well with the
ELISPOT results earlier where IFNy-secreting cells can be detected in the spleen and
lungs for mirus immunized mice but not particle immunized mice.
Particles
Blank
60
t 0h
40
X18
t=6 h
Blank
t =24 h
_
2D
"PaKVflI
Figure 12-2. Intraperitoneal administration of lipid-coated PBAE particles.
Fluorescence image of mice immediately, at 6 and 24 h following intraperitoneal
administration with rhodamine-lipid labeled particles.
101
Naked
Mirus
Particles
MDLN
Spleen
MLN
MDLN
Spleen
MLN
MDLN =
Medlastinal
Lymph Nodes
MLN= Mesenteric Lymph Nodes
Figure 12-3. Transfected tissues in dissected mice imaged post-mortem.
Representative bioluminescence images (top) and corresponding photographs (bottom) of
dissected mice 6 h following intraperitoneal mRNA administration. (Note the unlabeled
bright circular spots are due to autoluminescence from the pins used to hold tissues in
place and should be disregarded.)
Although naked mRNA injection in vivo can give rise to high levels of protein
expression in live mice, we observed that the transfection levels also tend to be highly
variable and lack consistency compared to nanoparticles (data not shown). In addition,
the protein expression obtained with parenteral administration of naked mRNA in mice is
also likely an artifact of the pressure generated by the injection of a relatively large
volume of fluid into a small animal which forces unprotected mRNA into cells directly
thereby circumventing nuclease degradation prevalent in the extracellular environment.
This phenomenon however is unlikely to translate to bigger animals such as primates or
human, and accounts at least in part for the poor translation of DNA vaccines from small
animal models to primate or human.
Quantifying the imaging results in Figure 12-4, we see that despite the weak
immune response generated in the immunization experiment, a good level of transfection
in the local tissue, higher than that obtained with mirus complexes, is attained with the
particles. This suggest that in addition to the requirement of achieving efficient
transfection with mRNA in tissues, the generation of a immune response is also
dependent on the anatomical location and therefore the cell type being transfected. In the
case of the particles, although the cells in the local tissue are transfected efficiently, the
antigen is perhaps not being delivered efficiently to the cells/tissue necessary to initiate
an immune response. For example, the mRNA on particles may be mainly taken up by
non-APCs and since the translated antigen is not designed to be secreted, few if any
APCs will be able to gain access to the expressed antigen. Mirus on the other hand is
likely able to access APC since they are relatively flexible complexes compared to our
solid coreshell nanoparticle system and can drain more efficiently to draining lymph
102
nodes to transfect APCs there and prime an immune response locally which can then
traffic to other remote lymphoid tissues.
0
20000-
e
0
Naked
NP
Mirus
15000-
10000*
5000-
0.,
e
.2
to
n' -- A&
I
I~U-
S
S
-0
-
1\00
COO
Ate"
\00o
Figure 12-4. Distribution of bioluminescence at various tissue sites.
Bioluminescence signal at the various dissected tissues at 6 h quantified from groups of
mice administered naked, particles- (NP) or mirus-associated mRNA intraperitoneally. (n
-4 mice/group).
One way to get around this problem would be to deliver mRNA that encodes for a
secreted antigen to increase the chance for APCs to encounter antigen by easing the
constraints of direct transfection of APCs currently necessary for this to occur. Another
way would be to directly target particles to appropriate APCs via targeting ligands which
can be incorporated into the lipid shell. Both strategies will provide ample room for
future work to move the particles towards a successful platform for immunization.
103
13. Appendix E: Cytosolic delivery of soluble molecules in
tumor cells
As shown in chapter 5, lipid-coated PBAE nanoparticles were able to deliver
calcein to the cell cytosol of DC2.4 efficiently. In these experiments, we also observed a
significant increase in the uptake of soluble calcein in the case of cells co-administered
calcein with nanoparticles relative to cells given calcein alone. We therefore became
interested in the possibility of using particles to chaperone the uptake of soluble
molecules from the surrounding extracellular environment. Such a strategy could be
useful in the context of cancer therapy where particles could be used to enhance uptake,
as well as subsequent cytosolic delivery, of unbound cancer therapeutics simply by codelivering them with particles.
As a first step towards this, we began by checking to see if particles are capable of
mediating endosomal lysis and cytosolic delivery in tumor cells by performing the same
calcein delivery experiments described in chapter 5, using the membrane-impermeable
fluorescent dye calcein as a tracer of endosome disruption. Indeed, we were able to
confirm this in 2 distinct tumor cell lines, Bl6F10, a murine melanoma cell line, as well
as A20, a lymphoma cell line. From Figure 13-1, we see that in cells co-incubated with
calcein (150 tg/mL) and particles (75 ptg/mL), calcein fluorescence was detected
throughout the cell with a significant increase in the cytosolic uptake of fluorescent
calcein relative to cells that were incubated with calcein only.
104
Figure 13-1. Cytosolic delivery of calcein by lipid-coated PBAE particles in tumor
cell lines.
pH-responsive lipid-enveloped PBAE particles chaperone the delivery of the membraneimpermeable dye molecule calcein into the cytosol of tumor cells. Tumor cells were
incubated for 1 h with calcein alone or calcein and lipid-coated poly-PBAE particles,
washed to remove unbound particles, then imaged live by confocal microscopy.
Representative confocal images of B1610 cells incubated with calcein (green) alone (A)
or calcein and lipid-coated PBAE particles (B) and A20 cells incubated with calcein
alone (C) or calcein and lipid-coated PBAE particles (D). Scale bars 20 Rm.
We next designed experiments to further understand the mechanism underlying
this process. In particular, we would like to verify if co-internalization of molecules with
the nanoparticles into the same endosomes is necessary for the enhancement in cytosolic
uptake that was observed. If the core-shell nanoparticle promoted calcein escape to the
cytosol via a transient osmotic pressure which was generated only under the acidic
conditions in endolysosomes, we expect that calcein and nanoparticles will be required to
localize in the same endolysosomal compartment. Nanoparticles internalized into discrete
vesicles from calcein should be unable to promote calcein access to the cytosol. To test
this hypothesis, we conducted a simple pulse-chase experiment where we delivered
calcein and nanoparticles separately. B16F10 cells were incubated for 1 h with calcein
(150 pg/mL), extensively washed to remove any remaining free calcein, and then
incubated for a second hour with medium alone, calcein mixed with particles (75 gg/mL),
or particles alone. The cells were then washed and imaged live via fluorescent confocal
microscopy. As shown in Figure 13-2A, cells treated first with calcein only showed
minimal uptake of calcein which remained in a punctuate distribution, corresponding to
endosomes containing calcein internalized by the cells. As expected, cells serially
105
incubated first with calcein, washed, and then incubated in a second step with particles
for 1 h also showed a similar trend where little or no cytosolic uptake of calcein was
detected (Figure 13-2B). In contrast, when the cells were co-incubated with calcein and
particles in the second hour, higher levels of calcein internalization was observed with
calcein distributed throughout the cytosol, due to co-internalization of calcein molecules
and particles into common endosomes, which were then ruptured by the pH-sensitive
particles (Figure 13-2C). When particles were given to cells first followed by calcein, the
results were qualitatively identical to Figure 13-2C (data not shown). We were able to
replicate the same results delivering a membrane-impermeable toxin, phalloidin (10 pM,
5 mol% alexa-488 conjugate) in B16F1O cells, where only co-delivery of phalloidin and
nanoparticle permitted the intracellular staining of the cytoskeletal elements found at the
cell periphery, demonstrating that this requirement is not limited to calcein but is likely to
be necessary for any cargo molecule of interest. (Figure 13-2D-F) These results
confirmed that to facilitate the cytosolic uptake of calcein, nanoparticles had to be cointernalized by cells and present in the same endolysosomes with calcein; the particles do
not disrupt other vesicles in cells following escape from their initial endolysosomal
compartment. Thus, maximal cargo delivery to the cytosol would only be expected if the
particles and cargo are administered simultaneously to ensure that all cargo-containing
endosomes are disrupted.
Figure 13-2. Cargo molecules must be localized in the same endolysosome as
particles for increased cytosolic uptake.
B16F1O cells were incubated with calcein in medium containing 10% FBS for 1 h at 37
*C. After 3 washes with warm medium, cells were incubated a second hour at 37 *C with
complete medium (A), lipid-coated PBAE nanoparticles only (B), or calcein and lipidcoated PBAE nanoparticles (C). The same setup was further tested using a toxin
molecule, phalloidin and images are shown for cells incubated a second time for 3 h at 37
106
'C with complete medium (D), lipid-coated PBAE nanoparticles only (E), or phalloidin
and lipid-coated PBAE nanoparticles (F). Cells were washed after the second incubation
and imaged live via confocal microscopy. Images show fluorescence overlays of calcein
or phalloidin alexa 488 conjugate (green) and rhodamine lipid-labeled particles (red).
Scale bars 20 pm.
Finally, we would also like to investigate the properties of the cargo molecules
necessary for the enhancement in cytosolic uptake following co-delivery with particles to
take place. We postulate that the increased uptake observed with particle co-delivery is
likely to be a consequence of increased endocytic activity associated with particle uptake
by cells. In order for particles to chaperone the uptake of cargo molecules co-internalized
in the same endosome, we hypothesize that there may be an upper limit to the molecular
size of the cargo molecule being co-delivered, where particles are no longer able to
efficiently shuttle co-delivered molecules into cell cytosol. To probe this, we carry out
experiments where we incubated B16F 10 cells with a model cargo molecule dextran of
increasing molecular weight, from 3 kDa to 40 kDa, with and without particles and
qualitatively compared the resultant cytosolic uptake via confocal microscopy. We also
examined the effect of the charge of cargo molecule being delivered by comparing
anionic and neutral version of the dextran molecule. The range of dextran tested was all
available commercially from Invitrogen and was fluorescently labeled with a conjugate
dye in the red channel. In this study, B16F1O cells were incubated with labeled dextran
(150 g/mL) with and without nanoparticle (75 pg/mL) for 1 h at 37 'C before washing
to remove excess molecules and particles and imaging under a confocal microscope. We
see in Figure 13-3 that for 3 kDa dextran, both the anionic and neutral dextran displayed
enhanced uptake and a cytosolic distribution when co-delivered with particles. For 10
kDa dextran, although increased cytosolic uptake was observed with nanoparticle coincubation for both anionic and neutral dextran, the degree of enhancement was reduced
in the case of neutral dextran. At 40 kDa, we no longer see cytosolic delivery or enhanced
uptake of soluble dextran following co-treatment of cells with particles. This suggests
that the chaperone effect observed will likely work best for lower molecular weight
molecules up till ~10 kDa where with increasing molecular weight, the enhancement
effect is greater for anionic compared to neutral cargo molecules. The increase in
enhancement observed with anionic molecules is not surprising since it is likely that
anionic cargoes may exhibit greater interaction with cationic particles, facilitating cointernalization into the same endosome for enhanced cytosolic delivery, where the degree
of enhancement is likely to become more critical as a bigger cargo molecule is to be
delivered. The data above provides support for testing the membrane-impermeable toxin,
phalloidin (788 Da) as a potential candidate cargo to be chaperoned by particles into
tumor cells since it is well within the range of molecular weight defined in the above
study where enhancement in cytosolic uptake should occur. This was examined and
described in chapter 5. For cargo molecules greater than 10 kDa, corresponding to the
majority of therapeutic biologics such as proteins, it is likely necessary to increase the
association between target cells and cargo molecules by additional means in order for
particles to subsequently trigger an increase in endocytic uptake and cytosolic delivery.
This could be the case when delivering immunotoxins where the cargo molecule can bind
107
to cells via their targeting moieties and particles can then be delivered to facilitate
cytosolic uptake. This was explored and documented in chapter 5.
108
Dextran
Dextran
+ NP
Dextran
Dextran
+ NP
Dextran
Dextran
anionic
neutral
10kDa
anionic
neutral
40kIEa
anionic
neutral
109
+ NP
Figure 13-3. Effect of molecular weight and charge of cargo molecules on cytosolic
uptake with lipid-coated PBAE particles.
B 16F 10 cells were incubated with dextran (red, 150 gg/mL) of various molecular weight
and charge with or without lipid-coated PBAE nanoparticles (unlabeled, 75 pg/mL) in
medium containing 10% FBS for 1 h at 37 0 C. Cells were washed and imaged live via
confocal microscopy. Scale bars 20 pm.
110
14.
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