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Integrin-mediated targeting of protein polymer nanoparticles carrying a cytostatic macrolide

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Integrin-mediated targeting of protein polymer nanoparticles carrying a cytostatic macrolide

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INTEGRIN-MEDIATED TARGETING OF PROTEIN POLYMER NANOPARTICLES
CARRYING A CYTOSTATIC MACROLIDE
By

Pu Shi

A Thesis presented to the
Department of Pharmacology and Pharmaceutical Sciences
University of Southern California
In partial fulfillment of the
Requirements for the degree
DOCTOR OF PHILOSOPHY
PHARMACEUTICAL SCIENCES

August 2014

Copyright 2014 Pu Shi

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. J. Andrew Mackay for his
support and patience. I would also like to thank all of my committee members Dr. Sarah
Hamm-Alvarez, Dr. Wei-Chiang Shen, Dr. Curtis Okamoto and Dr. Alan Epstein for their
feedback and time spent in reviewing my thesis. I would like to thank all my lab member
and friends for their support.
iv
Table of Contents

Dedication……………………………………………………………… ii
Acknowledgements…………………………………………………... iii
List of Tables………………………………………………………….. ix
List of Figures…………………………………………………………. x
Abbreviations………………………………………………………….. xiii
Abstract…………………………………………………………………. xiv
1.0. Chapter 1: Genetically engineered nanocarriers…………..............
for drug delivery
1.1. Introduction…………………………………………………….....
1.2. Genetically engineered polymeric drug carriers and their
nanostructures……………………………………………………
1.2.1. ELPs…………………………………………………………...
1.2.2. SLPs…………………………………..……..……..……..…..
1.2.3. SELPs…………………………….……………………………
1.2.4. Other polymeric genetically engineered drug carriers…...
1.3. Polymeric nanocarriers that mediate drug delivery…………...
1.3.1. ELP-mediated drug delivery………………………………….
1.3.2. SLP-mediated drug delivery………………………………….
1.3.3. SELP-mediated drug delivery……………………………......
1.4. Non-polymeric drug carriers and their structures…….……….
1.4.1. Vault protein………………………………………………….
1

1

6
6
10
10
11
12
12
17
18
22
22
v
1.4.2. Viral proteins……………………………………………….
1.5. Drug delivery using non-polymeric protein nanoparticles…….
1.5.1. Vault protein mediated drug delivery……………………….
1.5.2. Viral protein-medicated drug delivery……………………..
1.6. Discussion…………………………………………………………….
1.7. Conclusion…………………………………………………………..
2.0. Chapter 2: Elastin-based protein polymer nanoparticles carrying
drug at both corona and core suppress tumor growth
in vivo………………………………………………………

2.1. Introduction……………………………………………………….
2.2. Materials and methods……………………………………………..
2.2.1. Materials and reagents…………………..……………………..
2.2.2. Biosynthesis of ELPs…........................................................
2.2.3. Optical density measurement of ELP phase transition…….
2.2.4. Dynamic light scattering (DLS) analysis and zeta potential
measurement..……….......................................................
2.2.5. Transmission electron microscopy (TEM) and cryogenic
transmission electron microscopy (Cryo-TEM)……………

2.2.6. ELP-mediated encapsulation of a model drug,
Rose Bengal……………………………………………………

2.2.7. Thin-film encapsulation of an anti-proliferative drug, Rapa
23
23
24
28
29
31

33
33

36

36

36

39

39

40

40

43

2.2.8. Rapa encapsulation and release using FSI nanoparticles
2.2.9. Cell proliferation assay……………………………………….
2.2.10. In vivo evaluation of FSI/Rapa in human breast cancer
43
44
44
vi
Xenografts……………………………………………………….
2.3. Results and discussion…………………………………………..
2.3.1. Characterization of ELP diblock copolymer nanoparticles...
2.3.2. Nanoparticle assembly is required for entrapment of Rose
Bengal into plain ELP nanoparticles………………………...
2.3.3. The water insoluble drug, Rapa encapsulates into unmodified
ELP diblock copolymer nanoparticles……………………….
2.3.4. Decoration of ELP nanoparticles with FKBP protein and loading
with Rapa minimally influences nanoparticle dimensions……

2.3.5. Decoration of ELP nanoparticles with the FKBP domain prolongs
the release of Rapa …………..................................................
2.3.6. FSI Rapa is as potent as free Rapa in MTS cell proliferation assay
2.3.7. FSI Rapa nanoparticles have greater anti-tumor efficacy and lower
toxicity than free drug……………………………………………..
2.4. Conclusion……………………………………………………….
2.5. Acknowledgements……………………………………………..
47
45
45
50
53
56

61
64
67
73
73
3.0. Chapter 3: Triggered Sorting and Co-Assembly of Genetically
Engineered Protein Microdomains in the Cytoplasm….

3.1. Introduction………………………………………………………
3.2. Materials and Methods…………………………………………
3.2.1. Biosynthesis and Characterization of ELPs………………..
3.2.2. In Vitro Co-Assembly and Sorting…………………………..
3.2.3. Lentivirus (rLV) Production and Generation of Stable Cell
Line Hek-DsRed-V96 Expressing DsRed-V96……………..
75
75
78
78
78
79
vii
3.2.4. Intracellular Co-Assembly and Sorting……………………...
3.2.5. Statistical Analysis …………………………………………….
3.3. Results and discussion……………………………………………

3.3.1. Assembly of various genetically engineered protein
microdomains from ELPs……………………………………

3.3.2. In vitro co-assembly and sorting of genetically engineered
protein microdomains…………………………………………..

3.3.3. Intracellular co-assembly and sorting of genetically
engineered protein microdomains…………………………….

3.4. Conclusion……………………………………………………….

3.5. Acknowledgements……………………………………………..

4.0. Chapter 4: A co-assembled multi-functional protein-polymer
nanocarrier delivers specific chemotherapeutics, actively
targets tumor and suppresses its growth…………………

4.1. Introduction……………………………………………………….
4.2. Materials & Methods ……………………………………………..
4.2.1. Biosynthesis and characterization of ELPs………………....
4.2.2. Co-assembly of ELP micelle nanoparticles…………………
4.2.3. Integrin-mediated cell targeting assay and
fluorescence-activated cell sorting (FACS)………..………...
4.2.4. Evaluation of Rapamycin (Rapa) encapsulation and release
4.2.5. In vivo tumor regression studies using human breast cancer
xenografts………………..…………….....................................

4.3. Results and discussion………………………………………………

79
80
80
80
87
94
97
97
99
99
103
103
103
104
105
105
106

viii
4.3.1. Co-assembly of multi-functional ELP micelle nanoparticle
4.3.2. In vitro tumor targeting of the co-assembled multi-functional
ELP micelle nanoparticle……………………………………..
4.3.3. In vitro evaluation of drug (Rapa) loading and release……
4.3.4. The multi-functional nanocarrier delivers Rapa, targets
tumor and inhibit its growth with greater efficiency compared
to other formulations in vivo…………………………………….
4.4. Conclusion…………………………………………………………

5.0 Summary and Future Directions……………………………………….

6.0. References…………………………………………………………………

106

111
115

118

124

125

129

ix
List of Tables
Table 1: ELP protein polymers evaluated in Chapter 2…………………………….. 38
Table 2: ELP protein polymers of different molecular weight, hydrophobicity, and
polymer architecture with similar assembly temperatures…………………..

83
Table 3: ELPs evaluated in Chapter 4………………………………………………... 102

x
List of Figures
Figure 1: Design of genetically engineered drug carriers……………………..... 4
Figure 2: TEM of ELP nanoparticles. Diblock copolymers composed of ELPs
with various guest residues assemble micelles………………………. 8
Figure 3: Rapa encapsulated by FKBP-decorated nanoparticles has both
anticancer and immunosuppressive efficacy …………………………. 15
Figure 4: Silk-elastin-like protein polymers with different ratios of silk to elastin
self-assemble into various nanostructures ………………………………

20
Figure 5: Vault protein engineered for hydrophobic drug delivery……………..... 26
Figure 6: The differential association of selected small molecules to the
hydrophobic core ELP of I48S48 nanoparticles………………………. ..

42
Figure 7: Design of ELP nanoparticles that carry anti-proliferative drugs —
decoration with protein drug receptors minimally influences
assembly……………………………………………………………………. 48
Figure 8: Assembly of ELP nanoparticles slows the release of a water soluble
model drug, Rose Bengal…………………………………………………..

52
Figure 9: Thin film hydration with ELP nanoparticles promotes encapsulation of
the hydrophobic drug, Rapa ………………………………………………. 54
Figure 10: Morphology of ELP nanoparticles is minimally influenced by fusion of
FKBP or encapsulation of Rapa………………………………………….

58
Figure 11: TEM and cryo-TEM images of I48S48 micelles……………………….. 59
Figure 12: TEM images of FSI before and after Rapa encapsulation …………… 60
Figure 13: The FKBP domain prolongs the release of Rapa from FSI
nanoparticles …………………………………………………….………... 63
Figure 14: FSI-encapsulated Rapa inhibits cell viability in an mTOR dependent
cell line MDA-MB-468 but not in an insensitive cell line, MDA-MB-231 65
Figure 15: The relationship between the critical micelle concentration (CMC) and
the critical micelle temperature (CMT) for FSI nanoparticles…………... 66
xi
Figure 16: FSI-encapsulated Rapa has better anti-tumor efficacy and lower
toxicity than free Rapa in the mTOR dependent MDA-MB-468 breast
cancer xenograft…………………………………………………………… 68
Figure 17: Co-assembly versus self-sorting of Genetically Engineered Protein
Microdomains (GEPMs) in eukaryotes………………………………….. 77
Figure 18: Tunable assembly of micro-structures from hydrophobic ELP
monoblocks and amphiphilic ELP diblock copolymers ……………….. 84
Figure 19: Micron-scale GEPMs assembled by monoblock ELPs can be
observed using negative staining TEM imaging……………………….. 86
Figure 20: Only monoblock and diblock copolymers spatially sort into distinct
GEPMs………………………………………………………………………
.
90
Figure 21: Confocal imaging of mixtures of Rho and CF labeled ELPs below the
matched transition temperature………………………………………….. 92
Figure 22: Self-sorting between monoblock and diblock GEPMs is concentration
Independent………………………………………………………………… 93
Figure 23: Fluorescent ELP fusion proteins in the eukaryotic cytosol can co-
assemble or self-sort GEPMs…………………………………………….. 96
Figure 24: Co-assembly of multi-functional ELP micelle nanoparticle................... 108
Figure 25: Physicochemical properties of ISR and FSI……………………………. 109
Figure 26: Optical density measurement and DLS analysis of ISR and FSI
Mixtures…………………………………………………………………….. 110
Figure 27: Integrin-mediated cell targeting assay using co-assembled mixed
micelle nanoparticles …………………………………………………….. 113
Figure 28: Competitive binding assay using cyclic-RGD (c-RGDfK) and ISR
demonstrated the direct binding of ISR to the integrin on cell surface 114

Figure 29: ISR/FSI co-assembled mixed micelle nanoparticle has great Rapa
loading capacity and releases the drug similarly to FSI………………. 117

Figure 30: Mouse tumor regression studies demonstrated actively tumor
targeting ISR/FSI Rapa had greater anti-tumor efficacy than passively
tumor targeting FSI Rapa and free Rapa in MDA-MB-468 breast
xii
tumor xenografts…………………………………………………………… 121

Figure 31: Stable mouse body weights during tumor regression studies
indicated that the formulations did not have obvious cytotoxicity…… 123
xiii
Abbreviations
Elastin-like polypeptides (ELPs), silk-like polypeptides (SLPs), extended recombinant
polypeptide (XTEN), silk-elastin-like polypeptides (SELPs), deoxyribonucleic acid
(DNA), rapamycin (Rapa), positron emission tomography (PET), tumor-homing peptides
(THPs), small interfering RNA (siRNA), major vault protein (MVP), Virus-like particles
(VLPs), Virus-like particles (VLPs), cowpea mosaic virus (CPMV), human cervical
cancer cell line (HeLa), mammalian target of rapamycin (mTOR), inverse transition
cycling (ITC), dynamic light scattering (DLS), transmission electron microscopy (TEM),
cryogenic-transmission electron microscopy (Cryo-TEM), FK506 binding protein
(FKBP), 3-(4,5-dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-
tetrazolium (MTS), critical micelle concentration (CMC), genetically engineered protein
microdomains (GEPMs), rhodamine (Rho), carboxyfluorescein (CF), N-
Hydroxysuccinimide (NHS), enhanced permeability and retention (EPR), transition
temperature (Tt), 4',6-diamidino-2-phenylindole (DAPI), Institutional Animal Care and
Use Committee (IACUC), fluorescence-activated cell sorting (FACS), dialyzed slow-
release (Ds), human umbilical vein endothelial cell (HUVEC)

xiv
Abstract
Chapter 1: Cytotoxicity, low water solubility, rapid clearance from circulation, and off-
target side-effects are common drawbacks of conventional small-molecule drugs. To
overcome these shortcomings, many multifunctional nanocarriers have been proposed
to enhance drug delivery. In concept, multifunctional nanoparticles might carry multiple
agents, control release rate, biodegrade, and utilize target-mediated drug delivery;
however, the design of these particles presents many challenges at the stage of
pharmaceutical development. An emerging solution to improve control over these
particles is to turn to genetic engineering. Genetically engineered nanocarriers are
precisely controlled in size and structure and can provide specific control over sites for
chemical attachment of drugs. Genetically engineered drug carriers that assemble
nanostructures including nanoparticles and nanofibers can be polymeric or non-
polymeric. This chapter summarizes the recent development of applications in drug and
gene delivery utilizing nanostructures of polymeric genetically engineered drug carriers
such as elastin-like polypeptides, silk-like polypeptides, and silk-elastin-like protein
polymers, and non-polymeric genetically engineered drug carriers such as vault proteins
and viral proteins.

Chapter 2: Numerous nanocarriers of small molecules depend on either non-specific
physical encapsulation or direct covalent linkage. In contrast, this chapter explores an
alternative encapsulation strategy based on high-specificity avidity between a small
molecule drug and its cognate protein target fused to the corona of protein polymer
nanoparticles. With the new strategy, the drug associates tightly to the carrier and
xv
releases slowly, which may decrease toxicity and promote tumor accumulation via the
enhanced permeability and retention effect. To test this hypothesis, the drug Rapamycin
(Rapa) was selected for its potent anti-proliferative properties, which give it
immunosuppressant and anti-tumor activity. Despite its potency, Rapa has low
solubility, low oral bioavailability, and rapid systemic clearance, which make it an
excellent candidate for nanoparticulate drug delivery. To explore this approach,
genetically engineered diblock copolymers were constructed from elastin-like
polypeptides (ELPs) that assemble small (b100 nm) nanoparticles. ELPs are protein
polymers of the sequence (Val-Pro-Gly-Xaa-Gly)n, where the identity of Xaa and n
determine their assembly properties. Initially, a screening assay for model drug
encapsulation in ELP nanoparticles was developed, which showed that Rose Bengal
and Rapa have high non-specific encapsulation in the core of ELP nanoparticles with a
sequence where Xaa = Ile or Phe. While excellent at entrapping these drugs, their
release was relatively fast (2.2 h half-life) compared to their intended mean residence
time in the human body. Having determined that Rapa can be non-specifically
entrapped in the core of ELP nanoparticles, FK506 binding protein 12 (FKBP), which is
the cognate protein target of Rapa, was genetically fused to the surface of these
nanoparticles (FSI) to enhance their avidity towards Rapa. The fusion of FKBP to these
nanoparticles slowed the terminal half-life of drug release to 57.8 h. To determine if this
class of drug carriers has potential applications in vivo, FSI/Rapa was administered to
mice carrying a human breast cancer model (MDA-MB-468). Compared to free drug,
FSI encapsulation significantly decreased gross toxicity and enhanced the anti-cancer
activity. In conclusion, protein polymer nanoparticles decorated with the cognate
xvi
receptor of a high potency, low solubility drug (Rapa) efficiently improved drug loading
capacity and its release. This approach has applications to the delivery of Rapa and its
analogs; furthermore, this strategy has broader applications in the encapsulation,
targeting, and release of other potent small molecules.
Chapter 3: Elastin-like polypeptides (ELPs) are genetically encoded protein polymers
that reversibly phase separate in response to stimuli. They respond sharply to small
shifts in temperature and form dense microdomains in the living eukaryotic cytosol. This
chapter illustrates how to tune the ELP sequence and architecture for either co-
assembly or sorting of distinct proteins into microdomains within a living cell.

Chapter 4: Passive tumor targeting utilizing enhanced permeability and retention (EPR)
effect has limited efficiency in targeting non-leaky tumors such as MDA-MB-468 breast
tumor; however, an RGD tri-peptide decorated micelle nanoparticle can effectively
accumulate in tumor site via integrin-mediated active tumor targeting. Different from
inefficient and cytotoxic chemical linkage reactions, an elastin-based multi-functional
nanocarrier can be assembled by genetic protein fusion and micelle co-assembly
technology. The novel drug carrier contains the cognate Rapamycin (Rapa) receptor –
FK506 binding protein (FKBP) as the high-avidity drug binding domain and an RGD
peptide as the active tumor targeting domain. Here we show that by co-assembling
FKBP and RGD contained protein polymers into mixed micelle nanoparticles, they not
only competently targeted endothelial and tumor cells in cell assays, but specifically
delivered the drug with a slow release half-life of 38h. It was demonstrated that the
xvii
active tumor targeting formulation of Rapa more effectively suppressed tumor growth
compared to the passive tumor targeting formulation and free drug in tumor regression
studies of mouse MDA-MB-468 xenografts. We believe that the exciting results will
provide a new tool for the development of next-generation “smart” multi-functional drug
carriers.

1
Chapter 1
Genetically engineered nanocarriers for drug delivery
1.1 Introduction
Drug delivery systems are designed to lower toxicity and improve
pharmacokinetic/pharmacodynamic profiles of conventional drugs (Langer, 1998).
Following intensive development by many groups, numerous drug carriers have been
successfully developed. For the purposes of this chapter, we consider two general
classes of drug carriers: chemically synthesized carriers, and genetically engineered
carriers (Rabotyagova et al., 2011). Focusing on the latter, this chapter compares these
two different types of carriers in various aspects including features of the carriers,
carrier synthesis, and cytotoxicity, for example. Significant progress has been made in
the field of synthetic polymers to increase the efficiency of polymerization techniques
and lower polydispersities (Rabotyagova et al., 2011, Hecht, 2005). Despite this
progress, genetic engineering provides unparalleled control over the component
macromolecules used to build nanoparticulate carriers (Janib et al., 2014b). This
capability allows unique characteristics such as specific biodegradation profiles and fully
customized polymer and nanocarrier architectures to be engineered and modified as
needed for specific applications. Unlike chemically synthesized drug carriers, current
research into genetically engineered carriers only scratches the surface of potential
applications. During its emergence, it already produced multiple carriers that show
unique potential for clinical application. Examples of such technologies include a highly
potent yet side-effect-limiting doxorubicin formulation (MacKay et al., 2009), a
genetically engineered nanoparticle which effectively targets the coxsackievirus and
2
adenovirus receptor (Sun et al., 2011), and naturally derived carrier proteins which
entrap and allow delivery of hydrophobic drugs (Buehler et al., 2011).

Unlike chemically synthesized carriers, proteins offer unique opportunities to form
nanostructures based on the well-established secondary, tertiary, and quaternary
structures commonly found in natural proteins. Well studied secondary structures such
as -helices and -sheets can be used to bind together micro- or nanoparticle
structures (Koehl and Delarue, 1994, Elemans et al., 2003). In addition to unique
structural opportunities, genetically engineered drug carriers have hierarchical
structures (Janib et al., 2014b, Sun et al., 2011), on which structure–function studies
might be accomplished by site-directed mutagenesis at the primary amino acid
sequence. At the current time, genetically engineered drug carriers can be divided into
two categories in the consideration of primary amino-acid sequences: 1) polymeric
genetically engineered drug carriers, and 2) non-polymeric genetically engineered drug
carriers. The relationship between vehicles for drug delivery, nanostructure formation,
and protein polymers is visually conveyed in Figure 1, in which intersections 1, 2, and 3
represent polymeric drug carriers with nanostructures, non-polymeric carriers with
nanostructures, and polymeric carriers without nanostructures, respectively. Examples
of well-developed genetically engineered drug carriers include protein polymers
composed from elastin-like polypeptides (ELPs) (Urry, 1997), silk-like polypeptides
(SLPs) (Valluzzi et al., 2002), extended recombinant polypeptide (XTEN) polymers
(Schellenberger et al., 2009), and silk-elastin-like polypeptides (SELPs) (Megeed et al.,
2002). Alternatively, non-polymeric genetically engineered drug carriers with defined
3
tertiary and quaternary structure have been developed from viral proteins and vault
proteins (Georgens et al., 2005, Buehler et al., 2011). From the perspective of
sequence–structure relations, genetically engineered drug carriers present varying
macro-, micro-, or nanoscale properties, with differences in length and composition of
amino-acid sequences. In this chapter, we primarily aim to provide a summary of
polymeric and non-polymeric genetically engineered drug carriers, and focus on their
drug-delivery applications using various nanostructures.

For the purposes of this chapter, a “polypeptide” is defined as a repetitive amino acid
sequence built from a short motif. The term “protein polymer” is defined as an amino
acid sequence that fulfills roles (eg, electrostatic or steric repulsion) filled by synthetic
polymers. A protein polymer may (ELP) or may not (XTEN) be a polypeptide and may or
may not produce secondary structure. In contrast, the term “protein” is defined as a non-
repetitive amino acid sequence that generates tertiary and quaternary structures (vault
and viral particles), which produce specific molecular functions.

4

Figure 1 Design of genetically engineered drug carriers.
The field of biological nanomedicine (aka “BioNano”) is emerging at the intersections
between genetically engineered biomaterials, nano-assembly, and protein polymers. At
intersection 1, nanomedicines are being developed from protein polymers (eg, ELP,
SLP, and SELP). At intersection 2, protein-based materials (eg, viral capsids and vault
proteins) are being developed as platforms for assembly of nanostructures. At
intersection 3, proteins that avoid structure formation (eg, intrinsically disordered
proteins and XTEN fusion proteins) are being explored for their ability to alter
biodistribution and efficacy.
Adapted with permission from Galaway FA, Stockley PG. MS2 viruslike particles: a
robust, semisynthetic targeted drug delivery platform. Mol Pharm. 2013;10(1): 59–68.55
Copyright 2013 American Chemical Society, and Buehler DC, Toso DB, Kickhoefer VA,
5
Zhou ZH, Rome LH. Vaults engineered for hydrophobic drug delivery. Small.
2011;7(10):1432–1439.7 Copyright 2011 WILEY-VCH Verlag GmbH & Co. KGaA,
Weinheim.

6
1.2 Genetically engineered polymeric drug carriers and their nanostructures
Protein polymers consist of natural or unnatural repetitive amino-acid sequences and
are generally biosynthesized in cells, either prokaryotes or eukaryotes. Because protein
polymers can be engineered at the genetic level, their sequences can be precisely
controlled (Frandsen and Ghandehari, 2012, Urry et al., 2010, Cappello et al., 1990).
One significant advantage of protein polymers is that by changing several amino acids
in the repetitive sequences, libraries of polymers with different charges, hydrophobicity,
or secondary structures can be created to perform structure–function studies (Frandsen
and Ghandehari, 2012, Urry et al., 2010, Shah et al., 2012b). Compared with
conventional polymers, protein polymers may cause lower cytotoxicity, which may be
due to the fact that they have biologically relevant mechanisms for proteolysis into
relatively inert amino acids (Urry et al., 2010, Gustafson et al., 2010). Since 1986, when
Ferrari et al (Ferrari et al., Issued 1993, application date 1986) reported the first protein
polymer designed to be a potent drug carrier, a number of different protein polymers
have been developed for use as drug carriers, such as ELPs (MacKay et al., 2009, Aluri
et al., 2012, Pastuszka et al., 2012a, Janib et al., 2013, Ravi et al., 2012), SLPs (Elia et
al., 2011, Numata et al., 2012, Numata et al., 2011), and SELPs (Xia et al., 2011).
1.2.1 ELPs
Elastin is a major extracellular matrix protein that provides resilience and elasticity in
tissues and organs of many higher animals. ELPs are protein polymers which consist of
repeats of amino acid sequence Val-Pro-Gly-Xaa-Gly ([VPGXG]n), derived from a highly
7
conserved repeat sequence in mammalian tropoelastin (Urry, 1997). In natural elastin,
the guest residue X is frequently valine, alanine, or isoleucine. When the identity of X is
changed in the context of ELPs, many interesting properties can be imparted and
precisely tuned, for example, reversible phase-separations in aqueous solutions
(Rabotyagova et al., 2011, Frandsen and Ghandehari, 2012, Kim et al., 2011, Amiram
et al., 2011). One intriguing use of this guest residue-dependent modification of polymer
properties has been the creation of ELP block copolymers, which have been
constructed by genetically linking a hydrophobic block and a hydrophilic block together,
for example, [VPGIG]n1-[VPGSG]n2 (Janib et al., 2014b). These block copolymers have
been verified to form stable nanoparticle structures ranging from 50–90 nm in diameter,
which have various functions in drug delivery, and the formation of which is dependent
on the difference between the transition properties of the hydrophilic and hydrophobic
blocks (Rabotyagova et al., 2011, Sun et al., 2011, Urry et al., 2010, Shah et al., 2012b,
Janib et al., 2013, Amiram et al., 2011, Simnick et al., 2011, Fluegel et al., 2011). Figure
2 illustrates a series of ELP micelle nanoparticles formed by repetitive amino-acid
sequences with different guest residues in hydrophobic and hydrophilic blocks (Janib et
al., 2014b, Janib et al., 2013, Dreher et al., 2007, Garcia-Arevalo et al., 2013).

8

Figure 2 TEM of ELP nanoparticles. Diblock copolymers composed of ELPs with
various guest residues assemble micelles.
(A) TEM image of A96I96, which has a hydrophilic (Xaa = Ala, n=96, N-terminus)
and a hydrophobic (Xaa = Ile, n=96, C-terminus) block. Scale bar 50 nm. From
Janib SM, Liu S, Park R, et al. Kinetic quantification of protein polymer
nanoparticles using non-invasive imaging. Integr Biol (Camb). 2013;5(1):183–
194.23 Reproduced by permission of The Royal Society of Chemistry. (B) TEM
image of ELP I96S96, which has a hydrophilic (Xaa = Ser, n=96, C-terminus)
and a hydrophobic (Xaa = Ile, n=96, N-terminus) block. Scale bar 200 nm. From
Janib SM, Pastuszka MF, Aluri S, et al. A quantitative recipe for engineering
protein polymer nanoparticles. Polym Chem. 2014;5(5):1614–1625.4
Reproduced by permission of The Royal Society of Chemistry. (C) Cryo-TEM
9
image of ELP E50I60, which has a hydrophilic (Xaa = Val:Glu [4:1], n=50, N-
terminus) and a hydrophobic (Xaa = Ile, n=60, C-terminus) block. Scale bar 100
nm. Reproduced with permission from García-Arévalo C, Bermejo-Martín JF,
Rico L, et al. Immunomodulatory nanoparticles from elastin-like recombinamers:
single-molecules for tuberculosis vaccine development. Mol Pharm.
2013;10(2):586–597.34 Copyright 2013 American Chemical Society. (D) Cryo-
TEM image of ELP-96/90, which has a hydrophilic (Xaa = Val:Ala:Gly [1:8:7],
n=96, N-terminus) and a hydrophobic (Xaa = Val, n=90, C-terminus) block. Scale
bar 20 nm. Reprinted with permission from Dreher MR, Simnick AJ, Fischer K, et
al. Temperature triggered self-assembly of polypeptides into multivalent
spherical micelles. J Am Chem Soc. 2008;130(2):687–694.33 Copyright 2008
American Chemical Society.

10
1.2.2 SLPs
Silk proteins are natural polymers produced by either silkworms or spiders. Silkworm
silk fibroin from Bombyx mori and spider silk fibroin from Nephila clavipes are two of the
most well studied silk proteins at present (Lewis, 2006, Pritchard and Kaplan, 2011).
They generally are considered to be block copolymers with highly conserved repeats of
short side-chain amino acids as hydrophobic blocks and short sequences of larger side-
chain or charged amino acids as hydrophilic blocks (Rabotyagova et al., 2011, Numata
and Kaplan, 2010). The most common amino-acid sequence of SLPs derived from
Bombyx mori silkworms is the [GAGAGS]n repeat, while the most common spider silk
SLP is [GRGGLGGQGAGAAAAAGGAGQGGYGGLGSQG]n, derived from Nephila
clavipes. With the incorporation of cationic polylysine and/or polyarginine
deoxyribonucleic acid (DNA)-binding domains, nanofibers, and nanoparticles formed by
SLPs have been successfully applied to the field of gene delivery. Spider silk-based
nanoparticles containing tumor-homing peptides such as F3
(KDEPQRRSARLSAKPAPPKPEPKPKKAPAKK), Lyp1 (CGNKRTRGC), and CGKRK
and poly(L-lysine) domains have been demonstrated to deliver target-specific plasmid
DNA (pDNA) to the tumor cells (MDA-MB-435 and MDA-MB-231) with low cytotoxicity
and high efficiency. Therefore, these nanoparticles may have potential to be utilized as
DNA carriers in cancer gene therapy (Numata et al., 2012, Numata and Kaplan, 2010).

1.2.3 SELPs
SELPs have both motifs from the silkworm silk sequence [GAGAGS]n and mammalian
tropoelastin sequence [VPGVG]n. Because the silk blocks of SELPs tend to form -
11
sheet structures with intensive inter- and intramolecular hydrogen-bond interactions,
SELPs with high silk content precipitate out of aqueous solution at relatively low
concentration (Cappello et al., 1990). However, with the increasing repeat number of
elastin blocks that disorder the formation of crystalline silk structures, the entire SELP
block becomes water soluble at low temperature. This property is critical for protein
polymer purification and drug-delivery formulation (Cappello et al., 1998). SELP
hydrogels are formed after an irreversible phase transition, which makes them
amenable to the development of solvent-free injectable depots. As such, SELPS are
one of the most exciting emerging carriers for drugs and gene therapy (Price et al.,
2012). It has been speculated that the formation of a micelle core by hydrophobic
interactions of the silk blocks is the driving force to assemble SELP nanoparticle
structures that have potential drug-delivery applications (Xia et al., 2011).
1.2.4 Other polymeric genetically engineered drug carriers
Besides ELPs, SLPs, and SELPs, there are many other types of polymeric genetically
engineered drug carriers. For example, Farmer and Kiick have created alanine-rich
helical proteins. This repeating helical protein, which contains glutamine and glutamic
acid, can form nanofibril structures and is being developed as a multivalent drug
nanocarrier (Farmer and Kiick, 2005). Amunix Inc., a biotechnology company focused
on protein polymers, has genetically engineered a long protein polymer termed XTEN,
which is composed of hydrophilic and negatively charged residues. It has been
demonstrated that the half-life of many protein therapeutics can be drastically increased
by attaching XTEN, and the half-life can be tuned by varying the length of the XTEN
polymer (Schellenberger et al., 2009). The exact mechanism of how XTEN increases
12
protein half-life has not been fully delineated; however, it is plausible that slow
proteolytic biodegradation, high molecular weight, and anionic electrostatics repel the
extracellular matrix in the glomerulus. By reducing the rate of renal clearance, XTEN
polymers may prolong protein half-life similarly to a synthetic polymer (eg, polyethylene
glycol, HPMA [N-(2-hydroxypropyl)methacrylamide], and dextran).

1.3 Polymeric nanocarriers that mediate drug delivery
Protein polymers such as ELPs, SLPs, and SELPs can form nanoparticle structures
under certain conditions. In the last 2–3 years, multiple articles have been published
focusing on these nanocarriers in the delivery of genes and drugs (Sun et al., 2011,
Numata et al., 2012, Numata et al., 2011, Kim et al., 2011, Simnick et al., 2011,
Dhandhukia et al., 2013, Shah et al., 2013, Shi et al., 2013b).
1.3.1 ELP-mediated drug delivery
In the ELP field, Sun et al utilized fusion protein technology to decorate the corona of an
ELP block copolymer G[VPGSG]48-[VPGIG]48 with several different useful proteins
(Sun et al., 2011). The attachment of the knob domain of adenovirus serotype 5 fiber
protein to the serine block of this polymer was one example of this technology. Using
dynamic light scattering, the genetically modified knob-ELP fusion protein was
measured to form nanoparticles with a 40 nm diameter with the knob domain on the
surface. Cellular uptake studies using a coxsackievirus and adenovirus receptor-
expressing hepatocyte cell line revealed that the knob-ELP fusion protein presented
significantly stronger colocalization to lysosomes inside hepatocytes than plain ELPs,
which indicated that the knob domain of adenovirus serotype 5 fiber protein was the
13
critical factor to facilitate targeted cellular internalization of the fusion protein
nanoparticles (Sun et al., 2011). Moreover, FKBP (FK506-binding protein), the cognate
receptor of an antiproliferative drug rapamycin (Rapa) has also been genetically fused
onto the corona of micelles assembled from the ELP block copolymer (Figure 3A)
(Dhandhukia et al., 2013, Shi et al., 2013b). Because of high-avidity binding of the drug
to the receptor, the new fusion protein (FKBP-ELP [FSI]) slowed the terminal half-life of
drug release to 57.8 hours (Figure 3B) (Shi et al., 2013b). The in vivo antitumor and
immunosuppressant applications of the new Rapa formulation (FSI-Rapa) were
respectively examined in a MDA-MB-468 breast cancer xenograft nude mouse model
and Sjögren’s syndrome non-obese diabetic mouse model. It was discovered that FSI-
Rapa showed not only significantly less cytotoxicity but greater efficacy in tumor
regression and autoimmune response suppression than the free drug, respectively, in
the two models (Figure 3C and D) (Shi et al., 2013b, Shah et al., 2013). In addition, the
blood half-lives of ELPs in mice were estimated by a multi-compartmental
pharmacokinetic model using the data from noninvasive micro-PET (positron emission
tomography) imaging. Depending on molecular weight and assembled structure, the
blood half-lives of ELPs vary from 2 to 6 hours in vivo (Figure 3E) (Janib et al., 2013).
Because the half-life of Rapa release is much longer than the blood half-lives of ELPs, it
has been speculated that the drug will remain associated with the carrier blood
circulation with minimal detachment, which may reduce systemic side effects.

The Simnick et al study employed the same fusion protein technology to genetically
revise the corona of ELP amphiphilic block copolymers with the NGR (Asn-Gly-Arg)
14
tripeptide ligand targeting the CD13 receptor (Simnick et al., 2011). NGR-decorated
ELP amphiphilic block copolymers were designed to competently target CD13 receptors
expressed highly in tumor vasculature and perivascular cells. The results showed NGR-
decorated ELP amphiphilic block copolymers formed particles 25–30 nm in radius
above the critical micelle temperature, and NGR micelles achieved greater vascular
retention and extravascular accumulation in tumor tissue compared with normal tissue
in an intravital laser scanning confocal fluorescence microscopy study (Simnick et al.,
2011). These successful studies obviously reveal the advantages of using protein
polymer nanocarriers such as ELPs – precise molecular modification at the genetic
level. It remains challenging to chemically attach a complex protein such as a drug
receptor to synthetic carriers; however, using protein fusion technology, the modification
on genetically engineered drug carriers can be completed seamlessly.
Another example of ELP-mediated drug delivery using nanoconstructs is elastin-mimetic
amphiphilic diblock copolymer. These materials are based on the sequence
[(VPGVG)(VPGEG)(VPGVG)(VPGEG)(VPGVG)]10-([Glu2]10) as the hydrophilic block
and [(IPGVG)2VPGYG(IPGVG)2]15-([Tyr]15) as the hydrophobic block. This block
copolymer has been confirmed to form stable micelles and efficiently solubilize and
encapsulate dipyridamole, a model drug with anti-inflammatory activity. In vitro and in
vivo drug release experiments have verified that the retention time of dipyridamole
inside the micelle core reduces with the decrease of the length of the hydrophobic block
repeats. It also has been discovered that dipyridamole encapsulation effectively
suppresses in vivo recruitment of neutrophils in the presence of an inflammatory
stimulus (Kim et al., 2011).
15

Figure 3 Rapa encapsulated by FKBP-decorated nanoparticles has both
anticancer and immunosuppressive efficacy.
Notes: (A) ELP nanoparticles fused genetically to the cognate receptor of Rapamycin-
FKBP can specifically carry the drug with high avidity. (B) FSI significantly prolongs drug
release compared with plain ELP (SI). (C) FSI-Rapa has lower cytotoxicity and greater
16
antitumor efficacy than free Rapa in an MDA-MB-468 breast tumor xenograft mouse
model. Compared with free Rapa group, which showed severe cytotoxicity (15%
bodyweight loss by day 23), no obvious systemic cytotoxicity was observed in the FSI-
Rapa group. (A), (B), (C) Reproduced from J Control Release,171(3), Shi P, Aluri S, Lin
YA, et al. Elastin-based protein polymer nanoparticles carrying drug at both corona and
core suppress tumor growth in vivo, 330–338,42 Copyright 2013, with permission from
Elsevier. (D) FSI-Rapa suppresses transcription and expression of the protease
cathepsin-S (CATS), a biomarker of lacrimal gland autoimmune dacryoadenitis, better
than free Rapa in a mouse model of Sjögren’s syndrome. Reproduced from J Control
Release, 171(3), Shah M, Edman MC, Janga SR, et al, A rapamycin-binding protein
polymer nanoparticle shows potent therapeutic activity in suppressing autoimmune
dacryoadenitis in a mouse model of Sjogren’s syndrome, 269–279,43 Copyright 2013
with permission from Elsevier. (E) The blood half-lives of four ELPs estimated by
pharmacokinetic modeling in mice based on micro-PET imaging. ***indicates a P-value
of <0.001; **indicates a P-value of <0.01 by one-way ANOVA with Tukey’s multiple
comparison test. From Janib SM, Liu S, Park R, et al. Kinetic quantification of protein
polymer nanoparticles using non-invasive imaging. Integr Biol (Camb). 2013;5(1):183–
194.23 Reproduced by permission of The Royal Society of Chemistry.

17

1.3.2 SLP-mediated drug delivery
There are a number of SLP-based materials being developed as nanocarriers for gene
delivery. Numata, Mieszawska-Czajkowska et al and Numata, Reagan et al have
employed natural silk sequence to fabricate genetically engineered silk-like recombinant
protein including poly(L-lysine) domains and tumor-homing peptides (THPs) (Numata et
al., 2012, Numata et al., 2011). pDNA can interact with poly(L-lysine) domains of the
silk-like recombinant protein and together form globular pDNA-silk nanocomplexes with
diameters from 100 to 250 nm. So far, four different types of THPs (F3, Lyp1,
monomeric CGKRK, and dimeric CGKRK) have been genetically engineered to two
different types of silk-like proteins. MDA-MB-435 melanoma cells and highly metastatic
human breast tumor MDA-MB-231 cells were used to test the binding of THP-targeted
pDNA-silk nanocomplexes to specific tumorigenic cells. In vitro and in vivo transfection
experiments have demonstrated that all types of pDNA-silk nanocomplexes with
different THPs (F3, Lyp1, monomeric CGKRK, and dimeric CGKRK) have specifically
targeted tumorigenic cells, while no targeting was found when MCF10A non-tumorigenic
mammary breast epithelial cells were used. Field emission scanning electron
microscopy was applied to investigate the mechanism of the specificity of pDNA-silk
nanocomplexes to tumorigenic cells. pDNA-silk nanocomplex with F3 THP was
observed to be absorbed through the surface of MDA-MB-231 cells but not through
MCF10A cells, which indicated that the targeting specificity might be caused by THP-
tumor cell surface receptor-mediated absorption (Numata et al., 2012, Numata et al.,
2011).
18

1.3.3 SELP-mediated drug delivery
SELPs have previously been used to form hydrogels that are used for localized gene
delivery (Gustafson and Ghandehari, 2010). So far, no study has been published
reporting direct SELP nanoparticle-based drug delivery. However, efforts have been
made to synthesize many promising SELP nanostructures that are potentially excellent
carriers for different drugs. Recently, Xia et al reported different structures such as
nanoparticles, hydrogels, and nanofibers could be reversibly or irreversibly assembled
by precisely tuning the ratio of silk to elastin of SELPs (Xia et al., 2011). As illustrated in
Figure 4A, three different SELPs (SE8Y, S2E8Y, and S4E8Y) were biosynthesized with
various silk to elastin ratios, and their morphological changes upon heating to 60°C
were studied by atomic force microscopy (Figure 4B). At 60°C, both SE8Y and S2E8Y
assembled spherical nanoparticles with a hydrodynamic radius of 241±13 and 212±16
nm, respectively; however, the silk blocks of S4E8Y underwent crosslinking to form gel
states instead of obvious nanoparticles. Distinct structures were observed in the
aqueous solutions of the three SELPs in the cooling-down process (Figure 4C). At
20°C, small micelle-like nanoparticles were observed in SE8Y solution, worm-like
nanostructures composed of small spherical particles were assembled by S2E8Y, and
large, polydisperse aggregates appeared in S4E8Y solution. Moreover, aligned
nanofibers by the crosslinking of the silk blocks were also observed for S2E8Y and
S4E8Y. The study demonstrated the formation of various nanostructures self-
assembled by SELPs which may be potentially used as nanocarriers for controlled drug
delivery (Xia et al., 2011). Another advantage of genetically engineered carriers over
19
chemically synthetic carriers is that structure–activity studies can be easily and
accurately performed with modifications in primary amino-acid sequences. By changing
DNA sequences, various SELPs can be biosynthesized with different elastin to silk
ratios (Xia et al., 2011).

20

Figure 4 Silk-elastin-like protein polymers with different ratios of silk to elastin
self-assemble into various nanostructures.
(A) SELP constructs SE8Y, S2E8Y, and S4E8Y, which contain varying ratios of the silk-
to-elastin blocks in each monomer repeat. (B and C) Atomic force microscopy images
present the nanostructures self-assembled from SE8Y, S2E8Y, and S4E8Y at 60°C (B)
and 20°C (C).
Adapted with permission from Xia XX, Xu Q, Hu X, Qin G, Kaplan DL. Tunable self-
assembly of genetically engineered silk – elastin-like protein polymers.
21
Biomacromolecules. 2011;12(11):3844–3850.28 Copyright 2011 American Chemical
Society.

22
1.4 Non-polymeric drug carriers and their structures
Non-polymeric proteinaceous drug carriers are characterized by their lack of repetitive
amino-acid sequences, and thus rely on the formation of self-assembled quaternary
structures to act as drug carriers. To date, there are many non-polymeric genetically
engineered drug carriers that have been developed and applied extensively in gene and
drug delivery (Buehler et al., 2011, Georgens et al., 2005, Han et al., 2011, Kar et al.,
2011, Pokorski and Steinmetz, 2011, Wu et al., 2012, Yildiz et al., 2011). In this chapter
of nanometer-scale drug carriers, we highlight two main categories of non-polymeric
genetically engineered drug carriers that form useful nanostructures: vault proteins and
viral proteins.

1.4.1 Vault protein
About 25 years ago, vault protein was discovered as the most bulky ribonucleoprotein
complex, with a size of 13 MDa (Kedersha and Rome, 1986). Vault protein itself forms
71 nm × 42 nm × 42 nm nanoparticles and is abundant and conserved in most
eukaryotes (Kedersha and Rome, 1986, Suprenant, 2002). Previous studies have
revealed the broad cellular functions of vault protein, including nuclear-cytoplasmic
transport, mRNA (messenger ribonucleic acid) localization, drug resistance, cell
signaling, nuclear pore assembly, and innate immunity (Berger et al., 2009, Rome and
Kickhoefer, 2013). The large ribonucleoprotein complex consists of three different types
of proteins: major vault protein (MVP), vault poly(adenosine diphosphate-ribose)
polymerase, and telomerase associated protein 1. Because MVP comprises 75% of the
native vault protein mass and is sufficient to form vault nanoparticles on its own, in most
23
studies only MVP was expressed.7 Vault nanoparticles are regarded to be promising
drug carriers because they 1) have a spacious internal volume (5 × 104 nm3), which is
adequate for the encapsulation of bioactive molecules; 2) consist of naturally occurring
amino-acid sequences which have no known cytotoxicity and immunogenicity and are
easy to modify at the genetic level; and 3) form a “dynamic” nanostructure which can
dissociate into halves in a low pH environment, improving their utility for drug release
(Buehler et al., 2011, Han et al., 2011, Kar et al., 2011, Goldsmith et al., 2007).

1.4.2 Viral proteins
Viral nanoparticle assemblies are another prevalent strategy in the field of non-
polymeric nanocarriers for drug delivery. Naturally, viruses can infect plants and animals
effectively and transfer their genetic materials (DNA, RNA, or proteins) to the host cells
(Yildiz et al., 2011). Virus-like particles (VLPs) take advantage of this highly evolved and
efficient transfer strategy to deliver their cargos by mimicking the natural process of
viruses. VLPs have their own advantages: 1) milligram quantities of VLPs can be
produced quickly and efficiently, which allows easy scale-up; 2) VLPs tend to be very
robust because of their protein capsids and are stable in a range of solvents; and 3)
VLPs possess great cell membrane penetration ability because of the viral features.
Bioconjugations and encapsulations have been performed on VLPs to achieve decent
gene and drug delivery in many studies (Georgens et al., 2005, Pokorski and Steinmetz,
2011, Wu et al., 2012, Yildiz et al., 2011).

24
1.5 Drug delivery using non-polymeric protein nanoparticles
Non-polymeric genetically engineered drug carriers such as vault protein and viral
proteins have nanometer-range structures. They have their own advantages serving as
drug nanocarriers. According to recent publications, many modifications over their
primary sequences of these nanocarriers have been accomplished to improve their
capability to specifically target disease cells or to efficiently encapsulate bioactive
molecules (Buehler et al., 2011, Han et al., 2011, Kar et al., 2011, Wu et al., 2012, Yildiz
et al., 2011, Choi et al., 2011, Galaway and Stockley, 2013).

1.5.1 Vault protein mediated drug delivery
Kar et al have been exploring vault protein-mediated drug delivery using several
different strategies (Kar et al., 2011). Recently, they have successfully loaded CCL21, a
lymphoid chemokine into vault nanostructures. CCL21 can naturally bind to CCR7, a
cellular chemokine receptor. Therefore, CCL21 can attract cells that highly express
CCR7 such as dendritic cells, naïve and memory T-cells, and natural killer and natural
killer T-cells to effectively kill cancer cells (Kar et al., 2011). CCL21 has been efficiently
encapsulated into vault nanoparticles by genetically fusing onto a vault-targeting domain
named INT (for vault INTeraction). It was discovered that the administration of CCL21-
INT-Vault complex into lung cancer mice enhanced the recruitment of leukocytic
infiltrates, inhibited tumor growth, and reduced immune-suppressive cell frequencies
(Kar et al., 2011).

25
Another example of vault protein technology was the use of nanodisk (ND)
nanoparticles to construct ND-INT complexes (NDI) to facilitate the encapsulation of
therapeutics into vault nanoparticles (Figure 5A) (Buehler et al., 2011, Han et al., 2011,
Kar et al., 2011). NDs were 10–20 nm lipid nanoparticles consisting of small discoidal
lipid bilayer fragments derived from apolipoprotein-AI. With the genetic fusion of INT to
ND, NDI acquired the ability to entrap a wider range of therapeutics into vault particles
than standard INT (Buehler et al., 2011). A gene transcription regulator, all-trans retinoic
acid (ATRA), which bound to the retinoid acid receptor and retinoid X receptor and
altered functional genes on proliferation, differentiation, and apoptosis was tested and
demonstrated to be encapsulated into NDI-Vault nanocomplex (Figure 5B). ATRA
bioactivity in NDI-Vault nanocomplexes was also confirmed by MTT (3-(4,5-
dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide) assay in a malignant hepatoma
cell line (HepG2) showing that NDI-Vault complexed with ATRA decreased 42% of
HepG2 cell viability, while free ATRA induced only 18% cell death (Figure 5C) (Buehler
et al., 2011).The results demonstrate the highly insoluble hydrophobic drug ATRA can
be efficiently packaged into NDI-Vault complex and remain biologically active.
Compared with the free drug that is cleared off rapidly, ATRA is slowly released from
the dynamic vault structure and achieves greater drug efficacy (Buehler et al., 2011).

26

Figure 5 Vault protein engineered for hydrophobic drug delivery.
(A) Schematic diagram representing NDI-ATRA formation and encapsulation by the
vault nanoparticle. Components are not drawn to scale. (B) CP-MVP + NDI-ATRA
electron microscopic tomography slice showing NDI-ATRA vault packaging. (C) HepG2
cell viability assay. NDI-ATRA and CP-MVP + NDI-ATRA both display increased toxicity
than free ATRA over the course of 120 hours.
27
Adapted with permission from Buehler DC, Toso DB, Kickhoefer VA, Zhou ZH, Rome
LH. Vaults engineered for hydrophobic drug delivery. Small. 2011;7(10):1432–1439.7
Copyright 2011 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim.

28
1.5.2 Viral protein-medicated drug delivery
Yildiz et al employed viral nanoparticles from a plant virus named cowpea mosaic virus
(CPMV) to accomplish efficient intracellular delivery (Yildiz et al., 2011). In their study,
the surface of CPMV nanoparticles was modified with polyarginine (R5) cell-penetrating
peptides using bioconjugation techniques with the help of a hydrazone linker. Cell
uptake efficiency was examined using both CPMV-R5 VLPs and plain CPMV particles.
The result demonstrated that CPMV-R5 could be more efficiently taken up by a human
cervical cancer cell line (HeLa) than plain CPMV particles. The result also indicated that
higher R5 peptide density on the surface of CPMV-R5 particles determined greater
efficiency of HeLa cell uptake. The surface cell-penetrating peptide modification also
altered intracellular trafficking of CPMV nanoparticles. It was observed that plain CPMV
particles were mainly trapped in endolysosomes after cell uptake, while 30%–50% of
CPMV-R5 nanoparticles escaped from the endosome and trafficked to other cellular
compartments (Yildiz et al., 2011, Wu et al., 2012). This finding provided CPMV-R5
nanoparticles with a promising future to encapsulate bioactive molecules and deliver
them to different compartments within the cytoplasm.

There are two recent studies focusing on small interfering RNA (siRNA) delivery using
engineered viral nanoparticles (Choi et al., 2011, Galaway and Stockley, 2013).
Galaway and Stockley assembled a new VLP with the RNA bacteriophage MS2 coat
protein and an RNA conjugate of an siRNA and a capsid assembly signal (Galaway and
Stockley, 2013). The nanoparticle efficiently entered HeLa cells via receptor-mediated
endocytosis and had significant siRNA effects at nanomolar concentrations (Galaway
29
and Stockley, 2013). The study conducted by Choi et al utilized a capsid shell, integrin-
targeting peptide, and p19 RNA-binding protein to assemble a nanocarrier for siRNA
delivery (Choi et al., 2011). The capsid nanocarriers had affinity both for siRNA on the
interior and cellular integrin on the exterior. It was discovered that RGD (Arg-Gly-Asp)
peptides on the surface enabled the capsid nanoparticles to target cancer cells that had
high v3 integrin expression and deliver siRNA to the cytosol of the targeted cells
(Choi et al., 2011). In both studies, engineered VLPs protected siRNA from the external
nucleases and facilitated the endocytosis of the entire nanocarriers with the payload of
siRNA. Besides the precise and seamless modification at the genetic level, genetically
engineered carriers have another advantage over chemically synthetic carriers, which is
the capability to produce large quantities of identical carriers by a one-step biosynthesis.
Carriers genetically engineered from both vault protein and viral proteins can be scaled
up from the milligram to gram scale relatively quickly, which is not always easy to
accomplish using multistep formulations of nanoparticles prepared from synthetic
materials.

1.6 Discussion
As an emerging class of efficient drug carriers, genetically engineered nanocarriers
have been successfully evaluated in the delivery of a large number of drugs. Potential
biocompatibility and controlled immunogenicity make them attractive technologies
compared with chemically synthesized carriers; however, significant work remains to be
done in this area (Mackay and Chilkoti, 2008). Urry et al tested a -irradiated ELP
monoblock (VPGVG)120 and discovered no acute systemic toxicity in mice
30
(intraperitoneal and intravenous), no systemic antigenicity in guinea pigs (intravenous)
(Urry et al., 1991). Moreover, it was found that subcutaneously injected ELPs could not
generate antibodies unless complete Freund’s adjuvant was added (Mackay and
Chilkoti, 2008, Domb et al., 1997). Similarly to ELPs, Cappello et al also found low
immunogenicity of a SELP with the sequence of [(GVGVP)8(GAGAGS)2]18 when
evaluated in rabbits (Cappello et al., 1998). The SELP polymer was injected at time
zero (10 mg), 6 weeks (0.5 mg), and 8 weeks (0.5 mg) without adjuvant, and then sera
samples were tested by enzyme-linked immunosorbent assay after 9 weeks. No
reactivity (titer ,2) of the samples was discovered for binding to pure (VPGVG)8
sequence. On the contrary, sera samples collected from SELP with complete Freund’s
adjuvant group showed a serum titer of 480-fold (Domb et al., 1997, Mackay and
Chilkoti, 2008). As these materials move towards translational studies, more tests are
needed to demonstrate the safety of these genetically engineered nanocarriers in vivo.
Thus far, published data reveal that no antibody response has been stimulated with the
administration of these protein polymers alone, suggesting that genetically engineered
nanocarriers may be promising platforms for the development of new drug-delivery
systems (Mackay and Chilkoti, 2008).

Recombinant protein fusion technology is one of the major advantages in genetically
engineered nanocarrier delivery and has been widely used in many applications
(Simnick et al., 2011, Shah et al., 2013, Shi et al., 2013b). Compared with chemically
synthesized carriers that require complicated and low-efficiency chemical conjugation
reactions with many byproducts, genetically engineered nanocarriers utilize fast and
31
efficient molecular cloning technique to link the drug or other functional domains onto
the carriers at the DNA level. When biosynthesized, the resulting fusion products have
nearly perfect homogeneity and monodispersity in large (mg to g) quantities (Shi et al.,
2013b, Sun et al., 2011). Drugs with few modification sites, low stability in organic
solvents, and/or poor chemical conjugation efficiency are amenable to delivery by
genetically engineered nanoparticles. Furthermore, Shi et al reported a small screening
assay for suitable drug molecules for ELP micelle encapsulation and found that drug
molecules with high hydrophobicity (log P-values) and/or large numbers of hydrogen-
bond donors and acceptors had higher ELP encapsulation efficiency than others (Shi et
al., 2013b). Therefore, drugs that are currently difficult to formulate using more
conventional delivery vehicles might be good candidates for delivery by genetically
engineered nanoparticles.
1.7 Conclusion
Recently, major innovations in the field of drug delivery have resulted from
advancements in the use of genetic engineering to biosynthesize genetically engineered
biological nanocarriers, which were the focus of this c. The advantages of genetically
engineered carriers over chemically synthesized carriers are related both to the precise
control of the chain length and monodispersity, due to the ability to seamlessly introduce
precise modifications to their structures and biosynthesis at the genetic level.
Genetically engineered polymeric drug carriers can be designed to assemble into
nanoparticles or nanofibers. These nanostructures can be modified with multiple
functional groups such as targeting moieties, imaging agents, or attachment sites for the
32
purpose of drug and gene delivery. Similar to polymeric genetically engineered drug
carriers, non-polymeric genetically engineered drug carriers such as vault proteins and
viral proteins also form nanostructures, which are being explored for the delivery of
genes and drugs. Delivery using these genetically engineered nanocarriers has yet to
be translated aggressively to use in humans. At the current time, the understanding of
these materials remains in its infancy. New ideas and perspectives are needed to
advance genetically engineered nanocarriers into the clinic; however, their numerous
preclinical applications suggest that they may provide a powerful new approach for
creating nanomedicines.

33
Chapter 2
Elastin-based protein polymer nanoparticles carrying drug at both corona and
core suppress tumor growth in vivo
2.1 Introduction
Rapamycin (Rapa) is a cyclic and hydrophobic macrolide antibiotic which was
discovered from a product of Streptomyces hygroscopicus in a sample of soil from
Easter Island (Vezina et al., 1975). Because Rapa has great potency in suppressing
immune response by inhibiting proliferation of lymphocytes, its clinical applications have
shifted from anti-fungal to anti-transplant rejection formulations such as Rapamune (Ho
et al., 1996). Recently, Rapa's anti-proliferation properties have been explored, which
have led to the clinical observation of anti-tumor efficacy in malignancy of the breast,
prostate, and colon (Yu et al., 2001, Luan et al., 2003, Majumder et al., 2004, Seeliger
et al., 2004). Rapa's anti-proliferation mechanism has also been revealed — inhibition of
mTOR (mammalian target of rapamycin) pathway. When bound to its cognate receptor
FKBP (Kd = 0.2 nM) (Bierer et al., 1990), Rapa inhibits the mTOR pathway and then
sequesters cancer cells in G1 phase (Bjornsti and Houghton, 2004). mTOR has
essential functions in cell proliferation and growth. Screening studies confirmed that
cancer cell lines having overexpression of S6K1 and expression of phosphorylated Akt
e.g. MDA-MB-468 breast cancer cell are sensitive to Rapa treatment (Noh et al., 2004).

Although Rapa is extremely potent in cancer treatment, a number of drawbacks such as
severe cytotoxicity, low bioavailability and rapid clearance limit wider usage of free
34
Rapa. Recent studies have shown that Rapa and other macrolide mTOR inhibitors have
serious lung toxicity by causing interstitial pneumonitis (Chhajed et al., 2006). Free
Rapa has poor bioavailability because of its high hydrophobicity and low water solubility
(ca. 2.6 g/mL) (Simamora et al., 2001). As a result, organic solvents such as DMSO,
polyethylene glycol (PEG) 400 and ethanol are presently used to deliver free Rapa
(Yanez et al., 2008). However, most of these organic solvents are cytotoxic to the liver
and kidney, and they may also cause hemolysis and acute hypersensitivity reactions
(Hennenfent and Govindan, 2006, Gaucher et al., 2010). It has also been determined
that free Rapa has high tendency to partition into the erythrocytes which makes it more
difficult to reach intratumoral targets (Yatscoff et al., 1995). Therefore, a well-designed
Rapa formulation is currently in demand to overcome the limitations of this potent drug.

Derived from human tropoelastin, elastin-like polypeptides (ELPs) are repetitive protein
polymers with the sequence of (Val-Pro-Gly-Xaa-Gly)n, where Xaa is the guest residue
and n is the length of the repetitive units (Mackay and Chilkoti, 2008). ELPs undergo an
inverse phase transition, which can be used to promote temperature-dependent self-
assembly (Urry, 1997). Below a tunable transition temperature (Tt), these ELPs are
highly soluble. Above Tt they coacervate into a secondary aqueous phase, akin to a
lower critical solution temperature. This phase separation can be used to purify ELPs
and their fusion proteins by a process named inverse transition cycling (ITC). Here we
explore two ELP diblock copolymers with a hydrophobic to hydrophilic length of n = 1:1
that form stable nanoparticles e.g. G(Val-Pro-Gly-Ile-Gly)48 (Val-Pro-Gly-Ser-Gly)48Y
and G(Val-Pro-Gly-Phe-Gly)24(Val-Pro-Gly-Ser-Gly)24Y, which are named I48S48 and
35
F24S24 respectively. These diblock copolymers form nanoparticles that are potentially
excellent drug carriers because: i) they are genetically engineered, which enables
precise modification and fusion to proteins; ii) they can be biosynthesized efficiently in
E. coli; iii) they form monodisperse nanoparticles that do not require electrostatic
stabilization; iv) they are biodegradable into non-cytotoxic amino acid components. For
example, the ELP nanoparticle I48S48 is effectively biodegraded in murine hepatocytes
by elastase and collagenase endopeptidases without any obvious cytotoxicity (Shah et
al., 2012a). With these advantages, we raised the hypothesis that drug (Rapa)
encapsulation and release from ELP nanoparticles can be modulated by high-specificity
avidity binding to the cognate protein receptor for the drug (FKBP) decorated at the
nanoparticle surface. In such a formulation the drug can be encapsulated into the
nanoparticle core via hydrophobic interactions, similar to other micelle delivery systems;
however, it would also have specific drug binding capacity at its corona. The FKBP-
bound Rapa was expected to remain tightly associated with decorated nanoparticles,
which may promote tumor accumulation via the enhanced permeability and retention
effect. To discover the capability of drug encapsulation using ELP nanoparticles, four
different fluorescent small molecules were first screened for efficient encapsulation into
the hydrophobic core of the I48S48 nanoparticle. Next, encapsulation and release
experiments of the model drug Rose Bengal and the clinically-approved drug Rapa
were performed using ELP micelles I48S48 and F24S24. Rapa has been previously
demonstrated to bind competitively to an FKBP domain on an FKBP-ELP fusion protein
(Dhandhukia et al., 2013). To further enhance drug-specific encapsulation and improve
drug release, the FKBP domain was genetically fused onto the corona of the
36
nanoparticles formed fromSI. This optimized construct FSI was then examined for Rapa
encapsulation and release. Finally, cell proliferation assays and in vivo tumor regression
studies were performed using FSI with encapsulated Rapa (FSI Rapa) and free Rapa in
solvent (DMSO) to evaluate their relative toxicity and anti-tumor efficacy. These studies
reveal an exciting new strategy for drug delivery and targeted encapsulation using
genetically engineered nanoparticles.

2.2 Materials and methods
2.2.1 Materials and reagents
Rose Bengal, copper chloride, phosphate buffered saline (PBS) tablets,
polyethylenimine (PEI), Congo Red, Thioflavin and Erythrosin were purchased from
Sigma-Aldrich (St. Louis, MO). Rapamycin was ordered from LC Laboratories (Woburn,
MA). TB dry growth medium was obtained from MO BIO Laboratories, Inc. (Carlsbad,
CA). pET25b(+) vector and BLR (DE3) E. coli cell were ordered from Novagen Inc.
(Madison, WI). MDA-MB-468 and MDA-MB-231 cells were purchased from the
American Type Tissue Culture Collection. MDA-MB-468 cells were cultured at 37 °C
humidified in 5% CO2 in Dulbecco's modified Eagle's medium (DMEM)/F12 medium
with 10% fetal bovine serum. MDA-MB-231 cells were cultured at 37 °C humidified in
5% CO2 in Dulbecco's modified Eagle's medium (DMEM) with 10% fetal bovine serum.

2.2.2 Biosynthesis of ELPs
To generate the ELPs evaluated during this study, synthetic genes encoding for both
ELP mono and diblock copolymers were constructed (Table 1). BLR (DE3) E. coli cells
37
were transformed with recombinant pET25b(+) vectors containing ELP genes
(Dhandhukia et al., 2013, Sun et al., 2011). The E. coli cells were incubated using 50
mL of TB dry growth medium with 100 g/mL ampicillin overnight at a 37 °C orbital
shaker. The culture was then centrifuged at 4000 rpm for 10 min and the pellet was re-
suspended with 5 mL fresh TB dry growth medium. 500 L of the re-suspended culture
media was inoculated in 1 L TB dry growth medium with 100 g/mL ampicillin and then
incubated for 24 h at a 37 °C orbital shaker. The E. coli cells were harvested by being
centrifuged at 4000 rpm for 15 min and then re-suspended in PBS and lysed by
ultrasonication. The lysate was centrifuged at 12,000 rpm for 15 min to remove
insoluble cell debris, and nucleic acids were precipitated by PEI (0.5% w/v final
concentration) and removed by 12,000 rpm 15 min centrifugation. Inverse transition
cycling (ITC), which has been described previously, was used to further purify the cell
lysate containing ELPs (Dhandhukia et al., 2013, Sun et al., 2011). In brief, ELP was
heated to 37 °C and supplemented with NaCl (~1–3 M) to trigger phase separation. ELP
aggregates were isolated by 4000 rpm 10 min centrifugation at 37 °C. The ELP pellet
was re-suspended with PBS on ice. After ELPs re-dissolved (usually taking about 5 to
20 min with gentle agitation), other insoluble proteins and contaminants were removed
by 4000 rpm 10 min centrifugation at 4 °C. Generally two to six cycles of ITC are
needed to obtain pure ELP samples (Banki et al., 2005, Meyer and Chilkoti, 1999, Wu
et al., 2006). After purification, about 30–50 mg of ELP was obtained from 1 L of BLR
(DE3) E. coli culture. The purity and molecular weight of ELP were examined by SDS-
PAGE 4–20% gradient gel. 20–40 g of ELP samples were loaded onto an SDS-PAGE
gel and then stained by 10% (w/v) CuCl2 staining solution.
38
Table 1 ELP protein polymers evaluated in Chapter 2
ELP
Nomenclature
a
Amino acid sequence
b
Tt (
o
C)
c
Calculated
ELP
MW (Da)
d
Observed
ELP
MW (Da)
I48 G(VPGIG)48Y 22.0 20,566.9 20,309.9
S192 G(VPGSG)192Y 56.5 80,000.8 79,779.5
I48S48 G(VPGIG)48(VPGSG)48Y 27.0 39,643.6 39,435.5
F24S24 G(VPGFG)24(VPGSG)24Y n.a. 20,757.3 20,493.9
SI G(VPGSG)48(VPGIG)48Y 27.0 39,643.6 39,445.0
FSI
e
FKBP-
G(VPGSG)48(VPGIG)48Y
24.5 51,445.2 51,446.8
a
ELP gene sequences confirmed by DNA sequencing from N and C terminal and
diagnostic digestions.
b
Transition temperature (Tt) (25μM, pH 7.4) for I48, S192; Critical micelle temperature
(CMT) for I48S48, SI, and FSI; The CMT of F24S24 was not applicable (n.a.) as
nanoparticles form below 4
o
C
c
Estimated from open reading frame excluding methionine start codon
d
Results from matrix assisted laser desorption ion time of flight (MALDI-ToF) mass
spectrometry
e
FKBP amino acid sequence:
“MGVQVETISPGDGRTFPKRGQTCVVHYTGMLEDGKKFDSSRDRNKPFKFMLGKQEV
IRGWEEGVAQMSVGQRAKLTISPDYAYGATGHPGIIPPHATLVFDVELLKLE”

39
2.2.3 Optical density measurement of ELP phase transition
Optical density was used to monitor the thermal phase behavior for all ELPs evaluated
as a function of concentration. The optical density was measured by absorbance at 350
nm in DU800 UV–vis spectrophotometer (Beckman Coulter, CA) under a temperature
gradient of 1 °C/min. The ELP transition temperature was defined at the maximum first
derivative of the optical density at 350 nm. For ELP diblock copolymers, two phase
transitions were observed. The first transition temperature at which small
nanostructures assemble was defined as critical micelle temperature (CMT) and the
second transition temperature at which the secondmajor phase transition occurs was
defined as bulk transition temperature.

2.2.4 Dynamic light scattering (DLS) analysis and zeta potential measurement
To estimate nanoparticle hydrodynamic radii and stability, pure 25 M ELP samples in
PBS were passed through 20 nm membrane filters at 4 °C. 80 L of sample was applied
to a pre-chilled 384 well plate and covered with 20 L of mineral oil. Samples were
measured by Wyatt Dynapro plate reader (Santa Barbara, CA) at a temperature interval
of 1 °C. The experiment was performed in triplicate and the data was presented as
mean ± SD. To evaluate the surface charge of ELP nanoparticles, zeta potentials of FSI
and SI were measured using a Malvern Zetasizer Nano ZS90 (Worcs, U.K.). Pure 100
M FSI and SI samples in low salt PBS (10 mM NaCl, 1 mM Na2HPO4) were passed
through 20 nm membrane filters at 4 °C and measured for their zeta potentials at 20 °C
and 37 °C. The experiment was performed in triplicate and the data was presented as
the mean value of the three measurements.
40
2.2.5 Transmission electron microscopy (TEM) and cryogenic-transmission
electron microscopy (Cryo-TEM)
TEM was used to observe the dominant nanoparticle morphology for ELP nanoparticles.
For TEM procedure, 5 L of sample was pipetted on a Ted Pella carbon/formvar grid
(Redding, CA). 1% uranyl acetate was added to stain the sample and the excess liquid
was removed by filter paper. The grid was dried in an incubator at 37 °C and placed into
a JEOL JEM 2100 laB6 microscope (Tokyo, Japan). All images were captured under
200 kV accelerating voltage. In order to observe particle morphology in solution, cryo-
TEM specimens were prepared using an FEI Vitrobot (Hillsboro, OR). ELP solutions
were kept in an ice bath (4 °C) before processing and then raised to 37 °C immediately
prior to blotting. A typical procedure involves first loading ~6 L of the sample on a TEM
grid coated with a lacey carbon film (LC325-Cu, Electron Microscopy Sciences). Then,
the specimen was carefully blotted under 95% humidity following blotting parameters
that were preset depending on the viscosity and concentration of the studied sample.
The blotted grid was immediately transferred into liquid ethane, and stored in liquid
nitrogen environment. Micrographs were acquired using FEI Tecnai 12 TWIN
transmission electron microscope equipped with 16 bit 2K × 2K FEI eagle bottom mount
camera (Hillsboro, OR). All images were captured under 100 kV accelerating voltage
and processed using ImageJ (NIH, USA).

2.2.6 ELP-mediated encapsulation of a model drug, Rose Bengal
Having first determined that Rose Bengal associates with the core ELP, I48 (Figure 6),
Rose Bengal was then encapsulated in ELP diblock copolymers, I48S48 and F24S24
41
(250 M) by mixing at a 1:1 ratio by mol and dialyzed at a MWCO = 10,000 g/mol
cassette (Pierce Inc.). Rose Bengal without ELP was also added as a control. Dialysis
cassettes were divided into two groups: the first group was placed in 2 L PBS (pH = 7.4)
and kept at 37 °C and the second group were kept at 4 °C. The cassettes were slowly
stirred and samples were withdrawn during release. The dialysis buffer (PBS) was
changed every 2 h to maintain the sink conditions. The concentrations of Rose Bengal
were determined by absorbance at 559 nm (MEC = 90,400 M1 cm1). The experiment
was performed in triplicate and the data was presented as mean ± SD.

42

Figure 6 The differential association of selected small molecules to the
hydrophobic core ELP of I48S48 nanoparticles. Congo Red and Rose Bengal were
demonstrated to have high association with the ELP I48, the core of I48S48, FSI, and SI
nanoparticles; whereas Thioflavin T and Erythrosin had much lower association. Mean ±
SD, N = 3.

43
2.2.7 Thin-film encapsulation of an anti-proliferative drug, Rapa
To study the encapsulation of a clinically-relevant drug, diblock copolymers I48S48 and
F24S24 or a soluble control S192 were desalted by dialysis in deionized H2O. Samples
were lyophilized, and mixtures of ELP and Rapa were obtained by co-dissolving them in
acetonitrile. Thin films were prepared by evaporating acetonitrile at 50 °C using a rotary
evaporator. Films were re-hydrated using sterile deionized H2O. Insoluble Rapa was
removed by 10 min 13,000 rpm centrifugation and the samples were then filtered (0.2
m membrane) before analytical Reverse Phase HPLC analysis. Samples were injected
into a C-18 analytical reverse phase HPLC column (Waters Inc.) and run at a H2O:
acetonitrile gradient from 30% to 100% acetonitrile. The amount of Rapa entrapped into
ELP samples was determined using a standard curve. Statistical two-way ANOVA
analysis was performed to compare the experimental groups (F24S24 and I48S48 with
Rapa) with the control (S192 with Rapa). The experiment was performed in triplicate
and the data was presented as mean ± SD.

2.2.8 Rapa encapsulation and release using FSI nanoparticles
Thin film encapsulation was not feasible due to precipitation of the FKBP domain in
response to drying; therefore, a two-phase solvent evaporation method was developed
to encapsulate Rapa into FSI nanoparticles. An aqueous phase PBS (sterile-filtered)
containing FSI was mixed with an organic phase hexane/EtOH containing Rapa in a
glass vial (Molar ratio of FSI:Rapa = 1:1). The vial was stirred and heated to about 5 °C
higher than the CMT. Anhydrous nitrogen gas was applied to facilitate the evaporation
of the hexane/EtOH phase. A 13,000 rpm 10 min centrifugation (above the CMT) was
44
performed to remove any insoluble Rapa from the remaining aqueous phase. 100 L of
the sample was filtered (0.2 m membrane) and injected into a C-18 reverse phase
HPLC column (Waters Inc.) to analyze the amount of the Rapa in each solution. Rapa
release experiments were performed by dialysis. Samples were collected in the dialysis
cassette from 0 to 48 h. Prism software was used to plot Rapa release from FSI and SI,
identify different Rapa release phases, and calculate the half-life of each phase using a
two-phase exponential decay model. The experiment was performed in triplicate and
the data was presented as mean ± SD.

2.2.9 Cell proliferation assay
Cell proliferation was assessed by using 3-(4,5-dimethylthiazol-2-yl)-5-(3-
carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium (MTS) cell proliferation assay.
Cells (~3000 cells per well) were plated in 96 well plates with 100 L culture medium.
The cells were cultured for 24 h to adhere, and Rapa encapsulated by FSI or free Rapa
(0–10 M) were added to the culture medium and incubated for 3 days. Cell proliferation
was then examined by following standard protocol of CellTiler-96 Non-Radioactive Cell
Proliferation Assay Kit (Promega, Madison, WI). IC50 of FSI Rapa and free Rapa in
MDA-MB-468 cells were evaluated using a non-linear regression model in software
Prism. The experiment was performed with N = 6 per group, and the data was
presented as mean ± SD.
45
2.2.10 In vivo evaluation of FSI/Rapa in human breast cancer xenografts
All the animal experiments have been performed according to the guidelines of the
American Association of Laboratory Animal Care under an approved protocol. 3 × 106
MDA-MB-468 cells were injected into themammary fat pads of 7-week-old female
athymic nude (nu/nu) mice (Harlan, Inc.) First treatment started on the 11th day after
tumor implantation when the average tumor size reached 40 mm3. Themice were
randomly divided into three groups. Groups 1 and 2were received PBS and FSI Rapa
(0.75 mg/kg) three times a week intravenously. Group 3 was initially received free Rapa
(DMSO, 0.75 mg/kg) three times a week; however, due to severe weight loss after the
first two treatments, free Rapa was then dosed only once a week for the rest of the
study. Tumors were measured with an electronic caliper, and the sizes were calculated
based on the formula a2 × b × /6 (a, b are the width and length of the tumor
respectively). The experiment was performed with N = 9 per group. The data of tumor
sizes was presented as median ± interquartile range, and the data of body weights was
presented as mean ± SD.

2.3 Results and discussion
2.3.1 Characterization of ELP diblock copolymer nanoparticles
This study successfully developed a new approach to encapsulate drugs in
proteinacious nanoparticles through high-avidity interactions with their cognate-human
drug target decorated at their surface. To accomplish this, ELP diblock copolymers were
selected to assemble these nanoparticles (Figure 7a). To determine the critical
parameters necessary for this approach, a small library of ELPs were successfully
46
biosynthesized (Figure 7b). ELP diblock copolymers such as I48S48 and F24S24
contain approximately 50% by mass hydrophilic blocks and 50% hydrophobic blocks.
Optical density measurements and DLS analysis were performed to examine their
thermal and concentration dependent assembly. I48S48 displays two different phase
transitions: i) the formation of nanoparticles at the first transition at 27 °C; ii) the
formation of larger aggregates at the second transition at 85 °C (Figure 7c). The lower
transition temperature is defined as the CMT. Depending on concentration, the CMT
varies from 22 °C (250 M) to 27 °C (25 M), which is well below physiological
temperatures. Alternatively, F24S24 was used to explore the properties of an ELP
nanoparticle that remains assembled at all temperatures observed. The hydrophobic
core ELP of F24S24 contains phenylalanine, which gives it a CMT below 4 °C (Figure
7d). To further confirm the size of the nanoparticles, DLS analysis was used to estimate
the hydrodynamic radius (Figure 7d). For I48S48, unimeric ELP molecules were present
at temperatures below the CMT; however, above its CMT (27 °C) 24 nm radius
nanoparticles assembled. F24S24 nanoparticles were assembled even at 4 °C;
furthermore, their hydrodynamic radii were similar to that observed for I48S48. To
decorate the surface of the nanoparticles with the FKBP domain, it was necessary to
genetically engineer the folded protein domain at the amino terminal end of the ELP
diblock copolymer (FSI). An additional control ELP diblock SI was prepared as a control
(Table 1). Surprisingly, neither orientation nor fusion to the FKBP domain had a
substantial effect on the CMT or nanoparticle radius compared with I48S48 (Figure 7e).
To study the surface charge of FSI and SI, zeta potentials of these two nanoparticles
were measured at 20 °C (not assembled) and 37 °C (assembled). The zeta potentials of
47
FSI were 0.41 mV (20 °C) and 1.97 mV (37 °C) while the zeta potentials of SI were
4.31 mV (20 °C) and 3.74 mV (37 °C). Even in buffer with low ionic strength, the zeta
potentials of both ELP nanoparticles were close to neutral in charge.

48

Figure 7. Design of ELP nanoparticles that carry anti-proliferative drugs —
decoration with protein drug receptors minimally influences assembly. a.
Schematic showing high-avidity interaction between a drug and its cognate human
target decorated at the nanoparticle surface, while the nanoparticle core may facilitate
lower-avidity drug affinity. b. SDS-PAGE of ELP library stained with copper chloride
49
(Table 1). c. Representative optical density (I48S48) used to determine ELP-mediated
assembly as a function of temperature and concentration. At 25 M, the CMT is 27 °C,
while a bulk phase transition occurs at 85 °C. d. DLS analysis of I48S48 and F24S24.
Above CMT
(27 °C for 25 M), I48S48 forms stable micelles with a hydrodynamic radius of 24 nm.
For F24S24, nanoparticles have already assembled below 4 °C. (Mean ± SD, N = 3) e.
DLS analysis of SI and FSI. Above their CMT, Both SI and FSI (25 M) form stable
nanoparticles with similar hydrodynamic radii of 24 nm (Mean ± SD, N = 3).

50
2.3.2 Nanoparticle assembly is required for entrapment of Rose Bengal into plain
ELP nanoparticles
Prior to the decoration of nanoparticles with FKBP, we explored how model small
molecules associate with the core of plain ELP nanoparticles. Rose Bengal was
selected as a model drug because it showed high association with ELP nanoparticle
core I48 (Figure 6). Therefore, it seemed plausible that Rose Bengal can be effectively
entrapped into ELP nanoparticles. To confirm this hypothesis, ELP diblock copolymers
I48S48 and F24S24 were mixed with Rose Bengal at different ratios and dialyzed at 37
°C and 4 °C. Both I48S48 and F24S24 efficiently encapsulated Rose Bengal (Figure
8a); however, I48S48 nanoparticles only retained the drug in the dialysis cassette when
incubated at 37 °C, above their CMT. The release of Rose Bengal from I48S48 at 4 °C
was similar to that observed when no ELP was added (Figure 8a). This indicates that
ELP nanoparticle assembly is essential for Rose Bengal encapsulation. For F24S24,
nanoparticles are present at both temperatures. Therefore, the time-absorbance curves
of 37 °C and 4 °C showed minimal differences between each other (Figure 8a). After 6 h
of dialysis, 48% of Rose Bengal was retained inside the I48S48 nanoparticles at
temperature above its CMT; however, when incubated below its CMT, less than 10% of
the Rose Bengal remained (Figure 8b). For F24S24, 76% of Rose Bengal remained
entrapped by the nanoparticles, and this level was unaffected by low temperature
dialysis (Figure 8b). While both nanoparticles were able to encapsulate Rose Bengal,
F24S24 nanoparticles has slower release (50% loss after 24 h) compared to I48S48
nanoparticles (50% loss after 6 h). One possibility is that the higher hydrophobicity of
F24S24 nanoparticle core might slow drug release due to an increased affinity for the
51
drug. Therefore, optimization of the core of ELP nanoparticles may be a possible
strategy to control the rate of drug release.

52

Figure 8. Assembly of ELP nanoparticles slows the release of a water soluble
model drug, Rose Bengal. a. The absorbance was measured over time at 37 °C and 4
°C. F24S24 forms nanoparticles at both temperatures; however, I48S48 only forms
nanoparticles at 37 °C. Nanoparticle formation was associated with significantly delayed
release of drug. Mean ± SD, N = 3. b. Comparison of Rose Bengal encapsulation after 6
h dialysis. I48S48 retained significantly more Rose Bengal at 37 °C than 4 °C (**p =
0.005) while no difference was observed in F24S24. Mean ± SD, N = 3.

53
2.3.3 The water insoluble drug, Rapa encapsulates into unmodified ELP
diblock copolymer nanoparticles
After identifying the high efficiency encapsulation of the water soluble small molecule,
Rose Bengal (Figure 8), we investigated the encapsulation of a clinically viable drug
with low water solubility, Rapa. Unfortunately, the encapsulation method used for Rose
Bengal could not be used for Rapa because it has poor water solubility (~2.8 M)
(Simamora et al., 2001). As an alternative, film hydration was optimized for Rapa
encapsulation into plain ELP diblock copolymers I48S48 and F24S24 with monomer
S192 as the control (Figure 9). In the case of F24S24:Rapa = 1:1, 50.6% Rapa was
encapsulated after the film hydration. Following hydration, dialysis was used to exam
the stability of the formulation. The results showed that after 6 h, 22.9% Rapa remained
associated with the nanoparticles, while the soluble ELP control S192 had undetectable
levels of drug (Figure 9a). When the ratio of F24S24 to Rapa was increased to 5:1,
58.8% initial Rapa encapsulation was obtained. A two-way ANOVA was performed to
examine the differences between F24S24 groups and S192 control at 2 h. There was a
significant difference between F24S24:Rapa of 1:1 and 5:1 and the S192 control
respectively (p < 0.0001). The failure of S192 to solubilize and retain Rapa strongly
suggests the necessity of a hydrophobic ELP core to encapsulate Rapa. The same
experiment was carried for I48S48 nanoparticles formed and dialyzed above their CMT
(Figure 9b). Similar to F24S24, I48S48 nanoparticles display reasonably good
encapsulation of Rapa. By increasing the molar ratio of I48S48 to Rapa to 10:1, the
initial Rapa encapsulation reached 92.9%. This encapsulation was stable against 6 h of
dialysis, whereby 66.1% of the initial drug was retained (Figure 9b). ANOVA
54
demonstrated that Rapa encapsulation in I48S48 was significantly higher than for S192.
(p < 0.0001). In conclusion, similar to Rose Bengal, Rapa can be efficiently entrapped
into the nanoparticles of ELP I48S48 and F24S24.

Figure 9. Thin film hydration with ELP nanoparticles promotes encapsulation of
the hydrophobic drug, Rapa. Different ELP:Rapa ratios were assayed for Rapa
encapsulation using thin film hydration. Time 0 h depicts the initial percent of
encapsulated Rapa in the supernatant. S192 was used as a control because it cannot
form nanoparticles. a. Two-way ANOVA demonstrated statistically significant
differences between F24S24 and S192 groups. Mean ± SD, N=3. b. Alternatively, the
diblock copolymer I48S48was used to entrap Rapa at 37 °C, above its CMT. At high
55
ratios of ELP: Rapa, drug encapsulation in the nanoparticle approaches 90%. Two-way
ANOVA manifested statistically significant differences between I48S48 and S192
groups. Mean ± SD, N = 3. (****p = 0.0001).

56
2.3.4 Decoration of ELP nanoparticles with FKBP protein and loading with Rapa
minimally influences nanoparticle dimensions
Severe side effects of Rapa are thought to result from rapid partitioning into
erythrocytes and endothelial cells (Yatscoff et al., 1995). Therefore, we hypothesized
that ELP nanoparticles with an enhanced avidity for Rapa might prevent the rapid
diffusion of drug into non-target intracellular environments, thereby increasing the
tolerated dose for this drug. As a unique solution to this problem, the human cognate
binding domain of Rapa–FKBP was genetically fused onto the corona of SI
nanoparticles (Figure 7a), which are similar in composition to I48S48 (Table 1). The
DNA sequence of FKBP was inserted to the N-terminus of the ELP, where superior
protein activity has been reported (Christensen et al., 2009). DLS demonstrated that FSI
fusion proteins assembled particles with a 23.8 nm hydrodynamic radius (Figure 7e)
above their CMT (24.5 °C, 25 M) (Table 1). Compared to SI, the CMT of FSI shifted
slightly from 27 to 24.5 °C due to the FKBP fusion; however, hydrodynamic radius and
stability appeared to be nearly unaffected by the fusion of FKBP. Cryo-TEM imaging
verified that the radius of FSI nanoparticles (18.5 ± 1.3 nm) was slightly larger than SI
nanoparticles (15.0 ± 2.3 nm) (Figure 10a, b). The morphology of both FSI and SI is
consistent with I48S48 (Figure 11). Similar to plain SI or I48S48 nanoparticles, FSI
nanoparticles have a hydrophobic core (isoleucine ELP) and a hydrophilic corona
(serine ELP), in addition to their FKBP domains. Thus, we hypothesized FSI
nanoparticles may have two distinct sites for Rapa encapsulation, at their core and at
their corona (Figure 7a).

57
Unlike plain ELP diblock copolymers (SI, I48S48), FSI has limited stability in organic
solvents, which is presumably due to denaturation of the folded FKBP domain. For this
reason, the film hydration method could not be used for FSI Rapa encapsulation.
Instead, a new method of two-phase solvent evaporation was developed specifically for
FSI Rapa encapsulation. To confirm that drug loading did not influence particle
morphology, FSI before and after Rapa encapsulation were imaged by TEM (Figure 12).
Similarly, their hydrodynamic radii were measured by DLS. Both TEM and DLS revealed
that the radii of FSI nanoparticles after Rapa encapsulation were only slightly larger
than those before encapsulation (Figure 10c). These observations suggest that Rapa
encapsulation only minimally influences the particle size and morphology of FSI.

58

Figure 10. Morphology of ELP nanoparticles is minimally influenced by fusion of
FKBP or encapsulation of Rapa. a, b. Cryo-TEM imaging of ELP nanoparticles with
andwithout the FKBP domain. a. The average radius of FSI is 18.5 ± 1.3 nm. b. The
average radius of SI nanoparticles is 15.0±2.3 nm. Mean±SD (n=10). Bar length=200
nm. c. A comparison of FSI nanoparticle radius determined using TEM and DLS before
and after Rapa encapsulation suggests that presentation and loading of the FKBP
domain minimally affects nanoparticle morphology. Mean ± SD (n = 10). (**p = 0.005
and *p = 0.05).

59

Figure 11. TEM and cryo-TEM images of I48S48 micelles. a. TEM image of I48S48
micelles. The average diameter of I48S48 micelles in TEM image was 37.0 ± 3.5 nm. b.
Cryo-TEM image of I48S48 micelles. The average diameter of I48S48 micelles in cryo-
TEM image was 30.9 ± 3.7 nm. Mean ± SD, N = 10. Bar length = 100nm.

60

Figure 12. TEM images of FSI before and after Rapa encapsulation. a. The
nanoparticle radius before encapsulation was 21.1 ± 1.9 nm. b. After Rapa
encapsulation the nanoparticle radius was 27.9 ± 2.2 nm. Mean ± SD, N = 10. Bar
length = 200nm.

61
2.3.5 Decoration of ELP nanoparticles with the FKBP domain prolongs the release
of Rapa
Having confirmed that Rapa encapsulation does not greatly affect the properties of the
FSI nanoparticles, FSI, SI and S192 were evaluated for encapsulation and release
(Figure 13). Using the two-phase solvent evaporation method to load these three
formulations, both FSI and SI nanoparticles had about 75% efficiency encapsulation of
Rapa; however, S192 – a linear ELP mono-block that does not form nanostructures –
had very low efficiency (less than 10%) of Rapa association (Figure 13a). Release
experiments were conducted to compare the release of drug from ELP nanoparticles
with and without the FKBP domain (Figure 13b). Again, dialysis under sink conditions
was used to track the release of Rapa from FSI and SI nanoparticles. Encouragingly,
FSI and SI nanoparticles exhibited significantly different profiles of Rapa release. For
FSI, there were two exponential decay phases of Rapa release: a fast phase with a half-
life of 1.9 h and a slow phase with a half-life of 57.8 h. In contrast only one exponential
decay phase was observed for SI–Rapa. This release had a half-life of 2.2 h, which is
similar to the fast phase of FSI Rapa release (Figure 13b) and Rapa release from
I48S48 mixed at a 1:1 ratio (Figure 9b). After 12 h of dialysis, SI–Rapa samples
retained undetectable levels of Rapa. In contrast, approximately 30% of the drug
remained associated with FSI nanoparticles. Because the association between FKBP
and Rapa is much stronger than that between ELP nanoparticle core and Rapa, we
propose that the population of Rapa that is encapsulated into ELP nanoparticle core
mainly contributes to the fast phase of FSI Rapa release while the population of Rapa
that is bound to FKBP primarily contributes to the slow phase. Therefore, it might be
62
inferred that about 70% Rapa is associated into the FSI nanoparticle core, which
displays rapid release. In contrast, the remaining 30% is bound to FKBP and releases
much slower (Figure 7a). With a longer terminal release half-life, Rapa may remain
associated with FSI nanoparticles more stably, promoting its in vivo tolerability at higher
doses and potentially better anti-tumor efficacy.

63

Figure 13. The FKBP domain prolongs the release of Rapa from FSI
nanoparticles. a. Both FSI and SI nanoparticles efficiently encapsulate Rapa using a
two-phase solvent evaporation method. In contrast, an ELP S192, which does not form
nanoparticles, was unable to encapsulate the drug. (***p = 0.001) Mean ± SD (n = 3) b.
Dialysis was used to track the loss of Rapa from the nanoparticles. FSI–Rapa exhibited
bi-exponential drug release (Half-life fast = 1.9 h, Half-life slow = 57.8 h). SI–Rapa
release followed a mono-exponential release curve (Half-life = 2.2 h) down to the limit of
detection of 2 M. After 12 h, only background was detected for SI–Rapa; however,
approximately 30% of the drug remained associated with FSI. Mean ± SD (n = 3).

64
2.3.6 FSI Rapa is as potent as free Rapa in MTS cell proliferation assay
To test and compare in vitro anti-proliferative efficacy of FSI Rapa and free Rapa, MTS
cell proliferation assay was performed using two breast cancer cell lines. MDA-MB-468
cells are Rapa sensitive, while MDA-MB-231 cells are Rapa insensitive (Figure 14). In
Rapa sensitive MDA-MB-468 cells, both FSI Rapa and free Rapa efficiently decreased
cell viability. In addition, the additional solubility of FSI Rapa enabled a higher dose at
10 M and a further reduction in cell viability to 15%. Free Rapa could not reach such a
high concentration because of its limited water solubility. In contrast, in Rapa insensitive
MDA-MB-231 cells, neither FSI Rapa nor free Rapa reduced cell viability more than
50%, even at the highest concentrations. The inhibitory concentration yielding 50%
viability, IC50, of FSI Rapa and free Rapa in MDA-MB-468 was estimated using non-
linear regression. The IC50 of FSI Rapa was 0.28 nM and IC50 of free Rapa was 0.27
nM. The critical micelle concentration (CMC) for FSI was estimated at 0.17 nM (Figure
15); furthermore, since the ratio of Rapa to FSI is approximately 1:1, then above 1 nM
Rapa the nanoparticles remain intact. Even though the nanoparticles remain
assembled, the similar viabilities for FSI and Free Rapa reflect the assay's three-day
incubation period. Figure 13b shows that after 48 h, over 80% of the FSI Rapa has been
released, which allows released Rapa to influence cell viability similar to free Rapa.
These results demonstrate that the Rapa encapsulated in FSI nanoparticles was as
potent as free drug in decreasing the viability of MDA-MB-468 breast cancer cells. In
addition, the FSI nanoparticle had the additional benefit of extending the solubility of
Rapa at least 10 fold compared to free Rapa, which could facilitate their utility in vivo.

65

Figure 14. FSI-encapsulated Rapa inhibits cell viability in an mTOR dependent cell
line MDA-MB-468 but not in an insensitive cell line, MDA-MB-231. FSI Rapa is as
potent as free Rapa in decreasing the viability of Rapa sensitive MDA-MB-468 cells;
however, FSI encapsulation increased the Rapa solubility at least 10 fold, to about 10
M. Mean ± SD (n = 6).

66
Figure 15. The relationship between the critical micelle concentration (CMC) and
the critical micelle temperature (CMT) for FSI nanoparticles. a. Optical density of
FSI at 350 nm over a temperature gradient was obtained at different concentrations.
The data shows two inflections, whereby the lower inflection is defined as the CMT. b. A
correlation plot of CMT vs. the ln(CMC) was prepared. The solid line represents the
best-fit line in the linear regression, and the dotted curves represent the 95% confidence
band of the best-fit line. The best-fit model enables the estimation of CMC at any
temperature.

67
2.3.7 FSI Rapa nanoparticles have greater anti-tumor efficacy and lower toxicity
than free drug
Having optimized the ELP nanoparticle formulation to improve Rapa solubility and
extend drug release through combined encapsulation in both the core and corona
(Figure 7a), this formulation was compared head-to-head with free Rapa in the MDA-
MB-468 breast cancer xenograft mouse model (Figure 16a). The reported maximum
tolerated dose for free drug 0.75 mg/kg body weight was chosen for initial dosing in both
FSI Rapa and free drug groups three times a week. After the first two injections, the
subjects in the free Rapa group lost almost 10% body weight (Figure 16b). Due to this
toxicity, the administration frequency of free drug was then reduced to once a week.
Even after reducing the dose frequency, free Rapa still showed strong accumulative
toxicity. 23 days after the first treatment, all mice administered free Rapa lost more than
15% of their body weight and were removed from the study. In contrast, the FSI Rapa
group showed no signs of behavioral changes or body weight loss (Figure 16b).
Compared to the PBS treatment group, FSI Rapa effectively halted tumor growth
(Figure 16a, c). In summary, compared to free Rapa group, FSI Rapa nanoparticles
showed better anti-tumor efficacy and much lower overt toxicity in treating this mTOR
dependent breast cancer xenograft mouse model.

68

Figure 16. FSI-encapsulated Rapa has better anti-tumor efficacy and lower
toxicity than free Rapa in the mTOR dependent MDA-MB-468 breast cancer
xenograft. a. Tumor volumes of FSI Rapa group are significantly smaller than PBS
group (median ± interquartile range p b 0.0001, N = 9). b. FSI Rapa group also shows
69
less body weight loss than free Rapa group. c. Kaplan–Meier survival analysis
demonstrates that FSI Rapa has better anti-tumor efficacy than free Rapa. (p = 0.004).

70
Based on these findings, we propose that Rapa delivered by FSI nanoparticles more
effectively retains in the central blood compartment. These particles have several
potentially beneficial effects: i) increased aqueous solubility for Rapa; ii) interaction with
the FKBP domain provides a slow off-rate on the order of a 60 h half-life; iii) the slow
off-rate enables this formulation to passively accumulate in the tumor; and iv) the
protein polymer is entirely composed from biodegradable polypeptides, which may
release active drug following cellular internalization in the tumor (Dreher et al., 2003);
and v) upon proteolysis, the free Rapa relocates to the cytoplasm and inhibits mTOR-
dependent cell proliferation. Despite these encouraging findings, this study has several
limitations. First, it remains unknown how effectively the Rapa is retained by the
nanoparticles in vivo. It is encouraging to note that the FSI Rapa formulation retains its
cargo under in vitro sink conditions (Figure 13b) for at least as long as the ELP
nanoparticles have been reported to circulate in mice (Janib et al., 2013). Second, it
remains unknown how much the dose can be decreased, while maintaining the same
level of efficacy. Third, it remains unknown whether the current levels of efficacy result
from the fast Rapa release, the slow Rapa release, or some optimal combination of the
two. The level of Rapa efficacy towards mTOR dependent tumors is consistent with that
observed using synthetic polymeric systems, such as poly(lactic-co-glycolic acid)
(PLGA) (Zou et al., 2011), poly(ethylene glycol)-blockpoly(2-methyl-2-benzoxycarbonyl-
propylene carbonate) (PEG-b-PBC) (Lu et al., 2011), poly(ethylene glycol)-b-
poly(caprolactone) (PEG–PCL) (Forrest et al., 2006) and acetylated dextran
microparticles (Kauffman et al., 2012). Compared with these delivery systems, ELP
nanoparticles have unique advantages. First, they can be engineered consisting of only
71
five naturally occurring amino acids, Val, Pro, Gly, Xaa and Gly derived from human
tropoelastin (Dreher et al., 2008). Unlike the byproducts of some synthetic polymers,
ELP components are non-toxic and biodegradable in vivo. Second, the size of ELP
nanoparticles is optimal compared to other delivery systems. The hydrodynamic radii of
I48S48, F24S24, SI, and FSI nanoparticles are all less 30 nm. This size range brings
them significantly above the renal filtration cutoff without producing a significant
population of particles greater than 50 nm in diameter where hepatic uptake
accelerates. This property has enabled other ELP nanoparticles to display
pharmacokinetics in a mouse on the order of a 6 h half-life (Janib et al., 2013). Third,
ELP nanoparticles can be easily modified to display functional folded proteins via
biosynthesis. This avoids the common requirement of optimizing bioconjugation
chemistry, which can reduce the activity of sensitive protein domains. Fourth, the purity
and homogeneity of ELP protein polymers are among the best achievable using
synthetic polymerization strategies (Koehl and Delarue, 1994, Rabotyagova et al.,
2011). Lastly, the FSI nanoparticle has been engineered both at its core and corona for
specific encapsulation and delivery of Rapa — one of FKBP's natural ligands. FSI may
be useful to deliver other Rapa analogs; furthermore, by changing the fusion domain
decorating these nanoparticles, this strategy may enable the specialized engineering of
sustained release nanocarriers for other classes of drugs.

Having obtained evidence that the release of drug froman ELP nanoparticle can be
modulated by incorporation of its target protein, FKBP, there remain unknowns to this
approach. First, the relationship between the pharmacokinetic clearance of FSI and the
72
release of Rapa from FSI remains unknown. It was previously found that other ELP
nanoparticles have a blood half-life in mice in the 6 h range (Janib et al., 2013);
however, the effect of FKBP modification on nanoparticle clearance is unknown. For
FSI, the nanoparticle corona remains hydrophilic and neutral in zeta potential, which
may promote the steric stabilization of these nanoparticles. During circulation, these
properties may maintain low clearance of FSI from the central blood pool. While Rapa
remains bound, its clearance may also be reduced. Interestingly, FSI Rapa has both a
fast release half-life of 1.9 h (core) and a slow release half-life of 57.8 h (corona). While
Rapa released quickly may undergo rapid re-distribution in the body, the significant
fraction of drug that is released slowly is expected to remain associated with FSI during
tumor accumulation. Based on the current study, it is unknown if both slow release and
fast release are required for a therapeutic response (Figure 16), and this needs to be
assessed in future studies. Second, the mechanism for transfer of Rapa to the tumor
remains unclear. The FSI formulation is capable of releasing Rapa both prior to and
after accumulation in a tumor; furthermore, both free and encapsulated drugs may be
acting to halt tumor proliferation. Since this formulation is effective without targeting cell-
surface receptors, tumor targeting of FSI is most likely via passive tumor accumulation.
The size of the tumors treated (around 40 mm3) in this study suggests that they are
undergoing angiogenesis, which initiates when the size of the tumor reaches 2 mm3
(Bergers and Benjamin, 2003, Danhier et al., 2010). During the formation of new blood
vessels, remodeling of immature vasculature often results in irregular shaped, dilated,
tortuous and leaky tumor blood vessels, which has been demonstrated for various drug
carriers using multiple in vivo tumor models. Our group has previously obtained
73
evidence that these ELP nanoparticles are biodegradable (Shah et al., 2012a);
therefore, it is feasible that intra or extracellular proteases act upon FSI as a mechanism
to release active Rapa. Despite these unknowns, this novel formulation poses a
newapproach to drug encapsulation that is effective in vivo and warrants additional
study.

2.4 Conclusions
ELP diblock copolymers form stable, monodisperse and biocompatible nanoparticles.
For the first time, we report successful drug encapsulation into both the core and corona
of these emerging drug carriers. Both Rose Bengal and Rapa have high association
with the cores of I48S48 and F24S24 nanoparticles. With the genetic fusion of Rapa's
cognate receptor FKBP to nanoparticle corona, the terminal release half-life of
encapsulated Rapa has been prolonged from 2.2 h to 57.8 h. With a long Rapa terminal
half-life, FSI nanoparticles may stably retain a significant fraction of drug during the
circulation in the bloodstream, which may reduce toxicity and promote delivery to
tumors. This study shows that FSI Rapa achieves greater anti-tumor efficacy and lower
cytotoxicity than free Rapa in treating a breast cancer xenograft mouse model.
Moreover, Rapa-specific encapsulation and delivery using FKBP modified nanoparticles
is proof-of-concept for a new strategy to develop biodegradable drug-specific carriers.
2.5 Acknowledgments
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
74
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, and the USC Whittier Foundation. Dr. Herbert Meiselman
of USC generously made available his instrument for making Zeta Potential
measurements.

75
Chapter 3
Triggered Sorting and Co-Assembly of Genetically Engineered Protein
Microdomains in the Cytoplasm
3.1 Introduction
The various highly compartmentalized, membrane-bound organelles are essential for
cellular metabolism in eukaryotic cells (van Vliet et al., 2003, Rothman, 1994). The
specialized composition of proteins and lipids in different organelles enables them to
accomplish specialized processes, and sophisticated sorting mechanisms direct
molecules to specific intracellular locations to maintain cellular functions (Rothman,
1994, Mellman and Nelson, 2008, Pearse and Robinson, 1990). Although it is
ubiquitous to life that cells routinely generate and sort nanostructures, including protein
complexes, organelles and chromosomes, our ability to similarly engineer and sort
synthetic organelles in vivo remains primitive (Rothman, 1994, Lee et al., 2004, Palade,
1975). To address this challenge, this chapter describes an innovative approach for co-
assembly and self-sorting of materials inside living cells using genetically-encoded
protein polymers – elastin-like polypeptides (ELPs) (Sun et al., 2011). ELPs are
repetitive polypeptides with the sequence of (Val-Pro-Gly-Xaa-Gly)n derived from
human tropoelastin, where Xaa and n represent the ‘guest residue’ identity and number
of repeat units, respectively (Mackay and Chilkoti, 2008, Shah et al., 2012a). ELPs
mediate self-assembly by temperature-triggered phase separation above their transition
temperature (Tt) (Urry, 1997). Depending on the composition and arrangement of guest
residues, monoblock ELPs form micro-structures distinct from diblock ELPs after
76
assembly (Dreher et al., 2008, Janib et al., 2013, Wright and Conticello, 2002). While
monoblock ELPs form large protein coacervates, amphiphilic ELP diblock copolymers
sometimes assemble nanoparticles of less than 100 nm in diameter (Janib et al.,
2014b, Shi et al., 2013a). Prior studies revealed that ELP fusion proteins assemble
genetically engineered protein microdomains (GEPMs) in living eukaryotic or E. coli
cells (Ge et al., 2009, Pastuszka et al., 2012b); however, this chapter is the first report
exploring how ELPs can sort or assemble two distinct proteins. The hypothesis is that
different monoblock ELPs with similar transition temperatures may spatially coassemble
into mixed GEPMs, and that these will spatially sort from ELP diblock copolymers
(Figure 17). To test this hypothesis, three different monoblock ELPs and one diblock
ELP were biosynthesized and purified, and their micro-structures were identified and
characterized. The capability of these ELPs to spatially co-assemble and self-sort was
evaluated using confocal laser scanning microscopy both in vitro and in the eukaryotic
cytosol. These findings reveal a potentially powerful strategy for intracellular co-
assembly and sorting of GEPMs, and may have utility in organizing synthetic organelles
enriched in distinct functional proteins.

77

Figure 17. Co-assembly versus self-sorting of Genetically Engineered Protein
Microdomains (GEPMs) in eukaryotes. a) Two different soluble ELP monoblocks are
uniformly distributed in the eukaryotic cytosol when below their Tt. When induced to
phase separate, structurally similar ELP monoblocks mix and co-assemble into micron
sized mixed GEPMs. b) Structurally distinct ELP monoblock and diblock copolymers are
homogeneously mixed below Tt. After transition, the nanoparticles assembled by ELP
diblocks self-sort into GEPMs that are spatially separate from the monoblock ELP
GEPMs in the eukaryotic cell. This process is tunable, switchable, and reversible;
furthermore, because it is based on genetically- encoded protein polymers, it may be
useful to drive assembly and sorting of functional proteins in living cells.
78
3.2 Materials and Methods
3.2.1. Biosynthesis and Characterization of ELPs
The recombinant pET25b(+) vectors with ELP gene insertions were used for ELP
expression in BLR (DE3) E. coli . Inverse transition cycling was used to purify ELP
samples from the cell lysate. The detailed ELP expression and purification procedures
were described in the previous publications of our group (Dhandhukia et al., 2013, Shi
et al., 2013a). The optical density of ELPs (OD 350 nm, temperature gradient of 1 ° C
min
1
) was measured using DU800 UV-Vis spectrophotometer (Beckman Coulter, CA).
The temperature at the maximum first derivative of the optical density at 350 nm was
defined as ELP transition temperature (Tt). Dynamic light scattering (DLS) was used to
estimate the hydrodynamic radii of ELP nanoparticles. Pure ELP samples (25 M, in
PBS) were filtered (20 nm membrane filter, 4°C) and loaded onto a pre-chilled 384 well
plate. A Wyatt Dynapro plate reader (Santa Barbara, CA) was used to measure ELP
hydrodynamic radii (1°C temperature interval). Cryogenic-transmission electron
microscopy (Cryo-TEM) and negative staining TEM were used to observe nanoparticle
morphology in solution and in the dried-down state respectively. The detailed
preparation procedure was described in our previous publication (Shi et al., 2013a).

3.2.2. In Vitro Co-Assembly and Sorting
ELP samples were labeled with rhodamine (Rho) or carboxyfl uorescein (CF) using N-
Hydroxysuccinimide (NHS) chemistry. The labeled ELPs were diluted to the required
concentrations in each designated group to obtain the same Tt. The two ELPs in each
group (Rho and CF labeled) were mixed at 1: 1 (v/v) ratio in a 35 mm glass bottom dish
79
(MatTek, MA) and imaged using a Zeiss LSM 510 Meta NLO confocal microscopy
(Thornwood, NY) with an Instec HCS60 temperature control stage (Denver, CO). All
images were captured under a Plan-Apochromat 63x oil immersion lens with a working
distance of 0.19 mm.

3.2.3 Lentivirus (rLV) Production and Generation of Stable Cell Line Hek-DsRed-
V96 Expressing DsRed-V96
DsRed-V96 gene was inserted into pLVX-N1 lentiviral expression vector (Clontech
Laboratories, CA). The construct was then mixed with Lenti-X HT proprietary packaging
mix and used with Lentiphos HT transfection system (Clontech Laboratories, CA) to
transfect 293T cells for production of VSV-G pseudotyped, replication-incompetent rLV.
The supernatants containing the rLV were collected and concentrated by centrifugation
(1500 g, 45 min) and LentiX concentrator (Clontech Laboratories, CA) after 48 h post-
transfection. Viral titers were determined using p24 antigen ELISA (Invitrogen, CA). To
generate a stable cell line, Hek cells were infected with 104 infectious units/ml rLV
encoding DsRed-V96. To enrich the tissue culture, the cells were split at 1:10 and 1:100
ratios and continued to grow in the presence of selective agent puromycin. A few
resultant single colonies with bright DsRed fl uorescence levels were selected using a
pipet tip and allowed for cell culture growth (named Hek-DsRed-V96 cells).

3.2.4. Intracellular Co-Assembly and Sorting
ELP genes were inserted to the downstream of NT-GFP-pcDNA3.1 plasmid (Invitrogen,
CA) to produce mammalian GFP-ELP fusions. The GFP-ELP constructs were used to
80
transfect Hek-DsRed-V96 cells in plain Dulbecco's modified Eagle’s medium (DMEM)
using Turbofect (Fermentas, MA). The cells were incubated with Turbofect-DNA mixture
for 6 h, washed with PBS, and cultured in fresh DMEM for 48 h before confocal imaging
(Pastuszka et al., 2012b). Confocal images were captured using a Zeiss LSM 510 Meta
NLO confocal microscopy (Thornwood, NY) with an Instec HCS60 temperature control
stage (Denver, CO) under a Plan-Apochromat 63x oil immersion lens with a working
distance of 0.19 mm.

3.2.5. Statistical Analysis
Confocal images were converted from RGB-color to 8-bit and analyzed by JACoP in
ImageJ (NIH). Scatter plots of the red and green pixels, Pearson’s coefficients and
overlap coefficients were obtained from JACoP analysis to evaluate colocalization of
two ELP microdomains (Bolte and Cordelieres, 2006). The slope and r2 value in the
linear regression were used to predict the correlation between red and green pixels in
the scatter plots. Software LSM 510 (Thornwood, NY) was also used to acquire
Pearson’s coefficients from confocal images of each group with a manually determined
threshold of noise.

3.3. Results and discussion

3.3.1. Assembly of various genetically engineered protein microdomains from
ELPs

Four ELPs have been biosynthesized and purified (Table 2) in order to mimic the
natural protein sorting process in eukaryotes. Among these four ELPs, I24, V96 and
81
V192 are hydrophobic monoblock ELPs with isoleucine or valine as guest residues,
while S48I48 is an amphiphilic diblock ELP copolymer with serine and isoleucine as the
guest residues of hydrophilic and hydrophobic blocks respectively. The micro-structures
formed after their temperature-sensitive transitions were assessed by optical density
measurement, dynamic light scattering (DLS), cryogenic-transmission electron
microscopy (Cryo-TEM) and negative staining TEM (Figure 18). Optical density
measurement using a range of ELP concentrations from 25 M to 250 M confirmed a
negative correlation between the concentration and transition temperature (Fig 18a).
Using this fit, a common transition temperature of about 28 °C was identified for the four
ELPs by adjusting the concentration. Another optical density measurement of these four
ELPs transitioned at the same temperature (28 °C) with different concentrations
suggested the formation of micro-structures with widely differing optical density (Fig 1
8b). Above the assembly temperature, the high A 350 optical densities of I24, V96 and
V192 monoblock ELPs were consistent with the formation of large microparticles, while
the relatively low optical density of S48I48 was consistent with the formation of
nanoparticles. To further explore these micro-structures, DLS, cryo-TEM, and negative
staining TEM were performed to measure the hydrodynamic radii and observe
morphology of the nanostructures respectively (Fig 18c-e). DLS analysis determined
that amphiphilic S48I48 diblock ELP assembled micelle nanoparticles with a
hydrodynamic radius of 24 nm, while the other ELP monoblock coacervates formed
large microparticles (hydrodynamic radii from 0.4 to 1.3 m) above their Tt (Fig 18c).
Consistent with DLS data, negative staining TEM imaging also confirmed the assembly
of large GEPMs from monoblock ELPs I24, V96 and V192 (Fig 19). Moreover, the
82
morphology of S48I48 micelle nanoparticles was confirmed by cryo-TEM (Fig 18d) and
negative staining TEM (Fig 18e) imaging with an average measured radius of 14.3 ± 1.4
nm and 17.1 ± 1.4 nm, respectively.

83
Table 2. ELP protein polymers of different molecular weight, hydrophobicity, and
polymer architecture with similar assembly temperatures

ELP
Nomenclature
a
Amino acid
sequence
b
Tt
[
o
C]
c
Molecular
Weight
[Da]
d
Morphology
above Tt
e
Particle radius
[nm]
S48I48
G(VPGSG) 48
(VPGIG) 48Y
28.0 39,644 nanoparticle 23.7 ± 0.3
I24 G(VPGIG) 24Y 32.0 10,596 microparticle 437.2 ± 15.5
V96 G(VPGVG) 96Y 31.9 40,992 microparticle 918.8 ± 76.2
V192 G(VPGVG) 192Y 28.0 81,984 microparticle 1291.4 ± 12.7

[a] ELP gene sequences were confirmed by DNA sequencing and diagnostic digestions.
[b] Transition temperature by optical density (25μM, pH 7.4) is indicated for I24, V96
and V192, while the critical micelle temperature (CMT) is indicated for S48I48.
[c] Molecular weight was estimated from open reading frame excluding methionine start
codon.
[d] Morphology of ELPs after transition was confirmed by dynamic light scattering (DLS)
and Cryo-TEM imaging.
[e] Particle radii of ELP samples (25μM, pH 7.4) were measured by DLS and presented
as Mean ± SD (n = 3).

84

Figure 18. Tunable assembly of micro-structures from hydrophobic ELP
monoblocks and amphiphilic ELP diblock copolymers. Four ELPs are considered in
this study with varying molecular weight, ELP guest residue, and polymer architecture
(Table 2). a) The concentration-temperature phase diagram for these distinct polymers
enables the selection of concentrations that assemble at a common temperature. All
four ELPs transition at 28 ° C under different concentrations: S48I48 25 M, V96 250
M, V192 25 M, I24 100 M. b) Above this temperature, optical density measurements
show that ELP monoblocks I24, V96 and V192 form highly turbid particle suspensions,
while the ELP diblock S48I48 assembles particles with a minimal increase in optical
density. Under optimal conditions, all four constructs assemble at nearly identical
85
temperatures. c) DLS analysis demonstrates that S48I48 assembles into nanoparticles
with a hydrodynamic radius of 24 nm, while ELPs monoblocks form larger microparticles
ranging from 0.4 to 1.3 m in radius. d) Cryo-TEM imaging confi rms the formation of
S48I48 into small nanoparticles with a radius of 14.3 ± 1.4 nm (Mean ± SD, n = 10). Bar
length = 100 nm. e) Negative staining TEM image of S48I48 nanoparticles. The
measured radius is 17.1 ± 1.4 nm (Mean ± SD, n = 10). Bar length = 100 nm.

86

Figure 19 Micron-scale GEPMs assembled by monoblock ELPs can be observed
using negative staining TEM imaging. Using 2% uranyl acetate negative staining,
large GEPMs formed by ELP monoblocks were imaged by regular TEM imaging. a, b
and c are I24, V96 and V192 (100μM, in DI H2O) samples respectively. Bar length =
500nm.

87
3.3.2. In vitro co-assembly and sorting of genetically engineered protein
microdomains
Since distinct micro-structures above Tt have been identified, the capability of
monoblock and diblock ELPs to spatially co-assemble and self-sort was first examined
in vitro (Figure 20). The assay was developed in a glass bottom dish with a temperature
control stage. Using confocal microscopy, ELPs labeled with different fluorophores
could be distinguished and analyzed by quantitation of colocalization. Four groups of
ELPs were assigned based on their differences in the repetitive unit lengths (n), guest
residues (Xaa) and block composition (monoblock vs. diblock). The two ELPs in each
group were site-specifically labeled at the amino terminus with rhodamine (Rho) or
carboxyfluorescein (CF), mixed in the dish and imaged above their Tt (Fig 20a). Below
Tt, each group was uniformly mixed and highly colocalized (Figure 21). Above Tt, large
microdomains colocalized in the mixtures of V96-V96, V96-V192 and V96-I24. As
expected, mixtures of ELPs with the same guest residue (Xaa = Val) and different
molecular weights (n = 96 and 192) yielded nearly optimal spatial co-assembly above
their common Tt. Unexpectedly, in the V96-I24 mixture a fraction of I24-CF coacervate
forms ‘red free’ domains that are inaccessible to the surrounding V96-Rho coacervate
(Fig 20a). This may be due to I24’s higher hydrophobicity (Xaa = Ile vs. Val) and lower
molecular weight (1/4 of V96) compared to V96, which may allow it to form a crystalline
phase within the amorphous V96 melt. Thus, even though I24 and V96 spatially co-
assemble, they form heterogeneous microdomains containing subfeatures. In contrast,
the assembled microdomains of V96-Rho are spatially separate from those of S48I48-
CF in the V96-S48I48 mixture. Interestingly, in the V96-S48I48 mixture the size of V96
88
coacervates was significantly smaller than in the other mixtures. Despite their apparent
spatial sorting, S48I48 nanoparticles appear to mediate growth of smaller V96
microdomains, which is a phenomenon that will be explored in future studies.
Colocalization analysis was performed using software JACoP in ImageJ and LSM510
for all confocal images (Fig 20b-f). Scatter plots displayed dissimilar correlations
between red and green pixels in each group (Fig 20b, e). Both positive slopes of the
linear regression lines and tight distribution of data points along the lines indicated a
positive correlation between the two pixels in the mixtures of V96-V96 (Slope = 0.54, r
2

= 0.9939), V96-V192 (Slope = 0.54, r
2
= 0.9106) and V96-I24 (Slope = 1.23, r
2
=
0.9514) (Fig 20b-d). However, a negative slope and near-axial distribution of data
points was observed in the V96-S48I48 mixture, which may be interpreted as a
measure of spatial sorting between these two ELPs (Slope = 0.19, r
2
= 0.0361) (Fig
20e). To quantitatively evaluate the degree of colocalization between two ELPs in each
group, Pearson's coefficients (PC) were calculated by both JACoP and LSM510, and
overlap coefficients (OC) were obtained from JACoP (Fig 20f). A difference in PC
between JACoP and LSM510 is noticeable because LSM510 subtracts manually
determined background intensity from the image while JACoP does not. The PC for the
groups of V96-V96, V96-V192 and V96-I24 were close to 1, which revealed strong
colocalization. In contrast, the PC of V96-S48I48 group from both JACoP and LSM510
were close to 0 or even negative, which suggest that these two ELPs self-sort. The OC
quantifies the overlap of the two pixels (range from 0 to 1) also suggests strong co-
assembly of the three monoblock ELP groups and self-sorting between V96 and
S48I48. To further study the relation between self-sorting of monoblock and diblock
89
GEPMs and ELP concentration, different mixing ratios of V96 and S48I48 were tested
for in vitro self-sorting. It was discovered that the spatial self-sorting of monoblock and
diblock GEPMs was independent of ELP concentration (Figure 22).

90

Figure 20. Only monoblock and diblock copolymers spatially sort into distinct
GEPMs. a) Confocal microscopy imaging was used to characterize ELP micro-
structures above the matched transition temperature. On a glass bottom dish, two
purified ELPs labeled respectively with rhodamine (Rho) and carboxyfluorescein (CF)
were mixed. Below their transition temperature both colors remain mixed and diffuse
(Figure 21). The ‘red free’ domains in the V96-I24 mixture are indicated with red arrows.
Scale bar = 10 m. b-e) Scatter plots of red and green pixels were generated from
converted (8-bit) confocal images of ELP mixtures. Linear regression lines are shown in
red for each mixture. Noise in each image was determined as Red < 5 and Green < 5
and excluded from the scatter plots. f) For each mixture, Pearson’s coefficients (PC1,
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PC2) and overlap coefficients (OC) were estimated. PC1 and OC were generated using
JACoP software. PC2 was generated from LSM 510 software.

92
Figure 21. Confocal imaging of mixtures of Rho and CF labeled ELPs below the
matched transition temperature. Mixed in a glass bottom dish, the Rho and CF
labeled ELP samples emitted diffuse red and green fluorescence respectively lacking
any formation of GEPM structures. The yellow overlay fluorescence confirms the two
samples were uniformly mixed prior to heating (Figure 20). Scale bar = 50 μm.

93
Figure 22. Self-sorting between monoblock and diblock GEPMs is concentration
independent. Confocal imaging was used to observe mixtures of ELPs in vitro. While
holding the concentration of the diblock S48I48-CF constant (25 μM), three
concentrations of V96-Rho (100, 250, and 350 μM) were evaluated. In all three mixing
ratios, monoblock and diblock ELPs assembled into distinct GEPMs, which suggests
that the spatial sorting properties of these ELPs are relatively concentration
independent. Scale bar = 10 μm.

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3.3.3. Intracellular co-assembly and sorting of genetically engineered protein
microdomains
Having obtained evidence of in vitro spatial co-assembly and self-sorting, we
hypothesized that the genes encoding for ELP microdomains would have similar
behavior when observed in a complex biological environment such as the cytosol of a
living eukaryote. A human embryonic kidney (Hek) cell line (Hek-DsRed-V96) was
established to stably express DsRed-V96 – a fusion of a red fl uorescent protein and
monoblock ELP. Mammalian GFP-ELP fusions – GFP-V96 and GFP-S48I48 were
cloned and transiently transfected into Hek-DsRed-V96 cells respectively. Confocal
microscopy was utilized to image the cells containing both DsRed and GFP
fluorescence (Figure 23a). With a temperature control stage, it was determined that all
fusion ELPs had a similar intracellular transition temperature (DsRed-V96 26.3 °C,
GFP-V96 26.9 °C, and GFP-S48I48 25.6 °C). Therefore, 37 °C and 10 °C were selected
to represent intracellular co-assembly and self-sorting after and before microdomain
formation. In both groups of DsRed-V96/GFP-V96 and DsRed-V96/GFP-S48I48 at 10 °
C, GFP and DsRed fluorescence was uniformly distributed throughout the cell, which
was reflected by strong colocalization between the two soluble monomers. ELP
microdomains (red and green puncta) assembled in both groups when the temperature
was raised to 37 °C. The microdomains of DsRed-V96 and GFP-V96 extensively
colocalized with each other while spatially separated microdomains were observed in
cells with both DsRed-V96 and GFP-S48I48. Similar to in vitro data analysis, the
intracellular confocal images were quantitatively analyzed for pixel colocalization using
JACoP and LSM510 software. Consistent with the confocal images, the statistical
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analysis demonstrated intracellular co-assembly of DsRed-V96 and GFP-V96 and self-
sorting between DsRed-V96 and GFP-S48I48 (Figure 23b-f). While remaining soluble at
10 °C, both DsRed-V96/GFP-V96 and DsRed-V96/GFP-S48I48 groups showed positive
slopes and high (close to 1) r
2
values of the linear regression lines in scatter plots
(DsRed-V96/GFP-V96 slope = 1.24, r
2
= 0.8958; DsRed-V96/GFP-S48I48 slope = 1.12,
r2 = 0.6315) (Figure 23 b and d) After phase transition at 37 °C, the value of r
2
slightly
increased in the group of DsRed-V96/GFP-V96 (r
2
= 0.9642) while a significant
decrease was observed in both values of the slope and r
2
in DsRed-V96/GFP-S48I48
group (slope = 0.47, r
2
= 0.2362) (Figure 23 c and e). These data indicated that after
phase transition, the DsRed-V96/GFP-V96 group assembled microdomains with a
higher degree of colocalization. In contrast, self-sorting DsRed-V96 and GFP-S48I48
microdomains had a much lower the degree of pixel colocalization. Moreover, the high
(close to 1) values of PC and OC confirmed co-assembly of the microdomains of
DsRed-V96 and GFP-V96. In contrast, the coefficient values dramatically decreased
with the formation of DsRed-V96 and GFP-S48I48 microdomains indicating that these
two ELP fusion proteins can spatially self-sort in living cells (Figure 23f). In the
intracellular assay, PC values from JACoP and LSM510 differ because of the relatively
high background intensity resulting from the overexpression of DsRed-V96. Without
removing DsRed background, PC from JACoP included colocalization of low intensity
background pixels in DsRed-V96 and GFP-S48I48 37 °C group which resulted in a
slightly positive value; however, a negative value obtained from LSM510 after removing
those unwanted pixels provided stronger quantifi cation of intracellular sorting of DsRed-
V96 and GFP-S48I48 microdomains.
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Figure 23. Fluorescent ELP fusion proteins in the eukaryotic cytosol can co-
assemble or self-sort GEPMs. a) Confocal microscopy imaging of GFP-ELP and
DsRed-ELP fusions in Hek-DsRed-V96 cells. GFP-V96 and GFP-S48I48 were
respectively transfected and expressed in stable Hek-DsRed-V96 cells. The images
were captured before (10 ° C) or after (37 ° C) GEPM assembly. Scale bar = 5 m or
10 m. b-e) Scatter plots of green and red pixels from converted (8-bit) confocal
images. Linear regression lines are shown in red for 10 ° C and 37 ° C groups. f) For co-
transfected cells, Pearson’s coefficients (PC1, PC2) and overlap coeffi cients (OC) were
estimated. PC1 and OC were generated using JACoP software. PC2 was generated
from LSM 510 software. A, B, C and D represent the correlation of DsRed-V96/GFP-
V96 at 10 ° C, DsRed-V96/GFP-V96 at 37 ° C, DsRed-V96/GFP-S48I48 at 10 ° C and
DsRed-V96/GFP-S48I48 at 37 ° C, respectively.
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3.4. Conclusion
In summary, this chapter reports the biosynthesis of three different ELP monoblocks
and one diblock and their distinct microdomains assembled after ELP-mediated phase
separation. It was confirmed that ELP monoblocks with similar Tt assemble relatively
large protein coacervates which could coassemble in vitro; however, these coacervates
self-sorted from the nanoparticles assembled by ELP diblock copolymers. It is also
possible that some mixtures of ELP monoblocks will also sort; however, our data
suggests that when the transition temperatures are matched, ELP molecular weight and
sequence yield co-assembly. Most importantly, an intracellular assay demonstrated co-
assembly of overexpressed DsRed-V96 and transfected GFP-V96 and self-sorting
between DsRed-V96 coacervates and GFP-S48I48 nanoparticles in eukaryotic cells.
These encouraging in vitro and intracellular findings demonstrate that ELP gene
products can be induced to either spatially sort or co-assemble functional proteins
(GFP, DsRed) within the cytosol. For the first time, this simple strategy enables
advanced control over the organization of micro-structures in the cytosol, which may
promote the development of synthetic organelles.

3.5 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,
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the USC Ming Hsieh Institute, and the USC Whittier Foundation. We thank Drs. Liana
Asatryan and Daryl Davies at the Lentiviral core of the USC School of Pharmacy for the
production of recombinant lentivirus.

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Chapter 4
A co-assembled multi-functional protein-polymer nanocarrier delivers specific
chemotherapeutics, actively targets tumor and suppresses its growth
4.1. Introduction
As the second most common type of cancer behind lung cancer with over 1.3 million
newly diagnosed cases worldwide each year, breast cancer is one of the hottest areas
of oncology research these days (Grayson, 2012). Over four decades, many
chemotherapeutic drugs have been developed to significantly improve the survival rates
of breast cancer from 35% to 77%; however, there are still nearly half-a-million deaths
from this disease each year (Grayson, 2012, Perry and Wiseman, 1999, Jordan, 1993,
Demers, 1994). Typical breast cancer drugs like Trastuzumab and Tamoxifen have
obvious side effects such as cardiovascular cytotoxicity and endometrial changes that
limit their utility (Seidman et al., 2002, Grilli, 2006). Furthermore, the drug resistance on
certain types of patients also hinders broader application of these drugs (Griner et al.,
2013, Johansson et al., 2013). Therefore, new strategies for the treatment of breast
cancer are being required at present.

Our group previously developed a novel protein polymer-based, drug-specific
nanocarrier utilizing elastin-like polypeptides (ELPs) (Shi et al., 2013b). ELPs are
human-elastin-derived repetitive protein polymers with amino acid sequence of (Val-
Pro-Gly-Xaa-Gly)n, where Xaa represents the guest residue and n represents the length
of the repetitive units (Shi et al., 2013b, Mackay and Chilkoti, 2008). ELPs can promote
100
temperature-dependent self-assembly by undergoing an inverse phase transition above
the transition temperature (Tt) (Urry, 1997, Shi et al., 2013b). It has been demonstrated
that ELP di-block copolymers with the sequences of G(Val-Pro-Gly-Ile-Gly)48(Val-Pro-
Gly-Ser-Gly)48Y (named I48S48) and G(Val-Pro-Gly-Ser-Gly)48(Val-Pro-Gly-Ile-Gly)48Y
(named S48I48) can assemble stable micelle nanoparticles after their phase transition
(Table 3) (Shi et al., 2013b, Janib et al., 2014a, Shi et al., 2014). To specifically deliver
an anti-proliferative therapeutic drug Rapamycin (Rapa), the cognate protein target of
Rapa - FK506 binding protein 12 (FKBP) was genetically conjugated onto the surface of
S48I48 nanoparticles (FSI) to enhance the avidity towards Rapa (Shi et al., 2013b,
Bierer et al., 1990, Bjornsti and Houghton, 2004). FSI was confirmed to form stable
micelles that carry the drug at both surface and core (Table 3) (Shi et al., 2013b). The
novel nanocarrier slowed the terminal half-life of Rapa release to 57.8 h, and
significantly improved anti-tumor efficacy and reduced cytotoxicity compared to free
drug in a human breast cancer (MDA-MB-468) mouse model (Shi et al., 2013b).

In general, there are two routes for nanocarriers to reach tumor site – passive targeting
by enhanced permeability and retention (EPR) effect and active targeting by tumor
specific ligand – receptor binding (Danhier et al., 2010, Matsumura and Maeda, 1986).
Different from EPR effect which utilizes the size of nanocarriers and the leaky tumor
endothelia to allow passive accumulation in tumor microenvironment, nanocarriers with
active tumor targeting ligands effectively locate tumor site via high-avidity receptor
binding resulting higher tumor targeting efficiency (Haley and Frenkel, 2008, Allen,
2002). RGD (Arg–Gly-Asp) is a well-known peptide containing three essential amino
101
acids derived from extracellular matrix proteins and has affinity to bind v integrins that
are highly expressed in many types of tumor cells such as breast, colon and lung
cancer (Ruoslahti and Pierschbacher, 1986, Humphries et al., 2006, Goodman et al.,
2012). In this study, RGD was genetically fused onto an ELP di-block copolymer, and
the fusion protein was mixed and co-assembled with structurally similar FSI. We
expected the co-assembly of mixed ELP micelle nanoparticles possessing both specific
drug delivery and active tumor targeting. Both functionalities were examined by in vitro
cell-based integrin-binding assays and Rapa loading and release experiments.
Furthermore, positron emission tomography (PET) imaging was performed to examine
the effective tumor targeting with the introduction of RGD. Finally, in vivo tumor
regression studies were conducted in the comparison of Rapa formulations of passive
and active tumor targeting to evaluate their relative cytotoxicity and anti-tumor efficacy.

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Table 3 ELPs evaluated in Chapter 4
ELP
Nomenclature
a
Amino acid sequence
b
Tt (
o
C)
c
Calculated
ELP
MW (Da)
d
Particle
radius (nm)
I48S48 G(VPGIG)48(VPGSG)48Y 27.0 39,643.6 23.9 ± 0.4
S48I48 G(VPGSG)48(VPGIG)48Y 27.0 39,643.6 23.3 ± 0.4
ISR G(VPGIG)48(VPGSG)48Y-
GRGDGG
26.5 40,161.1 24.2 ± 0.3
FSI
e
FKBP-G(VPGSG)48(VPGIG)48Y 24.5 51,445.2 23.7 ± 0.1
a
ELP gene sequences confirmed by DNA sequencing and diagnostic digestions.

b
Transition temperature (Tt) is ELP critical micelle temperature (CMT) measured at
25μM, pH 7.4.
c
Estimated from open reading frame excluding methionine start codon.
d
Particle radii were measured by DLS (25 M, pH 7.4) and presented as Mean ± SD (n
= 10).
e
FKBP amino acid sequence:
“MGVQVETISPGDGRTFPKRGQTCVVHYTGMLEDGKKFDSSRDRNKPFKFMLGKQEV
IRGWEEGVAQMSVGQRAKLTISPDYAYGATGHPGIIPPHATLVFDVELLKLE”.

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4.2. Materials & Methods
4.2.1. Biosynthesis and characterization of ELPs
Synthetic genes encoding ELP di-block copolymers and fusion proteins were
constructed to generate the ELPs evaluated in this study (Table 3). ELPs were
expressed using recombinant pET25b(+) vectors with ELP gene insertions in BLR
(DE3) E. coli. Inverse transition cycling (ITC) was applied for ELP sample purification
from the cell lysate. The detailed procedures of ELP expression and purification were
described in the previous publications of our group (Shi et al., 2013b, Shi et al., 2014).
To determine ELP transition temperature (Tt), DU800 UV-Vis spectrophotometer
(Beckman Coulter, CA) was used to measure the optical density of ELPs (OD 350 nm,
temperature gradient of 1 °C min
-1
). The temperature at the maximum first derivative of
OD 350nm was defined as ELP Tt. Dynamic light scattering (DLS) was utilized to
estimate the hydrodynamic radii of ELP nanoparticles. ELP samples (25 M, pH 7.4)
were filtered (20 nm filter, 4 °C) and then measured using a Wyatt Dynapro plate reader
(Santa Barbara, CA) in a pre-chilled 384 well plate. The sample was measured at 1 °C
temperature interval. Cryogenic-transmission electron microscopy (Cryo-TEM) was
performed to observe the morphology of ELP nanoparticles in solution. The detailed
procedure of sample preparation was described in our previous publication(Shi et al.,
2013b).
4.2.2. Co-assembly of ELP micelle nanoparticles
FSI and ISR samples were respectively labeled with carboxyfluorescein (CF) or
rhodamine (Rho) using N-hydroxysuccinimide (NHS) chemistry. The two ELPs were
104
mixed at 1: 1 molar ratio in a 35 mm glass bottom dish (MatTek, MA) and imaged using
a Zeiss LSM 510 Meta NLO confocal microscopy (Thornwood, NY) with an Instec
HCS60 temperature control stage (Denver, CO). JACoP in ImageJ (NIH) was used to
analyze confocal images where scatter plots of the red and green pixels, Pearson’s
coefficients and overlap coefficients were obtained to evaluate colocalization of ELP
micelle nanoparticles (Bolte and Cordelieres, 2006). The slope and r
2
of the linear
regression analysis were utilized to determine the correlation of red and green pixels in
the scatter plots (Shi et al., 2014). Cryo-TEM, optical density measurement and DLS
were also utilized to demonstrate co-assembly of FSI and ISR micelle nanoparticles.

4.2.3. Integrin-mediated cell targeting assay and fluorescence-activated cell
sorting (FACS)
Three different groups Rho labeled ISR (100% ISR) and FSI (100% FSI), and Rho
labeled FSI mixed with plain ISR (1: 1 molar mixing ratio, 50% ISR 50% FSI) were
incubated for 1h with human umbilical vein endothelial cell (HUVEC) and MDA-MB-468
breast cancer cells respectively. The cells were then treated with 4',6-diamidino-2-
phenylindole (DAPI) for nucleus staining, and the fixed cells were imaged using a Zeiss
LSM 510 Meta NLO confocal microscopy (Thornwood, NY). All the images were
analyzed by ImageJ (NIH) for the comparison of pixel intensity. In the assay of
fluorescence-activated cell sorting (FACS), HUVEC and MDA-MB-468 cells were
treated with the mixtures of Rho labeled FSI and plain ISR (%ISR from 5% to 50%) for
1h. 100% Rho-FSI and 100% Rho-ISR were used as the negative and positive controls
for this assay. The cells were then washed, re-suspended in 500 μl DPBS and analyzed
105
on an Attune® acoustic focusing cytometer (Life Technologies, CA). The data files were
exported as .fcs and analyzed on FlowJo. The negative cells were gated for baseline
fluorescence and all cells above the set gate are positively stained with Rho labeled
FSI/ plain ISR mixtures.
4.2.4. Evaluation of Rapamycin (Rapa) encapsulation and release
A two-phase solvent evaporation method has been developed for Rapa encapsulation
using ELP nanoparticles. The detailed procedures can be found in our previous
publication (Shi et al., 2013b). In brief, FSI, ISR, 50% FSI 50% ISR co-assembled
micelles were used to encapsulate Rapa using the two-phase solvent evaporation
method while mono-block S192 was included as a negative control. Dialysis analysis
was utilized to perform Rapa release experiments where samples in the dialysis
cassette were collected from 0 to 48 h. Using Prism software, Rapa release curves of
FSI, ISR and 50% FSI 50% ISR co-assembled micelles were plotted, and different
release phases and release half-lives were predicted using mono-phase or bi-phase
exponential decay models.

4.2.5 In vivo tumor regression studies using human breast cancer xenografts
All animal procedures were approved by the Institutional Animal Care and Use
Committee (IACUC) at the University of Southern California. 6 × 10
6
MDA-MB-468
human breast cancer cells were injected into the mammary fat pads of 7-week-old
female athymic nude (nu/nu) mice (Harlan, Inc.). The first treatment started after the
average tumor size reached around 50 mm
3
, and then the mice were randomly divided
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into different groups in three individual studies. There were four groups PBS, S48I48
Rapa (0.25mg/kg), dialyzed slow-release (Ds) FSI Rapa (0.25mg/kg) and free Rapa
(DMSO, 0.25mg/kg) in the first study; five groups PBS, DsFSI Rapa 0.25mg/kg, DsFSI
Rapa 0.075mg/kg, DsFSI Rapa 0.025mg/kg and DsFSI Rapa 7.5μg/kg in the second
study; and four groups PBS, DsISR/FSI (co-assembled) Rapa 0.075mg/kg, DsFSI Rapa
0.075mg/kg and free Rapa (DMSO, 0.075mg/kg), respectively. The doses were injected
intravenously three times a week through mouse tail veins, the widths and lengths of the
tumors were measured using an electronic caliper, and the tumor sizes were estimated
based on a formula a
2
× b × /6 (a, b are the width and length of the tumor
respectively).

4.3. Results and discussion
4.3.1. Co-assembly of multi-functional ELP micelle nanoparticle
To introduce active tumor targeting strategy, RGD tri-peptide was genetically
conjugated onto the C-terminus of ELP di-block copolymer I48S48, generating a new
functional ELP fusion protein – I48S48-RGD (ISR) (Figure 24 and Table 3). The
physicochemical and morphological properties of ISR were studied using optical density
measurement and dynamic light scattering, and cryogenic-transmission electron
microscopy (Cryo-TEM) imaging respectively (Figure 25a, b and Figure 24d). There
was only a slight difference (up to 2
o
C) in transition temperatures (Tt) of ISR and FSI
due to the variation in protein molecular weight (Figure 25a and Table 3). Both ISR and
FSI assembled stable micelle nanoparticles above their Tt with similar hydrodynamic
radius of 23 nm (Figure 25b). The similarity in the physicochemical properties and
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assembled nanoparticle structures of ISR and FSI ascertained the co-assembly strategy
of generating a novel bi-functional mixed micelle nanoparticle (Figure 24a). To verify
the efficient co-assembly, ISR and FSI were fluorescently labeled with rhodamine (Rho)
and carboxyfluorescein (CF) respectively. Confocal images illustrated high degree
colocalization of these two ELP fusions (Figure 24b). JACoP colocalization analysis of
red (ISR) and green (FSI) pixels demonstrated that ISR and FSI had near-perfect
colocalization with Pearson's Coefficient r=0.976 and Overlap Coefficient r=0.982. The
positive slope of the linear regression line (slope = 0.898) and close distribution along
the line (r
2
= 0.953) in the pixel plot further emphasized the efficient co-assembly of the
two ELPs (Figure 24c). Moreover, Cryo-TEM, optical density measurement and DLS
analysis revealed that the co-assembled mixed micelle nanoparticles possessed parallel
morphology, Tt and hydrodynamic radius with either ISR or FSI (Figure 24d and Figure
26a and b).

108

Figure 24. Co-assembly of multi-functional ELP micelle nanoparticle. a. Because
of the structural similarity, FSI and ISR can co-assemble into mixed micelle
nanoparticles with both functions of specific drug delivery and active tumor targeting. b.
In vitro micelle co-assembly assay using fluorescently labeled FSI and ISR showed
spatial colocalization of two ELPs. c. JACoP colocalization analysis of red and green
pixels of the confocal images in the co-assembly assay (b) indicated strong
colocalization of the two colors. d. Cryo-TEM images of ISR (left) and 50% FSI 50% ISR
(right).
109

Figure 25. Physicochemical properties of ISR and FSI. a. ELP transition temperature
(Tt) measurement using UV-Vis spectrophotometer. There is only a slight difference in
Tt of ISR and FSI (up tp 2
o
C). b. Hydrohynamic radius measurement using dynamic
light scattering (DLS) analysis. The measurement showed the assembly of micelle
nanoparticles with a radius of 23nm from both ISR and FSI.

110

Figure 26. Optical density measurement and DLS analysis of ISR and FSI
mixtures. a. The optical density measurement revealed similar transition features (Tt
and solution turbidity) of all ISR and FSI mixtures. b. DLS analysis demonstrated all ISR
and FSI mixtures assembled micelle nanoparticles with comparable sizes.

111

4.3.2. In vitro tumor targeting of the co-assembled multi-functional ELP micelle
nanoparticle
Integrin-mediated cell targeting assay was used to evaluate the functionality of active
tumor targeting of the co-assembled mixed micelle nanoparticles in vitro. Human
umbilical vein endothelial cell (HUVEC) and MDA-MB-468 human breast cancer cell
have been previously reported to have high expression level of various types of v
integrins and therefore were excellent candidates for this assay (Goodman et al., 2012,
Meyer et al., 1998). In both types of cells, confocal imaging of Rhodamine (Rho) labeled
ISR illustrated efficient cell targeting by displaying bright red fluorescence while Rho
labeled FSI did not have specific cell targeting at all (Figure 27a and b). However, the
fluorescence intensity indicating integrin-mediated cell targeting was elevated from
almost none to nearly 50% of that of Rho labeled ISR when Rho labeled FSI was mixed
with non-labeled ISR to co-assemble into 50% ISR 50% FSI mixed micelles (Figure
27a, b, c and d). Therefore, the co-assembled mixed micelle nanoparticles were
demonstrated to have significant functionality of specific integrin-mediated cell targeting
using this cell-based assay. To further quantitatively determine the relationship between
the mixing ratio of co-assembly and ability of cell targeting, fluorescence-activated cell
sorting (FACS) assay was performed using both HUVEC and MDA-MB-468 cells. A
trend of increasing number of positive cells with increasing percentage of ISR in the
mixing ratios of co-assembly was discovered indicating the critical role of ISR in
integrin-mediated cell targeting (Figure 27e). It was observed that 50% ISR 50% FSI
co-assembled micelles effectively targeted about 90% cells which was comparable to
112
the positive control 100% ISR. Additionally, a competitive cell-based integrin binding
assay was performed in the comparison of a cyclic form of RGD - c-RGDfK and ISR
using MDA-MD-468 cells. Confocal imaging and a quantitative image analysis using
ImageJ both showed that high concentrations of c-RGDfK could totally block the cell
targeting of ISR due to a higher integrin binding affinity of c-RGDfK comparing to ISR
(Figure 28a and b). It confirmed the specificity of ISR cell targeting through cell surface
binding of v integrins and validated the novel ELP micelle co-assembly strategy to
introduce integrin-mediated active tumor targeting.

113

Figure 27. Integrin-mediated cell targeting assay using co-assembled mixed
micelle nanoparticles. a, b. HUVEC cells (a) and MDA-MB-468 cells (b) were
incubated with rhodamine labeled ISR, FSI and 50%ISR 50%FSI, respectively. The
images were captured using a confocal microscope. c, d. Fluorescence quantitation
analysis of confocal images of HUVEC (c) and MDA-MB-468 (d) targeting assays
demonstrated that 50% ISR 50% FSI co-assembled micelles efficiently targeted both
cell types. e. FACS assays using different mixing ratios of ISR and FSI showed an
increasing cell targeting effect with an increasing percentage of ISR in the mixture in
both HUVEC and MDA-MB-468 cells.

114

Figure 28. Competitive binding assay using cyclic-RGD (c-RGDfK) and ISR
demonstrated the direct binding of ISR to the integrin on cell surface. a. Rho
labeled ISR was added to MDA-MB-468 cells after 1h pre-incubation of unlabeled
cyclic-RGD (c-RGDfK) and washing steps. Confocal images were captured after 1h
incubation with a confocal microscope. b. Fluorescence quantitation analysis of confocal
images of MDA-MB-468 cell competitive binding assay. The result confirmed that c-
RGDfK could completely block cell targeting of rhodamined labeled ISR which indicated
the direct binding of ISR to the integrin on the cell surface.

115

4.3.3. In vitro evaluation of drug (Rapa) loading and release
One of the advantages of ISR/FSI co-assembled mixed micelle is that by adjusting the
mixing ratio of ISR and FSI for the co-assembly, the functionality of the mixed micelle
can be prone to either tumor targeting or specific drug delivery (Figure 29a). It has been
noticed that 50% ISR 50% FSI co-assembled micelles efficiently targeted 90% cells in
the previous FACS assay using both HUVEC and MDA-MB-468 cells. Therefore, this
co-assembled micelle (50FSI/50ISR) was evaluated for its drug loading and release
capability based on its responsible FSI domains. Using the same approaches that were
used in our previous studies, an anti-proliferative cancer drug Rapamycin (Rapa) was
loaded to FSI (100%), ISR (100%) and 50FSI/50ISR micelle nanoparticles respectively
(Shi et al., 2013b). The Rapa loading efficiency of 50FSI/50ISR was 58% which lied in
the middle of FSI (73%) and ISR (29%), reflecting the reasonable drug loading capacity
with intermediate FSI content (Figure 29b). The negative control S192 only had
baseline level drug loading because the mono-block ELP did not assemble any micelle
nanostructures. Rapa release profiles of FSI, ISR and 50FSI/50ISR were examined
using sink-condition dialysis assay. FSI and 50FSI/50ISR possessed completely
different drug release kinetics from ISR due to specific drug binding of FSI domain
(Figure 29c). FSI and 50FSI/50ISR both displayed two exponential decay phases of
Rapa release: a fast release phase with half-lives of 1.9h (FSI) and 1.4h (50FSI/50ISR)
and a slow release phase with half-lives of 57.8h (FSI) and 38.4h (50FSI/50ISR). On the
contrary, only one exponential decay phase was detected for ISR with a half-life of 1.5h
similar to the fast release phase of FSI and 50FSI/50ISR (Figure 29c). The release
116
profile of 50FSI/50ISR was comparable to FSI with slightly smaller half-lives for both
release phases indicating slightly less tight drug-specific binding through FSI domains.
The results concluded that 50% FSI in the co-assembled mixed micelle was
demonstrated to be sufficient to specifically deliver Rapa with relatively large drug
loading capacity and slow controlled release comparing with 100% FSI.

117

Figure 29. ISR/FSI co-assembled mixed micelle nanoparticle has great Rapa
loading capacity and releases the drug similarly to FSI. a. The mixing ratio of ISR
and FSI to co-assemble into mixed micelle can be controled with a balance of functions
of specific drug delivery and active tumor tagerting. b. Rapa loading assay
demonstrated efficient drug loading capability of ISR/FSI mixed micelle (~60%)
compared to FSI, ISR and S192. c. ISR/FSI mixed micelle had similar two exponential
decay phases of Rapa release to FSI with a 1.4h half-life of fast release phase and a
38.4h half-life of slow release phase.
118

4.3.4. The multi-functional nanocarrier delivers Rapa, targets tumor and inhibit its
growth with greater efficiency compared to other formulations in vivo
In vivo tumor regression studies were performed using MDA-MB-468 breast tumor cell
mouse xenografts. The first study was to evaluate and compare the anti-tumor efficacy
of fast-released Rapa and slow-released Rapa. During the encapsulation process of
Rapa using FKBP domain contained ELP micelles (FSI or 50FSI/50ISR), the small
molecule drug can be either associated with FKBP through specific receptor – ligand
binding or with ELP micelle core through non-specific hydrophobic interactions (Shi et
al., 2013b). Because the specific binding between Rapa and FKBP is much stronger
than the non-specific hydrophobic interactions, it has been assumed in our previous
publication that Rapa associated with micelle core was released first (fast-phase
release) followed by Rapa bound to FKBP domains (slow-phase release) (Shi et al.,
2013b). To simplify the complication and study the two release phase separately,
dialysis (12 h) was performed following initial drug loading to remove fast-released
Rapa in micelle core. The dialyzed sample only containing slow-released Rapa on
FKBP domains was named dialyzed-slow released FSI Rapa (DsFSI Rapa). Moreover,
ELP di-block S48I48 (SI) with no FKBP domain loaded with 100% fast-released Rapa
and was compared with DsFSI Rapa. The results revealed the elevated anti-tumor
efficacy of DsFSI Rapa comparing to SI Rapa and the free drug (Figure 30a). The
advantage of slow-released Rapa formulation DsFSI Rapa in suppressing MDA-MB-468
breast tumor growth might be explained by the longer circulation time in the central
blood system and thereby higher tumor accumulation via enhanced permeability and
119
retention (EPR) effect. The initial study was immediately followed by a dose de-
escalation study of DsFSI Rapa to determine the lowest dose that remained effective
anti-tumor efficacy. Besides the initial dose (50μM, 0.25mg/kg), DsFSI Rapa doses 15
μM (0.075mg/kg), 5 μM (0.025mg/kg) and 1.5 μM (7.5μg/kg) were tested in different
groups of mouse tumor xenografts. A decreasing trend of anti-tumor efficacy was
confirmed with the decrease of the doses (Figure 30b). It was noticed that DsFSI Rapa
dose 15 μM (0.075mg/kg) had an intermediate anti-tumor efficacy between the highest
dose (50μM, 0.25mg/kg) and the blank (PBS); therefore, this dose was selected to be
used in the next study which included the active tumor targeting formulation of Rapa.
The in vivo study was performed to examine and compare the efficacy of the passive
tumor targeting formulation of Rapa (DsFSI Rapa) and the active tumor targeting
formulation (dialyzed-slow released Rapa encapsulated in 50FSI/50ISR co-assembled
micelles, DsISR/FSI Rapa) in suppressing breast tumor growth with PBS group and free
Rapa group as the controls. Excitingly, it was discovered that the active tumor targeting
formulation of Rapa (DsISR/FSI Rapa) successfully inhibited the tumor growth at the
dose of 15 μM (0.075mg/kg), and the anti-tumor efficacy was much greater than DsFSI
Rapa and free Rapa at the same dose (Figure 30c). The Kaplan-Meier survival curves
also revealed that DsISR/FSI Rapa significantly prolonged mouse survival time
compared to DsFSI Rapa, free Rapa and PBS groups (Figure 30d). Furthermore, the
monitored mouse weight diagrams indicated no obvious cytotoxicity from all Rapa
formulations in all three studies (Figure 31a, b and c). Therefore, the new active tumor
targeting strategy introduced by co-assembling RGD and FKBP containing ELP di-
blocks into bi-functional mixed micelles has been demonstrated to significantly improve
120
anti-tumor efficacy of the passive targeting formulation DsFSI Rapa and lower the
effective dose to 15 μM (0.075mg/kg). The novel co-assembled 50FSI/50ISR mixed
micelles efficiently reached tumor sites via high-avidity RGD-v integrin binding and
specifically delivered Rapa to achieve noticeable tumor growth inhibition.

121

Figure 30. Mouse tumor regression studies demonstrated actively tumor targeting
ISR/FSI Rapa had greater anti-tumor efficacy than passively tumor targeting FSI
Rapa and free Rapa in MDA-MB-468 breast tumor xenografts. a. Dialyzed slow-
release FSI Rapa (DsFSI Rapa) had better tumor growth inhibition effect than SI Rapa
(fast release) and free Rapa at the dose of 50M (0.25mg/kg). b. Dose de-escalation
study indicated an decreasing tumor suppression efficacy with decreasing doses of
DsFSI Rapa. c. Compared to DsFSI Rapa and free Rapa, DsISR/FSI Rapa actively
targeted tumor and more effectively inhibited its growth at the dose of 15M (75g/kg).
122
d. Kaplan-Meier survival analysis showed that active tumor targeting Rapa formulation
(DsISR/FSI Rapa) significantly prolonged mouse survival time compared to other
formulations.
123

Figure 31. Stable mouse body weights during tumor regression studies indicated
that the formulations did not have obvious cytotoxicity. No dramatic mouse body
weight changes were observed in all three tumor regression studies (a, b and c)
indicating all the formulations used in the studies were non-cytotoxic.
124
4.4. Conclusion
Instead of using complex and cytotoxic chemical conjugation techniques for chemically
synthetic polymers, our group had accomplished introducing differerent functional
groups onto protein polymers and assembled multi-functional nanocarriers by genentic
protein polymer engineering and micelle co-assembly strategy. The functions of novel
protein-based nanocarrier containing FKBP for specific Rapa delivery and RGD for
integrin-mediated active tumor targeting have been examined and confirmed by in vitro
and in vivo assays. Despite the limited efficiency of passive drug delivery in non-leaky
tumors, RGD contained ISR/FSI co-assembled micelle nanoparticle effectively reached
MDA-MB-468 tumor site via high-avidity binding of RGD to highly expressed v integrin
on endothelial and tumor cells. Compared to passive targeting Rapa formulation and
free drug, active tumor targeting Rapa formulation ISR/FSI Rapa was demonstrated to
have a greater anti-tumor efficacy at the dose of 75μg/kg with no obvious cytotoxicity.
This reported new stratrgy to assemble multi-functional drug carrier by genetic
engineering and micelle co-assembly provides a useful tool in the development of next-
generation “smart” multi-functional drug carriers.

125
5. Summary and Future Directions
Three major research goals have been achieved in the four years of my PhD study
period. First of all, using amphiphilic elastin-based protein polymers and a drug specific
binding domain, a novel high-avidity drug-specific nanocarrier has been developed as
described in Chapter 2. Large cyclic hydrophobic drugs (i.e. Rapamycin) easily partition
into certain cells (i.e. red blood cells) and induce cytotoxicity.It is therefore very
important to maximize the release half-life of the drug and avoid the quick partition that
causes severe cytotoxicity (Yanez et al., 2008). Using in vitro drug release assay, the
actual release half-life of Rapamycin from the high-avidity drug-specific nanocarrier
FKBP-S48I48 (FSI) was about 58 hours.This was significantly longer than non high-
avidity carriers that mainly use hydrophobic interactions. It was also discovered in a
PET imaging study that an in vivo circulation half-life of the nanocarrier itself was about
5-6 hours. It is assumed that the drug will remain associated with the carrier during its
entire circulation in vivo. It is believed that nanoparticles use EPR effect to accumulate
in the tumor site; therefore, the prolonged drug release half-life may effectively increase
drug accumulation (Danhier et al., 2010, Matsumura and Maeda, 1986). An in vivo
tumor regression assay demonstrated that 0.75 mg/kg Rapamycin delivered by FSI had
greater anti-tumor efficacy compared to the free drug without observed cytotoxicity. The
effective dose of Rapamycin is more than 10 times lower than the doses used in other
published studies (Namba et al., 2006). This strategy of genetic fusion with a congnate
drug receptor can be widely applied to the delivery of other drugs. The similar genetic
fusion can be completed by conjugating a well identified drug binding domain and a
126
similar ELP block copolymer. This new fusion protein is assumed to have high-avidity
binding ability to its ligand. Second, intracellular co-assembly and sorting of genetically
engineered protein microdomains have been studied. Based on the difference in the
composition of ELPs, either large micrometer-size microparticles or small nanoparticles
can be assembled in PBS solution. Before this study, our understanding of how ELPs
assemble protein microdomains in the cytosol – a more complex intraceullar
environment was very limited. Discussed in Chapter 3, a cell line that stably expresses
DsRed-V96 has been established, which enables us to assemble two different ELPs in
the small cell using manual DNA transfection. It was discovered that ELPs with similar
composition co-assembled into large coacervates, while ELPs with different
compositions self-sorted into distinct protein microdomains. This unexpected finding
provides strong evidence for the efficient co-assembly of similar ELP fusions in vivo that
serves as one of the footstones in the research discussed in Chapter 4. The
demonstration of intracellular co-assembly and the self-sorting of ELP protein
microdomains also enlighten us to the development of synthetic organelles. Similar to
the study described in Chapter 3, ELP fusions that contain essential functional domains
may co-assemble to form a synthetic organelle, compensating the functionality loss in
many metabolic diseases. The final goal that has been achieved is active tumor
targeting using multi-functional ELP nanoparticles. Although EPR effect plays an
important role in the tumor targeting strategy of nanoparticles, it is not applicable in
many cases because some solid tumors do not grow leaky vasculatures (Danhier et al.,
2010). Therefore, an active rather than passive tumor targeting strategy is required to
treat non-leaky tumors (Allen, 2002). As described in Chapter 4, RGD peptide was
127
genetically conjugated onto an ELP block copolymer, generating a new ELP fusion:
ISR. Based on the demonstration of the co-assembly of similar ELPs in Chapter 3, ISR
and FSI co-assembled into mixed nanoparticles that have two important functions:
specific drug loading,release, and active tumor targeting. Both in vitro and in vivo
assays were performed to demonstrate effective tumor targeting, intact drug loading,
and release capability. Tumor regression studies revealed that the active tumor
targeting formulation of Rapa had a greater anti-tumor efficacy than the passive tumor
targeting formulation and the free drug. The new strategy further reduced Rapa dose to
75g/kg, which indirectly indicated more efficient drug accumulation in the tumor.

The fabrication of multi-functional ELP nanoparticles is believed to be a milestone in the
field of nanoparticle-based drug delivery. The functional ELP fusions in our study can be
transformed for delivery of other drugs. This can be done by changing the active
targeting domain and/or drug binding domain using simple genetic fusion
technology.With this technology, new multi-functional nanoparticles can be easily
fabricated. Furthermore, a screening assay can be designed and perfomed by
genetically fusing different active tumor targeting domains such as cyclic-RGD/iRGD,
DGEA, CPPs, NGR, etc. and testing the in vitro and in vivo tumor accumulation of the
drug carriers.

The specific high-avidity ligand-receptor binding enables the nanocarrier to be highly
specific to the ligand of the receptor on the ELP fusion, and assures an extended drug
release half-life. The specificity does not compromise the flexibility of the carrier
128
because new functional ELP fusions are fabricated relatively easily and rapidly.
Compared to complicated chemical reactions and multiple by-products in the synthesis
of chemically synthetic polymers, producing different drug-specific nanocarriers requires
only simple-step molecular cloning, protein expression, and purification.

Multi-functional nanocarriers might be the next-generation drug carriers for the future.
There is a large pool of potential functional domains for genetic engineering of multi-
functional ELP nanoparticles. Besides drug binding and tumor targeting domains, other
functional groups, such as temperature and/or pH sensitive cleavable domains, might
also be good candidates.

Inspired by the exciting results obtained from in vivo tumor regression studies, our
group is highly motivated to move the study of Rapa active and passive tumor targeting
formulation (ISR/FSI Rapa and FSI Rapa) forward into a pre-clinical safety study and
clinical Phase I study. Although limited correlation has been established between animal
and human oncology studies, it is still strongly believed by our group that these novel
Rapa formulations, using innovative approaches and perspectives, will be a great
success in the clinical studies.

129
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Abstract (if available)

Abstract Chapter 1: Cytotoxicity, low water solubility, rapid clearance from circulation, and off‐target side‐effects are common drawbacks of conventional small‐molecule drugs. To overcome these shortcomings, many multifunctional nanocarriers have been proposed to enhance drug delivery. In concept, multifunctional nanoparticles might carry multiple agents, control release rate, biodegrade, and utilize target‐mediated drug delivery

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Asset Metadata

Creator Shi, Pu (author)

Core Title Integrin-mediated targeting of protein polymer nanoparticles carrying a cytostatic macrolide

School School of Pharmacy

Degree Doctor of Philosophy

Degree Program Pharmaceutical Sciences

Publication Date 08/05/2014

Defense Date 05/27/2014

Publisher University of Southern California (original), University of Southern California. Libraries (digital)

Tag drug delivery,elastin-like polypeptide,nanoparticle,OAI-PMH Harvest,tumor targeting

Format application/pdf (imt)

Language English

Contributor Electronically uploaded by the author (provenance)

Advisor Mackay, John Andrew (committee chair), Epstein, Alan L. (committee member), Hamm-Alvarez, Sarah F. (committee member), Okamoto, Curtis Toshio (committee member), Shen, Wei-Chiang (committee member)

Creator Email pushi@usc.edu,sheenpearl@gmail.com

Permanent Link (DOI) https://doi.org/10.25549/usctheses-c3-453262

Unique identifier UC11287153

Identifier etd-ShiPu-2768.pdf (filename),usctheses-c3-453262 (legacy record id)

Legacy Identifier etd-ShiPu-2768.pdf

Dmrecord 453262

Document Type Dissertation

Format application/pdf (imt)

Rights Shi, Pu

Type texts

Source University of Southern California (contributing entity), University of Southern California Dissertations and Theses (collection)

Access Conditions The author retains rights to his/her dissertation, thesis or other graduate work according to U.S. copyright law. Electronic access is being provided by the USC Libraries in agreement with the a...

Repository Name University of Southern California Digital Library

Repository Location USC Digital Library, University of Southern California, University Park Campus MC 2810, 3434 South Grand Avenue, 2nd Floor, Los Angeles, California 90089-2810, USA