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Polypeptide based drug carriers for anti cancer applications
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Polypeptide based drug carriers for anti cancer applications
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Content
POLYPEPTIDE BASED DRUG CARRIERS FOR ANTI CANCER
APPLICATIONS
By
Suhaas Rayudu Aluri
A Thesis Presented to the Faculty of the
USC Graduate School
University of Southern California
In partial fulfillment of the
Requirements for the degree
DOCTOR OF PHILOSOPHY
PHARMACEUTICAL SCIENCES
August 2013
Copyright 2013 Suhaas Rayudu Aluri
ii
Dedication
This thesis is dedicated to all of my family, mentors, and friends.
iii
Acknowledgements
I would like to acknowledge my gratitude to my mentors Dr. John A. Mackay and
Dr. Alan L. Epstein for their support and patience. I would also like to thank all of
my committee members Dr. Sarah Hamm-Alvarez, Dr. Julio Camarero, and Dr.
Bogdan Olenyuk for their feedback and time spent in reviewing my thesis. I
would like to thank all my lab mates Siti Janib, Pu Shi, Wan Wang, Martha
Pastuzka, Mihir Shah, and Howard Chang for all their help and input. I would also
like to that members of the Epstein Lab Dr. Peisheng Hu, Mandy Han, Julie Jang
and Saman Karimi for their help and guidance.
iv
Table of Contents
Dedication……………………………………………………………… ii
Acknowledgements…………………………………………………... iii
List of Tables………………………………………………………….. ix
List of Figures…………………………………………………………. x
Abbreviations………………………………………………………….. xii
Abstract…………………………………………………………………. xiii
1.0. Chapter 1: Environmentally responsive peptides............................
as anticancer drug carriers
1.1. Introduction…………………………………………………….....
1.2. Mechanisms of peptide-mediated tumor targeting.................
1.2.1. Enhanced permeability and retention………………………
1.2.2. Ligand-mediated targeting…………………………………..
1.2.3. Temperature-mediated targeting……………………………
1.2.3.1. Application of hyperthermia……………………….
1.2.3.2. Temperature mediated release…………………...
1.2.4. Redox-mediated targeting…………………………………..
1.2.5. pH-mediated targeting……………………………………….
1.3. Discussion………………………………………………………...
1.4. Conclusions……………………………………………………….
1
1
4
6
9
18
18
20
27
30
37
38
v
2.0. Chapter 2: Elastin-like Peptide-Amphiphiles form…………………
nanofibers with tunable length environmentally
responsive polypeptides
2.1. Introduction……………………………………………………….
2.2. Materials & Methods……………………………………………..
2.2.1. Materials………………………………………………………..
2.2.2. Synthesis and purification of ELP biomaterials …..............
2.2.3. Determination of critical micellar concentration (CMC)…..
2.2.4. Preparation and particle size of ELPAs …………..………...
2.2.5. LCST determination of ELPAs………...................................
2.2.6. Secondary structure determination using …………………..
circular dichroism (CD)
2.2.7. Transmission electron microscopy (TEM)…………………..
40
40
44
44
44
46
46
47
47
48
2.2.8. Atomic force microscopy (AFM)……………………………..
2.2.9. Cell binding/uptake studies…………………………………..
2.2.10. Paclitaxel encapsulation and delivery studies…………...
2.3. Results……………………………………………………………
2.3.1. Synthesis and purification of ELPAs ……………………….
2.3.2. ELPAs assemble nanofibers with a low CMC …………….
2.3.3. ELP nanofibers exhibit LCST behavior ……………………
2.3.4. β-turn formation associated with approach of ELPA………
transition temperature
2.3.5. Inclusion of DOPE promotes ELPA cell uptake …………...
2.3.6. Encapsulated PAX reduces tumor cell viability ……………
48
49
50
51
51
51
57
57
61
64
vi
2.4. Discussion………………………………………………………..
2.5. Conclusion……………………………………………………….
2.6. Acknowledgements……………………………………………..
66
72
72
3.0. Chapter 3: Antibody-core protein polymer nanoworms……………
(ACPPNs) potentiate apoptosis better
than a monoclonal antibody
3.1. Introduction………………………………………………………
3.2. Materials and Methods…………………………………………
3.2.1. Materials………………………………………………………..
3.2.2. Expression and purification of scFv ELP fusions………….
3.2.3. Determination of purity and transition temperature………..
(Tt) of scFv assemblies
3.2.4. Light scattering analysis of scFv ELP fusions……………...
3.2.5. Electron microscopy of scFv ELP fusion……………………
3.2.6. Secondary structure determination using…………………..
circular dichrosim (CD)
3.2.7. In vitro CD20 recognition using laser………………………..
confocal microscopy
3.2.8. Cell viability assays……………………………………………
3.2.9. Detection of apoptosis through flow cytometry…………….
3.2.10. In vivo tumor regression and biodistribution studies……..
3.3. Results……………………………………………………………
3.3.1. Purity and biophysical properties of scFv driven…………..
assemblies
3.3.2. scFv ELP renaturation reduces particle size and………….
stabilizes secondary structure
73
73
77
77
78
81
82
82
83
84
84
85
87
89
85
91
vii
3.3.3. ACPPNs competitively bind CD20 cell surface…………..
Receptor
3.3.4. ACPPNs selectively reduce CD20+ cell viability……….
by inducing apoptosis
3.3.5. ACPPNs accumulate in xenografted tumors and………….
successfully retard tumor growth in mice
3.4. Discussion………………………………………………………..
3.5. Conclusion……………………………………………………….
3.6. Acknowledgements……………………………………………..
4.0. Chapter 4: ELP conformation determines liposome release…
and cellular uptake
4.1. Introduction……………………………………………………….
4.2. Materials & Methods……………………………………………..
4.2.1. Materials……………….........................................................
4.2.2. Synthesis and purification of dpKXn…………………………
4.2.3. LCST determination of dpKXn……………………..………...
4.2.4. Preparation and characterization of dpKXn liposomes…...
4.2.5. Liposome release assay…………….....................................
4.2.6. Liposome fusion assay………………………………………..
4.2.7. Biophysical characterization of dpKXn shielded……………
liposomes
4.2.8. Cell uptake of dpKXn and DOX loaded liposomes………...
4.3. Results……………………………………………………………
4.3.1. Purity and yield of synthesized dpKXn ……………….........
4.3.2. Preparation of dpKXn liposomes …………………………….
95
97
100
105
107
108
109
109
109
109
114
114
115
116
118
119
119
120
120
120
viii
4.3.3. Effect of dpKXn phase transition on content leakage..…...
4.3.4. Evaluation of dpKXn liposome release kinetics …………...
4.3.5. Contents release of dpKXn liposomes is due to……………
pore formation
4.3.6. Biophysical characterization of dpKXn surface…………….
conformation
4.3.7. Peptide conformation influences cell uptake……………….
4.4. Discussion………………………………………………………...
4.5. Conclusion…………………………………………………………
4.6. Acknowledgments………………………………………………..
5.0. References………………………………………………………………
124
126
128
128
132
134
137
137
138
ix
List of Tables
Table 1: Receptor & protein mediated targeting peptides………………… 17
Table 2: Thermally-responsive peptides……………………………………. 26
Table 3: pH-responsive peptides……………………………………………. 36
Table 4: Physico-chemical characterization of ELPAs……….…………… 53
Table 5: Biophysical characteristics of cloned scFv ELP fusions……….. 80
Table 6: Dry organ weights from treatment groups………………………..
103
Table 7: Characterization of the purified ELP-lipid conjugates…………..
122
Table 8: Physico-chemico properties of liposome formulations…………
123
Table 9: Relationship between grafted polymer conformation and……..
extent of release
131
x
List of Figures
Figure 1: Intra-tumoral vascular permeability permits entrapment……..
of drug carriers
8
Figure 2: Peptide-mediated targeting via ligand/cell-surface…………….
interactions
16
Figure 3: Strategies for temperature-mediated peptide targeting……….
25
Figure 4: Strategy for redox-mediated peptide targeting…………………. 29
Figure 5: Strategies for pH-mediated peptide targeting…………………... 35
Figure 6: Design of ELPAs with LCST behavior…….……………………..
43
Figure 7: ELPAs self-assemble micelles with a low critical micelle……..
concentration (CMC).
53
Figure 8: dpA3 self-assembles cylindrical fibers which form dense…….
networks
55
Figure 9: The length of the ELPA fibers can be controlled using………..
56
Figure 10: ELPAs (dpA3) and soluble ELPs (A192) both undergo………
LCST behavior
59
Figure 11: dpA3 forms temperature-dependent secondary………………
Structure
60
Figure 12: DOPE addition promotes cellular uptake of ELPAs…………..
63
Figure 13: dpA3 Encapsulated PAX reduces HeLa cell viability………...
65
Figure 14: Proposed mechanism for fiber-length by capping……………
Phospholipid
71
Figure 15: Antibody Core Protein Polymer NanoWorms………………..
(ACPPNs) enhance apoptotic signaling.
76
Figure 16: Single chain antibodies fused to protein polymers………….
form thermally-responsive nanoparticles
90
Figure 17: scFv refolding reduces coordination number forming………
ACPPNs
93
xi
Figure 18: ACPPNs competitively target CD20+ cells…………………
96
Figure 19: ACPPNs reduce viability of CD20+ human cell lines by….
inducing apoptosis
99
Figure 20: ACPPNs treatment shows relatively high tumor……………
accumulation and reduces tumor burden in
Raji xenograft.
102
Figure 21: Organ histology of RTXN and ACPPNs treated……………
animals show no marked changes in histology
104
Figure 22: Surface-graft density controls release from vesicles…………
stabilized by environmentally responsive polypeptides
112
Figure 23: Short ELPAs and high molecular weight ELPs have …………
similar phase behavior
125
Figure 24: ELP surface-graft density controls content release from…….
Vesicles
127
Figure 25: Low- surface density ELP vesicles have rapid release………
Kinetics
129
Figure 26: Polypeptide brushes prevent particle aggregation and………
contents release is due to pore formation
130
Figure 27: Polypeptide brushes reduce particle uptake in vitro………….
133
Figure 28: Peptide conformations do not effect DOX delivery…………...
from liposomal formulations
136
xii
Abbreviations
ELP,Elastin like polypeptides, dp-K,dipalmitoyl-lysine, A,Alanine, V,Valine,
RTXN, Rituxan, ELPAs, Elastin like peptide amphiphiles, scFv, Single chain
variable fragment, PAX, Paclitaxel, DOX, Doxorubicin, TEM, Transmission
electron microscope, AFM, Atomic force microscopy, EPR, Enhance permeability
and retention, ANXV, Annexin V, PI, Propidium iodide, TUNEL, Terminal
deoxynucleotidyl transferase dUTP nick end labeling, NHL, Non-hodgkin’s
lymphoma, PEG, Polyethylene glycol.
xiii
Abstract
Chapter 1: The tumor microenvironment provides multiple cues that may be
exploited to improve the efficacy of established chemotherapeutics; furthermore,
polypeptides are uniquely situated to capitalize on these signals. Peptides
provide: 1) a rich repertoire of biologically specific interactions to draw upon; 2)
environmentally-responsive phase behaviors, which may be tuned to respond to
signatures of disease; 3) opportunities to direct self-assembly; 4) control over
routes of biodegradation; 5) the option to seamlessly combine functionalities into
a single polymer via a one-step biosynthesis. As development of cancer-targeted
nanocarriers expands, peptides provide a unique source of functional units that
may target disease. This dissertation explores potential microenvironmental
physiology indicative of tumors and peptides that have demonstrated an ability to
target and deliver to these signals
Chapter 2: Peptide amphiphiles (PAs) self-assemble nanostructures with
potential applications in drug delivery and tissue engineering. Some PAs share
environmentally responsive behavior with their peptide components. Here we
report a new type of PAs biologically inspired from human tropoelastin. Above a
lower critical solution temperature (LCST), elastin-like polypeptides (ELPs)
undergo a reversible inverse phase transition. Similar to other PAs, elastin-like
PAs (ELPAs) assemble micelles with fiber-like nanostructures. Similar to ELPs,
ELPAs have inverse phase transition behavior. Here we demonstrate control
over ELPAs fiber-length and cellular uptake. In addition, we observed that both
peptide assembly and nanofiber phase separation are accompanied by a
xiv
distinctive secondary structure attributed primarily to a type-1 β turn. We also
demonstrate increased solubility of hydrophobic paclitaxel (PAX) in the presence
of ELPAs. Due to their biodegradability, biocompatibility, and environmental
responsiveness, elastin-inspired biopolymers are an emerging platform for drug
and cell delivery; furthermore, the discovery of ELPAs may provide a new and
useful approach to engineer these materials into stimuli-responsive gels and drug
carriers.
Chapter 3: B-cell lymphomas are widely occurring malignancies which have an
incidence rate of 27 per 100,000 subjects per year. A vital component of therapy
for B-cell lymphomas is a chimeric antibody, Rituxan® (RTXN), which has
appreciable clinical activity. RTXN targets B-cell surface receptor, CD20, and
induces tumor regression by a variety of mechanisms. One of the mechanisms of
action is through Fc crosslinking of RTXN by Fcγ expressing effector cells (NK
cells, Macrophages, etc). This phenomenon was confirmed in vitro by inducing
hyper-crosslinking via 2° antibodies and chemical m eans. To enhance the
apoptotic effects of RTXN in CD20+ B-cell lymphomas, we designed antibody
core protein nanoworms (ACPPNs) which take advantage of this phenomenon.
These novel ACPPNs are multimeric assemblies of anti-CD20 scFv-Elastin like
polypeptides (ELPs) fusions. These mutimeric ‘nanoworms’ are active and
competitively bind CD20 on B-cell lymphoma cell lines (Burkitt’s and Diffuse large
B-cell lymphomas). ACPPNs treatment shows a concentration dependent
reduction in CD20+ B-cell viability by selectively inducing apoptosis. Equivalent
scFv doses of ACPPNs show significantly higher induction of apoptosis than
xv
unconjugated RTXN and are as efficient as in vitro 2° goat anti-human Fc (GAH)
crosslinked RTXN. This potent in vitro activity of ACPPNs was successfully
translated in vivo in a Non-hodgkin’s lymphoma (NHL) xenograft model. Limited
ACPPNs treatment showed significant retardation in tumor growth and improved
survival in a human Burkitt’s lymphoma xenograft when compared to RTXN
alone. Microdistribution studies performed on rhodamine (RHD) labeled ACPPNs
showed appreciable liver and tumor accumulation. These results demonstrate
that novel recombinant ACPPNs are a first generation small molecule free
therapeutic for the treatment of B-cell related malignancies and disorders.
Chapter 4: While some amphipathic peptides interact directly with membranes,
many peptides have potentially useful environmentally-responsive behavior only
aqueous solution. We developed a simple approach to modify water-soluble
polypeptides via an amino terminal lysine using two palmitoyl lipids. As a model
polypeptide, we used our approach to graft short elastin like polypeptide (ELPs)
sequences to the surface of liposomes using Elastin like peptide amphiphiles
(ELPAs). High molecular weight ELPs undergo an environmentally responsive
phase transition above a solution critical temperature but the transition is not
sufficient to cause membrane rupture. Utilizing surface grafted ELPs we can both
stabilize and destabilize a lipid membrane in response to temperature. The
incorporation of two proximal lipid groups yielded an excellent degree of
stabilization to a lipid membrane. This low cost approach avoids chemistry using
expensive or pH-labile lipids, is compatible with solid phase chemistry, and is not
mediated by bulky peptide domains. Incorporation of this novel conjugate on a
xvi
lipid membrane causes rapid leakage of content when heated to 45° C. Leakage
of contents was confirmed to be due to membrane permeabilization. Surface
polymer conformation determines the extent of content leakage where low
grafting density or ‘mushroom’ conformation yields higher leakage than polymers
in ‘brush’ conformation. Also liposomes formulated with a ‘brush’ conformation
remained stable at all temperatures and behaved similar to PEG grafted
liposomes. The peptide conformation influences liposome uptake with the
‘mushroom’ conformation allowing for high cell uptake compared to the ‘brush’
conformation. Utilizing these liposomes, doxorubicin (DOX) was successfully
encapsulated and delivered to cancer cells.
1
Chapter 1
Environmentally responsive peptides as anticancer drug carriers
1.1 Introduction
According to the American Cancer Society, 7.6 million people died from cancer in
the world during 2007 (2008) , and about 565,650 Americans were expected to
die of cancer (2008) . Cancer is the second most common cause of death in the
US, exceeded only by heart disease (2008) . The sustained prevalence of cancer
continues to motivate the development of new therapies. Significant research
efforts have been directed towards targeting cancer drugs to tumors using
specialized drug carriers, and peptides have become an important component of
these targeting approaches. The contents of this review address the current
status of peptide-mediated targeting of drug carriers.
It is likely that drug carriers will play an increasing role in the treatment of
cancer. Cancer treatment is multi-pronged, consisting of surgery, radiation, and
drug-mediated chemotherapy, which varies depending on the nature of the tumor
(2008). Among these three modalities, improvements to chemotherapy offer
some of the most exciting opportunities to develop new approaches. Most
traditional chemotherapeutic agents have a therapeutic index close to one, and
they cause concentration dependent toxicity in non-cancerous tissues (Chabner
and Longo, 2001). The mechanism of dose-limiting toxicity varies from drug to
drug. For example, the administration of doxorubicin (DOX) is limited by
cardiomyopathy that arises from oxidative stress (Olson et al., 1981). In mice free
DOX has a therapeutic index of 1.2; however, this has not prevented its
2
widespread use over the past 4 decades (Di Marco et al., 1969). In contrast,
neuropathy and neutropenia are the dose-limiting toxicities for patients treated
with paclitaxel, a microtubule stabilizing agent (Seidman et al., 1998).
Presumably, the differences in mechanisms of action and toxicity between
chemotherapeutic agents will significantly impact the optimization of the drug
carrier strategy. In addition, the concentration dependent cytotoxicity of these
chemotherapeutics typically makes them unsuitable for administration routes that
produce high local concentrations, such as oral, transdermal, or subcutaneous
administration. To circumvent this localized toxicity near the site of
administration, many of these agents are delivered via intravenous administration
(Vokes and Golomb, 2003). After systemic administration only a small
percentage of the administered drug reaches the target site (Schilsky et al.,
1996); furthermore, encapsulation of chemotherapeutics inside drug carriers can
increase this percentage. For example, liposomal DOX accumulation in tumors is
3 to 15 fold higher than for free drug (Drummond et al., 1999). Ultimately
chemotherapeutics continue to produce dose-limiting toxicity by their interaction
with healthy tissues, diminishing their clinical efficacy; furthermore, a major
rationale for the development of nanoparticulate carriers has been to reduce drug
exposure to normal tissues.
Without a mechanism for releasing drug, carrier encapsulation is typically
insufficient to generate useful anti-tumor responses. To improve the therapeutic
index, carriers must have an optimal rate and mechanism of in vivo drug release.
For many carrier systems, the tumor drug concentration depends upon the
3
mechanism of encapsulation/attachment, as well as the physicochemical
properties of the drug. Most drugs differ in their rates of systemic and local
clearance; therefore, each drug must be evaluated on a case by case basis
(Allen and Cullis, 2004). With regards to the rate of drug release, these carriers
may fall into three possible regimes: 1) release is too slow; 2) release is too fast;
and 3) release is perfectly balanced (Lee et al., 2005). If the release of a drug
from a carrier is too rapid, then the drug may clear to the bloodstream prior to
reaching the tumor. While negligible rates of release of drug into the blood can
be tolerated, rapid and untargeted drug release is not a desirable property for a
reliable delivery system. Under this scenario, the carrier is unable to prevent
exposure at sites of toxicity, and may be unable to promote selective tumor
accumulation beyond that of free drug. Under the second scenario, where the
carrier release-rate is too slow, the local clearance of drug is faster than release
of a free drug. The resulting concentration of drug available in the tumor may
then be sub-therapeutic, even though the total concentration of encapsulated
drug in the tumor may be high (Allen and Cullis, 2004; Lee et al., 2005). In the
third scenario the rate of drug release from the carrier is perfectly balanced to
yield both optimal tumor accumulation and localized release; however, this
balance is difficult to achieve.
Various nanocarrier systems have been explored that approach optimal
rates of tumor accumulation and availability (Wong et al., 2007). Formulations
including Doxil
TM
(lipid-mediated) and Abraxane
TM
(peptide-mediated) have
partially overcome these barriers, and have been translated to the clinical setting
4
(Gabizon et al., 1998; Hawkins et al., 2008; Henderson and Bhatia, 2007;
Ranson et al., 2001). Falling short of its promise, a substantial limitation of the
liposomal formulation appears to be that it has a suboptimal rate of drug release
in the tumor (Kong et al., 2000a). Here we describe a range of mechanisms for
controlling localized accumulation and drug release; furthermore, we summarize
peptides with potentially useful behaviors that may enhance delivery to the tumor
microenvironment.
1.2 Mechanisms of peptide-mediated tumor targeting
Environmentally responsive delivery systems make use of tumor pathology to
trigger release of therapeutic agents at the target site. The tumor
microenvironment has been widely studied, generating a host of biomarkers
suitable for targeted delivery (Cook et al., 2004; Gerweck and Seetharaman,
1996; Mbeunkui and Johann, 2009). The list of potential biomarkers is extensive;
however, it is likely that only a subset of biomarkers can be engineered into
suitable triggers for targeted drug delivery. The list of potential biomarkers
provided by the tumor microenvironment can be broadly classified into physical
or molecular triggers. Physical triggers are activated by the nature of the tumor
microenvironment. For example, the tumor microenvironment contains regions of
reduced pH (Gerweck and Seetharaman, 1996) and increased oxidative potential
(Schafer and Buettner, 2001), and increased vascular wall permeability
(Matsumura and Maeda, 1986). On the other hand molecular triggers include the
target molecules that are upregulated in the tumor vasculature or within the
tumor cells. These targets include vascular endothelial growth factor (VEGF),
5
integrins, matrix metalloproteases and tumor necrosis factors (Mbeunkui and
Johann, 2009). Antibodies against these biomarkers have been successful in
tumor treatment studies (Ferrara et al., 2004), which have prompted more
research into pathways associated with these markers. Another example
biomarker in the tumor microenvironment is the upregulation of secreted
phospholipase A2 (Laye and Gill, 2003). Phospholipase A2 is involved in the
production of prostagladins (Laye and Gill, 2003). Phospholipase A2 mediates
carcinogenesis by two pathways: release of arachidonic acid, which produces
carcinogenic metabolites; and release of lysophospholipids, including
lysophosphatidic acids (LPA) that induce cell growth (Cummings, 2007). Both
physical and molecular biomarkers such as these are being explored to develop
an array of new nanocarriers (Shen et al., 2008; Tanaka et al., 2004).
These microenvironmental biomarkers such as these are being actively
explored for the ability to produce environmentally responsive drug release in the
tumor. One approach has been to develop environmentally sensitive polymers,
including peptides that respond to tumor microenvironment with targeted delivery
of drug (Dreher et al., 2007; Dreher et al., 2008a). Such approaches are intended
to either increase the accumulation of drug carrier in the tumor or increase the
release of active drugs from carriers that have already trafficked to the tumor.
Continued study of the tumor microenvironment is expected to reveal new cues,
which may be useful for controlled drug release. The remainder of this review
focuses on targeting mechanisms employed in current research and plausible
roles played by peptides/proteins.
6
1.2.1. Enhanced permeability and retention
Some of the most frequently applied mechanisms of tumor targeting utilize the
properties of the tumor vasculature. The tumor neovasculature is ‘chaotic’ in
nature, consisting of various loops, dead ends, and openings that lead directly
into the perivascular space (Carmeliet and Jain, 2000). These openings provide
a passive mechanism for targeting macromolecular or nanoparticulate
entrapment within the tumor. Additionally, lack of lymphatic drainage prevents
drug/carrier clearance from tumors. The combination of these two factors is
commonly referred to as the enhanced permeability and retention effect (EPR)
(Maeda et al., 2000) (Fig. 1). To obtain nutrients and oxygen, solid tumors grow
around existing blood vessels and produce new blood vessels to interconnect
with the existing vasculature, through a process known as angiogenesis. These
new vessels are characterized by high permeability and a haphazard
arrangement (Maeda et al., 2000; Yuan et al., 1994). Systemically administered
carriers must pass through this dense arrangement of vessels to reach the cells
of solid tumors. While tumors do recruit a blood supply, they fail to develop
functional lymphatic drainage (Carmeliet and Jain, 2000). One of the primary
functions of the lymphatic system is to provide a route for the clearance of
extravascular proteins, particulates, and white blood cells. Without a lymphatic
system, macromolecules and nanoparticulates that extravasate through the
‘leaky’ tumor vasculature accumulate to form a ‘depot’ in the perivascular space.
The pore sizes of some typical tumor blood vessels have been estimated to have
diameters around 400-600 nm (Yuan et al., 1995). Experiments have shown that
7
carriers with a diameter of less than around 100 nm (Ahmed et al., 2005; Kong et
al., 2000b) are ideal to target the tumor vasculature via EPR. Thus appropriately
sized particles that pass through the tumor can accumulate locally, which
sustains the total drug concentration compared to the parent small molecule.
This approach is relatively specific for tumor tissue, in addition to other tissues
with permeable endothelia, even though it does not require any specific
molecular interactions to drive accumulation.
EPR targeting suffers from several limitations, in that it provides no
mechanism for generating the active drug once the carrier localizes to the tumor
and that vascular perfusion is not uniform either between or within tumors
(Schroeder et al., 2005). For some drug formulations, such as Doxil
TM
(Ranson et
al., 2001), the rate of nonspecific drug release is balanced enough to reduce
tumor mass; however, the activity of novel EPR-targeted formulations should be
specifically optimized to address the drug diffusion and permeability in the tumor
and in the cell (Yuan et al., 1994). For carriers that target via the EPR effect, the
tumor microenvironment must play a critical role in mediating release from the
carrier. Typically mechanisms for achieving drug release are designed into the
carrier, including cleavable linkers, cell targeting ligands, or permeabilizing
agents, many of which are derived from polymers and peptides. Peptides in
particular provide a wealth of rational strategies for increasing the efficiency of
drug delivery.
8
Figure 1: Intra-tumoral vascular permeability permits entrapment of drug
carriers. (A) Intact normal vasculature. (B) Semi porous tumor vasculature. Drug
carriers selectively extravasate into tumor via pores (Carmeliet and Jain, 2000).
This EPR targeting strategy provides a mechanism for extracellular drug carrier
accumulation in the tumor, which may be further augmented via peptide-
mediated targeting.
9
1.2.2. Ligand-mediated targeting
Peptides and proteins can be used to target delivery of anticancer agents via
intracellular or extracellular release. These strategies typically benefit from
passive EPR targeting prior to initiation of targeted interactions. Similar to free
drug, extracellular release is influenced by cancer drug resistance mediated by:
1) activation of the P-gp proteins (Saeki et al., 2005); 2) activation of glutathione
detoxification system (Balendiran et al., 2004); and 3) by alterations in the genes
and the proteins involved with apoptosis, such as p53 and Bcl-2 (Stavrovskaya,
2000). To circumvent this barrier, peptides can also mediate the intracellular
delivery of drugs via targeting receptors at the cell surface (Table 1). For receptor
targeting strategies, including transferrin (Niitsu et al., 1987; Qian et al., 2002;
Tanaka et al., 2001), folate (Cho et al., 2005; Hilgenbrink and Low, 2005;
Leamon and Low, 1992, 1994; Leamon and Reddy, 2004; Lee and Huang, 1996;
Lee and Low, 1995; Lu and Low, 2002) and epidermal growth factor receptors (Li
et al., 2005; Song et al., 2008; Wartlick et al., 2004), the EPR mediated
accumulation of carrier is followed by improved internalization of the carriers into
target cells, a mechanism known as receptor-mediated endocytosis (RME).
Using RME strategies, tumors can be selectively targeted by increased
localization of the carrier to the tumor and also by enhancement of internalization
(Fig. 2) (Tanaka et al., 2004). Conjugation of drugs to these carriers can in turn
facilitate significant intracellular drug concentrations. Both imaging and
chemotherapeutic agents have been targeted to tumors via RME (Leamon and
Low, 1992, 1994; Leamon and Reddy, 2004; Mathias et al., 1996).
10
Receptor-mediated peptide strategies have been extensively studied.
Peptides have been used either to directly trigger drug release or to direct the
carrier to increase specificity and internalization. This can be achieved by: 1)
conjugating the ligand to carrier/drug; 2) conjugating an antibody against the
receptor to the carrier/drug; and 3) using cell-penetrating peptides (Wadia and
Dowdy, 2005) to promote nonspecific binding and internalization. For covalently
attached drugs, it is important to incorporate a cleavable bond between the drug
and carrier to facilitate drug-release. One recently successful linkage used in
polymeric drug-delivery is the hydrazone bond, which demonstrates release and
anti-tumor efficacy in a range of models (Hamann et al., 2002). One example of
this hydrazone strategy is the antibody conjugated formulation Mylotarg
TM
, which
is a calicheamicin hydrazide derivative attached to the oxidized carbohydrates of
the anti-CD33 antibody for the treatment of acute myeloid leukemia(Hamann et
al., 2002). Ketal and disulfide linkages have also shown appreciable efficacy in
intracellular drug delivery (Gillies et al., 2004; Saito et al., 2003).
Albumin was one of the first proteins to be used a drug carrier for a variety
of drugs, including anticancer agents (Kratz, 2008). Albumin has a wide range of
applications due to its versatility in binding to hydrophobic drugs. The first drug
conjugate to be evaluated was methotrexate. Though the methotrexate-albumin
conjugate wasn’t successful clinically (Kratz et al., 2007) it prompted further
research in albumin bound therapy, which culminated in the success of albumin
bound paclitaxel, Abraxane
TM
. Although paclitaxel is not chemically conjugated to
albumin, the hydrophobic binding affinity of the protein to drug is sufficient to be
11
useful as a drug carrier. In the clinic, Abraxane
TM
was shown to have a higher
anti-tumor activity than paclitaxel (Desai et al., 2006). Abraxane
TM
targets the
tumor by a combination of the EPR effect and albumin binding to Gp60 receptor
(Kratz, 2008) . This enhances the intratumoral concentration of drug and
therefore increases the efficacy of the system. Another significant improvement
of Abraxane
TM
over paclitaxel is the reduced incidence of hypersensitivity related
to the use of harsh surfactant vehicles, such as Cremophor-EL
TM
, which is
unnecessary for the albumin-drug complexes (Henderson and Bhatia, 2007).
Extensive work on transferrin-mediated targeting was done by Tanaka and
coworkers (Tanaka et al., 2001). Transferrin undergoes receptor-mediated
endocytosis in a broad range of cells; furthermore, when Mitomycin C is
conjugated to transferrin, the resulting conjugate demonstrates unaltered
receptor binding properties. The mechanism of internalization is the same for
normal and cancer cells; however, cancer cells achieved a higher concentration
of drug in the tumors due to increased surface expression of the transferrin
receptor (Niitsu et al., 1987). Similarly, for liposomes decorated with transferrin,
there was a higher concentration of drug observed in the tumor cells, which
significantly improved efficacy in a rat tumor model (Hisae Iinuma, 2002).
Another avenue to utilize the transferrin trafficking pathways was demonstrated
using anti-transferrin receptor antibodies (Huwyler et al., 1997; Suzuki et al.,
1997). The anti-transferrin receptor strategy boosted the accumulation of drug in
the tumor two-fold compared to control liposomes. Fluorescence experiments
12
suggested that these antibodies internalize into cells similar to transferrin via
receptor-mediated endocytosis (Huwyler et al., 1997).
The folate receptor has also been a useful target in the development of
anticancer peptide strategies. Conjugation of drug molecules to folic acid or other
ligands specific to the folate receptor can increase the localization and
internalization of the conjugate (Lu and Low, 2002). This effect has been
demonstrated with cargo varying in size from small molecule drugs to liposomes.
One of the first folate conjugated systems included a protein toxin called
momordin (Leamon and Low, 1992). The folate-momordin conjugate was shown
to be specific to tumor cells (Leamon and Low, 1994). Another early folate drug
carrier system was developed using a folate-phosphatidyl ethanolamine lipid
conjugate (Lee and Low, 1995). The lipid anchored the folate directly to a
liposome surface, and these targeted vesicles had higher cellular uptake and
higher cytotoxicity than the nontargeted drug, DOX. The folate targeting strategy
has been extended to enzymes (Cho et al., 2005), DNA (Lee and Huang, 1996)
and other non-lipid delivery systems.
Epidermal growth factor receptor (EGFR/HER2) expression is frequently
upregulated in tumor cells (Song et al., 2008). A successful antibody against
HER2 has been clinically approved, Herceptin
TM
(Chang, 2007), which has been
used to increase specificity of gelatin/albumin nanocarriers (Wartlick et al., 2004).
For this strategy, avidin is directly conjugated to the drug carrier, which is
subsequently modified by biotinylated HER2 (Wartlick et al., 2004). The authors
reportedly attached 370 antibodies onto the surface of each carrier (Wartlick et
13
al., 2004). The internalization of the carrier depends strongly on the
concentration of HER2 receptors on the cell surface; furthermore, these HER2
receptor targeted carriers demonstrated enhanced binding and internalization
compared to an untargeted control (Wartlick et al., 2004). An alternative use for
the anti-HER2 receptor antibodies has been the modification of DOX carrying
liposomes (Park et al., 2001). Interestingly, these antibody-grafted liposomes
showed higher internalization inside cells (6-fold) but did not substantially
increase the amount in the tumor vasculature (Kirpotin et al., 2006). A simpler
alternative to antibody targeting is to exploit short peptides with HER2 receptor
binding, such as the GE11 peptide (Table 1) (Li et al., 2005). GE11 modified
liposomes showed efficient transfection of cells, and increased cytotoxicity when
compared to either free DOX or unmodified DOX liposomes (Song et al., 2008).
Even short peptide motifs are capable of directing tumor targeting to specific
membrane receptors. The RGD (Arg-Gly-Asp) peptide is one of the most widely
studied motifs. RGD is recognized by integrins that promote endocytosis of the
polymer or particulate presenting the peptide. The RGD motif has been used
widely to targeting tumor cells with therapeutic drugs/proteins/liposomes/imaging
agents (Temming et al., 2005). For example, RGD conjugates have been used to
target the synthetic polymer hydroxypropyl methacrylamide (HPMA) specifically
to tumor epithelial cells (Mitra et al., 2005). An RGD-HPMA conjugate had 3.7
times higher tumor localization than compared to a control peptide conjugate
(Mitra et al., 2005). Alternatively, liposomes designed with the RGD motif can
decrease tumor growth; however, off-target interactions in the liver and spleen
14
were observed (Holig et al., 2004). Also incorporation of RGD lipopeptides in the
liposomes formulation reduced the circulation time drastically (Holig et al., 2004).
More recently, other short motifs have been identified that promote receptor-
mediated uptake including NGR (Asn-Gly-Arg) and GSL (Gly-Ser-Leu). These
short motifs are tumor specific because their specific targets are upregulated in
the tumor vasculature and not to the same extent in normal tissues (Ruoslahti,
2000). The NGR, RGD, and GSL motifs can home in to cancers in vivo (Arap et
al., 1998). The NGR motif has two distinct homing mechanisms: binding to
aminopeptidase N (CD13) which is a membrane bound matrix metalloproteinase;
and binding to vascular integrins (Corti et al., 2008). A major advantage of using
these short peptide motifs is that they home in to the tumor vasculature, which is
less dependent on the variability of receptors expressed directly on the tumor cell
surface (Arap et al., 1998; Ruoslahti, 2000).
A related approach where peptides have contributed to drug targeting is in
the use of enzymes that direct tissue or cell specific cleavage of active drug. For
many drugs, the carrier-drug conjugate is inactive, and release is a prerequisite
for activity. From a practical perspective drug cleavage should not occur during
circulation in the plasma. One elegant solution is to use the proteolytic capacity
confined within lysosomes. To gain access to these compartments, the
nanocarriers must be taken into cells and internalized to lysosomes, such as
through RME. Once in the lysosome, there are multiple enzymes that can
mediate degradation; however, the most frequently discussed is cathepsin B
(Kratz et al., 2007; Rejmanova et al., 1985; Schmid et al., 2007). Other
15
interesting enzyme targets include matrix metalloproteases (MMP’s). 26 different
kinds of MMP’s exist (Verma and Hansch, 2007), and studies on MMP-9 specific
release studies have shown encouraging results (Sarkar et al., 2008; Tauro and
Gemeinhart, 2005). MMP-9 and MMP-2 are therapeutically relevant, both being
upregulated in glioblastoma multiforme (Forsyth et al., 1999; Sawaya et al.,
1996). It is also interesting to note that MMP-2 and MMP-9 have common
substrates specificities derived from collagen (Hideaki Nagase and Gregg B.
Fields, 1996). A consensus sequence has been observed, GPQGaAGQR where
a= Leucine, Isoleucine, or Valine (Sarkar et al., 2008). By conjugating this
peptide to a lipids anchor, enzyme-dependent liposome rupture has been
demonstrated (Sarkar et al., 2008). Hence, by incorporating appropriate peptides
into a linkage between carrier and drug, it is possible to develop rapid release in
the presence of target enzymes without appreciably contributing to drug loss
during circulation in the central blood compartment.
16
Figure 2: Peptide-mediated targeting via ligand/cell-surface interactions.
These strategies use ligands grafted onto the carrier/drug surface to increase the
intracellular uptake of drug in target tissues (Table 1). A ligand is presented by
the carrier, which binds to a moiety on the cellular surface, like specific receptors
or nonspecific proteoglycans (Tyagi et al., 2001). Surface binding may be
followed by internalization through receptor-mediated endocytosis or other
uptake mechanisms. Once internalized, the carrier or drug may be further
processed via other cellular factors, such as pH or enzymatic activity.
17
Table 1. Receptor- & protein-mediated targeting peptides.
Peptides *Properties References
Transferrin
and anti-
transferrin
antibody
Iron binding proteins found in vertebrates.
Enables 2-fold increase in tissue
accumulation of non-specific antibody
conjugated liposome.
(Hisae Iinuma, 2002;
Huwyler et al., 1997;
Qian et al., 2002;
Suzuki et al., 1997;
Tanaka et al., 2001)
Herceptin
TM
/
Trastuzumab
Antibody against HER2. Showed
increased internalization of carrier and a
6-fold increase in internalization.
(Kirpotin et al., 2006;
Park et al., 2001;
Wartlick et al., 2004)
GE11 peptide Ligand for HER2 (human epidermal
growth factor receptor). Contains the
peptide sequence ‘YHWYGYTPQNVI.’
(Li et al., 2005; Song et
al., 2008)
Albumin Human plasma protein soluble in water.
Assists in transporting hydrophobic
molecules. Used as a carrier for
anticancer therapy.
(Henderson and
Bhatia, 2007; Kratz,
2008; Schmid et al.,
2007; Volk et al., 2008)
HIV TAT
peptide
The TAT protein enters cells when added
exogenously. The transduction domain
responsible for this property is
‘RKKRRQRRR,’ which can be appended
to other systems.
(Eguchi et al., 2001;
Gratton et al., 2003;
MacKay et al., 2008)
NGR peptide This motif has 2 distinct binding sites 1)
Aminopeptidase N (CD13) 2) α
v
β
3
Integrin. Plays a major role in cell
adhesion.
(Corti et al., 2008;
Ruoslahti, 2000)
RGD peptide This motif is a recognized by integrins
and is important for cell adhesion.
(Temming et al., 2006;
Temming et al., 2005)
GFLG peptide Acts as a cleavable peptide linker
sensitive to proteolyses by lysosomal
cathepsin B.
(Rejmanova et al.,
1985)
ALAL peptide Acts as a cleavable peptide linker
sensitive to proteolyses by lysosomal
cathepsin B.
(Schmid et al., 2007)
GPQGaAGQR
peptide
Where a = Leucine, Isoleucine or Valine.
This peptide sequence acts as a
substrate for MMP-2,9 and can be used
as a linker/trigger for drug release from
polymer conjugates or drug carriers.
(Hideaki Nagase and
Gregg B. Fields, 1996;
Sarkar et al., 2008)
*A=Alanine, F=Phenylalanine, G=Glycine, H=Histidine, I=Isoleucine,
K=Lysine, L=Leucine, N=Asparaginine, P=Proline, Q=Glutamine,
R=Arginine, T=Theronine, V=Valine, W=Tryptophan Y=Tyrosine.
18
1.2.3 Temperature-mediated targeting
Thermo-responsive polymers respond to their surrounding temperature by
altering their physicochemical properties; furthermore, as more advanced
methods for achieving local-regional deposition of heat become available these
polymers are becoming attractive targeting options to direct tumor-specific
delivery. Peptides make excellent candidates for thermo-responsive polymers;
furthermore, there have been exciting advances in the development of thermally
responsive peptides that confer temperature dependence onto drug carriers
(Dreher et al., 2007; Han et al., 2006). This section deals with the application of
hyperthermia as well as the use of thermally-sensitive peptides as triggers in
drug release (Figure 3).
1.2.3.1. Application of hyperthermia
To synergize with current hyperthermia research, thermo-responsive drug
carriers should be able to respond to mild hyperthermia conditions between
about 37
o
C and 42
o
C (Baronzio and Hager, 2006). Sustained temperatures
above this range induce protein denaturation and cell death without the need for
chemotherapy; however, the application of high temperature is often limited by
the need to spare critical anatomical structures. In contrast, mild hyperthermia
produces effects that complement drug delivery. For example, tumor cells are
more sensitive to hyperthermia-induced damage (Chang W. Song, 1980;
Gerweck, 1985). Also, hyperthermia increases vascular perfusion and
permeability, which contributes to improved passive targeting and anti-tumor
effect (Baronzio and Hager, 2006). Tumors can be heated either by direct or by
19
indirect heating (Baronzio and Hager, 2006). Under direct heating, energy is
distributed via conduction and convection (Baronzio and Hager, 2006). Due to
the efficiency of thermal homeostasis, this approach is only suitable to a limited
penetration depth (Baronzio and Hager, 2006). To achieve deep tissue
penetration, indirect heating is used to deposit energy via radiating waves from
either ultrasonic or electromagnetic sources. Whole body hyperthermia can be
induced by administration of pyrogens; however, pyrogenic hyperthermia is
perhaps undesirable compared to the other routes of localized heating (Baronzio
and Hager, 2006). One promising heat application technology permits both
localized heating and deep tissue penetration and is based upon high intensity
focused ultrasound (HIFU) (Kennedy, 2005). HIFU is a noninvasive procedure
that can be used to induce mild hyperthermia or tumor ablation (70 ºC). While
HIFU can heat deep seated tumors, ultrasonic waves have two main drawbacks:
1) they are unable to safely penetrate gas phases, which leads to inefficient
heating in the lung and bowel; and 2) they are strongly absorbed in bone, which
can damage the skeletal tissue (Kennedy, 2005). Despite these caveats,
ultrasound heating can be used as an adjuvant to nanocarrier chemotherapy
(Han et al., 2006). One alternative method to achieve indirect heating method is
to use strong magnetic fields in order to heat iron oxide carriers (Gupta et al.,
2007). Nanocarriers composed from both lipids (Han et al., 2006; Kong et al.,
2000a) or peptides (Dreher et al., 2007; Meyer et al., 2001a) have been
observed to increase exposure of the tumor to the cytotoxic agent. For example,
lipids that undergo sol-gel phase transitions including DPPC (Yatvin et al., 1978)
20
as well as polymers that display Lower Critical Solution Temperature (LCST) are
being developed for thermo-responsive delivery (Kono, 2001) and a few
interesting examples have been summarized in the next section. Recent reviews
(Baronzio and Hager, 2006; Mackay and Chilkoti, 2008) have summarized the
application of hyperthermia in more detail and will make interesting reading.
1.2.3.2. Temperature mediated release
Hyperthermia-based strategies that are under investigation explore both polymer-
drug conjugates or thermo-sensitive particulates that encapsulate drugs (Table
2). Some significant research into temperature-sensitive targeting focuses on
peptides that undergo unique biophysical behavior, and in several cases these
peptides can be reduced to repetitive amino acid sequences (Table 2). This
section focuses exclusively on the properties of some temperature dependent
peptides. The most widely studied thermally-responsive peptides are elastin-like
polypeptides (ELPs), collagen, leucine zippers, and silk (Mackay and Chilkoti,
2008). Each of these peptides display temperature dependent phase transitions,
with some being reversible and others irreversible. In most cases, the thermal
sensitivity of these peptides can be modulated by varying: 1) the MW of the
peptide; 2) the identity of specific amino acid residues; and 3) the concentration
(Meyer et al., 2001b). Leucine zippers are heptameric repeats of [abcdefg]
n
that
form α-helical structures, which can self-assemble (Table 2) (Hart and Gehrke,
2007; Mart et al., 2006). Leucine zippers are stable at low temperatures, and
dissociate under heating at a temperature that depends on their specific peptide
sequence. These zippers could potentially be used in delivery strategies that
21
drive assembly of particulate drug carriers that trap the drug in a carrier until
release. Also, leucine zippers can be used as switchable hydrogels, which serve
as matrices for controlled release under elevated temperatures (Petka et al.,
1998; Xu et al., 2005).
Another class of thermally responsive peptides, the ELPs, are biologically
inspired from human tropoelastin (Urry, 1997a). ELPs are pentameric repeats of
[aPGbG]
n
that exhibit a first order phase transition from a water soluble phase to
a two phase system. In vivo ELPs are sensitive to hyperthermia; furthermore, the
ELP phase behavior is reversible. The phase transition temperature can be
manipulated by modifying the guest residue in the a and b positions (Table 2).
The trend is that substitution at the b guest residue with increasingly hydrophobic
amino acids results in a lowering of the phase transition temperatures (Mackay
and Chilkoti, 2008; Urry, 1997a). Functionally, ELPs have a wide range of
applications: 1) they form temperature sensitive drug conjugates that promote
tumor accumulation upon cycling the tumor temperature around the phase
transition (Dreher et al., 2007); 2) ELP block copolymers can assemble
nanoparticles that may encapsulate drugs (Figure. 3A) (Dreher et al., 2008b);
and 3) ELP tags can be used to purify fusion proteins (Mackay and Chilkoti,
2008). Dreher and coworkers demonstrated that ELPs have a 2-fold increase in
tumor concentration under local hyperthermia (Dreher et al., 2007). ELP-DOX
conjugates also have demonstrated cellular uptake into low pH compartments,
which is enhanced by the thermally triggered phase transition (Figure. 3B)
(Dreher et al., 2003). The cytotoxicity of the ELP-DOX conjugate was also found
22
to be nearly equivalent to free DOX. It was proposed by both Dreher (Dreher et
al., 2007) and Bidwell (Bidwell et al., 2007a; Bidwell et al., 2007b) that the
mechanism of cytotoxicity differs from that of free DOX because the conjugate
does not localize into the nucleus as rapidly as free DOX.
Another strategy has been developed that combines a ligand-mediated
approach with a thermally responsive ELP (Bidwell et al., 2007b). This system
employs a cell penetrating peptide derived from the HIV trans-acting
transcriptional activator (TAT) to promote cellular uptake. TAT is an 86 amino
acid peptide that enters cells when introduced to the surrounding media;
furthermore, a short peptide from TAT is necessary and sufficient to impart this
behavior to other proteins via a non-specific uptake mechanism (Table 1) (Wadia
and Dowdy, 2005). An ELP linker is included between the DOX and TAT that
confers thermal sensitivity to the conjugate. The optimized conjugate had a
transition temperature of 40
o
C. Incorporation of TAT increased the binding by 3-
fold at 37
o
C and by 6-fold at 42
o
C, as compared to the same conjugate without
TAT. Cytotoxicity followed a trend similar to that observed for uptake, suggesting
that hyperthermia-responsive peptides may synergize with other peptide-ligand
approaches.
Other examples of temperature sensitive peptides have been derived from
silk and collagen repetitive motifs. Silk-like polypeptides (SLPs) are repeats of
[GAGAGS]
n
(Hart and Gehrke, 2007), though [GA]
n
, and [AA]
n
are capable of
similar behavior. SLPs form irreversible, aggregates on exposure to elevated
temperatures, which can be used for designing depot systems for drug release
23
(Table 2). The properties of these SLPs can be modulated by synthesizing block
copolymers with other motifs, such as ELPs (Mackay and Chilkoti, 2008). SLP’s
When mixed with gelatin SLP’s form a thermo-sensitive gel, which can be
stabilized at 37 ºC (Gil et al., 2005).
Collagen is a peptide polymer made from repetitive sequences that follow
a trend of [GPP
OH
]
n
(Hart and Gehrke, 2007; Mackay and Chilkoti, 2008)
.
The
second and third member of the collagen triplet can be nearly any amino acid;
however, there is a strong bias for proline in the second position and
hydroxyproline in the third position (Table 2). Collagen peptide motifs irreversibly
dissociate on heating, while cooling forms a gelatin hydrogel via crosslinking
between neighboring polypeptides (Hart and Gehrke, 2007). The stability of
collagen-like peptides (CLPs) depends on the degree of post-translational
conversion to hydroxyproline in the peptide; furthermore, bacteria lack the
required proline hydroxylase necessary to generate native collagens (Hart and
Gehrke, 2007). Both collagen and denatured collagen are widely used in the
pharmaceutical and food industries as gelling agents (Hart and Gehrke, 2007).
Having introduced, ELPs, CLPs, and SLPs, it is worth noting that a wide array of
block copolymers have been evaluated that bring elements of two more of these
peptide families together. The most studied thoroughly examined combination
are the SLP-ELP hybrids, which have shown potential for use in local thermal
delivery (Megeed et al., 2002).
The development of hybrid polymers between peptides and synthetic
polymers is also an active area for exploration. For example, dual-temperature
24
sensitive peptide polymers have been synthesized by Stoica (Stoica et al., 2008)
that utilize N-isopropylacrylamide-co-acrylamide (NIPAAM) conjugated to an
octapeptide ‘FEFEFKFK’. The ‘FEK16’ motif forms a hydrogel at room
temperature (Table 2). This conjugate has a NIPAAM-mediated LCST at 30ºC
and a FEK16 induced gel melting temperature at 75 ºC (Collier et al., 2001). Due
to their glutamic acid residues, both of the above polymers are pH sensitive, and
their properties can further be modified to release the encapsulated contents in a
pH dependent manner.
25
Figure 3: Strategies for temperature-mediated peptide targeting. The
application of local and regional hyperthermia provides an opportunity to change
the carrier behavior in the tumor region. (A) Temperature-directed assembly from
single peptides into multivalent nanoparticles. The multivalency may improve the
avidity for cell surface targets that promote cellular internalization (Dreher et al.,
2008a). (B) For peptides that transition from soluble to aggregated states under
hyperthermia, the sustained application of heat increases the accumulation of the
carrier in the tumor vasculature. Upon return to normal temperatures, the
dissolution of peptides produces a high local concentration that drives
extravascular accumulation of carrier, a ‘thermal-pump’ (Dreher et al., 2007).
26
Table 2: Thermally-responsive peptides
Peptide *Motif trend Properties References
Elastin-like
polypeptides
[aPGbG]
n
a = I , V
b ≠ P
n = 10 to 200
Reversible phase
separation above
adjustable transition
temperature, which
depends on
concentration, length ‘n’,
‘a’ and ‘b’, ionic
concentration.
Accompanied by a
change from a random
coil to β-turn spiral
conformation.
(Bidwell et al.,
2007a; Karle and
Urry, 2005; Mackay
and Chilkoti, 2008;
Meyer et al., 2001b;
Urry, 1997a)
Leucine
zippers
[abcdefg]
n
a-g ≠ P.
a = hydrophobic
d = L
e,g = charged
n = 5,6
Forms coiled-coil α-helical
structures in solution that
disassociate upon heating
to form random coil
structures.
(Mackay and
Chilkoti, 2008; Mart
et al., 2006; Petka et
al., 1998)
Silk-like
peptides
[GAGAGS]
n
n = 2 to 168
Irreversibly forms
aggregates on exposure
to high temperature.
Accompanied by the
formation of β-sheets.
(Gil et al., 2005;
Mackay and Chilkoti,
2008; Megeed et al.,
2002)
Collagen-like
peptides
[Gab]
n
a = often P
b = often P
OH
n = 100 to 500
Irreversibly dissociate on
exposure to high
temperature. Form triple
helices in solution but
dissociate to form fibrils
on heating.
(Hart and Gehrke,
2007; Long et al.,
1993; Mackay and
Chilkoti, 2008)
FEK16
peptide
[FEFEFKFK]
2
An indirect temperature
sensitive system. Ca
2+
dependent assembly of β-
sheet structures when
stimulated using
temperature responsive
liposomes.
(Collier et al., 2001;
Stoica et al., 2008)
*A=Alanine, E=Glutamic acid, F= Phenylalanine, G=Glycine, I=Isoleucine,
K= Lysine, L=Leucine, P=Proline, P
OH
=Hydroxyproline Q= Glutamine,
S=Serine, V=valine.
27
1.2.4. Redox-mediated targeting
Similar to pH mediated targeting, redox-mediated targeting can potentially be
triggered extracellularly or intracellularly. The extracellular oxidative potential is
maintained by redox-modulating proteins including NADPH oxidase, superoxide
dismutase (SOD), and thioredoxin (TRX-SH2)/oxidized thioredoxin (TRX-SS),
and free glutathione (GSH)/glutathione disulfide (GSSG) (Chaiswing et al., 2008;
Meister and Anderson, 1983). These enzymes are frequently used as markers to
study redox stress in cancer. For example, TRX levels in blood are markers for
oxidative stress in hepatocellular carcinoma (Miyazaki et al., 1998). In contrast,
the ratio of reduced glutathione to oxidized glutathione disulfide (GSH/GSSG) is
used to measure the intracellular redox state, where reduced glutathione is
abundantly available in the cell (Schafer and Buettner, 2001). Under normal
homeostasis, reactive oxygen species are kept in check by glutathione (GSH)
and superoxide dismutase (SOD)*. The intracellular levels of GSH in normal
tissues ranges from 1-10 mM (Hassan and Rechnitz, 1982) compared to that of
blood plasma which is 2μM (Jones et al., 1998). This makes GSH a reasonably
good target for intracellular delivery. In vitro the GSH levels in tumor cells were
shown to have a 7-10 fold increase in GSH concentrations* (Russo et al., 1986).
This combination of intracellular elevated GSH and the tumor-associated GSH
make redox triggers interesting candidates for control over peptide structure and
drug release (Ilangovan et al., 2002; Russo et al., 1986). Conversely SOD is
frequently found at reduced concentrations in tumors compared to normal
tissues*. Another issue observed with high levels of GSH is that it promotes
28
cancer growth and resistance to chemotherapy (Balendiran et al., 2004). Based
on the elevated GSH levels intracellular release of small molecules is possible by
using peptides stabilized by disulfide bonds (Saito et al., 2003). The disulfide
bond is broken down into two sulfhydryl moieties and the GSH is oxidized to
GSSG, which can be achieved either by cell surface protein disulfide isomerase
(PDI) or reductive cleavage in the cytoplasm (Figure. 4) (Saito et al., 2003).
Using non-peptide systems, multiple groups have demonstrated that the
glutathione couple has potential anticancer application (Carlisle et al., 2004;
Hong et al., 2006; Kommareddy and Amiji, 2005, 2007; Neu et al., 2007; Wang et
al., 2006). Also disulfide linkers can moderately increase carrier stability
(Kakizawa et al., 2001). Work on redox mediated polymers dates back to the
1970’s using early polyplexes between DNA and poly(l-lysine) (Laemmli, 1975).
Polymers ranging from chitosan to poly (ethylene imine) (PEI) have been studied
for redox-mediated targeting through disulfide linkages (Laemmli, 1975);
furthermore, redox targeting may be complementary with cystine peptide based
targeting strategies. Despite the potential for stabilizing/destabilizing peptide
secondary structures via redox targeting, relatively few studies have been
reported that utilize redox sensitive peptides with chemotherapeutic drug
delivery; furthermore, this may be an area of future opportunity.
29
Figure 4: Strategy for redox-mediated peptide targeting. These approaches
target the intratumoral ratio of GSH to GSSG. The disulfide bond confers redox
sensitivity peptide carriers. (A) The extracellular concentration of GSH is very low
but cleavage of disulfide bonds can be facilitated by the presence of membrane
protein disulfide isomerase (PDI) which can lead to extracellular release of drug
(Jiang et al., 1999). (B) The disulfide bond is cleaved by intracellular reductive
activity after cellular uptake. A limitation of this approach is the stability of the
disulfide linkage during circulation in the bloodstream (Ishida et al., 2001).
30
1.2.5 pH-mediated targeting
pH mediated triggering may be achieved by sensitizing nanocarriers to the
extracellular pH or the endosomal/lysosomal pH. The environment around the
tumor is at a lower pH than compared to normal tissue (Gerweck and
Seetharaman, 1996). Production of lactic acid and hydrolysis of ATP are the
major contributors to the acidic environment around the tumor (Tannock and
Rotin, 1989). In certain extracellular regions, tumors have a lower pH (~ 6.5) than
the blood (~ 7.4) (Gerweck and Seetharaman, 1996). Membrane ion transporters
are responsible for maintaining intracellular pH (Tannock and Rotin, 1989);
therefore, the cytoplasmic pH of tumor cells is similar to that for normal cells.
Also, similar to normal cells, the endosomal/lysosomal compartments within
tumor cells also have low pH that can be exploited to design pH sensitive
delivery systems to release drugs (Figure. 5). Within the tumor, acidic drugs with
pKa below the environmental pH or basic drugs with pKa above the
environmental pH are ionized and do not penetrate lipid bilayers as efficiently as
the uncharged species. To promote penetration across lipid bilayers, weakly
acidic drugs with pKa’s in the range of 4.5-6.5 are optimal (Gerweck et al., 2006).
Similarly, acidic and basic moieties can be used to promote pH dependent
accumulation or drug release from peptides, polymers, and nanocarriers (Figure.
5A-B). In order to achieve pH-mediated intracellular drug delivery, the carrier
must first be taken up into cells, as described previously (Figure. 2) (Har-el and
Kato, 2007). Cellular binding and uptake can be promoted using specific or
nonspecific ligands, including peptides (Table 1). Following internalization,
31
trafficking to low pH compartments can then be used to trigger drug release. The
early endosome has a pH of 6 (Murphy et al., 1984); furthermore, the pH
continues to decrease during trafficking as lysosomes fuse with the vesicles to
activate proteolytic degradation (Har-el and Kato, 2007).
Although widely explored, extravascular acidity has been difficult to target
in humans, partly due to the buffering capacity of the blood and the
inhomogeneous nature of the extravascular pH gradient. The extravascular pH in
tumors can range between 6.5 and 7.0; however, drugs and drug carriers must
diffuse on the order of 50 μm away from capillaries to reach regions with this pH
(Helmlinger et al., 1997). For a small molecule or drug, this penetration depth is
easily achieved; however, for macromolecules and nanoparticulate drug carriers
these distances may be unrealistic (Dreher et al., 2006). Alternatively, in regions
where the extracellular pH remains normal (~7.4), pH dependent mechanisms
that respond to cellular uptake have been successful. Cells within tumors
continuously sample molecules and particles from their environment, which are
often trafficked to low pH compartments for degradation (Dreher et al., 2003);
furthermore, the incorporation of pH-mediated strategies that detect cellular
internalization reduces the need to target extracellular drug release. While this
review focuses on peptides, many of the peptide-mediated delivery approaches
are developed in combination with lipids or synthetic polymers; therefore, we will
briefly summarize the pH responsive behavior of non-peptide systems.
Liposomes are one class of nanocarriers that have been extensively
modified with peptides and proteins to improve the delivery of their contents.
32
Liposomes share with many other nanocarriers an internalization pathway that
results in endosomal/lysosomal entrapment and degradation (Drummond et al.,
1999). To promote release of components into the cytoplasm, pH-sensitive
liposomes have been optimized that change phases between pH 7.4 and 5.0
(Straubinger et al., 1985). One way to achieve this is by preparing liposomes
from phosphatidyl ethanolamine (PE) lipid. In the absence of a stabilizing
amphiphile, PE lipids do not form a stable lamellar phase liposome at pH 7.4, but
instead adopt a hexagonal aggregate structure (Gruner and Jain, 1985; Harper et
al., 2001). Upon incorporation of amphiphiles, including peptides, polymers, or
charged lipids, into the bilayer PE liposomes can be stabilized by either steric or
electrostatic repulsion. Following cellular internalization and lysosomal trafficking,
the reduction in pH induces loss of stabilization that drives a transition from
lamellar to hexagonal phase and drug release (de Oliveira et al., 2000). The pH
sensitivity of this behavior is a function of the amphiphile identity, MW, and
concentration in the bilayer. Alternatively, Sudimack and coworkers described
approaches to prepare pH sensitive liposomes that do not rely upon hexagonal
phase lipids (Sudimack et al., 2002). For example, polymer-lipid conjugates can
fuse with membranes, such as the mixture of phosphatidylcholine (PC) with
succinylated poly(glycidol) (Kono et al., 1994). These formulations made from
mixtures of cationic/anionic lipid combinations were found to be efficient vehicles
for intracellular delivery. An alternative strategy to generate anionic pH-
responsive liposomes is to prepare vesicles from a mixture of diolein/cholesterol
hemisuccinate (6:4) (Guo et al., 2002). These formulations are generally stable at
33
physiological pH, but aggregate and release the encapsulated contents when the
pH decreases to around 5.0. Similarly, peptides that provide steric or electrostatic
stability to lipid vesicles may be useful for promoting drug release (Figure. 5B).
pH sensitive triggering is possible with many of the peptides that are temperature
sensitive, such as ELPs, leucine zippers, and CLPs (Table 2). Similar to how
temperature can drive these polypeptides to undergo conformational changes,
the protonation of acidic and basic peptide residues can also shift the peptides
from one conformation to another; therefore, many of these peptides can also be
redesigned with pH sensitivity. For example, leucine zippers are stable coiled coil
structures at low temperatures (Table 3) (Mackay and Chilkoti, 2008).
Substitution of an acidic group at the e and g position influences the stability of
the secondary structure, and protonation of these residues is a dynamic way to
control the stability of the complex. This property has been exploited by Stevens
(M. M. Stevens, 2004) and Ryadnov (Ryadnov et al., 2003); furthermore, a
similar principle applies to ELPs (Urry, 1997a).
The GALA peptide is another interesting pH-mediated system (Table 3).
GALA is composed from tetrameric repeats of [EALA] (Li et al., 2004b), which
under physiological pH adopts a random coil structure. When exposed to lower
pH (pH=5.7), the glutamic acid residues begin to neutralize, which enable the
formation of α-helical structures. By incorporating GALA into a liposome,
assembly of these helical structures results in pH dependent membrane
disruption (Figure. 5C) (Li et al., 2004b). Many variations on the GALA peptides
have been developed, such as YALA and GALAdel3E (D.H. Haas and R.M.
34
Murphy, 2004). Most of these have appreciable activity of a similar magnitude;
however, they can be optimized to transition at different pH (D.H. Haas and R.M.
Murphy, 2004). One variation of GALA is the cationic peptide, KALA, which has
lysine in place of glutamic acid (Wyman et al., 1997). The positive charge on the
peptide interacts with oligonucleotides and also promotes membrane disruption
and gene delivery (Li et al., 2004b).
Polyhistidine blocks are a simple example of pH responsive peptides
(Table 3). The poly(L-histidine) polymers synthesized by Asayama and
coworkers (Asayama et al., 2007) are sensitive to the endosomal/lysosomal pH.
The polymer is poorly soluble at physiological pH, but upon protonation in the
endosomes the peptide becomes water-soluble. Modifying the poly(L-histidine)
by carboxymethyl substitution on the polymer leads to the formation of anionic
charge (Asayama et al., 2008). Coating of this polymer to the PEI/DNA complex
forms a ternary complex that enhances gene expression by 300-fold;
furthermore, this peptide can deliver and express DNA in the nucleus, which has
been shown to inhibit cell division (Asayama et al., 2008).
35
Figure 5: Strategies for pH-mediated peptide targeting. (A) Weakly basic
peptides that protonate at endo-lysosomal pH may promote the lysosmotrophic
release of contents into the cytoplasm (Sonawane et al., 2003). (B) Vesicles with
surface-associated peptides may be ruptured via pH-dependent change in
peptide fold. Membrane disruption may release drug into endosomes, which can
then traffic into the cytoplasm. (C) Vesicle encapsulated peptides may undergo a
conformational change at endo-lysosomal pH, forming pores. These peptides
may rupture the membranes of both the carrier and endosomes, promoting
cytoplasmic drug release (Goormaghtigh et al., 1991; Subbarao et al., 1987).
36
Table 3: pH-responsive peptides.
Peptides *Motif trend Properties References
Leucine
zippers
[abcdefg]
n
a-g ≠ P.
a = hydrophobic
d = L
e,g = charged
n = 5,6
Acidic groups at ‘e’
and ‘g’ induce the
formation of a rigid
coiled coil structure
at pH 4.5; but on
increasing the pH (7-
11) the structure
disassembles.
(M. M. Stevens,
2004; Mackay and
Chilkoti, 2008;
Mart et al., 2006;
Stevens et al.,
2004)
Carboxymethy
l poly(l-
histidine)
[HOOCH
2
C-His]
n
Cationic peptide
polymer contains an
imidazole ring and
carboxymethyl group,
which helps in
conferring dual
functionality to the
polypeptide.
(Asayama et al.,
2007; Asayama et
al., 2008)
GALA peptide WEAALAEALAEALA
EHLAEALAEALEALA
A
Changes
conformation from a
random coil to an
amphipathic α-helix
when pH is lowered
from 7 to 5.
(D.H. Haas and
R.M. Murphy,
2004; Li et al.,
2004b; Mart et al.,
2006)
KALA peptide WEAKLAKALAKALA
KHLAKALAKALKAC
EA
Changes
conformation from a
random coil to an
amphipathic α-helix
when pH is increased
from 7 to 5.
(Wyman et al.,
1997)
*A=Alanine, E=Glutamic acid, H=Histidine, K= Lysine, L=Leucine,
P=Proline, W=Tryptophan.
37
1.3. Discussion
The triggers mentioned above are being explored as parts of tumor-targeted
delivery strategies. To complicate matters, most of these peptide-mediated
delivery strategies are difficult to classify under a single strategy. Instead, these
drug-carrier formulations are multifunctional, utilizing two or more targeting
strategies. Each combination of strategies has its advantages and
disadvantages. Some of the major limitations of the above mentioned strategies
include: 1) decreased sensitivity of receptor to the ligand; 2) elimination of
attached ligand during circulation in vivo; 3) off-target interactions; 4)
immunological response to foreign antigens; and 5) limitation of drug-conjugate
linker strategies. Using ligands to direct accumulation can lead to off target
effects since other tissues may express the receptors that bind to the specific
motif/ligand. When combined with cytotoxic chemotherapeutics, this could
produce toxic effects and even decrease the therapeutic effect of the carrier. In
addition, to overcome immunogenicity, it will become important to select peptide
sequences carefully, to minimize the number of antigenic epitopes, and perhaps
to use steric shielding to reduce the immuno-recognition where appropriate
(Torchilin et al., 2001). PEG shielding of protein therapeutics was shown to
markedly decrease the immunogenicity of the formulation and increase
circulation times (Veronese and Pasut, 2005). One of the drawbacks of PEG
modification is a loss of protein/peptide activity; therefore, the potential
improvements provided by polymeric modification need to be balanced with
decreases in activity (Veronese and Pasut, 2005). One exciting possibility is that
38
certain peptide sequences may have some of the properties of PEG;
furthermore, peptides of this nature could be integrated seamlessly into
genetically engineered constructs.
The tumor microenvironment provides multiple cues that may be exploited
to improve the efficacy of established chemotherapeutics; furthermore,
polypeptides are uniquely situated to capitalize on these signals. Peptides
provide: 1) a rich repertoire of biologically specific interactions to draw upon; 2)
multiple environmentally-responsive phase behaviors that can respond disease
signatures; 3) multiple opportunities to direct self-assembly; 4) extensive control
over routes of biodegradation; 5) the ability to seamlessly combine functionalities
into a single polymer or particle using biosynthesis. These opportunities are
significant, and it is reasonable to expect that peptides will play a major role in
the development of the next generation of environmentally responsive drug
carriers.
1.4. Conclusions
Peptides provide a level of molecular specificity that is naturally suited to the
development of environmentally responsive drug carriers. Most importantly,
peptide secondary and tertiary structures enable a degree of control and
functionality that surpasses what is easily achievable using lipids and
nonbiological polymers. Here we have discussed some of the aspects of peptide-
mediated drug delivery and present applications where peptides can be directed
by temperature, pH, and specific biological interactions. As the development of
39
cancer-targeted nanocarriers continues to expand, peptides are providing these
formulations with critical functionalities necessary to target disease.
40
Chapter 2
Elastin-like peptide-amphiphiles form nanofibers with tunable length.
2.1. Introduction
Peptide amphiphiles (PAs) have unique properties that arise from the
combination of their secondary structure and their ability to self-assemble via
hydrophobic interactions(Cui et al., 2010). Like other amphiphiles, PAs have
distinct hydrophobic and hydrophilic regions that drive nanostructure formation,
similar to those formed by detergents or phospholipids (Geng et al., 2005; Kim et
al., 2005; Paramonov et al., 2006a; Simone et al., 2009). The nature of the
hydrophobic region can be either peptide based or, more commonly, one or two
saturated lipid chains. Some of the most interesting observations of peptide
amphiphiles include their propensity to form cylindrical nanostructures
(Hartgerink et al., 2002; Hartgerink et al., 1996). Due to their unique self-
assembling properties, PAs are being explored as potential drug carriers and
substrates for tissue engineering (Huang et al., 2010; Song et al., 2003; Yu et al.,
1996). Because they are generated using peptides, PAs may be engineered to
be biodegradable and bioresponsive.
A host of peptide sequences are known to be responsive to their
environment. There are peptide sequences that are thermo-responsive, pH-
responsive, reductive-responsive, ion-responsive, and even responsive to
analytes (Aluri et al., 2009). To develop PAs that are responsive to their
microenvironment, we have focused on building PAs using peptides that display
an inverse phase transition temperature. For example, elastin-like polypeptides
41
(ELPs) are environmentally responsive peptides derived from the human protein,
tropoelastin (Urry, 1997a). Similar to some synthetic polymers, including Poly(N-
isopropylacrylamide) (pNIPAM), ELPs undergo an inverse phase transition that is
analogous to the lower critical solution temperature (LCST) observed for
synthetic peptides. ELPs consist of the motif (Val-Pro-Gly-Xaa-Gly)
n
, where the
identity of Xaa and n can be adjusted to control the phase behavior of these
peptides(Urry, 1997a). During the aggregation process, ELPs form additional
secondary structures, which have been attributed to β-turn spirals (Urry, 1997a).
The phase transition temperature of an ELP depends strongly on Xaa and n;
furthermore, most ELPs with LCST behavior have been reported with n between
20 to 200 repeat units (Fig. 6a).
Here we describe a new type of elastin-like peptide-amphiphiles (ELPAs)
composed from short ELPs (n=3), which are unable to phase separate. Our
surprising finding is that the addition of two saturated 16-carbon chains enables
these short oligopeptides, dpX3 (Fig. 6b) to phase separate at temperatures
similar to their higher molecular weight counterparts, A192 (Fig 6a).
Dipalmitoylation (vs. monopalmitoylation) was evaluated to promote co-
formulation with traditional phospholipids and to provide a stronger driving force
for hydrophobic assembly. To the best of our knowledge, this is the first report of
ELPAs; furthermore, this class of peptides have the following unique properties
that may promote their in vivo applications: i) they are charge neutral; ii) they
display temperature dependent phase separation (Fig. 6b), and iii) their aspect
ratio can be controlled by adjusting the ratio of a ‘capping’ lipid. In addition, we
42
present evidence that suggests that both nanofiber assembly and phase
separation are accompanied by significant alterations in peptide secondary
structure. We also demonstrate formulation of paclitaxel using ELPAs.
43
Figure 6: Design of ELPAs with LCST behavior. (a) A high molecular weight
ELP, A192. (b) A low molecular weight ELPA construct modified with a dialmitoyl
lysine, dpA3. The dipalmitoyl lipids drive assembly of cylindrical micelles with
inverse phase transition behavior.
44
2.2. Materials & Methods
2.2.1. Materials
1,2-Dioleoyl-sn-glycero-3-phosphoethanolamine (DOPE) was purchased from
Avanti polar lipids (Birmingham, AL). Rink amide MBHA resin,
Fluorenylmethyloxycarbonyl Valine (FMOC-Val), FMOC-Proline (FMOC-Pro),
FMOC-Glycine (FMOC-Gly), FMOC-Alanine (FMOC-Ala), FMOC-Isoleucine
(FMOC-Ile), O-Benzotriazole-N,N,N’,N’-tetramethyl-uronium-hexafluoro-
phosphate (HBTU), N-Methyl-2-pyrrolidone (NMP), and Acetonitrile (ACN) were
purchased from EMD chemicals (Gibbstown,NJ). Kaiser test kit (ninhydrin
assay), DiFMOC-Lysine (DiFMOC-Lys), Diethyl ether (Ether), Triisopropyl silane
(TIPS), N,N-Diisopropylethylamine (DIPEA), Chloroform (CHCl
3
), trifluoroacetic
acid (TFA), Methanol (MeOH), Ethanol (EtOH), 4-(2-hydroxyethyl)-1-
piperazineethanesulfonic acid (HEPES) and Palmitic acid purchased from Sigma
Aldrich (St. Louis, MO). 4-Methyl piperidine (MeP) was purchased from Fisher
Chemicals (Tustin, CA). The Waters C-18 and YMC C4 reverse phase semi prep
columns were purchased from Waters Inc. (Milford, MA). 4',6-Diamidino-2-
Phenylindole, Dihydrochloride (DAPI) and 1,1′-Dioctadecyl-3,3,3′,3′-
Tetramethylindocarbocyanine Perchlorate (DiI) were purchased from Invitrogen
(Grand Island, NY). Deionized water was obtained using a Barnstead purification
system (Asheville, NC) and used for all aqueous buffers.
2.2.2. Synthesis and purification of ELP biomaterials
Short sequences with three pentameric ELP repeats were synthesized using
standard FMOC chemistry with Ala (A3), Val (V3) and Ile (I3) as guest residues
45
(Fields et al., 1991). Briefly, 500 mg (0.6 meq/g substitution) of Rink amide
MBHA resin was allowed to swell in NMP for 30mins. The resin was deprotected
using 25% MeP. The amino acid coupling was performed for 45mins with 5 times
excess amino acid and HBTU under basic conditions. The generation and
consumption of free primary amine is monitored using the ninhydrin test. The
deprotection and coupling sequence is continued until the peptide of desired
length is obtained. The N-terminal is capped with DiFMOC-Lys (KA3, KV3 and
KI3) and the protecting groups are removed. A sample was collected for
qualitative analysis.
The deprotected N-termini were treated with molar excess of palmitic acid
under basic conditions to give dipalmitoylated ELP sequences (Fig. 6b). The
reaction was performed until all primary amines are exhausted. After completion
the conjugates (dpA3, dpV3 and dpI3) were cleaved using a cleavage mixture
made of 95% TFA in the presence of TIPS for 3 hrs at room temperature. The
cleaved product is precipitated using cold ether. The crude product was
dissolved in 100% HPLC grade MeOH.
The crude ELP sequences and lipid conjugates were qualitatively
analyzed at 214nm on a Perkin Elmer 200 series high performance liquid
chromatography (HPLC) system using a Waters C-18 and YMC C-4 reverse
phase column respectively. The ELP sequences were run on a H
2
O:MeOH
gradient with MeOH starting from 30% to 70%. The lipid conjugates were run on
a H
2
O:MeOH gradient starting at 80% to 100% MeOH. The purified fractions
were collected and mass confirmed using a DECA-LcQ ESI mass spec system
46
(Thermo scientific, Waltham, MA). After confirmation the products were purified
and lyophilized. Stocks of the lyophilized samples were made in 100% HPLC
grade MeOH.
2.2.3. Determination of critical micellar concentration (CMC)
CMC was calculated by the pyrene assay (Goddard et al., 1985). Briefly, 0.6
μmoles of pyrene was added to round-bottom test tubes and dried under a
stream of air. Increasing concentrations of the ELPA were added to the dried
pyrene and vigorously vortexed until pyrene dissolves. The fluorescence intensity
of the samples were analyzed at λ
ex
at 334 nm (slit width= 8 nm) and λ
em
recorded between 350 and 410 nm (slit width = 2 nm) using a Horiba Fluorolog
®
3 spectrofluorometer (Edison, NJ). The ratio of 373/384 (L1/L3) was calculated
and plotted against concentration. The concentration at which the slope change
is maximum (break point) is the CMC. All ELPA samples were prepared in
filtered PBS.
2.2.4. Preparation and particle size of ELPAs
All ELPAs were prepared using film hydration with an aqueous solution. Briefly,
the required amounts of ELPA stocks were mixed in a clean glass test tube and
evaporated on a Heidolph Laborota 4011 (Schwabach, Germany) to form an
even film. For cell binding/uptake studies, 0.2% mol of DiI was added to the film.
The film was placed under vacuum for 4 hrs to remove trace organic solvent. The
dry film was then hydrated with appropriate buffer and placed in a water bath at
75 ° C for 2 mins. The hydrated film was sonicated f or 5-10 secs and vortexed
gently for an additional 10-15 secs to ensure complete hydration of film. After
47
complete hydration, 50 μl of the sample was analyzed on a Wyatt Dynapro plate
reader (Santa Barbara, CA) using a 384 well Greiner Bio-One clear bottom plate
(Monroe, NC). Samples for light scattering measurements were hydrated in 10
mM HEPES buffered saline.
2.2.5. LCST determination of ELPAs
The temperature dependent phase transition of ELPAs was demonstrated using
optical density measurements at 350 nm over a set temperature range.
Increasing concentrations of constructs were added to 300 μl Beckman coulter
Tm microcells (Brea, CA) and the temperature was ramped at a rate of 1° C/min.
The optical density was plotted as a function of temperature, and the maximum
first derivative of this curve was defined as the phase transition temperature. All
ELPA samples were prepared in filtered PBS
2.2.6. Secondary structure determination using circular dichroism (CD)
CD was performed to determine the secondary structure adopted by the
synthesized constructs at different temperatures. Increasing concentrations of
the constructs were run on a Jasco J-815 CD spectrometer (Easton, MD) using a
quartz cuvette (path length ~ 1 mm). The ellipticity was monitored from 180-300
nm and the spectra of buffer subtracted post run. All the constructs were
prepared in filtered diH
2
O. All plots were exported using Jasco Spectral manager
V2 (Easton, MD) to Excel and deconvoluted assuming that the observed molar
ellipticity [θ] is a weighted linear sum of the ellipticity for known secondary
structures (Eq. 1). The data was fit using nonlinear regression on Microsoft
Excel.
48
….…(Eq. 1)
Where,
θ: Observed ellipticity
θ
std
: Ellipticity of standard
C
std
: Fraction of standard
2.2.7. Transmission electron microscopy (TEM)
A small drop of freshly prepared ELPA was pipetted onto a Ted Pella
carbon/formavar grid (Redding, CA). The excess liquid was wicked off using filter
paper and a drop of 1 % uranyl acetate was pipetted on. The excess uranyl
acetate was wicked off using filter paper and the grid was dried at 37 ° C. The grid
was carefully placed into a JEOL JEM 2100 laB6 microscope (Tokyo, Japan). All
samples were run using an accelerating voltage of 200 kV and processed using
ImageJ (NIH, USA). The end-to-end measurements of the ELPA nanofibers was
first made by setting the scale using the scale bar on the raw image. Once the
scale was set the length was measured using the ‘straight line’ function on
ImageJ.
2.2.8. Atomic force microscopy (AFM)
AFM was used as an alternate approach to control for procedural artifacts on the
TEM. Briefly, 2 μl of sample was pipetted onto a freshly stripped flat mica surface
and dried under a steady stream of air. The mica plate was then analyzed using
a tapping mode Veeco Nanoscope IIIa Atomic Force Microscope (Santa Barbara,
CA). Bruker AFM probes (Camarillo, CA) with a spring constant (k) between 20-
80 N/m and a frequency (f
o
) of 320-357 kHz were used. All samples used were
prepared in filtered diH
2
O to avoid formation of salt crystals. After image
49
acquisition the files were exported as jpegs and analyzed using ImageJ (NIH,
USA) using previously described procedure.
2.2.9. Cell binding/uptake studies
Cell uptake studies of ELPAs were performed on human embryonic kidney cells
(HEK 293) and cervical cancer cells (HeLa). ELPA samples for cell uptake
studies were prepared in 0.2 μm sterile-filtered diH
2
O. Glass coverslips (22mm X
22mm) were seeded with 3 x 10
5
cells and cultured in DMEM supplemented with
10% FBS until they were >80% confluent. Appropriate quantities of the
nanoparticles were added to the cells and incubated at 37˚C at 5% CO
2
for 5 hrs.
DAPI was added 2 hours prior to imaging. After incubation, coverslips were
washed with PBS and mounted onto glass slides using Fluoromount
®
(Sigma
Aldrich). The slides were allowed to dry and imaged on a Zeiss LSM 510
confocal microscope using appropriate settings for fluorophore used (DiI λ
ex
=543
nm, DAPI λ
ex
= 790 nm). The pinhole sizes used for DiI and DAPI channels are
240 μm and 504 μm respectively. The gain for all the images was kept constant
at 840 (red channel) and 1100 (blue channel).
Cellular uptake was compared between cells using red pixel area. The
pixel area was calculated using ImageJ. Briefly, the red channels are exported as
TIFFs and converted to 8-bit images. The image is thresholded for contrast of red
puncta. Puncta were selected using the ROI manager and area in (Pixel)
2
measured. The measured area is then converted to μm
2
based on the
dimensions of the original image. The data was compared using ANOVA/Tukey
analysis.
50
2.2.10. Paclitaxel encapsulation and delivery studies
To test optimal loading concentrations 5, 12.5, 25, 50 and 100 μM solutions of
Paclitaxel (PAX) in methanol were added to a fixed ELPA concentration (125
μM). The methanolic PAX:ELPAs were dried down to make even films and
hydrated as previously discussed. All films were hydrated with sterile diH
2
O for
HPLC analysis. After hydration the solution was spun at 10,000 rpm for 5 mins to
remove unencapsulated PAX. The supernatant was then run on an analytical
reverse phase C-18 column using a 20% to 100% ACN (0.1% TFA) gradient and
encapsulated PAX absorbance was followed using UV-Vis spectrophotometry at
227 nm. The concentration of encapsulated PAX was determined using a
standard curve run under identical conditions.
PAX delivery by ELPAs was tested by determining HeLa cell viability.
Briefly, 5 x 10
4
cells were seeded in a 96-well plate and grown for 24hrs or until
>80% confluence was achieved. PAX encapsulated nanoparticles were added to
the cells in increasing concentrations and incubated for 48 hrs. After incubation
the cells media was replaced and cell viability determined. The cell viability was
measured by treatment with water soluble MTS based Promega CellTiter
®
96
Cell Proliferation kit (Madison, WI) according to the manufacturer’s protocol. The
plates were then analyzed at 490 nm using a Biorad Benchmark plus plate
reader (Hercules, CA). For cell viability assays all nanoparticles were prepared in
sterile filtered diH
2
O and control PAX dissolved in 20% sterile filtered DMSO. The
amount of PAX encapsulated was determined using previously mentioned
reverse phase HPLC method before cellular incubation.
51
2.3. Results
2.3.1. Synthesis and purification of ELPAs
Solid phase synthesis yielded a relatively pure crude product, which was further
purified using reverse phase chromatography to produce pure ELPAs (~97%).
Mass spectrometry was used to characterize the synthesized ELPAs (Table. 4).
2.3.2. ELPAs assemble nanofibers with a low CMC
Using the pyrene fluorescence assay, the CMC of dpA3 was estimated to be 1.7
μM (Fig. 7a,b). In contrast, dpV3 and dpI3 were not soluble as binary mixtures
with water; therefore, their CMC was not assessed. Dynamic light scattering was
used to characterize the hydrodynamic radius of dpA3. Above the CMC, dpA3
formed particles with a broad size distribution of 94.2 ± 87.6 nm (Fig. 7c) in
radius. Their radius and polydispersity suggest that the ELPAs exist as
heterogeneous particles in aqueous buffer; furthermore, these properties are not
consistent with small, spherical micelles. While the > 10 nm hydrodynamic radius
provided some evidence that peptide dipalmitoylation produces nanostructures,
TEM and AFM were necessary to define the morphology and aspect ratio for
these structures (Fig. 8a-f). TEM images (Fig. 8a) show that dpA3 assembles
very long cylindrical micelles 1.7 ± 0.2 μm (approx.) length and 16.0 ± 2.0 nm in
width (Fig. 8b,c). AFM provided independent confirmation of the nanofiber
assembly. AFM images showed dense networks (Fig. 8d,e) composed of
bundles of cylindrical micelles (Fig. 8e). The average width of each individual
strand is 13.9 ± 1.4 nm (Fig. 8f,c). The control KA3 peptide did not demonstrate
structural assembly. The lengths on the AFM could not be assessed due to
52
imaging limitations that prevented the identification of both ends of a fiber.
Regardless, both microscopy techniques confirmed that dpA3 assembles thin
fibers with very high aspect ratios.
To control the length of the dpA3 nanofibers, phospholipids were mixed
with the ELPA to determine if this approach could adjust the average aspect
ratio. This attempted to create end-caps to the fibers, under the assumption that
the ratio of capping lipids to dpA3 may provide a simple approach to modulate
the length, and possibly cellular uptake, for PA nanofibers. As a capping lipid,
phospholipid DOPE, was selected because it cannot form stable lipid bilayers in
the absence of other lipid components(Yeagle and Sen, 1986). This allows for
easy identification of ELPA fibers. Since experiments were performed in neutral
conditions the adoption of an alpha-II conformation by DOPE is unlikely. It was
intended that DOPE would be less likely to form a second population of vesicles,
which may have resulted from mixtures with phosphocholine lipids. At ratios of
4:1 dpA3 to DOPE, the size distribution of the ELPAs was brought down to 282.7
± 68.4 nm. At [1:1] ratios of dpA3 to DOPE, the size of the ELPAs was further
reduced to 71.3 ± 13.2 nm (Fig. 9a,b). The addition of capping lipids had no
effect on the width of observed nanofibers. The widths by TEM imaging for
ELPAs at 4:1 and 1:1 dpA3 to DOPE are 12.5 ± 0.9 nm and 13.2 ± 2.2 nm.
Hence, we have successfully developed a method to independently modulate the
length for ELPAs, which may have utility in the design of peptide amphiphiles
using other peptide sequences.
53
Table. 4: Physico-chemical characterization of ELPAs.
ELPA
Sequence Purity
(%)
‡
Expected Mass
(Da)
†
Observed Mass
(Da)
KA3 K(VPGAG)
3
98 [M+H
+
]= 1290.4
[M+Na
+
]=1312.5
[M+H
+
]= 1289.8
[M+Na
+
]= N.O
KI3 K(VPGIG)
3
98 [M+H
+
]= 1416.7
[M+Na
+
]= 1438.
[M+H
+
]= 1416.1
[M+Na
+
]= N.O.
KV3 K(VPGVG)
3
98 [M+H
+
]= 1374.6
[M+Na
+
]= 1396.6
[M+H
+
]= 1375.1
[M+Na
+
]= N.O.
dpA3 dp-
K(VPGAG)
3
97 [M+H
+
]= 1767.2
[M+Na
+
]= 1789.2
[M+H
+
]= 1766.7
[M+Na
+
]= N.O.
dpI3 dp-
K(VPGIG)
3
96 [M+H
+
]= 1893.5
[M+Na
+
]= 1915.5
[M+H
+
]= 1892.9
[M+Na
+
]= 1915.1
dpV3 dp-
K(VPGVG)
3
98 [M+H
+
]= 1849.6
[M+Na
+
]= 1871.6
[M+H
+
]= 1850.9
[M+Na
+
]= N.O.
‡
Mass estimate based on ELPA sequence
†
Observed mass using ESI mass spectrometry.
54
Figure 7: ELPAs self-assemble micelles with a low critical micelle
concentration (CMC). dpA3 was assessed for micelle assembly using pyrene
and dynamic light scattering. (a) Emission spectra of pyrene as a function of
dpA3 concentration. A decrease in the intensity of spectra was observed with
decreasing concentrations of the conjugate. (b) L1/L3 was plotted against
log(concentration) of the dpA3 conjugate, where L1 and L3 are the intensity at
373, and 384 nm respectively. The CMC for dpA3 conjugate was determined to
be below 1.7 μM at the position indicated by the arrow. (c) Particle size
distribution of dpA3 at 1000 μM in 10mM HEPES buffered saline using dynamic
light scattering (DLS).
55
Figure 8: dpA3 self-assembles cylindrical fibers which form dense
networks. The dpA3 (1000 μM) was stained with 1% uranyl acetate and imaged
using transmission electron microscopy (TEM). (a) Numerous fiber-like cylindrical
micelle structures 1.7 ± 0.2 nm (mean ± SD, n=3) in length were observed. (b)
Distribution of fiber lengths. (c) Distribution of fiber widths. Scale bar = 50 nm.
(d)100 μM dpA3 was dried onto a freshly stripped mica surface and imaged using
atomic force microscopy (AFM). (e) dpA3 forms a network of bundled fibers. The
bundled fibers are composed of parallel cylindrical fibers (f) The widths of the
individual fibers are consistent with the measurements obtained from TEM (n=3).
Length measurements were not obtained because the fibers extended outside
the field of view. Scale bar = 250 nm.
56
Figure 9: The length of the ELPA fibers can be controlled using
phospholipids. (a) The cylindrical fibers were shortened in length by 10-fold to
71.0 ± 13.2 nm (mean ± SD, n=3) by mixture of equimolar quantities of
DOPE:dpA3 [1:1] (b) Distribution of fiber length. (c) The distribution of fiber
widths was unaffected by the addition of phospholipid. Scale bar = 50 nm.
57
2.3.3. ELP nanofibers exhibit LCST behavior
High molecular weight ELPs like A192 (Fig. 10a), phase separate above an
inverse phase transition temperature, which is similar to an LCST. This effect is
length dependent. Short ELPs like KA3 do not phase separate below 100 ° C at
atmospheric pressure (Meyer and Chilkoti, 2004). It was anticipated that ELPA
self-assembly into fibers would form a dense peptide brush that would have
LCST properties. When heated, the ELPA nanofibers did exhibit a concentration
dependent change in the transition temperature (Fig. 10b). The phase diagram of
dpA3 (2kD) was similar to that for a high molecular weight ELP, A192 (73.2 kD),
albeit at slightly higher molar concentrations (Fig. 10c). The plain peptide, KA3,
does not phase separate in this temperature range.
2.3.4. β-Turn formation associated with approach of ELPA transition
temperature
To investigate the mechanism for ELPA phase separation, CD was used to
characterize the peptide secondary structure as a function of concentration,
dipalmitoylation, and temperature. The CD spectra were deconvoluted using the
standard curves of (a) α-helix (Greenfield and Fasman, 1969) (b) β-sheet
(Greenfield and Fasman, 1969) (c) Random coil (Greenfield and Fasman, 1969)
(d) β-turn 1 (Perczel and Fasman, 1992) and (e) β-turn 2 (Brahms et al., 1977).
Preliminary CD studies of KA3 failed to show any concentration (data not shown)
or temperature (Fig. 11) dependent change in structure. In contrast, the addition
of two palmitic chains had a very significant impact on the secondary structure of
KA3 (Fig. 11a,b). By inducing peptide assembly into nanofibers, the palmitic
58
chains promote order in dpA3. At 25 ° C, the majorit y of dpA3 exists in either a
random coil or a β-sheet conformation (Fig. 11c). While approaching the phase
transition temperature, there is an increase in the β-type-1 turn content. Above
the phase transition temperature (~60 ° C), the stru cture becomes highly
disordered like the unmodified peptide. The control KA3 peptide has CD spectra
very similar to a random coil, which has been observed previously for ELPs
below their transition temperatures(Klok, 2008). Deconvolution of the spectra
showed that the majority of the KA3 peptide exists in a random coil / β-sheet with
a small population in a β-turn 1 conformation, which remains largely unchanged
at higher temperatures (Fig. 11d). These data confirm that nanofiber assembly of
dpA3, mediated by N-dipalmitoylation, produces a moderate degree of β type-1
turn structures; furthermore, increasing temperature stabilized this structure up
until above the bulk phase transition temperature.
59
Figure 10: ELPAs (dpA3) and soluble ELPs (A192) both undergo LCST
behavior. The optical density was monitored over a temperature gradient of 1 ° C
per min from 45 to 80 ° C. (a) A soluble ELP with a high MW (73.2 kD), A192 (100
μM=7.3 mg/ml). (b) ELPA nanofibers composed from the low molecular weight
(1.7 kD) dpA3 (1000 μM=1.7mg/ml). (c) The transition temperatures for both
constructs are inversely proportional to Log of concentration for both A192 (slope
= -8.9 ± 0.8 ° C (Log10[μM])-1 (mean ± 95% CI), r
2
= 0.9975, p < 0.01) and dpA3
(slope = -7.1 ± 1.8 ° C (Log10[μM])-1 (mean ± 95% CI ), r
2
= 0.98, p < 0.01).
60
Figure 11: dpA3 forms temperature-dependent secondary structure. (a)
Circular dichroism (CD) spectra for dpA3 (250 μM, red) and KA3 (250 μM,
orange) below the transition temperature (T = 55 ° C ). The spectrum for the
unmodified KA3 (R
2
=0.9930) peptide is considerably different from the
dipalmitoylated dpA3 (R
2
=0.9922), suggesting that nanofiber assembly induces
formation of partial secondary structure. (b) Above the transition temperature for
dpA3 (T = 65 ° C) there is a noticeable change in th e molar ellipticities (red,
R
2
=0.9832); however, the unmodified peptide KA3 (orange, R
2
=0.9913) remains
unchanged. (c) At lower temperatures the majority of the dpA3 (250 μM)
population adopts either a β-sheet or random coil conformation. On heating there
is an increase in the β-turn 1 component suggesting that β-turn 1 formation is
necessary for the phase separation of this class of ELPAs. (d) The major
conformation of unmodified KA3 at all temperatures is random coil. There is no
significant change in the secondary structure at higher temperatures.
61
2.3.5. Inclusion of DOPE promotes ELPA cell uptake
We next investigated the possibility that the inclusion of other lipids might change
the physico-chemico properties of dpA3 nanofibers, with the goal of tuning
cellular uptake. To accomplish this, we evaluated multiple phospholipids in
mixture with ELPAs, which included phosphocholines (POPC, DPPC) and
phosphoenthanolamines (DOPE). Only DOPE formulations produced stable
nanofibers with shortened lengths (< 50% by mol lipid); therefore, these
formulations were evaluated for their cellular uptake in cultured HEK 293 and
HeLa cells. ELPA uptake was studied using confocal microscopy, whereby
nanofibers were labeled with 0.2% by mol lipid DiI (red). DiI is a strong, non-
exchangeable marker for stable cell membranes and requires a hydrophobic
environment to fluoresce. Hence, DiI makes an excellent hydrophobic compound
to study ELPA nanofiber-mediated cellular uptake. DiI labeled DOPC liposomes
with equal quantities of DOPE were used as negative controls for the nonspecific
uptake of neutral liposomes. The uptake of pure dpA3 nanofibers was minimal in
both HEK 293 and HeLa cells (Fig. 12a-c,m-p). In contrast, the addition of DOPE
promoted cellular uptake of ELPAs in both cell lines. The inclusion of DOPE had
an easily observable effect on cellular binding and uptake. This effect may arise
from the negative curvature of DOPE lipids, which promotes fusion of pure
phospholipid membranes (Guo et al., 2003). Alternatively, evidence that DOPE
shortens the length of ELP nanofibers (Fig. 9) from microns to below 300 nm
raises the possibility that ELPA internalization (non-phagocytic) may be tuned by
controlling fiber aspect ratio. Interestingly, the addition of DOPE to DOPC
62
liposomes also enhanced the transfer of the DiI label to cells however to a much
lesser extent than for ELPA nanofibers Unformulated; free DiI showed minimal
staining of phospholipid cellular membranes. Hence the addition of DOPE may
mediate uptake through two mechanisms: i) membrane fusion and ii) reduction of
nanofiber length. Due to the inability of non-fusogenic phospholipids (DPPC,
POPC) to form stable and short ELPA nanofibers, we are currently unable to
distinguish between these two mechanisms. Regardless, DOPE incorporation
unambiguously modulates the transfer of a model hydrophobic (DiI) compound
between ELPA nanofibers and cells.
63
Figure 12: DOPE addition promotes cellular uptake of ELPAs. Particles at a
concentration of 16.7 μM were added to HEK293 cells and incubated for 5 hrs.
DiI was used to label nanoparticles (red). DAPI (Blue) was used to stain the
nucleus. All images were obtained with the same microscope settings. Image
panels (a) to (d) and (m) to (p) show minimal uptake of dpA3 in both cell lines
respectively. Conversely, image panels (e) to (l) and (q) to (x) show uptake of
dpA3:DOPE particles which suggests DOPE influence in cellular uptake. Control
DOPC liposomes did not show any cellular uptake. Scale bars represent 5 μm.
64
2.3.6. Encapsulated PAX reduces tumor cell viability
Film hydration showed no precipitation of PAX, suggesting high efficiency
encapsulation. Using dpA3, we have shown a significant increase in PAX
solubility in aqueous solvents i.e., from <0.1 μg/ml to 85 μg/ml (100μM)(Konno et
al., 2003) . PAX encapsulation is saturated above a 1:0.8 molar loading ratio of
dpA3 to PAX (Fig. 13a). Using encapsulated PAX we demonstrate significant
inhibition of HeLa cell growth (Fig. 13b). The inhibition of cell growth is similar to
soluble unencapsulated PAX delivered in DMSO. At the concentrations studied,
the inclusion of DOPE did not enhance the efficacy of encapsulated PAX (Fig.
13b). The study was not performed at lower dpA3 concentrations due to
concerns that below their CMC, the ELPA nanofibers may disassemble. The
maximum concentration of DMSO over the cells was 10%. These results suggest
that dpA3 nanofibers may have potential as drug delivery vehicles for PAX.
65
Figure 13: dpA3 Encapsulated PAX reduces HeLa cell viability. (a) At a
constant concentration of dpA3 (125 μM), increasing concentrations of PAX were
encapsulated using thin film hydration. Analytical Reverse Phase HPLC was
used to quantify PAX, and the area under the curve (AUC) for the appropriate
peak is presented. This trace shows saturable encapsulation above 50 μM PAX
at dpA3:PAX ratio of 1:0.8; therefore, dpA3:PAX nanoparticles approach ~29%
by mass drug loading. (b) A cytotoxity assay was used to determine if dpA3 was
able to deliver PAX to cells using the MTS assay. There is no difference in
efficacy of encapsulated PAX compared to free PAX solubilized in DMSO.
66
2.4. Discussion
Here we report the formation of peptide amphiphile nanofibers of tunable length
with some properties of elastin-like polypeptides. The dipalmitoylation of KA3
yielded a PA with a low CMC, similar to that found for other dipalmitoyl
phosphatidylcholine lipids using the same pyrene assay (Kanamoto et al., 1981).
A potential advantage of this chemistry is the production of stable amide bonds
between the peptide and the lipids, which lack the ester bonds found on
phospholipids. The palmitic acid chains reduce the transition temperature to well
below 100 ˚C, allowing this approach to be used in the induction of the phase
behavior of short hydrophilic ELPs. In addition to the stability of ELPAs (low
CMC, amide linkage), we have identified several key variables that enable tuning
of the aspect ratio (capping lipids) and cellular uptake (mixing ELPAs with
different fusogenic lipids). We have also uncovered new evidence that the ELPA
phase separation is mediated by the formation of distinct secondary structure. At
low temperatures, the CD spectra for dpA3 showed sheet like structures that
were similar to those reported for other PAs(Xu et al., 2010; Yokoyama et al.,
2001). CD spectra further revealed that dpA3 phase separation correlates with
the maturation of β-type-1 turns during heating. In contrast, the unmodified KA3,
remains in a random coil conformation and shows no temperature-dependent
change in secondary structure. Intriguingly, above its phase transition
temperature (>60 ° C), dpA3 (Fig. 11c) forms a rand om coil conformation;
suggesting that an entropic change from a highly ordered secondary structure
below the transition temperature to a disordered peptide backbone above the
67
transition temperature may partly contribute to differences in the free energy of
mixing responsible for phase separation. While this possibility warrants further
study, the ability to probe peptide structure above the bulk phase transition using
CD is limited by light scattering from large aggregates.
Similar to other reported PAs, dpA3 formed fiber-like micelles above its
CMC(Chow et al., 2010; Hartgerink et al., 2001). We confirmed the assembly of
these fibers using TEM and AFM. To further modify its assembly we incorporated
fusogenic lipids in our PA. Previously, the Hartgernik group demonstrated that
phospholipid inclusion modulates the gelation properties of Pas (Paramonov et
al., 2006b); however, this approach was never demonstrated to control the length
of any PA nanofibers. Here we used the same concept to cap the nanofiber
termini and control their lengths. Our proposed mechanism is that the
phospholipids arrest the growth of the nanofiber, possibly by local phase
separation to the termini with dpA3 occupying the trunk of the nanofiber (Fig. 14).
This approach to control aspect ratio is significant because it provides a method
to tune the length of peptide amphiphile nanofibers as needed to optimize cellular
uptake (Fig.14), mechanical properties, or pharmacokinetics. At equimolar ratios
of capping lipid DOPE to dpA3, length was reduced 100 fold with almost no
change in fiber width. To the best of our knowledge, this degree of control over
PA nanofiber length has never been reported. By further increasing the capping
lipid concentration to above 50% by mol lipid, the fibers were abolished and the
major species observed were round structures (Fluck, 1966). Under our
proposed mechanism, the cylindrical trunk of the nanofiber is predominantly
68
composed of dpA3 (Fig. 10a). Increasing percentages of the capping lipid,
DOPE, reduced the average nanofiber length observed using TEM (Fig. 10b).
The addition of DOPE also had significant effect on the cellular uptake of dpA3
(Fig. 10c). The poor uptake of dpA3 may be attributed to its extreme aspect ratio
and uncharged nature. By controlling the aspect ratio, lipids like DOPE may
shorten ELPA nanofibers into shorter fibers (Fig. 9), which are more suitable for
non-phagocytic mechanisms of cellular internalization, including potocytosis or
clathrin-mediated endocytosis. Alternatively, the inclusion of DOPE may promote
membrane fusion that mimics intracellular uptake. In either case, this lipid
capping approach may also useful for polarized functionalization of ELPA termini
with cargo or targeting moieties.
In addition to the delivery of a model hydrophobic fluorophore, DiI, the
hydrophobic core of the ELPA nanofibers demonstrated high efficiency
encapsulation of the hydrophobic chemotherapeutic PAX. PAX is commonly
marketed as Taxol
®
and is dissolved in a castor oil based Cremophor EL
®
system(Rowinsky et al., 1992). Cremophor EL
®
is a non-ionic surfactant which
helps emulsify PAX in aqueous solutions. In addition to severe discomfort during
infusion, the Cremophor EL
®
system is also reported to cause allergic reactions,
which can be fatal (Liebmann et al., 1993). The pure ELPA system can serve as
an ideal alternative to such surfactants since it is non-ionic and not easily taken
up into cells (as shown in Fig. 12. Further studies may be warranted in an effort
to translate ELPA carriers from the bench to the clinic. While genetically
engineered ELPs such as A192, have many potential applications (Chow et al.,
69
2008), ELPAs have three potential advantages as drug carriers: i) they have a
truly hydrophobic core capable of physical entrapment of hydrophobic drugs like
PAX; ii) they are not recombinant products, reducing their cost and complexity;
and iii) the observation that they formed nanofibers makes them a substrate with
a very different surface compared to large soluble ELPs or spherical ELP
nanoparticles.
One potential application of ELPAs will be in the development of
nanoparticles that are synergistic with hyperthermia-mediated drug delivery
(Andrew Mackay and Chilkoti, 2008). In this manuscript, we have focused on
developing ELPAs as soluble drug carriers with a tunable fiber length; however,
future studies will be required to tune their thermal responses closer to
physiological temperature. The most direct approach to modulate ELP transition
temperature is to adjust the hydrophobicity of the guest residue from Ala to Val or
Ile. In fact, we successfully purified both compounds (Table 4); however, neither
dpV3 nor dpI3 formed soluble, stable nanofibers in the presence of water. One
plausible explanation is that, similar to dipalmitoyl phospholipids, the
dipalmitoylated ELPAs have a lipid melting temperature slightly above body
temperature. Furthermore, it is plausible that if the transition temperature of the
ELP corona is below that of the lipid core, the ELPAs are unable to reorder
themselves into stable nanofibers upon hydration. Any effort to melt the lipid core
results in bulk phase separation of the peptide corona. Having now optimized the
drug loading and tunability of nanofiber length, future studies will address the
70
lipid core melting temperature and tunability of nanofiber transition temperature
to better utilize hyperthermia-mediated targeting.
PAs are an emerging nanomaterial with applications ranging from tissue
engineering to drug delivery. Stupp and coworkers have extensively studies PA
networks for cell repair and regeneration. For example, they developed PA
systems with tunable RGD presentation which have shown promise as cell
adhesion scaffolds (Storrie et al., 2007). The same group has also shown
neuronal regeneration using laminin-derived PAs, which have potential for
treating spinal cord injuries (Tysseling et al., 2010). Apart from tissue
regeneration, PAs have found use in conventional drug delivery aspects. PAs
have been used to encapsulate hydrophobic small molecules (Cisplatin,
Nabumetone) (Kim et al., 2009; Matson and Stupp, 2011) and as contrast agents
(Bull et al., 2005a; Bull et al., 2005b) in MRI imaging. As they are bioresponsive,
uncharged, and of tunable length, we propose that elastin-like PAs provide
multiple opportunities to facilitate the successful translation of PA-derived
therapies such as those above into clinical use.
71
Figure 14: Proposed mechanism for fiber-length by capping phospholipid.
(a) Unsaturated lipid chains of DOPE promote intra-fiber phase separation to the
termini of ELPA nanofibers assembled by saturated lipids. (b) The average
nanofiber length was plotted as a function of the % of DOPE capping lipid. (c)
Plain dpA3 fibers have low propensity for cellular uptake. ANOVA analysis shows
a significant difference between the 3 groups (P=0.0188). Tukey’s post-hoc
analysis shows a significant difference between dpA3 and formulation with 20%
& 50% DOPE (P<0.05). No significant difference in uptake between 20% & 50%
DOPE was observed. Hence, addition of DOPE increases uptake irrespective of
the cell line used. Values indicate the mean ± SD (n=3)
72
2.5. Conclusion
Herein we describe a new type of peptide amphiphiles prepared by
dipalmitoylation of a non-ionic polypeptide derived from human tropoelastin.
These ELPAs form cylindrical micelles in solution, which undergo inverse phase
transition behavior similar to high molecular weight ELPs. The lengths of these
ELPA nanofibers can be tuned by inclusion of other lipids. The addition of lipids
promoted cellular uptake of described ELPA nanofibers. We also have
preliminary evidence for its application as a drug delivery vehicle. This new class
of PAs may be a useful platform to develop nanostructures with potential in
targeted drug delivery applications.
2.6. Acknowledgments
This work was made possible by the University of Southern California, the
National Institute of Health R21EB012281, The Wright Foundation, The Stop
Cancer Foundation, the USC Ming Hsieh Institute, and the Whittier Foundation to
J.A.M., P30DK048522-15 to the USC Research Center for Liver Diseases,
P30CA014089 to the Norris Comprehensive Cancer Center and the Translational
Research Laboratory at the School of Pharmacy. CD and AFM data were
obtained at USC Nanobiophysics core facility. The TEM images were obtained at
the USC Cell and Tissue Imaging Core. We extend thanks to S. Louie for access
mass spectrometry equipment in the USC Pharmacoanalytical Core Facility.
73
Chapter 3
Antibody-core protein polymer nanoworms (ACPPNs) potentiate apoptosis
better than a monoclonal antibody.
3.1. Introduction
Antibody fragments are a promising approach to the treatment of several
diseases, and possess many advantages over intact antibodies (Holliger and
Hudson, 2005; Melmed et al., 2008). Single chain variable regions (scFv) are one
type of antibody fragment which has shown significant potential for therapeutic
applications (Nelson, 2010) for several reasons including their simple expression
and modification (Weisser and Hall, 2009) significant targeting capability in drug
delivery or imaging applications (high signal to noise) (Qian et al., 2008; Weisser
and Hall, 2009), small size compared to antibodies (King et al., 1994b; Rudnick
and Adams, 2009; Weisser and Hall, 2009; Yokota et al., 1992), and relatively
short half-lives (~4 hrs)(King et al., 1994a). The short half-life can be useful in
imaging applications, but detrimental to their application as functional
therapeutics as therapeutics generally require long circulation to produce
maximal uptake (Cumber et al., 1992; Schneider et al., 2009). The half-lives of
recombinant scFv fragments can be prolonged by simple conjugation to large
molecular weight hydrophilic polymers (Müller et al., 2007; Yang et al., 2003).
Using this principle, we evaluated a novel recombinant scFv based nanoparticle
for the management of B-cell lymphomas, which not only target tumors but also
retard tumor cell growth.
74
Lymphomas have an incidence rate of 27 in 100,000 people per
year(Howlader N, 2012). Non-Hodgkin’s Lymophomas (NHL), mostly of B-cell
origin (~85%), are the most prevalent with around 70,000 patients expected to be
diagnosed in 2012 (Howlader N, 2012). Therapy against NHL includes the use
of a regimen of combination chemotherapeutics (ex. CHOP, ABCVP) with or
without the antibody Rituxan® (RTXN) (Coiffier et al., 2002; Fitoussi et al., 2011;
Webber, 1998). RTXN is a chimeric antibody against the B-cell surface marker,
CD20, which is exclusively expressed in pre-B-cells to mature B-cells and has
been shown to be active as a single agent in many indolent lymphomas (Sousou
and Friedberg, 2010). As a single agent, RTXN is thought to function through (1)
direct induction of cell apoptosis (2) complement activation and (3) antibody
dependent cell cytotoxicity (Weiner, 2010). Interestingly, RTXN efficacy by
induction of apoptosis is significantly enhanced when crosslinked via anti Fc
antibodies (Unruh et al., 2005; Zhang et al., 2005) or Fcγ receptors (Roda and
Byrd, 2007). Crosslinking CD20 promotes translocation of the CD20 complex into
lipid rafts, causing inhibition of P38 MAPK and ERK1/2 survival pathways (Deans
et al., 2002; Jazirehi and Bonavida, 2005; Semac et al., 2003). This apoptotic
mechanism has been utilized by a variety of groups in designing therapeutic
systems utilizing multivalent Fabs (Rossi et al., 2008) to Fab polymer
conjugates(Johnson et al., 2008; Zhang et al., 2005). All these constructs have
shown potent activity in vitro and in vivo confirming the crosslinking phenomenon
to be an effective process.
75
In order to exploit the CD20 induced tumorcidal effect, recombinant Anti-
CD20 scFv was fused (Fig. 15a, Table 5) with elastin like polypeptide (ELP).
ELPs are hydrophilic biopolymers with pentameric repeats of [VPGXG]
n
, where X
can be any amino acid. ELPs undergo a characteristic reversible phase transition
above a certain critical temperature (LCST)(Urry, 1997b). The recombinant scFv
fusion was designed with the RTXN scFv fragment fused to the N-terminus of a
large molecular weight (MW) ELP (Fig. 15b). The large MW ELPs were chosen
for several reasons. ELP tags enable quick and efficient purification via inverse
temperature cycling (ITC) and serve as biodegradable carriers for scFvs,
improving their circulation time. Genetic engineering and biological synthesis
allows for accurate control over length and sequence, and by designing the
construct as a direct fusion of the scFv and ELP, chemical conjugation is
avoided. Additionally, the bacterial expression of these fusions allows for a
commercially viable product.
76
Figure 15: Antibody Core Protein Polymer NanoWorms (ACPPNs) enhance
apoptotic signaling. (a) Expression of a fusion between a single chain antibody
(scFv) and an environmentally-responsive protein polymer (i.e. ELPs) yields
stable nanoworms. The nanoworms target cell-surface CD20 receptor, inducing
apoptosis in B-cells and hence will be ideal for lymphoma therapies. (b) An anti-
CD20 scFv consisting of both a heavy and light chain was fused to the amino
terminus of an elastin-like polypeptide (ELP). The ELP protein polymer, A192,
was selected to promote solubility at physiological conditions and phase
separation upon binding the cell surface.
77
3.2. Materials and Methods
3.2.1. Materials
The DNA sequence for anti-CD20 scFv (Table. 5) was designed and purchased
from Integrated DNA technologies (Coralville, IA). Cloning vector (Pet25b(+)),
Top10, and Origami B(DE3) were purchased from Novagen (Darmstadt,
Germany). Terrific broth (TB) dry powder was purchased from Mo-bio
Laboratories (Carlsbad, CA). All Restriction enzymes were purchased from New
England Biolabs (Ipswich, MA). SYBR® safe DNA stain, low and high melting
point agarose, AnnexinV/PI apoptosis kit and TUNEL staining kit were purchased
from Invitrogen (Grand Island, NY). DNA mini prep and DNA purification kits
were purchased from Qiagen (Germantown, MD). Bacteriological grade Agar and
sodium chloride was purchased from Sigma Aldrich (St.Louis, MO). Non-
radioactive cell viability MTS assay kit was purchased from Promega (Madison,
WI). Precast 4-20% SDS PAGE gels were purchased from Lonza (Basel,
Switzerland). Raji, CEM, SU-DHL-7, RTXN, and chimeric Lym-1 (chLym-1)
antibodies were provided to us by Dr. Alan Epstein (USC, Los Angeles, CA).
Polyclonal goat anti human Fc antibody (2° GAH) was purchased from Thermo
scientific (Rockford, IL). Cell culture media, Roswell Park Memorial Institute
medium (RPMI 1640), was purchased from Corning (Tewksbury, MA). All cells
were cultured in RPMI 1640 supplemented with 10% FBS at 37° C humidified in
5% CO
2
.
78
3.2.2. Expression and purification of scFv ELP fusions
The anti-CD20 scFv was fused to ELPs using restriction enzyme digestion
followed by sticky end ligation. The expressed protein was purified from bacterial
lysate using inverse temperature cycling. Briefly, the anti-CD20 scFv sequence
(756 bp) was purchased in an ampicillin resistant proprietary pIDTsmart
TM
vector.
The scFv sequence was inserted into a pet25b(+) expression vector containing
the ELP sequences (Table 5) using restriction enzyme digestion. Sequence
confirmed plasmid was transformed into Origami B (DE3) Escherichia coli
(E.Coli) using heat shock at 42° C for 5 mins. The h eat shocked bacteria was
plated onto an ampicillin (100μg/l) agar plates and incubated at 37° C for 15-18
hrs and transformed colonies selected. The selected colonies were grown in 5 ml
TB culture media with 100 μg/l ampicillin for 15-18 hrs at 37° C. The cultures were
pelleted at 4,000 rpm for 15 mins and lysed to check for protein expression using
SDS-PAGE. A colony with high protein expression was selected and grown out in
a 50 ml starter culture with 100 μg/l ampicillin at 37° C. The bacterial culture was
then pelleted and inoculated into 1 liter TB media with 100μg/l ampicillin. The
cultures were grown for 24hrs and bacteria suspended in filtered PBS (4 lits of
culture in 25 ml of PBS) for downstream cell lysis. The bacteria were lysed using
ultrasonication to release expressed cytosolic fusion protein and bacterial DNA
was complexed out using polyethylenimine (50% w/v PEI) at 12,000 rpm for 15
mins. The supernatant containing the fusion protein was filtered through a 0.2 μm
filter before protein purification using inverse temperature cycling (ITC).
79
The DNA free supernatant was equilibrated to room temperature and ELP
phase transition induced by 3M NaCl (i.e. for 50 ml of supernatant 8 gms of
NaCl). The ELP coacervate was spun down at 25° C for 20 mins at 4000 rpm
(HOT SPIN). The supernatant was discarded and the pellet solubilized in cold
PBS. The solubilized pellet contains the ELP fusion with insoluble bacterial
proteins which were centrifuged out at 4° C at 12,00 0 rpm for 15 mins (COLD
SPIN). The hot and cold cycle was repeated twice and 6M Guanidine HCl added
to perform scFv refolding. Added guanidine is slowly dialyzed out to promote
scFv renaturation using a 20kD cut off dialysis cassette against cold PBS at 4° C.
Dialysis is carried out with a 100:1 sink condition with 4 changes of buffer. A cold
spin is performed on the dialyzed protein and a final temperature cycling step
performed to ensure complete removal of guanidine. The final protein stock is
filtered through a sterile 0.2 μm filter and protein concentration determined using
the molar extinction coefficient at 280 nm by:
………… (Eq. 2)
Where,
A
280
: Absorbance at 280 nm
A
350
: Absorbance at 350 nm
MEC (ε): Molar extinction coefficient (67,900 M
-1
C
-1
)
l: Path length (cm)
80
Table 5: Biophysical characteristics of cloned scFv ELP fusions
ELP
Nomenclatur
e
Amino acid sequence*
T
t
(
o
C)
†
ELP
behavior
#
Observed
ELP MW
(Da)
§
A192 G(VPGAG)
192
Y 55.1 Soluble 73,472.8
A96I96 G(VPGIG)
96
(VPGIG)
96
Y 56.5 Micelle 77,512.5
scFv-A192 scFv-G(VPGAG)
192
Y 42.0 Protein
core
particles
99,292.3
scFv A96I96 scFv-
G(VPGAG)
96
(VPGIG)
96
Y
n.a. n.a. n.a.
*Gene sequences confirmed by DNA sequencing from N and C terminal.
†
Transition temperature (Tt) (25 μM, pH 7.4 determined by optical density
measurements at 350 nm.
#
ELP behavior in PBS.
§
Molecular weight estimated using SDS-PAGE (n.a: not available).
‡
scFv sequence :
HWVKQTPGRGLEWIGAIYPGNGDTSYNQKFKGKATLTADKSSSTAYMQLSSLT
SEDSAVYYCARSTYYGGDWYFNVWGAGTTVTVSAGGGGSGGGGSGGGGSQ
IVLSQSPAILSASPGEKVTMTCRASSSVSYIHWFQQKPGSSPKPWIYATSNLAS
GVPVRFSGSGSGTSYSLTISRVEAEDAATYYCQQWTSNPPTFGGGTKLEIKRT
G’
81
3.2.3. Determination of purity and transition temperature (Tt) of scFv
assemblies
Purity of the constructs was determined using SDS-PAGE. Briefly, 10-15 μg of
protein is added to SDS page loading buffer and boiled at 95° C for 5 mins. The
sample is then run on 4-20% precast SDS-PAGE gel. After the samples are run
the gel is stained using 10ml of Coomassie Brilliant Blue staining solution. The
gel is imaged on a Biorad versadoc gel imager using white light. The purity of
samples was calculated using Image J. Briefly, pictures were imported into
ImageJ and converted to 8-bit files. Individual lanes are selected and an intensity
plot of each lane made. The peak areas were calculated and the purity was
determined using:
………… (Eq. 3)
Where,
A
peak
: Area of peak
A
tot
: Total area
The Tt is used to understand the effect of scFv fusion on the ELP. The Tt
of the fusions was determined using optical density measurements at 350 nm.
Briefly, increasing concentrations of constructs were added to 300 μl Beckman
coulter Tm microcells (Brea, CA) and the temperature was ramped at a rate of
1° C/min. The optical density was plotted as a funct ion of temperature, and the
maximum first derivative of this curve was defined as the Tt. The Tt for all
samples was determined in PBS.
82
3.2.4. Light scattering analysis of scFv ELP fusions
Light scattering was used to determine stability and assembly properties of the
scFv ELP fusions. To prevent detection of artifacts, all buffers used were sterile
filtered using 0.45 μm filter. Dynamic light scattering (DLS) was used to
determine the hydrodynamic radius (Rh), temperature stability and the
polydispersity of the protein in solution. Briefly, increasing concentrations of
ACPPNs were pipetted into a 384-well clear bottom plate and read on a Wyatt
DynaPro plate reader (Santa Barbara, CA) using a 830 nm laser and a 1° C/min
temperature ramp from 20° C-45° C. .
Multi angle light scattering (MALS) was used to determine the Rg,
molecular weight, and coordination number of the ACPPNs. The fusions were
analyzed using tandem size exclusion chromatography and multi angle light
scattering (SEC-MALS). Briefly, 250 μg of constructs were injected onto a
Shodex® size exclusion column using sterile filtered PBS at 0.5 ml/min. The
column eluents were analyzed on a Wyatt Helios system (Santa Barbara, CA)
and the data fit to a Debye plot to determine the R
g
and the molecular weight.
The coordination number for the assemblies was determined by dividing the
absolute molecular weight (M
abs
) by the calculated monomeric scFv fusion
molecular weight. The R
g
/R
h
ratio was used to determine the morphology of the
scFv ELP fusion.
3.2.5. Electron microscopy of scFv ELP fusion
Cryogenic TEM (cryoTEM) was performed to determine morphology in the
presence of aqueous buffer. Briefly, cryoTEM specimens were prepared using an
83
FEI Vitrobot (Hillsboro, OR). ELP solutions were kept in an ice bath (4° C) before
processing and then raised to 37 ° C immediately pri or to blotting. Six μL of
sample was pipetted onto a TEM grid coated with a lacey carbon film (LC325-Cu,
Electron Microscopy Sciences). The specimen was then blotted under 95%
humidity, immediately transferred into liquid ethane, and stored in liquid nitrogen
environment. Micrographs were acquired using FEI Tecnai 12 TWIN TEM
equipped with 16 bit 2Kx2K FEI eagle bottom mount camera (Hillsboro, OR). All
cryoTEM images were acquired at an accelerating voltage of 100 kV. Images
were analyzed using ImageJ (NIH,USA).
3.2.6. Secondary structure determination using circular dichrosim (CD)
CD was performed to determine the secondary structure of the scFv constructs.
The constructs were run on a Jasco J-815 CD spectrometer (Easton, MD) using
a quartz cuvette (path length~1 mm). The ellipticity was monitored from 185-250
nm and the spectra of buffer subtracted post run. All the constructs were
prepared in filtered diH
2
O. Deconvolution was performed under the assumption
that the observed molar ellipticity [θ] is a weighted linear sum of the ellipticity for
known secondary structures. The data was fit using nonlinear regression on
Microsoft Excel using
…………(Eq. 4)
Where,
θ: Observed ellipticity
θ
std
: Ellipticity of standard
C
std
: Fraction of standard
84
3.2.7. In vitro CD20 recognition using laser confocal microscopy
CD20 recognition was tested in CD20+ and CD20- cells. For CD20+ cells,
Burkitt’s (Raji) and diffuse large B-cell lymphoma (SU-DHL-7) cell lines were
evaluated against ACPPNs. T-acute lymphoblastic leukemia (T-ALL) cell, CEM,
was used a CD20- control. scFv CD20 recognition was performed using
rhodamine (RHD) labeled proteins under laser assisted confocal microscopy.
Briefly, 50μg of RHD labeled scFv ELP and RTXN were added to 1ml of 2 X 10
5
cells suspended in 1% BSA DPBS. The cells were incubated with the protein for
15 min at room temperature with occasional agitation. After incubation, cells were
transferred to 3ml test tubes and centrifuged at 750 rpm for 5 min to remove
unbound proteins. The cell pellets were washed twice with DPBS and suspended
in 100 μl 1% BSA DPBS. The cells were mounted onto glass slides and observed
under a Zeiss LSM510 confocal microscope with a 543 nm green excitation
laser. For RTXN competition studies, the cells were incubated with 1 mg of
unlabeled antibody for 15 min and washed prior to incubatation with RHD labeled
scFv constructs for a further 15 min. GAH Crosslinked RTXN was imaged is a
similar fashion to RHD RTXN treated cells but 10 μg of 2° GAH was added to the
washed cells and incubated for a further 15 min to induce crosslinking. After the
incubation, the cells were washed and imaged. Images were analyzed using
Image J.
3.2.8. Cell viability assays
A formazan based colorimetric assay was used to determine cell viability.
Viability assays were performed on CD20+ and CD20- cell lines used in CD20
85
binding assays. All assays were performed in 5% FBS RPMI 1640
supplemented with Pen-Strep
.
Briefly, 2 x 10
4
cells/well were pipetted in to 96-
well plates and serial dilutions of scFv ELP and RTXN were added in triplicates.
RPMI 1640 with appropriate protein dilution was used as blank control. The cells
were incubated with the protein for 24hrs, after which 30 μl of MTS /PMS was
added to determine the number of viable cells. The cells were further incubated
for 2 hrs and read at 490 nm using a Biorad benchmark plus® plate reader
(Hercules, CA). The % cell viability was calculated and plotted versus protein
concentration and viability determined by:
…………(Eq. 5)
Where,
A
treated
: Treated cell absorbance at 490 nm
A
cont
: Control absorbance at 490 nm with appropriate protein
concentration
A
untreated
: Untreated cell absorbance at 490 nm
A
0
: Control absorbance at 490 nm with no cells
3.2.9. Detection of apoptosis through flow cytometry
Induction of apoptosis was determined using early and late stage apoptotic
markers. Annexin V (ANXV)/PI staining was used to detect early induction of
apoptosis. ANXV is a cellular protein which has a strong affinity for
phosphotidylserine which is flipped to the exterior of the cell membrane during
apoptosis. Propidium iodide (PI) is used in combination with ANXV to detect DNA
content. Chimeric Lym-1 (chLym-1) is an antibody against HLA-Dr10 tumor cells
and was used a positive control for ADCC. The chLym-1 antibody binds HLA-
Dr10 expressing tumor cells inducing cell lysis. Briefly, 2 X 10
5
cells in 10% FBS
86
RPMI 1640 supplemented with 1000 units Penicillin and 1000 μg streptomycin
were added to each well in a 12 well plate. The cells were incubated with
equivalent scFv concentrations of ACPPNs, RTXN, RTXN + 2° GAH and chLym-
1 at 37° C with humidified 5% CO2 for 18 hrs. For AN XV+ve and PI+ve
compensation controls cells were treated with 50μg of paclitaxel. For 2° GAH
mediated crosslinking the cells were incubated with RTXN for 30 mins and
resuspended in fresh cell culture media. After washing, 50 μg of 2° GAH was
added the cells and incubated for 18 hrs. After incubation the cells were spun
down in 5ml FACS tubes at 750 rpm for 5 min. The cells were washed twice with
PBS and suspended in 100 μl ANXV staining buffer. The cells were stained with
ANXV and PI as per the manufacturer’s instructions i.e. 5 μl of Alexa Fluor
®
488-
ANXV stock and 1μl of 5-fold diluted PI stock were added to the cell and
incubated for 15 min. The volume of cells was made up to 500 μl with ANXV
binding buffer and analyzed on an AttuneTM acoustic focusing flow cytometer
(Life technologies, Grand Island, NY). The data were collected as .fcs files and
analyzed on Flowjo.
Late stage apoptosis was detected using terminal deoxynucleotidyl
transferase dUTP nick end labeling (TUNEL). TUNEL is used to detect apoptotic
DNA strand breaks by labeling the 3’-hydroxl ends with 5-bromo-2-deoxyuridine
5’-tri-phosphate (BrdUTP). The uridine label is detected using a fluorescently
labeled anti-Brdu antibody. PI was added to samples to detect different stages of
cell cycle. The labeling was performed as per the manufacturer’s protocol.
Briefly, 2 x 10
6
cells in 10% FBS RPMI 1640 supplemented with 1000 units
87
Penicillin and 1000 μg streptomycin were added to each well in a 12 well plate.
The cells were treated with equivalent scFv concentrations of ACPPNs, RTXN,
and RTXN + 2° GAH and incubated at 37° C with humidi fied 5% CO2 for 18 hrs.
The RTXN crosslinking by 2° GAH (50 μgs) was perfor med similar to ANXV/PI
staining procedure. After incubation the samples were fixed in in 1%
formaldehyde for 15 mins and dehydrated using 70% ethanol for 5 hrs on ice.
The fixed cells were washed transferred to a brdUTP labeling buffer and labeled
with brdU overnight at room temperature. After completion of the reaction,
Alexa488 labeled anti-brdUTP antibody was added and incubated for 2 hrs. PI
was added 30 mins before sample analysis and data collected as .fcs files.
Analysis of ,fcs files was performed on Flowjo.
3.2.10. In vivo tumor regression and biodistribution studies
Human Burkitts’ lymphoma xenografts (Raji) were used to determine in vivo
efficacy of scFv constructs. All procedures performed were in accordance to the
university approved IACUC protocol. Briefly, athymic nude mice were irradiated
using an X-ray irradiator (400 rads) to lower their NK cell population and allowed
to recover for 24 hrs. After recovery, a 200 μl inoculum of 5 x 10
6
Raji cells and
10
5
human fetal fibroblasts used to support early tumor gowth were implanted
subcutaneously on the right flank of the mouse. The mice were divided into 3
treatment groups (n=5): PBS, RTXN (1.7 mgs/dose) and ACPPNs (scFv-A192,
2.5 mgs/dose). RTX and ACPPNs were dosed at an equivalent scFv dose of 600
μg. Animal dosing was started once all tumors reached 150 mm
3
and the total
number of doses limited to 8 per mouse. The first two doses were administered
88
on consecutive days and the following six doses given every other day. The
weight of the mice and the tumor volumes were monitored and animals were
sacrificed after reaching the tumor volume end point (1000 mm
3
) or due to
occurrence of any adverse reactions to treatment. Organs from sacrificed
animals were harvested and fixed in zinc formalin for 18 hrs and dehydrated in
70% alcohol for 24 hrs before paraffin embedding. After dehydration the dry
weights of the liver, spleen, and tumor recorded. After paraffin embedding, fine
5μm slices of the organs were stained with Hematoxylin and eosin (H & E) and
studied for histological changes. The tumor volume for this study was calculated
using the following formula:
…………(Eq. 6)
Where,
w: Measured tumor width
l: Measure tumor length
In vivo biodistribution studies were performed using RHD labeled scFv-
A192 in Raji xenografted mice (n=3). A therapeutic dose of ACPPNs (2.5mgs)
was administered to the animals and the animals sacrificed after 8 hrs. The
organs of the animals were harvested and fixed in zinc formalin for 18 hrs and
dehydrated in 70% alcohol for a further 24 hrs prior to paraffin embedding.
Paraffin was removed and the sections permeabilized with 10% SDS. The slides
were then incubated in a 1:1000 dilution of DAPI for 1 hr. After incubation, the
sections were washed with 1% BSA PBS and slides prepared with antifade
reagent. The slides were dried overnight and imaged under a Zeiss LSM510
laser confocal microscopy to determine microdistribution using 543 nm and 790
89
nm excitation lasers. All images were taken under same microscope setting and
images processed on ImageJ.
3.3. Results
3.3.1. Purity and biophysical properties of scFv driven assemblies
Utilizing the ELP tag, the ACPPNs were purified from bacterial lysate using ITC
(Fig 16a.Table 5). The purity determined through Coomassie stained SDS-PAGE
was 91.4 ± 1.3 %. The yield of the fusion was estimated to be 20-30 mg/L of
bacterial culture. The purified fusion retained its phase transitioning property (Fig.
16b) but transitioned at a lower temperature due to the conjugation of the scFv
fragment (Fig. 16c). The high absorbance at 350 nm suggests formation of large
scFv-A192 assemblies even at room temperature. The scFv-A192 fusion
transitions at ~ 41° C when compared to the plain A1 92 ELP at ~ 55° C. DLS
confirmed formation of assemblies with a R
h
of 85.7 ± 16.5 nm (Fig. 16d). The R
h
for A192 is 6.7 ± 0.2 nm, suggesting that the fusion assembled particles. The
assembly of particles could be due to the scFv and it is likely that the scFv is
forming the core of these particles. Micellar scFv A96I96 showed extensive
degradation and hence was not purified.
90
Figure 16: Single chain antibodies fused to protein polymers form
thermally-responsive nanoparticles. (a) ACPPN, scFv-A192, purified and run
on a 4-20% SDS-PAGE shows a mol wt of ~100kD and a purity of 91.4 ± 1.3 %
(b) Optical density was used to characterize the phase behavior of scFv-A192.
(c) Compared to A192, scFv-A192 reduced the transition temperature. (d) scFv-
A192 assemblies have a hydrodynamic radius Rh of 85.7 ± 16.5 nm, which are
associated with the high optical density observed at 350 nm.
91
3.3.2. scFv renaturation reduces particle size and stabilizes secondary
structure
To address the unexpected particle assembly, 6M guanidine was added to aid
renaturation. The fusion protein before (raw) and after renaturation (refolded)
was analyzed using light scattering and circular dichroism. Multi-angle light
scattering (MALS) analysis performed on ‘raw’ particles confirmed assembly of
particles with an absolute molecular weight (M
abs
) of 25,490 kD (Fig. 17a). The
M
abs
showed that these particles are composed of ~250 scFv-A192 monomers
with an R
g
of 47.7 ± 0.1 nm (Fig. 17a). The R
g
/R
h
ratio of 0.56 hints at assembly
of spherical particles with a densely packed core (R
g
/R
h
ratio for polymeric
micelles <0.774). Microscopy utilizing cryoTEM confirmed monodisperse
spherical structures in solution with a diameter of 48.1 ± 11.8 nm (Fig. 17e). The
secondary structure of raw particles showed no characteristic spectra (Fig. 17c)
but deconvolution revealed a mixture of secondary structures (Fig. 17d). The
same analysis performed on the ‘refolded’ particles showed a significant
reduction in M
abs
giving rise to a mixture of 8,372 kD and 8,073 kD particles (Fig.
17b). The reduction in M
abs
translates to ~ 80 scFv-A192 monomers making up
each particle in both populations. Interestingly, the two particle populations
appear as a single population with an R
h
of 65.3 ± 15.5 nm. The two populations
have a R
g
of 45.2 ± 0.1 and 33.7 ± 0.1 nm in spite of having similar M
abs
(Fig.
17b). The R
g
/R
h
ratios for the two populations are 0.69 and 0.52, respectively,
which suggests that the particles have different packing densities and shape. The
refolded particles imaged using cryoTEM showed a major population of ‘worm-
92
like’ structures with lengths of 56.2 ± 15.9 nm and widths of 17.9 ± 3.5 nm (Fig.
17f). Minor populations of spherical particles with a diameter of 27.4 ± 7.5 nm
were also observed (Fig. 17f). Images from microscopy are consistent with the
light scattering experiments. The change in refolded protein secondary structure
was confirmed by a significant reduction in β-turn content (Fig. 17d), suggesting
a decrease in ELP aggregation. Hence these refolded particles are referred to as
ACPPNs. The particles when ramped on the DLS at a rate of 1° C/min showed
no change in R
h
. The renaturation process caused a decrease in β-turn
population. The refolded ACPPNs were therefore used to perform in vitro and in
vivo experiments.
93
Figure 17: scFv refolding reduces coordination number forming ACPPNs.
To reduce the number of polymers associating in each nanoparticle, the raw
scFv-A192 was denatured and then refolded. (a) Multi-angle light scattering
(MALS) analysis on raw particles show a molecular weight of 25,490 kD and a
radius of gyration Rg of 47.7 ± 0.1 nm. (b) MALS analysis on the refolded protein
showed a population of 8,372 kD (mass fit 1) and 8,073 kD (mass fit 2) with Rg of
45.2 ± 0.1 and 33.7 ± 0.1 nm, respectively. (c) Changes in protein structure
94
during refolding were characterized by circular dichroism, which show a
significant shift in spectra at 190 and 220 nm. (d) Deconvolution of the curves
showed a significant decrease in β-turn 2 content. The lowering of the β-turn
content suggests a change in the ELP secondary structure. (e) cryoTEM
performed on frozen raw fusion showed monodisperse spherical assemblies in
solutions. The diameter of particles was calculated to be 48.1 ± 11.8 nm. (f)
cryoTEM performed on refolded particles show a major population of ACPPNs
with lengths of 56.2 ± 15.9 nm. Scale bar represents 100 nm.
95
3.3.3. ACPPNs competitively bind CD20 cell surface receptor
RHD labeled RTXN and ACPPNs successfully recognized two CD20+ B-cell
lymphomas (Fig. 18a,b). RTXN efficiently bound CD20 with equal distribution of
CD20 on the cell surface to form a ring staining pattern (Fig. 18a,b). On addition
of goat anti-Fc antibody, the surface bound RTXN showed a speckled or
punctate pattern on the cell surface due to translocation of crosslinked RTXN into
lipid rafts (Fig. 18a,b). ACPPNs also bound CD20 forming a punctate pattern
similar in appearance to crosslinked RTXN. The binding of ACPPNs can be
blocked by pretreating both CD20+ cells with unlabeled RTXN suggesting
competitive binding of cell surface CD20. (Fig. 18b). Conversely, RTXN and
ACPPNs showed minimal binding to CD20- CEM cells. Unmodified A192 also
showed minimal binding of Raji and SU-DHL-7 cells (Fig. 18c).
96
Figure 18: ACPPNs competitively target CD20+ cells. (a) Panels i-iii and vii-ix
shows CD20 recognition by RHD labeled RTXN on both Raji and SU-DHL-7
cells. RHD labeled RTXN forms a ring pattern around the target cell. Crosslinking
surface bound RTXN by a 2° goat Anti human Fc (Pane ls iv-vi and x-xii) shows
the ring pattern shift to a more punctate appearance. (b) Panels i-iii show
recognition of surface CD20 by RHD labeled ACPPNs. ACPPNs binding also
forms a punctate appearance similar to crosslinked RTXN. (c) ACPPNs binding
was significantly higher than unmodified A192 and RTXN block (P=0.031*,
0.038**). Normalized intensity of RHD was calculated using image J (n=4 slides).
Scale bar represents 5 μm.
97
3.3.4. ACPPNs selectively reduce CD20+ cell viability by inducing apoptosis
Preliminary experiments with trypan blue exclusion show a significant increase in
trypan positive cells when CD20+ B-cells, Raji, and SU-DHL-7, were treated with
increasing concentration of ACPPNs (Fig. 19a).A formazan (MTS/PMS) based
colorimetric assay confirmed a concentration dependent reduction in cell viability
of CD20+ cells with an IC50 of 32 μM and 41 μM in Raji and SU-DHL-7 cells,
respectively (Fig. 19b). In contrast, RTXN treatment showed minimal changes in
cell viability of Raji cells. RTXN treatment reduced cell viability by 20-25% in Raji
cells irrespective of concentration used and therefore an IC50 could not be
calculated. Interestingly RTXN showed potent concentration dependent reduction
in viability of SU-DHL-7 cells with an IC50 of 4.8 μM. CD20– cells, CEM, did not
respond to RTXN treatment but showed a slight decrease in viability at higher
ACPPN concentrations (Fig. 19b). The IC50 of ACPPNs for CEM cells was 294
μM.
Induction of early and late stage apoptosis was detected by ANXV/PI
staining and TUNEL, respectively. ACPPN treatment (scFv dose-1.5 mg/ml)
significantly enhanced induction of early apoptosis in both CD20+ cell lines.
RTXN dosed at the same scFv concentration showed variable induction of
apoptosis in CD20+ cell lines with Raji cells responding better than SU-DHL-7
cells (Fig. 19c). On crosslinking RTXN with 2° GAH, both cell lines showed an
increase in early apoptosis. Since ACPPNs induce apoptosis on binding, a
positive control antibody, Chimeric Lym-1 (chLym-1), with similar mechanism of
action was used. The chLym-1 control is an anti HLA-Dr10 antibody which is an
98
effective inducer of apoptosis on direct cell binding (Tobin et al., 2007; Zhang et
al., 2007) (Fig. 19c). Treatment with an equi-scFv dose of chLym-1 performed
better than plain ACPPNs in Raji cells but was less effective in SU-DHL-7 cells.
The variable response could be due to lower expression of surface HLA-Dr10 on
SU-DHL-7 cells (Fig. 19c). Unlike chLym-1, ACPPNs were equally potent in both
cell lines. It is interesting to note that results for RTXN treated SU-DHL-7 cells
from the formazan based viability assay are contrary to findings from ANXV/PI
staining. CEM cells treated with ACPPNs showed minimal induction of apoptosis
and hence were not evaluated further (Fig 19c).
TUNEL was used to determine the induction of apoptosis since ANXV/PI
staining is known to detect only early apoptosis. ACPPN treatment significantly
enhances apoptosis compared to plain RTXN (P=0.0007*, P=0.0008**) dosed at
the same scFv concentration (scFv dose-2.5 mg) in both CD20+ cell lines (Fig.
19d). The efficacy of RTXN can be enhanced to the same extent as ACPPNs by
crosslinking with 2° GAH (Fig. 19d). Hence the ACPP Ns are effective in both B-
cell lymphoma cell lines and outperform RTXN in vitro.
99
Figure 19: ACPPNs reduce viability of CD20+ human lymphoma cell lines
by inducing apoptosis. (a) Trypan blue exclusion showed a significant increase
in trypan blue positive cells with increasing concentrations of ACPPNs. (b)
CD20+ cells, Raji and SU-DHL-7, show a concentration dependent reduction in
cell viability. The calculated IC50s for Raji and SU-DHL-7 are 32 and 41μM
respectively CD20- T-cell like lymphocytes, CEM, are less effected by ACPPNs
treatment. The IC50 for CEM cells is 294 μM which is ten times higher than
ACPPNs. (c) Raji and SU-DHL-7 cells both show an increase in ANXV/PI
staining after ACPPNs treatment. RTXN crosslinked by 2° GAH and apoptosis
control, chLym-1, both induce apoptosis. (d) TUNEL staining confirms ACPPNs
induction of apoptosis. ACPPNs outperform plain RTXN in both cell lines
(P=0.0007*, P=0.0008**). ACPPNs induce apoptosis to the same extent as 2°
GAH crosslinked RTXN.
100
3.3.5. ACPPNs accumulate in xenografted tumors and successfully retard
tumor growth in mice
RHD labeled ACPPNs injected in Raji xenografted athymic nude mice (n=3)
showed accumulation in various organs (Fig. 20a-h). RHD signal was seen in the
liver (Fig. 20a,e), spleen (Fig. 20b,f), tumor (Fig. 20c,g) and kidney (Fig. 20 d,h)
and minimal accumulation was observed in the heart and lungs. Tumor
regression studies were performed in mice with Raji xenografts implanted in the
right flank of athymic nude mice (n=5/group). The mice were dosed every other
day until 8 doses were administered, and tumor volume was monitored until the
tumor volume endpoint (1000 mm
3
) was reached. ACPPNs treatment
significantly retarded tumor growth compared to plain RTXN and PBS groups
(Fig. 20i). Repeated measures 1-way ANOVA performed on the mean tumor
volumes showed a significant difference between ACPPNs, RTXN, and PBS
treated groups (P=0.0011). Bonferroni post hoc analysis show a statistically
significant difference between ACPPNs, RTXN (P=0.0015) and PBS (P=0.018)
treatments. Tumor volumes of RTXN and PBS showed no statistically significant
difference (P=0.148). ACPPNs treatment significantly improved survival when
compared to RTXN and PBS treatment groups (P=0.013, Fig. 20j). The median
survival times for ACPPN, RTXN and PBS are 33, 19 and 25 days respectively.
The administered doses were adequately tolerated with weight loss (~20%)
observed in ACPPNs group after the first dose. The weight was recovered by day
13 with no causalities to treatment.
101
The dry weight of the organs between the three groups did not change
appreciably except for the spleen. A slight increase in dry spleen weight was
observed in RTXN and ACPPNs groups when compared to PBS treatment group
(Table 6). Similarly, organs collected showed no major histological changes
between the three groups except for the tumor (Fig. 21). The tumors in the RTXN
and ACPPNs treatment showed similar histology with prominent necrotic regions
compared to that seen in the PBS tumors.
102
Figure 20: ACPPNs treatment shows relatively high tumor accumulation
and reduces tumor burden in Raji xenografts. (a-h) ACPPNs microdistribution
(2.5 mgs/dose, n=3) using laser confocal microscopy shows accumulation of
RHD labeled ACPPNs in liver (a,e), spleen (b,f), tumor (c,g), and kidney (d,h).
There was minimal accumulation of these particles in the lungs and heart.
Interestingly particle accumulation can be seen in the spleen marginal zone (MZ)
and the bowman capsule (BC). Scale bar represents 20 μm. (i) As of day 25,
ACPPNs significantly reduced mean tumor burden when compared to PBS and
RTXN treated groups (n=5/group, P=0.0011). (j) ACPPNs treatment significantly
enhances survival when compared to PBS and RTXN control groups (P=0.013).
The highlighted tick marks indicate days of dose administration.
103
Table 6: Dry organ weights from treatment groups
#
Group Liver (μg) Spleen (μg) Tumor (μg)
PBS 1033.7 ± 209.2 91.8 ± 14.7 551.3 ± 195.4
RTXN 1039.3 ± 167.4 145.0 ± 49.9 474.0 ± 81.4
ACPPNs 964.7 ± 75.9 118.2 ± 41.8 451.0 ± 104.1
#
Organs collected after tumor volume end point is reached.
104
Figure 21: Tumor histology of RTXN and ACPPNs treated animals show
marked changes. Tumors harvested from animals were fixed in zinc formalin,
dried in 70% alcohol, and embedded in paraffin. The embedded tumors were cut
into 5 μm thick sections and imaged under a light microscope. Tumor from RTXN
and ACPPNs treated groups show larger necrotic areas (arrows) when compared
to PBS group.
105
3.4. Discussion
Utilizing simple genetic engineering, we were able to construct scFv based
therapeutics which were successfully purified from bacterial lysates using the
ELPs as the purification tag. To the best of our knowledge, this is the first
demonstration of ELPs being used as a purification tag for bacterially expressed
recombinant scFv. The ‘raw’ protein formed large spherical particles which could
be due to (1) high salt concentration used to induce ELP phase transition, which
could contribute to scFv denaturation leading to assembly or (2) recombinant
scFv multimerization forming dimeric, trimeric, and even hexameric
molecules(Dolezal et al., 2000; Kortt et al., 2001) which could trigger particle
assembly with a multimeric scFv core. The particle formation was reduced by
guanidine renaturation but the process led to the formation of recombinant
ACPPNs which efficiently targeted CD20 expressed on the surface of B-cell
lymphomas. The formation of ‘worms’ was confirmed using cryoTEM which
showed particles of 56.2 ± 15.9 nm in length. The worms are still assembles with
a scFv core but with a lower M
abs
and relatively constant R
g
. Due a lower mass
distributed in the same volume after refolding, the scFv core is more accessible,
allowing for CD20 recognition. The small population with a lower diameter could
have a less accessible core and may not contribute to the molecules efficacy.
In vitro activity was first confirmed by measuring cell viability which showed
selective killing of CD20+ cell albeit at a relatively high IC50 (32 μM and 41 μM).
Compared to the poor reduction of viability by RTXN, however, the IC50 seems
acceptable. The reduction in cell viability was confirmed to be due to the
106
induction of apoptosis using two separate techniques targeting different stages of
apoptosis. ACPPNs treatment greatly induced apoptosis (~60%) in both cell lines
and outperformed equi-scFv dose RTXN treatment. Also, 2° GAH crosslinked
RTXN showed the same efficacy of induction as single agent ACPPNs treatment.
An unexpected observation was that RTXN showed potent reduction in SU-DHL-
7 viability using formazan based assays but minimal cell staining apoptosis in
both apoptosis assays. This contradiction could have arisen due to the 100 fold
less cell concentrations used for the experiment.
The in vitro activity of ACPPNs was successfully translated in vivo using a
Raji cell xenograft. ACPPNs treatment showed a significant delay in tumor
growth when compared a RTXN dosed at the same equivalent scFv dose. A high
dose for ACPPNs was chosen based on its activity in vitro. Tumor accumulation
of ACPPNs was confirmed by injecting Raji cell xenografted nude mice with RHD
labeled reagent. The particles showed liver, spleen, tumor, and kidney
accumulation (Fig. 20a-h). The high liver and spleen uptake is mostly attributed
to the presence host reticuloendothelial system (Brigger et al., 2002; Buzea et
al., 2007). Kupffer cells in the liver are responsible for nanoparticle clearance
(Dobrovolskaia et al., 2009) and hence could be taking up ACPPNs from the
circulation. In the spleen, the major RHD signal was present in the marginal zone
(MZ) lining the white pulp region of the spleen (Fig. 20b,f). The marginal zone is
mainly populated by phagocytic macrophages and lymphocytes(Kraal, 1992;
Martin and Kearney, 2002) which play a significant role in filtering and clearing
nanoparticles from the body (Aichele et al., 2003; Demoy et al., 1999). Hence the
107
liver and spleen signal could be due to phagocytosis and clearance by effector
cells. This finding is consistent with the uptake of nanoparticles with similar size
(Moghimi et al., 2012). Interestingly, kidney sections also show accumulation of
ACPPN particles in the glomerulus confirming the high molecular weight
(glomerulus filtration cut off = 60,000 kD) of ACPPNs(Meibohm and Zhou, 2012).
In conclusion, the these novel first generation ACPPNs (1) outperform RTXN as
a single agent, (2) are biodegradable due to their peptidic nature, (3) are
genetically engineered to offer precise control over the sequence, (4) are
cheaper to produce than high molecular weight antibodies, and (5) represent a
simple platform to apply to various other scFv targets.
3.5. Conclusion
Herein we describe an exciting new approach utilizing scFv fusions as tools to
developing novel therapeutics for B-cell lymphomas. The ELP tag served as both
the carrier and purification tag for scFv. The fusions assemble ACPPNs which
still show selective CD20+ targeting in vitro. The accessible scFv core showed
appreciable activity in two different NHL B-cell lines, Burkitt’s (Raji) and Large
diffuse B-cell (SU-DHL-7) lymphoma. The in vivo tumor accumulation and
activity of ACPPNs was successfully confirmed in a CD20+ Raji cell tumor
xenograft where ACPPNs outperform RTXN dosed at the same scFv
concentration. Hence these particles offer a new platform for development of
small molecule free therapeutics for various cancers and immunology related
diseases.
108
3.6. Acknowledgements
This work was made possible by the University of Southern California, the
National Institute of Health R21EB012281 to J.A.M., and P30 CA014089 to the
Norris Comprehensive Cancer Center, the USC Molecular Imaging Center, the
USC Nanobiophysics Core Facility, the Translational Research Laboratory at the
School of Pharmacy, the American Cancer Society IRG-58-007-48, the Stop
Cancer Foundation, the USC Ming Hsieh Institute, the USC Whittier Foundation,
and the USC Wright Foundation.
109
4. ELP conformation determines liposome release and cellular uptake
4.1. Introduction
Amphipathic peptides have long been studied for their membrane rupturing
properties (Fernandez-Carneado et al., 2004). These peptides interact with
liposomal bilayers due to their amphipathic nature i.e through hydrophobic
domain on the peptide (Fernandez-Carneado et al., 2004). This interaction
causes membrane destabilization and promotes membrane mixing between
liposomes (Fernandez-Carneado et al., 2004). A perfect example of such a
peptide is GALA. GALA is a 30 amino acid pH sensitive peptide which changes
conformation from a random coil to α-helix in an acidic environment (pH=5) (Li et
al., 2004a). The hydrophobic α-helix inserts into the lipid bilayers promoting
fusion and rupture (Li et al., 2004a). Due their hydrophobic domains amphipathic
peptides are easier to incorporate onto a lipid membrane and hence their effect
on lipid bilayers is easier to study.
Non-amphipathic peptides do not possess a hydrophobic domain and
hence require a lipid anchor for membrane incorporation. There is limited
literature on non-amphipathic peptides and their effect on the liposome
membrane (Hayashi et al., 1996; Holig et al., 2004; Papahadjopoulos et al.,
1991). But these peptides can be incorporated into lipid membranes using single
lipid chains (ex. Stearic acid) (Riche et al., 2004) or double (ex.
Phosphatidylethanolamine) (Riche et al., 2004). The most commonly used non-
amphipathic peptides are usually targeting sequences ex. RGD (Holig et al.,
2004) which do not have a marked effect on the integrity of a liposome but
110
increase internalization (Holig et al., 2004). Non-amphipathic polymers like
NIPAAM have shown to affect the liposome in response to an external stimulus
(Hayashi et al., 1996). NIPAAM exhibits a reversible temperature dependent
phase transition causing it to precipitate out of solution when heated above its
transition temperature (Hayashi et al., 1996). This phase transitioning property
was utilized to design a thermally sensitive liposome in which NIPAAM was used
to stabilize fusogenic lipid bilayer. The same polymer ruptured the bilayer when
heated above its transition temperature (Hayashi et al., 1996).
The presented work focuses on incorporating water soluble non
amphipathic elastin like peptide amphiphiles (ELPAs) into lipid bilayers. ELPA’s
are lipidic conjugates of ELPs. ELPs are repetitive sequences of (VPGXG)
n
(where ‘X’ = any amino acid) and exhibit reversible temperature dependent
phase transition (Urry, 1997a). The phase transitioning behavior of the ELP’s
depend on 1) Length. 2) Nature of guest residue (X). 3) Concentration of ELP. 4)
Ionic concentration of solution (Urry, 1997a). Due to their non-amphipathic
nature, we used novel synthesized ELPA’s (Aluri et al., 2012) to graft these
peptides onto the liposome surface. This strategy provides a cheap and effective
means of using PAs lipid anchors without the use of expensive phospholipids
conjugation chemistry.
Here we present findings which show ELPs grafted on to liposome surface
stabilized fusogenic lipid bilayers and also induced membrane leakage when
heated (Fig. 22a,b). The surface grafted ELPs show similar phase transition
behavior to normal large molecular weight ELPs. We were also able to
111
characterize these peptides in very small confined spaces. By increasing surface
concentrations of ELPs we found liposomes to be stable with minimal leakage.
We see maximum leakage at low ELP grafting concentration and low leakage at
high density grafting. The grafting densities in turn influence the surface
conformation of the ELP. Furthermore we see a dramatic increase in permeability
under hyperthermic conditions (Fig. 22a,b). We also determined the mechanism
of release to be due to pore formation caused by ELP transitioning on the surface
of the liposome. In summary we show (1) ELPA stabilization of fusogenic lipids,
(2) Surface conformation dependent and temperature dependent contents
leakage, (3) Contents leakage due to pore formation, (4) Particle uptake
dependent on grafting density, and (5) Successful encapsulation of
chemotherapeutic (DOX).
112
Figure 22: Surface-graft density controls release from vesicles stabilized by
environmentally responsive polypeptides. (a) Dipalmitoylation of an N-
terminal lysine that enables attachment of ELPs to lipid bilayers. The peptide
hydrophobicity can be modified by adjusting amino acid hydrophobicity (R) and
the peptide molecular weight by adjusting length, n. R=Alanine, Isoleucine &
Valine, n=2 & 3. (b) Peptide phase transitions can be determined by disruption of
stabilized liposomes. The liposome destabilization is dependent on the surface
concentration of the grafted polymer. Our experiments suggest that content
release is observed when the polymer is in a ‘mushroom’ conformation whereas
the ‘brush’ conformation is relatively stable.
113
4.2. Materials & Methods
4.2.1. Materials
1,2-Dioleoyl-sn-glycero-3-phosphoethanolamine (DOPE), 1-palmitoyl-2-oleoyl-
sn-glycero-3-phosphoethanolamine (POPE), 1,2-distearoyl-sn-glycero-3-
phosphoethanolamine-N-(Methoxy(polyethylene glycol)-2000) (mPEG-2000-
DSPE) and the mini extruder kit were purchased from Avanti polar lipids
(Birmingham,AL). Rink amide MBHA resin, Fluorenylmethyloxycarbonyl Valine
(FMOC-Val), FMOC-Proline (FMOC-Pro), FMOC-Glycine (FMOC-Gly), FMOC-
Alanine (FMOC-Ala), FMOC-Isoleucine (FMOC-Ile), O-Benzotriazole-N,N,N’,N’-
tetramethyl-uronium-hexafluoro-phosphate (HBTU), N-Methyl-2-pyrrolidone
(NMP), and Acetonitrile (ACN) were purchased from EMD chemicals
(Gibbstown,NJ). Kaiser test kit (ninhydrin assay), DiFMOC-Lysine (DiFMOC-
Lys), Diethyl ether (Ether), Triisopropyl silane (TIPS), N,N-Diisopropylethylamine
(DIPEA), Chloroform (CHCl
3
), trifluoroacetic acid (TFA), Methanol (MeOH),
Ethanol (EtOH), 4-(2-hydroxyethyl)-1-piperazineethanesulfonic acid (HEPES),
Palmitic acid N-hydroxysuccinimide ester (NHS-Palimitic acid) and sodium
dodecyl sulfate (SDS) were purchased from Sigma Aldrich (St. Louis, MO). 1,1'-
Dioctadecyl-3,3,3',3'-Tetramethylindocarbocyanine Perchlorate (DiI), 8-
aminonaphthalene-1,3,6-trisulfonic acid disodium salt (ANTS) and p-xylene-bis-
pyridinium bromide (DPX) were purchased from Invitrogen (Carlsbad, CA).
Sodium chloride (NaCl) and 4-Methyl piperidine (MeP) was purchased from
fisher chemicals (Tustin, CA). Sephadex PD-10 desalting columns were
purchased from GE healthcare (Piscataway, NJ). The Waters C-18 and YMC C4
114
reverse phase semi prep columns were purchased from Waters inc (Milford, MA).
De ionized water from Barnstead purification system (Thermo scientific,
Asheville, NC) was used for all aqueous buffers.
4.2.2. Synthesis and purification of dpKXn
The ELP lipid conjugates were synthesized using the solid phase peptide
chemistry. The protocol followed is described in (Aluri et al., 2012). Briefly, short
ELPs were synthesized using solid phase peptide synthesis on Rink amide resin
using FMOC chemistry (Rink, 1987). The short ELP sequences where capped
with (FMOC)
2
lysine on the C-terminus. The lysine is deprotected using 20%
MeP and conjugated to HBTU activated palmitic acid under basic conditions. The
deprotection and conjugation reactions were monitored using ninhydrin. After
complete exhaustion of the primary amines the synthesized dpKXn was cleaved
of the resin using 95% TFA and precipitated in cold 100% ether. The crude
conjguates were dissolved in100% MeOH and purified on a semi prep reverse
phase C4 column at 214 nm on a Perkin Elmer 200 series high performance
liquid chromatography (HPLC). The conjugates were run on a H2O:MeOH
gradient starting at 50% to 100% MeOH. The peak of interest was collected and
mass confirmed using a DECA-LcQ ESI mass spec system (Thermo scientific,
Waltham, MA). After confirmation the products were purified and lyophilized.
Stocks of the purified samples were made in 100% HPLC grade MeOH
4.2.3. LCST determination of dpKXn
The temperature dependent phase transitioning property of ELPs was
demonstrated using optical density (OD) measurements at 350 nm over a set
115
temperature range. The LCST was determined by calculating the first derivative.
Decreasing concentrations of polypeptides were added to 300 μl T
m
microcells
(Beckman Coulter, Brea, CA). The microcells were ramped from 45
o
C to 75
o
C on
a Beckman Coulter DU 800 UV-Vis spectrophotometer (Beckman Coulter, Brea,
CA).
4.2.4. Preparation and characterization of dpKXn liposomes
1mM stocks of appropriate lipids were made in CHCl
3
. The purified dpKXn was
dissolved in methanol. The dpKXn liposomes were made using thin film hydration
and sized through a 100nm PC membrane. Briefly, appropriate quantities of
DOPE, POPE and dpKXn were mixed in a round bottom test tube. The solvents
were evaporated on a Heidolph laborota 4011 (Heidolph Brinkmann, Elk Grove
Village, IL) at 50
o
C with a dry ice trap to form an even lipid film. For even mixing
of the stabilizing lipid, the dried film was again dissolved in 1ml CHCl
3
and dried
down again. The prepared film is then stored under vacuum to remove trace
organic solvent. The film is stored under argon at 4
o
C till further use.
Liposomes were prepared by film hydration with appropriate buffer. For
release assays the lipid films were hydrated with 1ml 50mM ANTS/DPX in
HEPES buffered saline (pH=7.4). The hydrated film is placed at 60° C for 15 secs
and sonicated for a further 30 secs to ensure complete hydration of film. The
liposomes are extruded at room temperature through a 100 nm PC membrane.
The extruded liposomes were then desalted using a PD-10 column to remove
unencapsulated ANTS/DPX using HEPES buffered saline (pH=7.4). For fusion
assays all liposomes were hydrated with HEPES buffered saline (pH=7.4).
116
Particle size and stability of liposomes were determined using dynamic light
scattering on a Wyatt DynaPro plate reader (Santa Barbara, CA). Briefly, 50ul of
prepared liposome stock was transferred to a 384 well clear bottom plate and the
hydrodynamic radius measured at 25
o
C. The stability of particles was determined
by monitoring particle size from 25
o
C to 60
o
C.
The total lipid in the liposome stock was calculated using the phosphate
assay. Briefly, 100ul of the liposomes stock was hydrolyzed with 6N H
2
SO
4
at
180
o
C for 3 hrs. The hydrolyzed lipid is allowed to cool and ascorbic acid and
ammomium molybdate were added. The absorbance of the molybdate complex
at 820 nm was determined and compared to known concentrations of NaHPO
4
standards to determine the lipid content in solution.
4.2.5. Liposome release assay
The release assay was performed on Horiba Jobin Yvon Fluorolog-3
spectrofluorometer (Edison, NJ). The assay is based on the dequenching of
water soluble fluorophore (ANTS) from the quenched ANTS/DPX complex. The
dequenching of ANTS was used to determine the LCST and evaluate the release
kinetics of the liposomes. The fluorescence signal was monitored using excitation
of 395 nm (5 nm) and emission of 513 nm (5 nm). For LCST determination the
liposomes were ramped from 25-60
o
C and the second derivative of the curve
was taken to determine the LCST for the liposome formulation. Briefly, known
molar concentrations of liposomes were added to the cuvette and the volume
made up to 3ml. The cuvette is stirred continuously for efficient heating of added
liposomes. The cuvette is ramped from 25
o
C to 60
o
C at a rate of 1
o
C/min. At the
117
end of the ramp the liposomes were lysed with 10% SDS to calculate total
encapsulated flurophore (C
100
). The total encapsulated content (C
100
) was used
to determine temperature dependent % leakage using Eq. 7.
100 …………(Eq. 7)
Where ‘t’ is the set temperature, ‘C
0
’
is the concentration at time zero, ‘C
t
’
is the concentration at the set temperature and ‘C
100
’ is the concentration of
ANTS/DPX after lysis. Simultaneously the prepared liposomes were ramped on
the DLS at 1° C/min to monitor changes in hydrodynam ic radius.
The encapsulated volume for liposomes was calculated using Eq. 8.
Volume (V) = ………… (Eq. 8)
Where V is the volume encapsulated, C
ANTS
is the concentration of ANTS
in liposomes (mM) and C
lipid
is the concentration of total lipid (mM).
The release kinetics of liposomes were studied under similar condition to
the previous experiment. Briefly, known molar concentrations of liposomes were
added to the cuvette and the temperature stepped to above and below the
transition temperature over a period of 5 mins at each temperature. Liposomes
were monitored at room temperature i.e. 25
o
C for 5 mins and the temperature
stepped to a 35
o
C and 45
o
C. At the end of the study the liposomes were lysed
and the total amount encapsulated calculated. The raw data files were stored in
ASCII format and analyzed on Windows Excel. The curve obtained were fit to the
following
…………(Eq. 9)
( )
( ) lipid
ANTS
C
C
×
×
50
1000
118
Where ‘C’ is the concentration at time ‘T’ in sec and ‘K’ is the rate constant.
4.2.6. Liposome fusion assay
Liposome fusion assay was performed using resonance energy transfer of
surface bound fluorophore. The assay was based on the procedure mentioned in
(Guo et al., 2003). Briefly, dpKXn liposomes were prepared with 1% mol each of
NBD-PE and RHD-PE. The formulation was then extruded through a 100nm PC
filter. The labeled liposomes were mixed with unlabeled liposomes at a 1:2 molar
ratio to give a total lipid content of 100 μM. The mixture was then ramped at
1° C/min to check for temperature dependent fusion. The fluorescence signal was
monitored using excitation at 467 nm (5 nm) and emission signal at 550 nm (18
nm). To determine the intensity at 100% fusion, liposomes with 0.33% mol each
of NBD and RHD-PE were made. The fusion data was normalized to prepared
0.33% mol control liposomes. The fluorescence intensity was normalized and
plotted as a % of initial intensity using
………… (Eq. 10)
Where ‘Ft’ is fluorescence intensity of test liposomes at time ‘t’, ‘F0’ is
fluorescence intensity at time ‘0’, ‘F’t’ is fluorescence intensity of control
liposomes at time ‘t’
4.2.7. Biophysical characterization of dpKXn
shielded liposomes
ELPs exist as a random coil in solution and adopt a β-turn conformation when
heated above the LCST (Urry, 1984). It is assumed that each amino acid is a
repeat unit. Based on our assumptions X2 would have 10 units and X3 would
119
have 15 repeat units. The Flory radius (Rf) of these polymers in solution is
calculated by
…………(Eq. 11)
Where ‘l’ is the average bond length, ‘n’ is the number of bonds, The
cross-sectional Flory area (Af) occupied by surface grafted polymer is
…………(Eq. 12)
Area available per polymer (Ap)
…………(Eq. 13)
Where, ‘A
lipid
’ is the area occupied by a single lipid (For DOPE ~ 0.65 nm
2
(Zuidam and Barenholz, 1997)), ‘f’ is the mol fraction of dpXn used
When Af ≥ Ap the polymers become constrained on the surface causing them to
adopt a brush conformation of the surface. Solving for ‘f’ determines the
conformation of the surface polymer for liposome formulation.
4.2.8. Cell uptake of dpKXn and DOX loaded liposomes
The cell uptake of dpKXn liposomes were evaluated in a MDA-231 breast cancer
cell lines. Briefly, 4 x 10
5
cells were grown on a 25 x 25 mm cover slip in 2 ml
DMEM supplemented with 10% FBS in humidified 5% CO
2
. Liposomes were
prepared with 0.1% mol DiI (red) and extruded through a 100 nm PC filter. The
liposomes were sized and 100 μl of stock added to the cells and incubated for 1
hr at 37 ° C. DOX loaded liposomes were made using f ilm hydration. Briefly, lipid
films were prepared with 100 μg of DOX. The lipid films were hydrated with 1 ml
HEPES buffered saline and the liposome extruded through a 100 nm PC filter.
120
The extruded liposomes were desalted using a PD10 column and the liposomes
collected and 250 μl of the stock added to the cell preparations.
After incubation the cells were stained with DAPI and fixed with 4%
paraformaldehyde before mounting onto a glass slide using flouromount. The
cells are observed using laser confocal microscopy using 488 nm (green), 543
nm (red), and 790 nm (blue) lasers. The images were exported using LSM image
browser and analyzed using ImageJ.
4.3. Results
4.3.1. Purity and yield of synthesized dpKXn
All peptides with Ala (KA2, KA3), Ile (KI2, KI3) & Val (KV2, KV3) as guest
residues were synthesized with high yield and purity. The peptide amphiphiles
(PAs) were efficiently synthesized by conjugating a dipalmitoyl lipid anchor
(Figure 22a). The PAs were further purified on the HPLC with a C4-reverse
phase column using a H
2
O:ACN gradient. The PAs were obtained in moderate
yield and high purity. All the synthesized conjugates were mass confirmed using
ESI mass spectrometry (Table 7).
4.3.2. Preparation of dpKXn
liposomes
Liposomes prepared with fusogenic lipids DOPE/POPE were stabilized by
addition of ELPAs, dpKXn
(Table 8). POPE was used to promote even mixing of
the lipids in the bilayer due to the presence of a saturated palmitoyl chain and
also an unsaturated oleoyl chain. dpKXn with Ala as the guest residue (dpKA2
and dpKA3) stabilized the liposomes at room temperature (25
o
C). Lipid films with
valine (dpKV2 & dpKV3) and isoleucine (dpKI2 & dpKI3) were found to be
121
unstable at room temperature and lipid aggregation was observed on hydration.
Formulations with dpKA2 and dpKA3 were easily hydrated and extruded at room
temperature. The radii for different formulations range from 70-120 nm in radius
when extruded through a 100nm filter (Table 8). The minimum amount of dpKA2
and dpKA3 required to stabilize the DOPE/POPE membrane is 5% and 1%,
respectively. Incorporation of dpKA3 to unstable formulations led to formation of
stable liposomes (Table 8). But dpKA3 was not able to stabilize dpKI3 liposome
formulations.
122
Table 7: Characterization of the purified ELP-lipid conjugates
Synthesis
(#)
Label
*
Lipid-ELP
†
Yield
(%)
‡
Purity
(%)
§
Observed Mass
(Da)
1 dpKA2 dp-
K(VPGAG)
2
70 98 [M+H
+
]= 1385.5
[M+Na
+
]= N.O.
2 dpKA3 dp-
K(VPGAG)
3
75 99 [M+H
+
]= 1766.7
[M+Na
+
]= N.O.
3 dpKI2 dp-
K(VPGIG)
2
49 99 [M+H
+
]= 1468.9
[M+Na
+
]= N.O.
4 dpKI3 dp-
K(VPGIG)
3
60 99 [M+H
+
]= 1892.9
[M+Na
+
]= 1915.1
5 dpKV2 dp-
K(VPGVG)
2
55 99 [M+H
+
]= 1440.8
[M+Na
+
]= 1464.0
6 dpKV3 dp-
K(VPGVG)
3
65 99 [M+H
+
]= 1850.9
[M+Na
+
]= N.O.
*
dipalmitoyl lysine (dp-K)
†
yield using solid phase chemistry
‡
purity of product by absorbance at 214 nm using reverse phase HPLC
§
mass of product characterized by electrospray mass spectrometry, not observed
(N.O.)
123
Table 8: Physico-chemico properties of liposome formulations
Prep
(#)
Label
Lipid Composition
*
R
h
± SD
(nm)
†
T
t
(
o
C)
1 10%
PEG
DOPE:POPE:PEG
[45:45:10]
74.7 ± 20.3 N.O.
2 5%
dpKA2
DOPE:POPE:dpKA2
[47.5:47.5:5]
93.5 ± 20.4 40
3 10%
dpKA2
DOPE:POPE:dpKA2
[45:45:10]
99.3 ± 23.9 39
4 20%
dpKA2
DOPE:POPE:dpKA2
[40:40:20]
89.5 ± 21.9 38
5 1%
dpKA3
DOPE:POPE:dpKA2
[49.5:49.5:1]
88.9 ± 19.1 42
6 5%
dpKA3
DOPE:POPE:dpKA3
[47.5:47.5:5]
78.3 ± 19.8 40
7 10%
dpKA3
DOPE:POPE:dpKA3
[45:45:10]
74.0 ± 17.8 38
8 20%
dpKA3
DOPE:POPE:dpKA3
[40:40:20]
78.3 ± 16.8 37
9 10%
dpKI2
DOPE:POPE:dpKI2
[45:45:10]
D.N.F. N.O.
10 10%
dpKI3
DOPE:POPE:dpKI3
[45:45:10]
D.N.F. N.O.
11 10%
dpKV2
DOPE:POPE:dpKV2
[45:45:10]
D.N.F. N.O.
12 10%
dpKV3
DOPE:POPE:dpKV3
[45:45:10]
D.N.F. N.O.
15 100%
dpKA3
dpKA3
[100]
83.6 ± 19.4 N.O.
*
hydrodynamic radius ± standard deviation, did not form particles (D.N.F.)
†
transition temperature determined by leakage of contents, not observed (N.O.)
124
4.3.3. Effect of dpKXn
phase transition on content leakage
The short dpKXn peptide chains were found to behave similar to higher MW ELP
A192 (Fig. 23, 24) in solution and also on a lipid bilayer. Both peptide variations
showed a concentration dependent decrease in transition temperature.
Interestingly dpKA3 showed a greater variation in transition temperature than
dpKA2. Liposomes prepared with hydrophobic ELPAs (X=I & V, n=2,3) did not
form stable liposomes and hence could not be evaluated. A decrease in liposome
leakage rate with increase in surface concentration of ELP was observed (Fig.
24). Liposomes with dpKA2 and dpKA3 have LCST’s between 37
o
C-40
o
C (Fig.
24). The surface grafted ELPAs showed a lower LCST than ELPAs in solution.
Ten percent PEG 2000 liposomes were run as thermally stable control liposomes
(Fig. 24).
125
Figure 23: Short ELPAs and high molecular weight ELPs have similar
phase behavior. (a) Increasing concentrations of ELP A192 were run on a UV
ramp (λ
abs
= 350 nm) from 45-80
o
C. The LCST of ELP A192 was found to be
inversely proportional to the concentration. (b) Increasing concentrations of
dpKA3 were run on a UV ramp (λ
abs
= 350 nm) from 45-80
o
C. (c) ELP A192 and
dpKA3 show a concentration dependent change in transition temperature (r
2
=
0.9966, 0.9498).
126
4.3.4. Evaluation of dpKXn liposome release kinetics
Liposome formulations with 5% dpKA2 and 1% dpKA3 liposomes were chosen
due to their leakage rates. The formulations where stepped from 25
o
C to 35
o
C
(5 min at each temperature) and a new run performed from 25
o
C to 45
o
C. Both
liposome formulation were found to be stable at 25
o
C and 35
o
C. Liposomes with
1% dpKA3 released most of their encapsulated contents (Fig. 25) when heated
rapidly (20 sec) to 45
o
C. Liposomes with 5% dpKA2 were able to release 50% of
their content when heated to 45
o
C (20 sec). Rapid content release was
observed in both formulations but release did not go to completion. When fit to
Eq. 9, the rate constant ‘K’ at 35 ° C is ~99 which reduced to ~ 35 when heated
to 45 ° C. The liposomes showed rapid release with a T
1/2
calculated at 45 ° C of
36 sec. Temperature ramps performed on 10% PEG 2000 showed no release at
both 35
o
C and 45
o
C. The decrease in intensity observed could be due to
bleaching of any trace fluorophore.
127
Figure 24: ELP surface-graft density controls content release from vesicles
(a) A2 liposomes (5% to 20%) showed a decrease in contents leakage was
observed with increasing concentrations of surface grafted ELP. (b) A3
liposomes (1% to 20%) showed a similar decrease in content leakage with
increasing concentrations of surface grafted ELP. (c) A concentration dependent
reduction in transition temperature calculated by second derivative was observed
with both dpKA2 and dpKA3 liposomes.
128
4.3.5. Contents release of dpKXn liposomes is due to pore formation
Liposome fusion assay performed via fluorescence resonance energy transfer
(FRET) showed minimal increase in NDB signal when compared to the 0.33%
mol control. The minimal increase in NDB signal suggests quenching by RHD
(Fig. 26). Hence contents release by 1% dpKA3 liposomes is through pore
formation and not through membrane fusion. Temperature ramps performed on
the DLS show formation of large aggregates by 1%dpKA3 liposomes but the
aggregation occurs at >50° C (Fig. 26). Preliminary temperature ramps performed
on 5% dpKA2 liposomes shows formation of large aggregates on the DLS above
40° C. Fusion assays could not be performed on dpKA2 liposomes due to their
inherent instability.
4.3.6. Biophysical characterization of dpKXn
surface conformation
The length of the synthesized conjugates was calculated based on average
amino acid bond lengths. The chain length for a single VPGXG repeat is 0.15
nm. The ‘f’ value calculated for dpKA2 and dpKA3 was determined to be 1% and
10% respectively. For dpKA2 formulations with more than 15% mol, the peptide
should exist in ‘brush’ conformation. Below 15% mol the peptide would exist in
‘mushroom’ conformation. Since dpKA3 conjugates are longer, <10% mol
surface grafting forms mushroom conformation. Any grafting density higher than
10% mol exists as a ‘brush’ conformation (Table 9). The model utilized for the
characterization does not account for poly peptide chain crowding and cis-proline
conformation. The model also does not account for the peptide distribution
between bilayers.
129
Figure 25: Low- surface density ELP vesicles have rapid release kinetics.
dpKA2 and dpKA3 liposomes were stepped from 25 ° C t o 35
o
C and 45
o
C
respectively. (a) 60 μM of dpKA2 liposomes also showed rapid leakage of
contents (~ 80%, above) when temperature was stepped to 45
o
C (% release ±
SD) but minimal leakage at 35 ° C (below). (b) Simil arly, 50 μM of dpKA3
liposomes showed rapid content leakage (above) when temperature was stepped
to 45
o
C (% release ± SD) but minimal leakage was observed at 35 ° C (below).
130
Figure 26: Polypeptide brushes prevent particle aggregation and contents
release is due to pore formation. (a) DLS ramps show formation of large
aggregates in liposomes with low surface grafting. Liposomes with a dense
‘brush’ remained stable throughout the ramp. (b) Contrary to DLS findings
suggesting formation of large aggregates low surface grafted liposomes show
minimal membrane fusion (% fusion ± SD) proving that contents release is due to
pore formation and not membrane fusion.
131
Table 9: Relationship between grafted polymer conformation and extent of
release
Prep
(#)
Label Polymer
density
(%)
*Polymer
MW
(Da)
†Predicted
conformation
Maximum
release
(%)
1 10%
PEG
10 2000 Brush None
2 5%
dpKA2
5 780 Mushroom 75
3 10%
dpKA2
10 780 Mushroom 63
4 20%
dpKA2
20 780 Brush 18
5 1%
dpKA3
1 1161 Mushroom 93
6 5%
dpKA3
5 1161 Brush 30
7 10%
dpKA3
10 1161 Brush 14
8 20%
dpKA3
20 1161 Brush 11
*molecular weight of surface-grafted polymer
†based on Eq 11,12 & 13
132
4.3.7. Peptide conformation influences cell uptake but not DOX delivery
Liposomes with mushroom (1% dpKA3) and brush (20% dpKA3 and 10% PEG
2000) confirmation were evaluated in the cell uptake assay. Liposomes with
‘mushroom’ conformation showed higher uptake than liposomes in MDA-231.
Liposomes with a ‘brush’ border (Fig. 27) showed minimal uptake cells with both
dpKA3 and PEG 2000. Interestingly, DOX loaded liposomes with both mushroom
and brush surface conformations successfully delivered DOX to cells (Fig. 28).
As evidenced with DiI, NBD labeled liposomes showed higher uptake with a
mushroom conformation than brush, but both formulations efficiently deliver DOX
intracellular to the nucleus.
133
Figure 27: Polypeptide brushes reduce particle uptake in vitro. MDA-231
cells were incubated with DiI (red) labeled liposomes for 1 hr to study particle
uptake. Liposomes with a ‘mushroom’ conformation,1% dpKA3, showed higher
cell uptake than both 20% dpKA3 and 10% PEG 2000 brush border liposomes.
Scale bar represents 20 μm.
134
4.4. Discussion
The CMC of the utilized ELPAs was determined to be 1.69 μM (Aluri et al., 2012)
and is similar to other PC lipids (Kanamoto et al., 1981). The LCST of dpKA2 and
dpKA3 is concentration dependent and behave similar to higher molecular weight
ELPs. Surface grafting of ELPAs decreased the apparent LCST causing
membrane permeability. The decrease in LCST is due to the high local surface
concentration on the liposome surface. ELPAs with hydrophobic guest residues
could not be evaluated due to their low LCST and the hydrophobicity of total
ELPA molecule. Also dpKA2 liposomes are unstable and precipitate if kept
longer than 2 days.
Surface grafted hydrophilic ELPA permeabilized liposomes when heated.
Increasing mol% of ELPA reduced contents leakage. The decrease in contents
release is due to the conformation of the surface polymer. The different surface
conformations are due to various mol% of added ELPAs. ELPAs in ‘mushroom’
conformation had higher contents leakage than ELPAs in ‘brush’ conformation.
ELPAs grafted at high mol% exist in ‘brush’ conformation leading to rigid packing
of the ELPs. The packed surface inhibits the ELPs ability to form a mature β-
spiral which is essential for liposome permeabilization. In ‘mushroom’
conformation the low grafting densities promotes β-spiral formation which causes
content leakage. The peptide grafting concentrations also influences the cellular
uptake with ‘mushroom’ conformation having higher cellular uptake than
liposomes with a ‘brush’ border. The exact mechanism of cellular uptake has to
be investigated further. This higher cellular uptake maybe due to the increased
135
accessibility of fusogenic lipids that promote membrane fusion. In spite of having
higher cellular uptake, the mushroom conformation showed similar DOX delivery
potential as the liposomes in brush conformation. Hence by using both DiI and
DOX we have shown delivery of two different lipophilic drugs using the
aforementioned liposomal formulations.
136
Figure 28: Peptide conformations do not effect DOX delivery from
liposomal formulations. DOX (red) encapsulated in NBD (green) labeled
liposomal formulations efficiently deliver DOX to 231 cell nucleus (blue). The
peptide conformation had no significant effect on the delivery of DOX. Scale bar
represents 20 μm.
137
4.5. Conclusion.
We have extensively studied the effect of ELPs on the surface of liposomes. The
ELP phase transition depends on the surface conformation of the peptide
polymer. Liposomes can effectively be used as a tool to study peptide
interactions at very small distances. ELPs grafted liposome application as drug
delivery vehicles was demonstrated using two different lipophilic molecules in DiI
and DOX. The grafting concentration also influences uptake making the system
tunable in two different properties. Hence by fine tuning the properties of these
liposomes can selectively cause drug release and enhance particle uptake locally
at tumor site. This strategy can also reduce off-target toxicity and increase tumor
therapy efficacy.
4.6. Acknowledgments
The authors would like to thank USC School of pharmacy and NIH/NIDDK for
their financial support.
138
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Abstract (if available)
Abstract
Chapter 1: The tumor microenvironment provides multiple cues that may be exploited to improve the efficacy of established chemotherapeutics
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Creator
Aluri, Suhaas Rayudu
(author)
Core Title
Polypeptide based drug carriers for anti cancer applications
School
School of Pharmacy
Degree
Doctor of Philosophy
Degree Program
Pharmaceutical Sciences
Publication Date
07/26/2014
Defense Date
06/14/2013
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(original),
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Tag
cancer,immunology,liposomes,nanoparticles,nanoworms,OAI-PMH Harvest,tumor targetting
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Mackay, John Andrew (
committee chair
), Camarero, Julio A. (
committee member
), Epstein, Alan L. (
committee member
), Hamm-Alvarez, Sarah F. (
committee member
), Olenyuk, Bogdan (
committee member
)
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saluri@usc.edu,suhaas.rayudu@gmail.com
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Tags
immunology
liposomes
nanoparticles
nanoworms
tumor targetting