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Centre for Drug Research
Division of Biopharmaceutics and Pharmacokinetics
Faculty of Pharmacy
University of Helsinki
Finland
In vitro, in vivo, and in silico investigations of
polymer and lipid based nanocarriers
for drug and gene delivery
Julia Lehtinen
ACADEMIC DISSERTATION
To be presented, with the permission of the Faculty of Pharmacy of the University of
Helsinki, for public examination in Auditorium 2 at Viikki Korona Information Centre,
Viikinkaari 11, on 7th
September 2013, at 12 noon.
Helsinki 2013
Supervisors: Professor Arto Urtti, Ph.D.
Centre for Drug Research
Faculty of Pharmacy
University of Helsinki
Finland
Professor Heike Bunjes, Ph.D.
Institute of Pharmaceutical Technology
Technische Universität Braunschweig
Germany
Mathias Bergman, Ph.D.
Karyon Ltd.
Helsinki
Finland
Reviewers: Docent Juha Holopainen, M.D., Ph.D.
Institute of Clinical Medicine
Faculty of Medicine
University of Helsinki
Finland
Professor Stefaan de Smedt, Ph.D.
Laboratory of General Biochemistry & Physical Pharmacy
Faculty of Pharmaceutical Sciences
University of Ghent
Belgium
Opponent: Academic Rector, Professor Jukka Mönkkönen, Ph.D.
University of Eastern Finland
Finland
© Julia Lehtinen 2013
ISBN 978-952-10-9039-4 (pbk.)
ISBN 978-952-10-9040-0 (PDF, http://ethesis.helsinki.fi)
ISSN 1779-7372
Helsinki University Print
Helsinki, Finland 2013
3
Abstract
Nanomedicine research has expanded rapidly in the last decades. Several nanoparticle
formulations are accepted in clinical use, e.g. for the treatment of cancer, infections and
eye diseases, and also for diagnostics. Nanoparticle mediated drug delivery has many
potential advantages over the free drug, such as better pharmacokinetic profile, lowered
toxicity, and its possible use for cell-specific targeting and intracellular drug release.
Therapeutic genes can also be packed into nanocarriers to protect them from enzymatic
degradation and to mediate their cellular entry. The transfection efficacy of these synthetic
vectors is modest when compared to viral vectors, but they are considered to be safer.
Nonetheless, even though nanoparticles have so many advantages, there are many
extracellular and intracellular barriers to overcome before achieving successful drug or
gene delivery.
The focus of this research work was the formation and physico-chemical features of
lipid and polymer based nanoparticles for drug and gene delivery. In addition, two classes
of cancer cell targeting approaches were evaluated in biological and physical studies. First,
the effect of the polymeric gene carrier composition and structure on DNA condensation
efficacy, transgene expression, and cellular toxicity was examined. The linear architecture
and flexibility of poly(2-(dimethylamino)ethyl methacrylate) (PDMAEMA)-based block
co-polymers clearly enhanced DNA condensation and transfection efficiency. In addition,
by conjugating a membrane active protein, hydrophobin (HFBI) to DNA-binding cationic
dendrons, the transfection efficacy was increased compared to plain dendron. However,
cationic polymer-DNA complexes are prone to disruption by polyanions such as
glycosaminoglycans (GAG) in the extracellular space. We coated poly(ethyleneimine)
PEI/DNA complexes with anionic lipid mixture. The coating was able to protect the
contents against GAGs, and it could respond to the change of endosomal pH and release
the cargo inside the cells.
The next studies aimed to evaluate targeted liposomal cancer drug carriers in
physicochemical studies and in cancer cell models in vitro and in mice. A promising
activated endothelium targeting peptide (AETP) failed to target the liposomes to the cells.
Molecular modeling revealed that hydrophobic AETP was hidden in the PEG shield of the
liposomal surface thus it was not accessible for the target receptors. The last study
describes applicability of pre-targeting and local intraperitoneal administration of
liposomes for drug targeting to tumors located in peritoneal cavity. Epithelial growth
factor receptor (EGFR)-targeted liposomes bound specifically to ovarian cancer cells in
vitro. In the animal study, increased accumulation of liposomes in the xenograft tumors of
the mice was seen after intraperitoneal administration compared to intravenous
administration.
In conclusion, the composition and architecture of nanocarriers have a crucial impact
on DNA condensation, stability of the complexes and transfection efficacy. In liposomal
cancer drug targeting, polyethylene glycol (PEG) shield may hinder the targeting
efficiency of small molecular peptides. Intraperitoneal administration of liposomal drugs
seems to be promising route for targeting to tumors located in the peritoneal cavity.
4
Acknowledgements
This study was carried out at the University of Kuopio, in the Department of
Pharmaceutics, (currently University of Eastern Finland, School of Pharmacy) during the
years 2003-2005 and it was continued at the University of Helsinki, Division of
Biopharmaceutics and Pharmacokinetics and the Centre for Drug Research (CDR) during
the years 2006-2013. This work has been financially supported by the Academy of
Finland, the Association of Finnish Pharmacies, the Finnish Cultural Foundation, the
Finnish Pharmaceutical Society, the National Agency of Technology (TEKES Finland),
the Science Foundation of Orion-Farmos, and the University of Helsinki. All financial
support is greatly acknowledged.
I want to express my deepest gratitude to my principal supervisor, Professor Arto Urtti
for his continuous support and optimistic attitude during these years. I am also very
grateful to my other supervisors, Professor Heike Bunjes for introducing me to the world
of nanoparticles and Mathias Bergman, Ph.D., for his skilful advice and guidance in
peptide targeting.
Professor Stefaan de Smedt and Docent Juha Holopainen are greatly acknowledged
for critical reading of this dissertation and for their valuable comments. I am honored that
Professor Jukka Mönkkönen has accepted the invitation to be my opponent in the public
defense of this thesis.
I would like to thank the current and former Deans of Faculty of Pharmacy in Kuopio
and in Helsinki, and Heads of Department of Pharmaceutics and Heads of Division of
Biopharmaceutics and Pharmacokinetics for providing excellent working facilities.
I wish to warmly thank my co-authors: Anu Alhoranta, M.Sc., Kim Bergström, Ph.D.,
Alex Bunker, Ph.D., Annukka Hiltunen, M.Sc., Zanna Hyvönen, Ph.D., Professor Olli
Ikkala, Raimo Ketola, Ph.D., Mauri Kostiainen, Ph.D., Katariina Lehtinen, M.Sc., Huamin
Liang, Ph.D., Aniket Magarkar, M.Sc., Ann-Marie Määttä, Ph.D., Jere Pikkarainen, Ph.D.,
Sari Pitkänen, M.Sc., Mari Raki, Ph.D., Tomasz Róg, Ph.D., Michał Stepniewski, M.Sc.,
Astrid Subrizi, M.Sc., Professor Heikki Tenhu, Päivi Uutela, Ph.D., Thomas Wirth, Ph.D.
and Professor Marjo Yliperttula for their valuable contribution to this work. It has been a
pleasure to collaborate with you all. I am also very grateful to Lea Pirskanen in Kuopio
and Leena Pietilä in Helsinki for their skilful and friendly assistance in laboratory. I also
wish to thank the personnel of Karyon ltd for welcoming me to do part of my Ph.D. work
in their laboratory.
My sincere thanks go to my friends and colleagues in the Faculty of Pharmacy in
Kuopio, and in the CDR and Drug Delivery and Nanotechnology (DDN) group in
Helsinki. Special thanks to the girls of the girls’ room: Astrid, Heidi, Johanna, Jonna,
Kati-Sisko, Mari, Marika, Martina, Melina and Polina for their friendship, joyful
company, and refreshing conversions.
Finally, I want to warmly thank my friends and relatives for their support and presence
5
during these years. I owe my dearest gratitude to my husband and colleague Mika for his
love and support but also for scientific advice, and to our wonderful children Viljam and
Hilda for brightening up our everyday life.
Helsinki, July 2013
Julia Lehtinen
6
Contents
Abstract 3
Acknowledgements 4
List of original publications 8
Abbreviations 9
1 Introduction 11
2 Review of the literature 13
2.1 Nanoparticles as drug and gene carriers 13
2.1.1 Liposomes 13
2.1.2 Polymeric carriers 15
2.1.3 Hybrid particles 17
2.2 Challenges in efficient drug and gene delivery 18
2.2.1 Stability of the vector 18
2.2.2 Tissue distribution and elimination 19
2.2.3 Cellular uptake 20
2.2.4 Intracellular distribution and cargo release 22
2.2.5 Diffusion in cytoplasm and nuclear import 23
2.3 Targeted cancer therapy 25
2.3.1 Passive targeting 25
2.3.2 Active targeting 26
2.3.2.1 Cancer cell targeting in solid tumors 27
2.3.2.2 Targeting to the tumor vasculature 27
3 Aims of the study 30
4 Overview of the methods 31
10 Summary of the main results 34
11 General discussion 36
7
11.1 Structure-activity relationship of polymeric DNA carriers on DNA-complex
formation, transfection efficacy, and toxicity 36
11.2 Lipid-coated DNA-complexes as stable gene delivery vectors 37
11.3 Hindering effect of liposomal PEG on the targeting efficiency of a small
hydrophobic peptide, AETP 39
11.4 Pre-targeting and local administration of liposomes as potential approaches in
tumor targeting 40
12 Conclusions 43
13 Future prospects 44
Nanoparticles – drugs of the future? 44
References 45
8
List of original publications
This thesis is based on the following publications:
I Anu M. Alhoranta, Julia K. Lehtinen, Arto O. Urtti, Sarah J. Butcher,
Vladimir O. Aseyev and Heikki J. Tenhu. Cationic amphiphilic star and
linear block copolymers: synthesis, self-assembly, and in vitro gene
transfection. Biomacromolecules 2011, 12, 3213-3222
II Mauri A. Kostiainen, Géza R. Szilvay, Julia Lehtinen, David K. Smith,
Markus B. Linder, Arto Urtti and Olli Ikkala. Precisely defined protein-
polymer conjugates: construction of synthetic DNA binding domains of
proteins by using multivalent dendrons. ACS NANO 2007, 1, 103-113
III Julia Lehtinen, Zanna Hyvönen, Astrid Subrizi, Heike Bunjes and Arto Urtti.
Glycosaminoglycan-resistant and pH-sensitive lipid-coated DNA complexes
produced by detergent removal method. Journal of Controlled Release 2008,
131, 145-149
IV Julia Lehtinen, Aniket Magarkar, Michał Stepniewski, Satu Hakola, Mathias
Bergman, Tomasz Róg, Marjo Yliperttula, Arto Urtti and Alex Bunker.
Analysis of course of failure of new targeting peptide in PEGylated
liposome: molecular modeling as a rationale design tool for nanomedicine.
European Journal of Pharmaceutical Sciences 2012, 46, 121-130
V Julia Lehtinen, Mari Raki, Kim A. Bergström, Päivi Uutela, Katariina
Lehtinen, Annukka Hiltunen, Jere Pikkarainen, Huamin Liang, Sari
Pitkänen, Ann-Marie Määttä, Raimo A. Ketola, Marjo Yliperttula, Thomas
Wirth and Arto Urtti. Pre-targeting and direct immunotargeting of liposomal
drug carriers to ovarian carcinoma. PLoS ONE 2012, 7(7):e41410
The publications are referred to in the text by their roman numerals
9
Abbreviations
Ab antibody
AETP activated endothelium targeting peptide
αvβ3, αvβ5 integrins upregulated in proliferating endothelial cells
bp base pair
BSA bovine serum albumin
CHEMS cholesteryl hemisuccinate
CMC critical micelle concentration
CPP cell penetrating peptide
CMV cytomegalovirus
DMPG 1,2-dimyristoyl-sn-glycero-3-phospho-(1'-rac-glycerol)
DNA deoxyribonucleic acid
DOPE 1,2-dioleyl-sn-glycerol-3-phosphoethanolamine
DOTAP 1,2-dioleyl-3-trimethylammonium-propane
DPPC 1,2-dipalmitoyl-sn-glycero-3-phosphocholine
DSPE-PEG 1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-amino(polyethylene
glycol)
DSPG 1,2-distearoyl-sn-glycero-3-phospho-(1'-rac-glycerol)
EGFR epidermal growth factor receptor
Egg PC egg phosphatidyl choline
Egg SM egg sphingomyelin
EMA-DNA ethidiummonoazide
EPR enhanced permeability and retention
Fab´ antigen-binding fragment of antibody
FACS fluorescence-activated cell sorting
FITC fluorescein isothiocyanate
Fmoc 9-fluorenylmethyloxycarbonyl
GAG glycosaminoglycan
HII inverted hexagonal structure of lipid membrane
HDL high density lipoprotein
HER-2 human epidermal growth factor receptor 2
HFBI hydrophobin
HIV-1 human immunodeficiency virus 1
HPMA N-(2-hydroxypropyl)-methacrylamide copolymer
HSA human serum albumin
HSPC fully hydrogenated phosphatidyl choline
HUVEC human umbilical vein endothelial cell
IgG immunoglobulin G
IgM immunoglobulin M
Kd dissociation constant
Lα multilamellar structure or liquid crystalline phase of lipid membranes
Lβ gel phase of lipid membranes
LCDC lipid-coated DNA complexes
10
LC-MS liquid chromatography - mass spectrometry
LDL low density lipoprotein
LUV large unilamellar vesicle
mAb monoclonal antibody
miRNA microRNA
MLV multilamellar vesicles
MPS mononuclear phagocyte system
mRNA messenger RNA
MTT (3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide
NCE new chemical entity
n/p nitrogen/phosphate ratio
ONPG ortho-nitrophenyl-β-D-galactopyranoside
PAMAM poly(amidoamine)
PBuA poly(n-butyl acrylate)
PC phosphatidylcholine
PDMAEMA poly(2-(dimethylamino)ethyl methacrylate)
PDP N-(3´-(pyridyldithio)propionoylamino
PE phosphatidylethanolamine
PEG polyethylene glycol
PEI poly(ethylene imine)
PG phosphatidylglycerol
PGA poly-L-glutamic acid
PL phospholipid
PLL poly-L-lysine
PS polystyrene
RGD arginine, lysine and aspartic acid containing peptide
Rho rhodamine
RNA ribonucleic acid
scFv single chain variable fragment of antibody
siRNA small interfering RNA
SPECT-CT single-photon emission computed tomography - computed tomography
TAT trans-acting activator of transcription
TRF time-resolved fluorescence
VEGFR vascular endothelial growth factor receptor
11
1 Introduction
Discovery of new potent drug molecules has significantly improved the treatment of
serious illnesses, such as cancer and cardiovascular diseases. However, the development
of new compounds to serve as clinical drugs has become more and more difficult. This is
evident by the decreasing numbers of new chemical entities (NCE) that are introduced
annually for clinical use. Many new compounds face significant development challenges,
such as poor water-solubility or short half-life in the blood circulation. In many cases,
adverse effects do hamper treatment, particularly in the case of anti-cancer drugs. Drug
delivery systems can be used to modify the drug properties, for example by increasing
solubility, modulation of pharmacokinetics, and improving the safety of drug treatment.
Gene therapy was launched in the 1990s as an alternative to traditional drug therapy. It
presents a promising alternative for the correction of genetic deficiencies, e.g. haemophilia
or cystic fibrosis, but also for the treatment of acquired diseases, such as cancer and
cardiovascular conditions. Gene medicines can either induce protein translation in the
target cells (gene therapy) or silence the expression of the target protein (oligonucleotide
drugs, e.g. siRNA). These compounds (plasmid DNA, siRNA) cannot be delivered as
such, because they undergo rapid enzymatic degradation and do not reach target tissues.
The negative charges of DNA and RNA, and a large size of plasmid DNA prevent their
passage through cell membranes.
Nanocarriers are one possible solution to overcome the pharmacokinetic challenges in
drug and gene delivery. Nanocarriers, often generally referred to as nanoparticles, are
typically below one micrometer in diameter, usually consisting of lipids (e.g. liposomes,
lipid-DNA complexes), polymers (e.g. polymeric nanoparticles and micelles, polymer-
DNA complexes), peptides, proteins and/or metallic nanoparticles. Liposomes were
already developed in the 1960s by Alec Bangham and polymer nanoparticles in the 1970s
by Peter Speiser. Thereafter, the field of nanomedicines has expanded tremendously:
currently almost 20 000 publications are found in PubMed database with the search terms
‘nanoparticle and drug’.
By formulating a drug in a nanocarrier, the solubility and the pharmacokinetic profile
can be dramatically improved (Shi et al. 2010). Nanoparticles may also decrease the toxic
effects of the drug from off-target sites. For example, liposomal encapsulation of
doxorubicin lowered the risk of cardiotoxicity seven times compared to the free drug
(O'Brien et al. 2004). Nanocarriers can be tailored to release drugs in a controlled manner
or triggered by a change in environmental conditions and they can be targeted to desired
cell type expressing the target protein. In addition, they can be used to deliver several
drugs simultaneously as a combination therapy (Zhang et al. 2008). Complexation of DNA
and RNA into small nanoparticles masks their negative charges and protects them against
enzymatic degradation. When a massive molecule of plasmid DNA (mw of millions) is
condensed to a nanosized particle it is more suitable for systemic delivery. In this case,
nanoparticle has dual function of DNA protection from enzymatic catalysis and
augmenting cellular entry.
Around 40 nanoparticle formulations are now accepted in clinical use, mainly for the
treatment of cancer, but also for the treatment of infections, anemia, hypercholesterolemia,
12
hepatitis, age-related macular degeneration, and in diagnostics (Duncan, Gaspar 2011).
The marketed products are so-called first generation nanomedicines that are not targeted.
The second generation targeted nanomedicines can bind to specific cellular antigens, but
they have not reached the market. Although targeted therapeutics have a lot of potential,
there are plenty of challenges and risks when more complicated formulations are designed
(Cheng et al. 2012).
Efficient but safe gene delivery vectors are still under development. For 500 million
years, viruses have developed a very efficient way to carry genetic material into cells.
Viral vectors are effective in DNA delivery, but their safety has not been totally verified.
In clinical studies, viral vectors have shown severe immunological reactions and even
caused patient death (Marshall 1999, Giacca, Zacchigna 2012). To date, three virus-based
gene therapy products have received market authorization from regulatory agencies in
China (Gendicine® and Oncorine
® for the treatment of cancer) and Europe (Glybera
®, for
the treatment lipoprotein lipase deficiency). Although non-viral polymer- and lipid-based
gene carriers lack the efficiency of the viral vectors, they are considered to be safer. In
addition, they are easier to synthesize and produce in large scale, and their DNA loading
capacity is higher than in the viral vectors (Kreiss et al. 1999). By learning from viruses,
more efficient synthetic carriers might also be developed.
In this study, the effect of the composition and architecture of polymer- and lipid-
based gene carriers on DNA condensation efficacy, transgene expression and cellular
toxicity was investigated. Furthermore, two types of liposomal cancer cell targeting
approaches were evaluated in physical and biological studies.
13
2 Review of the literature
2.1 Nanoparticles as drug and gene carriers
The family of nanocarriers includes lipid-based carriers, such as liposomes and micelles,
polymer-based carriers, such as polymer conjugates, polymeric nanoparticles and
dendrimers, gold nanoparticles and carbon nanotubes. The size of the nanocarriers usually
varies between a few nanometers (polymer-drug conjugates, micelles and dendrimers) to
some hundreds of nanometers (liposomes and polymeric nanoparticles). Examples of
nanocarriers used for drug and gene delivery are presented in Figure 1. In the following
review, liposomes and polymeric nanocarriers for drug and gene delivery are discussed in
more detail.
Figure 1 Schematic illustration of different kinds of nanoparticles used for drug and gene
delivery.
2.1.1 Liposomes
Liposomes are spherical, self-assembling vesicles formed by one or several lipid bilayers
leaving an aqueous core inside. The lipid bilayer is composed of amphiphilic lipids,
derived from or based on the structure of biological membrane lipids. The hydrophobic
part of the lipid is usually formed of two hydrocarbon chains, which typically vary from 8
to 18 carbons in length and can be either saturated or non-saturated. Long and saturated
acyl chains form a membrane in gel phase (Lβ), resulting in increased stability and rigidity
of the liposomes. On the contrary, the use of short and/or unsaturated acyl chains results in
more fluid, liquid crystalline (Lα) bilayers. Incorporation of cholesterol into the lipid
bilayer minimizes the membrane permeability and improves the mechanical strength of
the liposomes. Surface charge of the liposome can be affected by varying the hydrophilic
head group of the lipid: being either zwitterionic (e.g. phosphatidylcholine (PC) and
14
phosphatidylethanolamine (PE)), negatively charged (e.g. phosphatidylglycerol (PG)), or
positively charged (e.g. 3-trimethylammonium-propane (TAP)) (Figure 2) (Ulrich 2002).
Figure 2 Chemical structures of some phospholipids (fully hydrogenated soy
phosphatidylcholine (HSPC), 1,2-dioleyl-sn-glycerol-3-phosphoethanolamine (DOPE), 1,2-
distearoyl-sn-glycero-3-phospho-(1'-rac-glycerol) (DSPG), 1,2-dioleyl-3-trimethylammonium-
propane (DOTAP)) and cholesterol.
In water, amphiphilic lipids tend to form bilayers since they are poorly water soluble
with a critical micelle concentration (CMC) of 10-12
to 10-8
M. Spontaneously formed
multilamellar vesicles (MLV) are very heterogenous in lamellarity and in size, ranging
from 500 to 5000 nm. More sophisticated, small unilamellar vesicles (SUV, <100 nm) and
large unilamellar vesicles (LUV, 100–800 nm) can be prepared by sonication or extrusion
(Ulrich 2002, Torchilin 2005).
Hydrophobic drug molecules can be entrapped passively in the liposome bilayer during
the preparation of the liposomes using the aforementioned methods. Hydrophilic drugs are
encapsulated in the aqueous core of the liposome (or in the aqueous phase between
bilayers in the case of MLVs) using passive loading procedures, such as reverse phase
evaporation (Szoka Jr., Papahadjopoulos 1978), dehydration-rehydration method (Shew,
Deamer 1985), or active loading involving pH-gradient across the liposome membrane
(Mayer, Bally & Cullis 1986, Hwang et al. 1999). Remote loading of doxorubicin into
preformed liposomes using ammonium sulfate gradient as a driving force results in the
efficient and stable entrapment of the drug (Bolotin et al. 1994). Some liposomal cancer
drugs that are currently employed clinically, utilise remote loading; including Caelyx® and
Myocet® loaded with doxorubicin, and Daunoxome
® with daunorubicin.
Cationic liposomes can be used for complexation of negatively charged DNA or RNA
molecules. The formed complexes are called lipoplexes. The size of the highly cationic
15
lipoplexes varies typically between 100 and 450 nm, whereas the lipoplexes carrying a
charge close to neutral are more heterogenic, varying from 350 to 1200 nm in diameter. In
lipoplexes, two types of structures have been observed; multilamellar structure (Lα), where
DNA is located as a monolayer between cationic membranes (Radler et al. 1997), or
inverted hexagonal structure (HII), where DNA is encapsulated within cationic lipid
monolayer tubes (Koltover et al. 1998) (Figure 3). To enhance gene delivery, so called
helper lipids, such as 1,2-dioleyl-sn-glycerol-3-phosphoethanolamine (DOPE), are often
mixed with cationic lipids to promote conversion of the lamellar phase into a hexagonal
structure (Hafez, Cullis 2000).
Figure 3 Schematic structures of lamellar (A) and inverted hexagonal phase (B) in the
cationic lipid/DNA complexes. Modified from Morille et al. (2008).
2.1.2 Polymeric carriers
Polymer based carriers can be divided into different categories by their structure; 1)
polymeric nanoparticles have a structure of a capsule or matrix, 2) polymeric micelles of
amphiphilic polymers with core-shell structure are spontaneously formed in water, 3)
polymersomes are polymeric vesicles, membrane bilayer constructed from amphiphilic
polymers, and 4) dendrimers are hyperbranched structures, composed of multiple
branched monomers emerging radially from the core (Cho et al. 2008, Brinkhuis, Rutjes &
Van Hest 2011). Drug is usually either linked covalently to a polymer or physically
entrapped into the polymer capsule or matrix (Rawat et al. 2006). Even though natural
polymers such as albumin, chitosan, and heparin have been used for the delivery of drugs
and genetic material, the synthetic polymers may be preferable because they can be
designed and synthesized to achieve required properties. Among various polymers tested
N-(2-hydroxypropyl)-methacrylamide copolymer (HPMA), polyethylene glycol (PEG),
and poly-L-glutamic acid (PGA) are examples of synthetic polymers used for drug
delivery (Cho et al. 2008). Albumin-bound paclitaxel (Abraxane®) is an example of a
nanoparticle formulation in clinical use for the treatment of metastatic breast cancer.
Cationic polymers are able to bind and condense DNA into polyplexes which are even
smaller than lipoplexes formed after condensation of DNA with cationic liposomes
(Dunlap et al. 1997). Poly(ethylene imine) (PEI), poly-L-lysine (PLL), and poly(2-
(dimethylamino)ethyl methacrylate) (PDMAEMA) are well known polymers for gene
delivery (Figure 4). Charge ratio between positive nitrogen atoms of polymer and negative
16
phosphate groups of nucleic acid (n/p ratio) is an important factor in mediation of
transfection and toxicity. An excess of positive charges, n/p ratio 2.5 in the case of PEI, is
needed for total DNA condensation (Boeckle et al. 2006). Increasing the n/p ratio of
PEI/DNA complexes from 2 to 20 results in a decrease in a particle size from 1000 nm to
100-200 nm and a simultaneous reduction in a polydispersity (Erbacher et al. 1999). High
cationic charge of polyplexes leads to enhanced transfection efficiency, but as a drawback,
to higher toxicity because of the free polymer in solution (Hanzlíková et al. 2011).
It has been demonstrated that polymer architecture has an impact on the DNA
condensation and gene transfection properties of the polyplexes. Männistö et al. (2002)
demonstrated lower transfection activity for dendritic PLL compared to linear PLL. They
reasoned it might be due to unfavourable shape and orientation of dendritic amines for
DNA binding. However, in the case of PDMAEMA having a star-shaped architecture,
which mimics the structure of dendrimer, the polymer showed enhanced transfection
efficacy and reduced cytotoxicity compared to linear PDMAEMA (Xu et al. 2009). It has
been shown that the molecular weight of PEI strongly influences on the transfection
efficiency (Godbey, Wu & Mikos 1999). Choosakoonkriang et al. (2003) showed that both
branched and linear PEI (25 kDa) mediated higher transgene expression than smaller,
branched PEI 2 kDa. Linear PEI 22 kDa (ExGen 500) has proven to be more effective
than branched PEI 25 kDa in mediating transfection in lung epithelial cells both in vitro
and in vivo (Wiseman et al. 2003). However, linear PEI forms large unstable aggregates in
salt-containing medium that might explain its high gene delivery ability in vivo
(Wightman et al. 2001).
Figure 4 Chemical structures of some essential polymeric DNA carriers. Branched
poly(ethylene imine) (PEI), poly-L-lysine (PLL), poly(2-(dimethylamino)ethyl methacrylate)
(PDMAEMA), and polyethylene glycol (PEG).
17
2.1.3 Hybrid particles
To design an optimal drug or gene delivery vehicle, a combination of lipids, polymers, and
proteins, forming different types of nanostructures, could be used. The materials can be
linked together covalently or mixed as a physical mixture. Probably the most well-known
modification of nanoparticles is the steric stabilization of the particle surface with a
hydrophilic polymer, most commonly PEG. PEG is a bio-compatible, water-soluble, and
chemically inert synthetic polymer. It generates a stealth effect on the surface of the
particles thus reducing aggregation, enhancing the stability, and prolonging the circulation
time in the body (Allen et al. 2002). The surface coverage of the particles is determined by
the molecular weight of the polymer as well as the graft density. PEG molecular weight of
2000–5000 has been shown to be the most effective in prolonging circulation half-life of
liposomes (Allen et al. 1991). It has been proposed that if the graft density of PEG2000 on
liposomes is below 5%, it takes the shape of a “mushroom” or a half sphere, and if the
density is high (> 5%), it takes an extended “brush” shape (Allen et al. 2002) (Figure 5).
The inclusion of 3-10% of PEG in liposomes has been shown to prolong their circulation
times (Allen et al. 1991).
The surface of the nanoparticles can be functionalized with targeting ligands, such as
antibodies, antibody fragments, or peptides that are able to bind to the specific cell types.
For the purpose of diagnostics and imaging, different kinds of markers such as radio
ligands or fluorescent markers can be used (Figure 5). Membrane-active, cationic peptides
on the surface of the particles have proven to enhance intracellular delivery (Kale,
Torchilin 2007, Ye et al. 2010). When the targeting ligands or membrane active peptides
are added to the surface of PEGylated particles, these ligands are usually coupled to the
end of the PEG chains rather than straight onto the particle surface (Hansen et al. 1995).
This minimizes the interference of the PEG shield thus enabling the interaction between
the ligand and the target cell or antigen (Shiokawa et al. 2005). The correct orientation of
the targeting moieties is important in order to achieve efficient interaction with the
receptors.
Figure 5 Functionalization of a liposome. Steric stabilization with PEG2000 at < 5 mol %
results in “a mushroom” shape of the PEG molecules (a), if the graft density is > 5 mol % PEG
takes“a brush” shape (b). Targeting antibody (c) and cell penetrating peptide (d) coupled to the
distal end of a PEG chain. Hydrophilic drugs or imaging agents can be encapsulated in the core
of the liposome (e) and lipophilic ones into the liposome bilayer (f).
18
2.2 Challenges in efficient drug and gene delivery
Despite the great potential of drug and gene carriers, they face multiple challenges on their
way from the vial to the site of action (Figure 6). The vector has to remain stable during
storage, and also in physiological conditions. It should not be cleared too fast from the
blood circulation or cause immunological reactions. Intravenously administered carrier
must pass from the vasculature to the target tissue, bind, and internalize to the target cells.
The drug or gene should be released from the carrier at the target site and, in certain cases,
be imported into the nucleus (Mastrobattista, Koning & Storm 1999, Wang, Upponi &
Torchilin 2011). These issues should be taken into account during the development of new
nanocarrier systems.
Figure 6 Critical steps for efficient drug and gene delivery. Storage stability (1), stability and
half-life in blood circulation (2), extravasation from the blood stream (3), specific binding and
internalization into the target cells (4), escape from the endosomes and intracellular trafficking
(5), and nuclear localization (6).
2.2.1 Stability of the vector
In order to have a good nanoparticle formulation for clinical use, it should first be stable
during storage. Uncoated particles are prone to aggregation, which results from many
factors, such as ionic strength and pH of the solution, the initial size distribution of the
particles and storage temperature (Lee, Mount & Ayazi Shamlou 2001). Charge-neutral
complexes or complexes formed at low n/p ratios tend to aggregate because of
hydrophobic interactions or van der Waals forces. Whereas higher surface charge reduces
aggregation because of electrostatic repulsion (Tros de Ilarduya, Sun & Düzgüneş 2010).
In addition to physical stability, chemical stability also has to be taken into account; for
example, the lipids may be hydrolysed, resulting in lysolipids, and especially unsaturated
lipids can be oxidized easily. Hydrolysis and oxidation finally lead to degradation of
lipidic carriers. To enhance chemical stability, antioxidants can be added to the
19
preparation, or liposomes can be stored as lyophilized powders, but the size distribution,
morphology, and entrapped cargo must be examined after reconstitution (Ulrich 2002).
Reserving the stability is even more difficult in physiological conditions. In blood
circulation there are serum proteins, mainly albumin, but also lipoproteins (high- and low
density lipoproteins, HDL and LDL) and many other proteins which may interact with
polymeric and lipidic particles. They can alter the complex diameter and zeta potential,
especially in the case of cationic complexes, and lead to premature release of encapsulated
material (Zelphati et al. 1998). Extracellular space and cell surface contain negatively
charged glycosaminoglycans (GAGs), components of connective tissues that are often
covalently linked to protein in the form of proteoglycans. Sulphated GAGs, such as
chondroitin sulphate and heparan sulphate are able to block the transfection of cationic
polyplexes and lipoplexes (Ruponen, Ylä-Herttuala & Urtti 1999, Ruponen et al. 2004).
To prevent aggregation and premature disruption in vivo, the carrier system should be
neutral in charge, and the particle size and shape should be optimal (Tao et al. 2011).
Interestingly, polymeric nanoparticles, having shapes like long cylinders (Geng et al.
2007) or elliptical discs (Muro et al. 2008), have demonstrated longer blood circulation
times than their spherical counterparts. Most of the current nanocarriers, however, are
cylindrical in shape, probably because of ease of manufacture. PEG-shield on the surface
of the particles can mask the possible charges and form a hydrated steric barrier against
aggregation (Tirosh et al. 1998, Erbacher et al. 1999).
2.2.2 Tissue distribution and elimination
To be able to find their targets in the body, therapeutic particles should remain long
enough in the blood circulation. Still, only a small fraction of the dose can reach the
tumor. In mice, 24 h from intravenous injection of PEGylated liposomes, roughly, only
0.5–5% of the injected dose has internalized the tumor xenograft, 10–20% still remains in
the blood circulation, 10–20% is up taken by the liver, and 2–5% is up taken by the spleen
(Chang et al. 2007, Chow et al. 2009, Lee et al. 2010). The defence mechanisms of the
body react rapidly against foreign material. The mononuclear phagocyte system (MPS)
consists of phagocytic cells in spleen, liver (Kuppfer cells), lungs, and lymph nodes.
Especially cationic or hydrophobic particles can interact with serum proteins, be
opsonised, and removed from the blood circulation by MPS (Allen et al. 1991, Dash et al.
1999). This can be seen as high accumulation of carrier systems in liver, spleen, and lungs.
Aggregation of the complexes may also lead to embolization in the lungs, which is
obviously life threatening (Morille et al. 2008). Opsonization can also activate the
complement system that induces phagocytosis and initiates inflammatory responses
against the foreign particles (Müller-Eberhard 1988). Vauthier et al. (2011) reported
binding of bovine serum albumin (BSA), fibrinogen, and a complement activating protein,
C3, on the nanoparticles consisting of poly(isobutylcyanoacrylate)-dextran co-polymers.
Adsorption of BSA on the nanoparticles following C3 protein binding activated the
complement cascade, while fibrinogen induced aggregation of the particles.
There are several approaches to reduce MPS recognition of the particles after i.v.
administration, referred in Harasym, Bally & Tardi (1998). The first method is to modify
20
the particle surface properties, e.g. by incorporation of hydrophilic polymers and size
(favourably 50–150 nm) to inhibit the recognition of the immune system. In the second
approach, the phagocytic cells are saturated by increasing the dose of the particles beyond
the level that is required for therapy (>100 mg lipid/kg in murine studies). Certain drugs,
such as liposomal doxorubicin, can also be used to block the phagocytic system. From
these methods, the first option is more relevant because of its safety. Coating the liposome
surface with inert, hydrophilic polymers (e.g. PEG) provides steric stabilization against
interactions with opsonic factors. PEGylated liposomes show longer circulation times and
reduced uptake by MPS compared to conventional, non-PEGylated ones (Allen et al.
1991, Lu et al. 2004). PEG forms a highly hydrated shield around the liposome that has
been thought to sterically inhibit both electrostatic and hydrophobic interactions with the
serum proteins (Mastrobattista, Koning & Storm 1999, Ogris et al. 1999). However, there
is increasing evidence suggesting that PEG does not inhibit plasma protein binding on the
liposome surface (Moghimi, Szebeni 2003, Dos Santos et al. 2007). In fact, PEG may
even enhance complement activation via binding of immunoglobulin M (IgM) and IgG
(Moghimi, Szebeni 2003). Instead, the mechanism for prolonged circulation time provided
by PEG could be prevention of aggregation of the liposomes (Dos Santos et al. 2007).
Attachment of targeting ligands to the nanoparticle surface may lower the stability
and alter the pharmacokinetics of the carrier. Especially antibody-coupled carriers are
rapidly recognized by the immune system and cleared from the blood circulation. The
higher the antibody density on the particles, the faster the clearance (Aragnol, Leserman
1986). Harding et al.´s (1997) pharmacokinetic study with repeated injections of antibody-
coupled liposomes showed even more rapid clearance after second and third injection
compared to initial administration, evidencing immunogenicity of the formulation.
Interestingly, they also demonstrated that antibodies coupled to liposomes are more
immunogenic than free antibodies. Using smaller antibody fragments (Fab´, scFv) instead
of the whole antibody molecule, the half-life can be prolonged to almost the same level
with PEGylated, non-targeted liposomes (Maruyama et al. 1997, Pastorino et al. 2003b). A
prolonged circulation time is prerequisite for efficient accumulation of the particles to the
target site. The mechanism of the liposomal accumulation from blood circulation into the
tumors is discussed in section 2.3.1.
2.2.3 Cellular uptake
When the nanocarrier reaches the target tissue, for example tumor, it is facing the next
barrier, the cell membrane. There are two main routes for cellular uptake: endocytic and
non-endocytic. Endocytic cell uptake can occur via several pathways: clathrin-mediated
endocytosis, caveolae-mediated endocytosis, macropinocytosis, or phagocytosis
(Hillaireau, Couvreur 2009). Fusion and penetration of the nanocarriers through the
cellular membrane are examples of non-endocytic pathways (Xiang et al. 2012).
It has been shown that endocytosis (Figure 7) is the predominant route for
internalization of polymeric and lipidic nanoparticles (Wang, Upponi & Torchilin 2011).
Cationic particles trigger endocytosis by interacting non-specifically with the negatively
charged cell surface via cell membrane associated proteoglycans (Mislick, Baldeschwieler
21
1996, Mounkes et al. 1998) or other anionic components on the cellular membrane. On the
other hand, it has been shown that cell-surface glycosaminoglycans can also inhibit
cellular uptake and gene expression of cationic DNA complexes (Ruponen et al. 2004).
Figure 7 Cellular uptake mechanisms of drug-loaded nanoparticles. Non-specific adsorption
and internalization via endocytosis (a), target-specific binding followed by receptor-mediated
endocytosis (b). If drug carrier binds to non-internalizing receptor, drug can be released outside
of the cell (c). Lipidic carriers can fuse with the cell membrane (d) or exchange the lipid
components with the cell membrane (e), leading to drug release inside the cell. Modified from
Torchilin (2005).
Receptor-mediated endocytosis can take place after specific binding of the targeting
ligand to its receptor. This is a very important mechanism by which the cells take up
nutrients and regulatory proteins, and it is nowadays also utilized in drug and gene
delivery. Binding to the receptor does not automatically mean rapid internalization into the
cell. In the case that the drug carrier is targeted to a non-internalizing receptor, the carrier
should release the drug outside the cell after which the free drug is taken up by the host
cell and also by the neighbouring cells (Figure 7). This kind of “bystander effect” might be
preferable in solid tumors where diffusion of large carrier systems is limited or all of the
cancer cells do not express the targeted antigens (Mastrobattista, Koning & Storm 1999,
Sapra, Allen 2003). However, liposomal drug carriers endocytosed via receptor binding
have been shown to have enhanced antitumoral efficacy over the carriers bound to non-
internalized receptors (Chuang et al. 2010).
Non-endocytic pathways are preferable for non-viral gene delivery because the
destructive effect of lysosomes is then usually avoided (Morille et al. 2008). To enhance
the internalization of drug and gene carriers, cationic membrane active peptides can be
coupled to the particle surface. These cell penetrating peptides (CPP), for example trans-
acting activator of transcription (TAT) peptide from HIV-1, can mediate intracellular
22
delivery via endocytic, and according to some studies, also via non-endocytic pathways
(Torchilin et al. 2001, Bolhassani 2011). Nonetheless, the non-endocytic mechanism of
cellular penetration of CPPs can be considered to be ambiguous. In the study of Subrizi et
al. (2012), only the endocytic uptake mechanism for CPPs could be seen.
Lipidic carriers are able to fuse with the cellular membrane and directly release the
contents to the cytoplasm before entering the endocytic pathways. However, the main
uptake route for lipoplexes is the endocytic pathway, whilst fusion plays an important role
in releasing the DNA in endosomes (Hafez, Maurer & Cullis 2001, Xiang et al. 2012).
2.2.4 Intracellular distribution and cargo release
Following internalization via endocytic pathway, endosome capture and subsequent
lysosomal degradation are the major obstacles to efficient gene delivery. To be effective,
the vector, or at least its contents must be released from the endosome before its
maturation into the lysosome. DNA and RNA degrade easily in the lysosomal
compartments by hydrolytic enzymes. The endosomal release should happen rather fast
since after endocytosis, the endosomal vesicles mature into lysosomes in 10–20 min
(Simões et al. 2004).
For polymer-based vectors, two possible escape mechanisms have been proposed. The
first one, physical disruption of the negatively charged endosomal membrane via
interaction with cationic polymers has been suggested by Zhang, Smith (2000). They
noticed that high generation poly(amidoamine) (PAMAM) dendrimers were much more
effective than PLL in inducing lipid mixing and leakage of the contents. This escape
mechanism seems to depend also on the composition of the cellular membrane (cell type).
The other, better known mechanism, “proton sponge” effect can be applied by PEI,
PDMAEMA and PAMAM which contain protonable secondary and tertiary amines
having pKa-value of 5–7, near to endosomal pH (Boussif et al. 1995). The proton-sponge
hypothesis is based on high buffering capacity of the polymers. Increase in the endosomal
pH causes transportation of protons into the endosome that then results in an influx of
counter ions (Cl-). This promotes osmotic swelling and finally rupture of the endosomal
membrane (Boussif et al. 1995, Sonawane, Szoka & Verkman 2003).
Cationic lipid-based carriers are able to destabilize the anionic endosomal membrane
via electrostatic interactions. So called flip-flop-mechanism has been described by Xu,
Szoka (1996) and Zelphati, Szoka (1996). After endocytosis, the cationic complex
destabilizes endosomal membrane resulting in flip-flop of anionic lipids. The anionic
lipids diffuse into the complex, forming a charge neutral ion pair with cationic lipids. As a
consequence, entrapped nucleic acids dissociate from the complex and are released into
the cytoplasm. Destabilization of the endosomal membrane can also occur after lipid
phase transition. Helper lipid, DOPE, as discussed earlier, is able to acquire an inverted
hexagonal phase (HII) which is unstable and rapidly fuses and releases DNA or drug upon
adhering to endosomal vesicles (Koltover et al. 1998, Mönkkönen, Urtti 1998).
To release the DNA or the drug in a controlled manner at the desired target site,
delivery systems that are sensitive to a certain signal have been developed. The pH inside
the endosomes is 5–6, which is more acidic compared to its environment. The low pH can
23
trigger the release of cargo from pH-sensitive carriers. A combination of
DOPE/cholesteryl hemisuccinate (CHEMS) is widely used in pH-sensitive formulations
(Kirchmeier et al. 2001, Simões et al. 2001, Shi et al. 2002b, Fattal, Couvreur & Dubernet
2004). In an acidic environment, anionic CHEMS becomes protonated, and this neutral
form induces formation of fusogenic hexagonal phase with DOPE. Whereas, at neutral or
alkaline pH, CHEMS stabilizes DOPE into a more stable lamellar phase (Hafez, Cullis
2000).
In addition to pH, increased temperature and enzymatic activity have been utilized to
trigger drug release. In tumors, the temperature is slightly higher than in healthy tissues,
but in practice the temperature difference is so small that it makes the controlled release
challenging. Drug release from thermosensitive carriers can be triggered by using
localized external heating. Kullberg, Mann & Owens (2009) used external heating up to
42 °C to trigger calcein release from temperature-sensitive 1,2-dipalmitoyl-sn-glycero-3-
phosphocholine (DPPC)-based immunoliposomes. The drug release is based on a sharp
gel to liquid crystalline phase transition of thermosensitive lipids at a certain temperature.
Paasonen et al. (2007) created a liposomal system where the contents of the liposomes
were released by UV-light induced heating of the gold nanoparticles incorporated into the
liposome bilayer. Local heating of the gold nanoparticles resulted in leakage of
thermosensitive liposomes. Enzymatically active carriers, for example human serum
albumin (HSA) nanoparticles, have shown degradation and drug release in the presence of
physically existing enzymes: trypsin, proteinase K, protease, pepsin, and intracellular
enzyme cathepsin B (Langer et al. 2008).
2.2.5 Diffusion in cytoplasm and nuclear import
Some drugs can act in the cytoplasm, but others, such as plasmid DNA and many
cytostatic drugs, must enter the nucleus to reach the site of action (Table 1). After escape
from the endosomes, the drug then faces the challenges of intracellular trafficking and
nuclear localization. For large molecular weight DNA in particular, these are difficult
barriers to overcome. Mobility of DNA in cytoplasm is slow because of the tight network
of cytoskeletal filaments, the presence of cell organelles, and high protein concentration
(Lechardeur, Verkman & Lukacs 2005). The diffusion rate of DNA depends strongly on
the size of the molecule. Plasmid DNA, containing 1 000–10 000 base pairs (bp), diffuses
much slower than DNA or RNA molecules under 250 bp (Dauty, Verkman 2005). Slow
diffusion makes DNA an easy target for cytoplasmic nucleases (Lechardeur, Verkman &
Lukacs 2005).
24
Table 1 Examples of target sites inside the cell of different therapeutic agents.
If DNA remains complexed in the cytosol, the resistance against nucleases can be
increased (Lechardeur et al. 1999). Pollard et al. (1998) showed that 1% of cytoplasmic
DNA/PEI complexes entered the nucleus, which was 10 times more than uptake of free
plasmid DNA. Since passive diffusion into the nucleus via nuclear pore complexes is
limited for particles less than 10 nm in diameter, the nuclear entry of plasmid DNA can
occur mainly during cell division when the nuclear envelope is reformed (Görlich, Mattaj
1996). Toropainen et al. (2007) demonstrated substantially higher transgene expression in
dividing human corneal epithelial (HCE) cells compared to differentiated HCE cells after
transfection with PEI/DNA and DOTAP/DOPE/DNA complexes. More specifically,
higher nuclear accumulation has been seen in the cells which are close to mitosis phase
compared to the cells in post-mitotic phase (Gap 1) (Männistö et al. 2007). Männistö et al.
(2007) also demonstrated that despite the high amount of imported transgene in the
nucleus, only 10-6
to 10-4
parts were totally released from the carrier and thus available for
transcription. Release of DNA from the carrier is thus a critical step since premature
disassembly can lead to DNA degradation while incomplete release impairs gene
expression.
Therapeutic agent Site of action in the cell
Genetic drugs
plasmid DNA nucleus
oligonucleotides (siRNA, miRNA) mRNA, cytoplasm/nucleus
Small molecular anticancer drugs
doxorubicin DNA, nucleus
paclitaxel microtubules, cytoplasm
camptothecin DNA enzyme topoisomerase I, nucleus
25
2.3 Targeted cancer therapy
Anti-cancer drugs are usually toxic for the cell, which is why they are undesirable in
healthy tissues. With targeted nanocarrier systems, the drug concentration could be
increased at the tumor site (ElBayoumi, Torchilin 2009) and thus the spreading of the
harmful drug to the normal tissues may be reduced. Targeting also improves
internalization of the drug into cancer cells (Mamot et al. 2003, Dubey et al. 2004). Tumor
targeting can be divided into two types; passive and active (Figure 8).
Figure 8 Passive targeting and active targeting of nanoparticles. A. Passive targeting is
based on leaky vasculature of the tumor and long circulation time of nanocarriers. B. In active
targeting, nanocarriers can be targeted to the receptors overexpressed either on tumor cells (1) or
on angiogenic endothelial cells (2). Modified from Danhier, Feron & Preat (2010).
2.3.1 Passive targeting
When tumor volume reaches 1–2 mm3, it starts to form new blood vessels in order to bring
oxygen and nutrients to the growing cells (Feron 2004). This blood vessel formation is
called angiogenesis. The morphology of tumor vasculature differs from the normal
vessels. The tumor vessel endothelium is malformed and leaky; having 100–600 nm gaps
between the endothelial cells, whereas normal endothelial cells form a continuous,
uniform monolayer (Yuan et al. 1995, Hashizume et al. 2000). Pericytes, the cells
surrounding the endothelial cells, are also malformed in angiogenic tumor vessels
(Morikawa et al. 2002). Thus, small 50–200 nm particles can enter the tumor. Moreover,
due to a non-functional, or absent, lymphatic drainage system, nanoparticles can be also
retained in the tumor interstitium. This phenomenon is called the “Enhanced permeability
and retention” (EPR) effect and is utilized in passive targeting of nanoparticles into tumor
(Maruyama 2011). Because of the EPR effect, it is possible to achieve even 10–50 fold
local concentrations of nanoparticles in tumor compared to normal tissues (Iyer et al.
2006).
The properties of the nanocarriers can influence on the EPR effect. The carriers should
have a long half-life in blood in order to have enough time for efficient tumor
26
accumulation. Optimally, the size of the particle should be more than 10 nm to avoid
filtration through the kidneys, but under 100 nm to avoid a capture by the liver (Danhier,
Feron & Preat 2010). In addition, they should be sterically stabilized to avoid aggregation
and rapid recognition by the MPS system.
The tumor environment also provides some barriers for successful therapy, such as
heterogenous blood flow, increased interstitial fluid pressure, and large transport distances
in the tumor interstitium (Jain 1990, Harrington et al. 2000). Harrington et al. (2000)
demonstrated the influence of tumor size on the uptake of PEGylated liposomes. In large
tumors, the uptake of liposomes is probably reduced due to higher osmotic pressure and
because of a relatively low vascular volume, reflecting areas of poor perfusion or even
necrotic areas.
2.3.2 Active targeting
To mediate active tumor targeting, cancer cell specific targeting ligands are attached to the
surface of the nanocarrier. The chosen ligand actively binds to the receptors that are either
selectively expressed or overexpressed in cancerous cells compared to normal cells (Sapra,
Allen 2003). The most commonly used ligands for liposome and nanoparticle targeting
are: monoclonal antibodies (mAb) (ElBayoumi, Torchilin 2009), fragments of the
antibodies (Fab or svFc) (Pastorino et al. 2003b, Iyer et al. 2011), growth factors (Lee et
al. 2010), peptides (Moreira et al. 2001, Temming et al. 2005, Xiong et al. 2005), small
molecule ligands (such as folate and transferrin) (Gabizon et al. 1999, Voinea et al. 2002,
Riviere et al. 2011), sugars (such as galactosamine, lactose, and trivalent galactose) (David
et al. 2004), and aptamers (Tong et al. 2010) (Table 2).
To attach the ligands on the sterically stabilized liposomes, the ligands are preferably
coupled to the termini of the PEG chains. When the ligands are attached on the bilayer of
the liposome, the PEG may serve as a steric hindrance for both ligand coupling and later
on for binding to the receptors, especially in the case of small molecular weight ligands
(Sapra, Allen 2003). The end group of the PEG-spacer can be functionalized for the
chemical ligand coupling, for example with maleimide (Kirpotin et al. 1997) or N-(3´-
(pyridyldithio)propionoylamino (PDP) (Allen et al. 1995) for the thiol-containing ligands,
or with biotin for avidin-coupled ligands (Loughrey, Bally & Cullis 1987). The amount of
targeting ligands attached on the liposomes is crucial since excessive ligand density leads
to rapid clearance of the liposomes, while insufficient ligand density fails to facilitate
satisfactory targeting efficiency. Only 10–20 molecules of whole targeting antibody or
Fab´ fragments per liposome are required for sufficient internalization to the target cell
(Park et al. 1997, Iden, Allen 2001). For antibody density in excess of 35
molecules/liposome, an increased rate of clearance has been reported (Allen et al. 1995).
In the case of small peptides, even 200–500 peptide molecules/liposome did not cause
highly elevated blood clearance compared to PEGylated, non-targeted, liposomes
(Zalipsky et al. 1995).
Active tumor targeting could be achieved by direct targeting, where the targeting
ligands are coupled straight on the drug carrier, or by a pre-targeting (multistep) approach.
In the pre-targeting method, ligands are not covalently linked to the carrier system;
27
instead, the target-specific ligand is administered as a first step. Once ligand has bound to
the target receptor, the ligand-binding drug-containing nanoparticles are administered.
Pre-targeting is commonly based either on biotin-avidin binding, that shows extremely
high affinity of Kd ~ 1015
M-1
, (Weber et al. 1989, Lesch et al. 2010) or on bispecific
antibodies (Sharkey et al. 2003). Pre-targeting has been utilized in targeting of polymeric
nanoparticles (Nobs et al. 2006, Pulkkinen et al. 2008) and liposomes (Xiao et al. 2002,
Pan et al. 2008).
2.3.2.1 Cancer cell targeting in solid tumors
To reach the cancer cells throughout the solid tumor, nanocarriers should extravasate from
blood circulation to tumor interstitial space and diffuse evenly around. Moreover, the
target receptors should be expressed homogenously on all targeted cells. The most
targeted receptors in solid tumors are: 1) human epidermal growth factor receptors (EGFR
and HER-2), overexpressed in many tumor types, e.g. in breast, colon, ovarian, pancreatic,
head and neck cancer, and non-small cell lung cancer; 2) transferrin receptor, which
participates in iron transfer, expressed in tumor cells 100-fold more than in normal cells;
and 3) folate receptor, which takes care of folic acid intake and is also overexpressed in
many human cancer types (Danhier, Feron & Preat 2010).
Even though a high affinity to the target receptor is desirable, it can also limit the
tumor penetration properties of the nanocarriers. Adams et al. (2001) showed that with
low affinity (Kd = 3.2 x 10-7
M), anti HER2/neu scFv exhibited broad diffusion from the
vasculature to the tumor, whereas the high affinity scFv (Kd = 1.5 x 10-11
M) failed to
traverse more than 2-3 cell diameters. This study was done with labelled scFv, (molecular
weight 27 kDa) without nanocarrier. After coupling these high affinity ligands onto
nanoparticles or liposomes (molecular weight of millions), even more restricted spreading
to the tumor would be expected. Interestingly, Sugahara et al. (2010) showed that co-
administration of non-conjugated tumor-penetrating peptide (iRGD) improved tumor
tissue penetration and therapeutic efficacy of free doxorubicin, nanoparticles
(Abraxane®), doxorubicin liposomes, and antibody trastuzumab. The mechanism of tumor
penetration for iRGD is distinct from the passive EPR-effect, since it is receptor-mediated
and energy-dependent.
2.3.2.2 Targeting to the tumor vasculature
Tumor vasculature targeting aims to obstruct the blood supply of the tumor. That leads to
a lack of nutrients and oxygen, in turn causing tumor cell starvation and death. When
compared to tumor cell targeting, vascular targeting has some advantages: 1) direct
accessibility to the endothelial cells from blood circulation avoiding the problems related
to poor extravasation and tumor tissue penetration; 2) high efficacy, since one tumor
capillary supplies hundreds of tumor cells; 3) avoidance of drug resistance, because
endothelial cells are genetically stable compared to tumor cells that may become resistant
28
to the therapy; 4) broad applicability, since most solid tumors are dependent on
neovascularization (Mastrobattista, Koning & Storm 1999, Feron 2004).
Besides during tumor growth, angiogenic vessels are formed also in some non-
malignant conditions such as in atherosclerosis, wound healing, psoriasis, and in certain
eye diseases, e.g. in the wet form of age-related macular degeneration and
neovascularisation of the cornea. Angiogenic vessels express markers such as vascular
endothelial growth factor receptor (VEGFR) and integrins (αvβ3 and αvβ5) that are not
present in the resting blood vessels of normal tissues (Ruoslahti 2002). The integrins are
also upregulated in different tumor cells, including metastic melanoma cells (Conforti et
al. 1992, Seftor, Seftor & Hendrix 1999). Integrins can be specifically recognized by
RGD-peptide, consisting of arginine, glycine, and aspartic acid. RGD-peptide was found
by screening of phage display peptide libraries (Pasqualini, Koivunen & Ruoslahti 1997)
and it is one of the most studied tumor vasculature homing peptides.
29
Table 2 Nanocarrier-based targeted therapeutics in clinical development and examples of
preclinical studies.
Targeting ligand Formulation Target Study
phase
Reference
Clinical trials
stomach cancer
specific GAH mAb
PEGylated liposomal
doxorubicin (MCC-
465)
tumor antigen in
stomach cancer
I Reviewed in
(Cheng et al.
2012)
anti-transferrin
receptor scFv
liposomal p53 plasmid
DNA (SGT-53)
transferrin receptor I
human transferrin liposomal oxaliplatin
MBP-426)
transferrin receptor I/II
human transferrin siRNA loaded
nanoparticles
(CALAA-01)
transferrin receptor I
peptide docetaxel-loaded
polymeric
nanoparticles (BIND-
014)
prostate specific
antigen
I
Preclinical studies
cancer cell specific
mAb 2C5
PEGylated liposomal
doxorubicin
cancer cell surface bound
nucleosomes
(ElBayoumi,
Torchilin 2009)
Fab´ fragment of
cetuximab
PEGylated liposomal
doxorubin/vinorelbin
epidermal growth factor
receptor
(Mamot et al.
2005)
mesothelioma
targeting scFv
(M1)
PEGylated 111
In-
labeled liposomes
surface antigens on human
mesothelioma tumor cells
(Iyer et al. 2011)
epidermal growth
factor
111In-labeled
polymeric micelles
epidermal growth factor
receptor
(Lee et al. 2010)
A10 aptamer polymer-paclitaxel
conjugates
prostate-specific membrane
antigen
(Tong et al.
2010)
folate PEGylated liposomal
doxorubicin
folate receptor (Riviere et al.
2011)
cyclic RGD-
peptide
PEGylated liposomal
5-fluorouracil
αvβ3 integrins (Dubey et al.
2004)
NGR-peptide PEGylated liposomal
doxorubicin
angiogenic endothelial cell
marker aminopeptidase N
(Pastorino et al.
2003a)
Fab´ fragment of
anti-VEGFR-2
PEGylated liposomal
doxorubicin
vascular endothelial growth
factor receptor
(Roth et al.
2007)
mAb anti-
disialoganglioside
and NGR-peptide
PEGylated liposomal
doxorubicin
disialoganglioside receptor and
aminopeptidase N
(Pastorino et al.
2006)
30
3 Aims of the study
The general objective of this study was to develop and evaluate lipid and polymer based
nanocarriers for targeted drug and gene delivery using in vitro, in vivo, and in silico
methods. The specific aims were:
1. To investigate the effects of the architecture and flexibility of cationic amphiphilic
star and linear PDMAEMA-based block copolymers on DNA complex formation,
in vitro transfection efficiency, and cytotoxicity.
2. To determine the DNA binding ability, in vitro transfection efficiency, and
cytotoxicity of novel BSA- and hydrophobin (HFBI)-dendron conjugates.
3. To develop an extracellularly stable gene delivery vector that can release its
contents at the acidic endosomal pH.
4. To investigate a targeted liposomal drug delivery system when novel activated
endothelium targeted peptide (AETP) is used as a targeting ligand.
5. To explore an epidermal growth factor receptor (EGFR) targeted liposomes using
direct targeting and pre-targeting approaches.
31
4 Overview of the methods
The materials and prepared formulations for gene delivery (I – III) and for targeted
liposomal drug delivery (IV, V) are summarized in Table 3. The cell lines used in the
original publications are shown in Table 4. General methods of physicochemical and
biological studies are shown in Tables 5 and 6, respectively. In addition, a computational
study combining molecular dynamics simulation and ligand-protein docking was
performed (IV). Materials and methods are described in detail in the publications.
Table 3 Summary of materials and prepared formulations in the publications.
Material/Formulation Source/Preparation method Publication
plasmid DNA (pCMVβ) encoding β-
galactosidase
amplification in E.coli, purification by
column separation
I - III
Polymeric materials
PDMAEMA block copolymers:
(PS-PDMAEMA)6,
(PBuA-PDMAEMA)6,
PDMAEMA-PS-PDMAEMA,
PDMAEMA-PBuA-PDMAEMA
synthesized, purified and characterized
in the Department of Chemistry,
Laboratory of Polymer Chemistry,
University of Helsinki
I
protein-polyamine dendron
conjugates:
HFBI, HFBI-G1, HFBI-G2,
BSA, BSA-G1, BSA-G2
synthesized, purified and characterized
in the Department of Engineering,
Physics, and Mathematics, and Center
for New Materials, Helsinki University
of Technology
II
other polymers:
PEI 25K, PLL
commercially available I - III
Lipidic materials
DOPE, CHEMS, Egg PC, Egg SM,
DMPG, cholesterol
commercially available III
HSPC, cholesterol, DSPE-PEG2000,
DSPE-PEG2000-maleimide,
DSPE-PEG2000-biotin
commercially available IV, V
Targeting ligands
AETP phage display and peptide synthesis
using Fmoc chemistry in Karyon Ltd.,
Finland
IV
cetuximab commercially available V
Formulations
cationic polyplexes self-assembling I - III
lipid-coated polyplexes detergent removal III
calcein containing liposomes reverse-phase evaporation III
PEGylated liposomes extrusion IV, V
32
AETP-targeted liposomes extrusion + post-insertion method of
AETP
IV
cetuximab-targeted liposomes extrusion + biotin-avidin-biotin binding
of cetuximab
V
doxorubicin- liposomes extrusion + remote-loading of
doxorubicin
IV, V
CMV = cytomegalovirus, PDAMEMA = poly(2-(dimethylamino)ethyl methacrylate), PS = polystyrene , PBuA = poly(n-butyl acrylate), HFBI = hydrophobin, BSA = bovine serum albumin, G1 = first generation, G2 = second generation, PEI = poly(ethylene imine), PLL = poly-L-lysine, DOPE = 1,2-dioleyl-sn-glycerol-3-phosphoethanolamine, CHEMS = cholesteryl hemisuccinate, Egg PC = egg phosphatidyl choline, Egg SM = egg sphingomyelin, DMPG = 1,2-dimyristoyl-sn-glycero-3-phospho-(1'-rac-glycerol), HSPC = fully hydrogenated phosphatidyl choline, DSPE-PEG2000 = 1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-[amino(polyethylene glycol)-2000], AETP = activated endothelium targeting peptide, Fmoc = 9-fluorenylmethyloxycarbonyl
Table 4 The cells used in the publications.
Table 5 Physicochemical characterization methods used in the publications.
Study objective Method Probe Publication
particle size dynamic light scattering - I, III - V
DNA-complex integrity ethidium bromide intercalation
assay
ethidium
bromide
I-III
pH-sensitivity calcein-dequenching assay calcein III
phospholipid (liposome)
concentration
fluorescence fluorescein-PE IV, V
liposomal drug
concentration
absorbance doxorubicin IV, V
ligand coupling efficiency fluorescence tryptophan IV, V
Cell line Type Species Publication
ARPE-19 retinal pigment epithelial cell line human I
C8161 melanoma cell line human IV
CV1 kidney fibroblast cell line monkey I-III, V
primary
HUVEC
umbilical vein endothelial cells human IV
KS1767 kaposi’s sarcoma cell line human IV
SKOV-3 ovarian adenocarcinoma cell line human V
SKOV3.ip1 ovarian adenocarcinoma cell line human V
SVEC4-10 lymph node endothelial cell line mouse IV
33
Table 6 Biological methods used in the publications.
Study objective Method Probe Publication
in vitro studies
cytotoxicity/therapeutic activity
of DNA-complexes and
liposomes
MTT-assay formazan II - IV
Alamar Blue®-assay resorufin I, V
transfection efficacy ONPG-assay ONPG I - III
cellular affinity/uptake of the
formulations
FACS-analysis EMA-DNA III
fluorescein IV - V
in vivo studies
pharmacokinetics: half-life of
liposomes
TRF Europium IV
biodistribution of liposomes TRF Europium IV
confocal microscopy Rho/FITC IV
LC-MS analysis doxorubicin V
SPECT-CT and
gammacounting
99mTechnetium V
MTT = 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide, ONPG = ortho-nitrophenyl-β-D-galactopyranoside, FACS = fluorescence-activated cell sorting, EMA-DNA = ethidiummonoazide, TRF = time-resolved fluorescence, Rho = rhodamin, FITC = fluorescein isothiocyanate, LC-MS = liquid chromatography - mass spectrometry, SPECT-CT = single-photon emission computed tomography - computed tomography
34
10 Summary of the main results
The main results of the experimental data and molecular modeling are presented in
Table 7.
Table 7 Summary of the main results.
Gene delivery vectors Publi-
cation
DNA-binding
properties
All PDMAEMA block co-polymers bound DNA. Only (PS-
PDMAEMA)6 was not able to condense DNA totally.
I
HFBI-G1, HFBI-G2 and BSA-G2 showed the highest binding
capacity towards DNA.
II
LCDCs condensed DNA completely. After treatment with
anionic dextran sulphate, LCDCs prepared by detergent dialysis
method were stable at least for 24 h, while LCDCs prepared by
ethanol injection method had released ~50% of DNA at 24 h
time point.
III
pH-sensitivity LCDCs were stable at neutral pH, but they were able to fuse with
calcein-containing endosome-mimicking liposomes at endosomal pH
(5-6).
III
Transfection
efficacy in
vitro
Linear PDMAEMA-PBuA-PDMAEMA was able to transfect ARPE-
19 cells with efficacy comparable to PEI, and CV-1 cells with
efficacy of 1/10 of that of PEI. The other carriers showed some
transfection efficacy in the following order: PDMAEMA-PS-
PDMAEMA ≈ (PBuA-PDMAEMA)6 > (PS-PDMAEMA)6.
I
Of the protein-dendron conjugates, only HFBI-G2 was able to
transfect CV-1 cells. Transfection efficacy was only 1/20 of that of
PEI.
II
Transfection efficiency of LCDCs was ~80% of that of PEI/DNA
polyplexes in the absence of serum. In the presence 10% serum, the
transgene expression level decreased dramatically.
III
Cytotoxicity
in vitro
All PDMAEMA block co-polymers showed over 80% viability in
ARPE-19 and CV-1 cells at low n/p ratios (0.5–4). With increasing
n/p ratio, the cell viability decreased gradually.
I
All protein-dendron conjugates showed over 80% viability in CV-1
cells at n/p low ratios (0.125–4). HFBI-G1 and HFBI-G2 were
observed to be slightly cytotoxic at higher n/p ratios.
II
Cancer-targeted liposomes
Cellular
affinity in
vitro
AETP did not increase the cellular affinity of liposomes in HUVEC-
cells.
IV
EGFR-targeted liposomes had 22–38 times higher affinity towards
SKOV-3 cells and 13–17 times higher affinity towards CV-1 cells
compared to affinity of non-targeted liposomes. Competition with
V
35
free cetuximab decreased the affinity of the targeted liposomes to the
level of the non-targeted liposomes. Pre-targeting showed lower
cellular association than direct targeting.
Cytotoxic
activity in
vitro
AETP did not increase the cytotoxic efficacy of doxorubicin-loaded
liposomes in HUVEC or SVEC4-10 cells.
IV
EGFR-targeted doxorubicin-loaded liposomes showed slightly higher
cytotoxicity than non-targeted liposomes in SKOV-3 cells. However,
free DXR was significantly more cytotoxic than both of the liposome
types.
V
Pharmacokine
tics in vivo
AETP-targeted liposomes showed similar elimination half-life (~7–
10 h) to non-targeted liposomes (~6–7 h).
IV
Cellular
affinity in
vivo
Significant differences in co-localization of liposomes and
endothelial cells could not be seen between the AETP-targeted and
non-targeted liposomes in confocal microscopy.
IV
Uptake in
tumor
No significant difference in tumoral uptake between AETP-targeted
and non-targeted liposomes could be seen.
IV
After direct targeting or pre-targeting, no significant difference in
uptake in tumor between EGFR-targeted and non-targeted liposomes
could be seen at 24 h time point. However, intraperitoneal
administration of the liposomes, targeted or not, led to faster and
higher tumoral accumulation of the liposomes than intravenous
administration.
V
Orientation of
the targeting
ligands in
silico
Computational modeling revealed that both AETP and RGD peptides
located in the PEG region of the PEGylated liposomes. However,
AETP was more covered by the PEG chains, while RGD was more
exposed to the solvent.
IV
Affinity
towards
serum
albumin in
silico
Protein-ligand docking showed HSA having 12 times stronger
binding affinity to PEG and 14 times stronger affinity to AETP than
to RGD peptide.
IV
36
11 General discussion
11.1 Structure-activity relationship of polymeric DNA carriers on DNA-complex formation, transfection efficacy, and toxicity
To be able to condense the DNA, polymer should have a sufficient density of positive
charges. On the other hand, high cationic charge density is often related to increased
toxicity. Electrostatic interactions with the plasma membrane are known to damage the
cell (Lv et al. 2006). An additional critical step of gene delivery is escape from the
endosomes. Because cationic polymers lack hydrophobic parts, they cannot destabilize the
endosomal membrane. However, some polymers do protonate at endosomal pH (e.g. PEI,
PAMAM or PDMAEMA) and they can act as swelling proton sponges. To modify the
properties of the cationic polymers, they can be synthesized in different lengths, with
different architecture (linear or branched), with different substitutions or additions of
functional groups (Tros de Ilarduya, Sun & Düzgünes 2010). Nonetheless, it seems that
there is no universal rule regarding how the architecture of a polymer effects its properties
as a gene delivery vehicle.
To investigate the effect of polymer composition and architecture on the ability to
condense DNA and mediate transgene expression, two different star block copolymers and
corresponding linear triblock copolymers were synthesized. The hydrophobic core of star-
like polymers was composed of either glassy polystyrene (PS) or rubbery poly(n-
butylacrylate) (PBuA). The outer block consisted of hydrophilic and cationic PDMAEMA.
All of the synthesized polymers were able to form spherical core-shell micelles in aqueous
solutions. The linear polymers and star-like (PBuA-PDMAEMA)6 could condense plasmid
DNA completely, while glassy (PS-PDMAEMA)6 formed loose DNA complexes. The
gene transfection efficiency was highest for linear PDMAEMA-PBuA-PDMAEMA at n/p
ratio 2 and 4, and lowest for star-like (PS-PDMAEMA)6. Cytotoxicity was increased at
charge ratios of n/p 8 and higher, (PS-DMAEMA)6 being least toxic.
The poor transfection efficiency of star-like (PS-PDMAEMA)6 may arise from glassy
polystyrene in the core of the star polymers which stiffens the structure of the polymer and
may restrict the interaction with plasmid DNA. On the contrary, rubbery pBuA core may
not limit the contact of PDMAEMA arms and the DNA. Linear polymer chain is more
flexible making the structure more favorable for DNA condensation and transfection
compared to star-like structures. In addition, the amine groups are more available for DNA
binding in the linear form of the polymer. Linear architecture and rubbery pBuA core
seem to be favorable features for mediating transfection. The positive effect of linear
architecture was also shown in the case of poly-L-lysine polymers (Männistö et al. 2002).
This is the first known investigation into the effects of rubbery and glassy states on gene
transfer.
This study has revealed evidence that polymer architecture and composition effect
gene packing ability and transfection efficiency of the studied block copolymers.
Accordingly, linear PDMAEMA-PBuA-PDMAEMA showed transfection efficiency close
to PEI 25 kDa in a retinal pigment epithelial cell line in vitro. However, similar to all
37
cationic polyplexes, PDMAEMA-based polyplexes are also liable to deactivation by
negatively charged serum proteins and other polyanions in vivo.
In another study, protein-dendron conjugates, consisting of either BSA or hydrophobin
(HFBI) protein and first-generation (G1) or second-generation (G2) polyamine dendrons,
were synthesized. BSA is a rather large (66.4 kDa) biocompatible protein that exhibits
long circulation times, while HFBI is a small (8.7 kDa) surface-active protein that is also
well known for its safety. Since both of the proteins lack DNA binding motifs, cationic
polyamine dendrons were conjugated to them to achieve sufficient DNA binding. All of
the dendron-conjugated proteins showed enhanced DNA binding over the naked proteins.
HFBI-conjugates expressed stronger affinity to DNA than corresponding BSA-conjugates,
especially BSA-G1. This is not surprising because the dendron (~1 kDa) is relatively small
compared to the size of BSA and plasmid DNA.
HFBI-G2 was the only conjugate with enhanced transfection activity, even though
BSA-G2 could bind to DNA with a similar affinity. This could be due to the amphiphilic
nature of HFBI that may lead to a favorable interaction with cell membrane structures.
Despite the successful use of BSA as a drug nanocarrier, it was not able to mediate
transfection when conjugated with dendrons. It has been shown earlier that the small
polyamine dendrons are relatively inefficient in mediating transfection on their own,
without endosomal disrupting agent, chloroquine (Hardy et al. 2006). Nonetheless, in
combination with hydrophobic HFBI, enhanced transgene activity without increased
toxicity could be achieved. The transfection efficiency of HFBI-G2 was, however, fairly
low compared to PEI 25 kDa. It has been shown that by increasing the size of the
PAMAM dendrimers, higher transfection could be achieved (Haensler, Szoka 1993). This
may also be the case for HFBI-dendrons. The endosomal escape of HFBI-dendrons would,
however, most likely persist as a bottleneck because of deficiency of protonable tertiary
amines in the structure of polyamine dendrons.
The concept of protein-dendron conjugates worked in the sense of efficient DNA
condensation, but the transfection efficiency was very modest compared to PEI 25 kDa.
The protein-dendron conjugates are not suitable for gene delivery as such; the concept
should be further developed.
11.2 Lipid-coated DNA-complexes as stable gene delivery vectors
Cationic lipoplexes and polyplexes are able to transfect cells in vitro, but in vivo they are
susceptible to deactivation by several polyanionic proteins and polysaccharides, such as
GAGs. Extracellular GAGs can bind to cationic complexes and are able to alter both
cellular uptake and intracellular behavior thus decreasing the gene transfer (Ruponen et al.
2004). Optimal gene delivery vectors would be stable in extracellular space but are
capable of releasing the cargo inside the cell. Coating of cationic DNA polyplexes with a
stabilizing layer of anionic lipids results in envelope-type particles (Guo, Lee 2000,
Mastrobattista et al. 2001, Nahde et al. 2001, Guo, Gosselin & Lee 2002, Khalil et al.
2007). The stability of such particles against GAGs has not been evaluated, however.
38
In this study, a method to coat the PEI/DNA complexes by anionic and fusogenic lipid
mixture was developed. In this detergent removal coating procedure, the polyplexes are
slowly coated by a mixed micellar solution of DOPE/CHEMS/octyl glucoside. The
resulting mixture is subsequently diluted with buffer that leads to extraction of the
surfactant from the mixed micelles and to the formation of a lipid bilayer. The lipid-coated
DNA complexes (LCDC) showed a good stability against a model GAG, dextran sulphate,
while uncoated polyplexes disintegrated rapidly in the presence of the GAGs. LCDCs
prepared by the detergent removal coating method showed superior stability compared to
the LCDCs prepared by ethanol injection procedure, in which the polyplexes were coated
with preformed liposomes. The addition of mixed micellar solution to the polyplex surface
apparently forms a more uniform coating compared to the addition of preformed
liposomes.
The coated complexes showed pH-sensitivity at low pH (5-6) by fusion and
aggregation. This means that the complexes are able to fuse with endosomes and release
their contents, which is important for efficient gene delivery. Because of the negative
surface charge, the cellular uptake was lower than the uptake of non-coated cationic
PEI/DNA complexes. This may explain why the transfection efficacy also remained lower
compared to PEI/DNA polyplexes. Cellular uptake of LCDCs was improved by
neutralizing the negative charges with PEGylation, but this modification reduced the pH-
sensitivity and lowered the level of transfection. PEG is ideal for preventing aggregation
and thus prolonging the time in blood circulation, but it becomes unnecessary after particle
internalization into the cells. To avoid the obstructing effects of PEG inside the cells,
cleavable, acid labile PEG chains have been generated (Guo, Szoka 2001, Romberg,
Hennink & Storm 2008).
Local DNA delivery, for example in the eye or in some surgical situations, may not
require PEGylation because the particles are not removed by the reticuloendothelial
system in such cases. The anionic lipid coat or PEG coating may be useful however, in
preventing the interactions with local anions, like hyaluronic acid in the vitreous body of
the eye. Kurosaki et al. (2013) showed higher transfection activity in the retina of rabbits
for PEI/DNA polyplexes coated with anionic polymer-coating than for uncoated cationic
PEI/DNA polyplexes. Because of the aggregation, cationic polyplexes were immobilized
in the vitreous and only a small amount could reach the target cells in the posterior eye. To
improve the cellular uptake of LCDCs, cell penetrating agents could be employed. This
kind of envelope-type structure, with a functionalized surface, could be one step towards a
virus-mimicking gene delivery vector.
In conclusion, LCDCs were stable against GAGs, but they were able to fuse with the
endosomal membranes at low pH. The overall cellular uptake of the LCDCs remained
low, however. By enhancing the cellular uptake with targeting peptides, the LCDCs might
reach higher cellular uptake and become suitable for local gene delivery, to locations such
as the eye.
39
11.3 Hindering effect of liposomal PEG on the targeting efficiency of a small hydrophobic peptide, AETP
Steric stabilization of therapeutic nanoparticles with PEG is important for maintaining
their stability and long blood circulation times, but, as discussed earlier, it can hamper
complex internalization and endosomal membrane diffusion (Holland et al. 1996, Shi et al.
2002a). In addition, PEG shielding may also impair the interaction between the targeting
moiety of the particle and the target antigen on the cell surface, as was hypothesized in the
study of activated endothelium targeting peptide (AETP)-targeted liposomes.
AETP was first discovered by phage display screening of activated human umbilical
vein endothelial cells (HUVEC) and then of human Kaposi’s sarcoma xenograft in mice. It
was thus expected to be an efficient targeting moiety towards neovascular endothelium.
Despite successful phage display screening, the AETP was not able to act as a targeting
moiety when conjugated to drugs (Bergman, M., personal communication). This was due
to the very low water solubility of the peptide. Therefore, it was assumed that AETP could
function as a targeting ligand on PEGylated liposomes. However, neither the cellular
affinity nor the cytotoxic efficacy of AETP-targeted doxorubicin-containing liposomes
was enhanced compared to the non-targeted ones. Pharmacokinetic investigation
demonstrated that AETP-targeted liposomes had a similar or longer blood half-life
compared to non-targeted, PEGylated liposomes. This was a positive result, since the
peptide did not disturb the stealth effect of the PEG on the liposomes, but it could also
mean that the peptide was covered with the PEG shield. AETP-targeted liposomes did not
show any enhanced accumulation in the tumor tissue or enhanced specific binding to the
endothelial cells in vivo when compared to non-targeted liposomes.
The interaction between AETP moieties and liposomal PEG was demonstrated by
molecular modeling. A well-known RGD-peptide that has been found to be an effective
targeting ligand was used as a comparison in simulations. As a significantly more
hydrophobic molecule, AETP was found to be covered by PEG while the surface of the
more hydrophilic RGD was more exposed to water. In addition, AETP was also located
deeper in the PEG shield than RGD peptide, rendering it unavailable to bind the receptors.
An obscuring effect of PEG on targeting efficiency of PEG2000-liposomes targeted with
PEG2000-folate has also been observed by Gabizon et al. (1999) and Shiokawa et al.
(2005). When folate was coupled to the end of a clearly longer PEG arm, PEG3500 instead
of PEG2000, clearly extruding from the PEG2000 shield, the cellular association of the
liposomes increased. Recently, Stefanick et al. (2013) revealed the importance of linker
length effect on cellular uptake of peptide-targeted PEGylated liposomes. However, using
longer PEG-spacers may not work in the case of hydrophobic AETP because it tends to
escape from the solvent and interact with the PEG molecules. PEG also has hydrophobic
properties, even though it is described to form a hydrophilic sheath around the liposome.
Hydrophobic interactions between AETP and PEG may explain the unfavorable
orientation of AETP molecules for targeting purposes.
Ligand docking studies showed a high binding affinity of human serum albumin
(HSA) to both AETP and PEG. Although the effect of protein binding on AETP targeting-
ability warrants further studies, molecular modeling shows very strong evidence that PEG
coating hinders target binding of AETP. Protein binding to PEG has been previously
40
revealed by others experimentally (Price, Cornelius & Brash 2001, Dos Santos et al.
2007).
Even though PEG presents many advantages and it is accepted for clinical use, it may
not be suitable for all cases of drug delivery. In the case of having hydrophobic targeting
ligands, PEG-coating on the liposomes could be replaced by more hydrophilic polymer to
avoid the interactions between the ligand and the steric coating. More rigid polymer may
also work, given that it can force the ligand out from the polymer cloud. But whether the
stealth effect is then lost due to a firmer polymer structure should be assessed. Some
possible alternatives for PEG have been suggested by Knop et al. (2010). When
PEGylated nanosystems are used it might be beneficial to choose a hydrophilic targeting
ligand.
In a summary, AETP, a promising targeting candidate, failed to show any targeting
efficiency in these studies. This was most likely due to the interference of liposomal PEG
shield that may have prevented the interaction of the peptide with the target receptors.
Nevertheless, cancer targeting is a complicated task, and many other factors may influence
the efficacy of drug targeting.
11.4 Pre-targeting and local administration of liposomes as potential approaches in tumor targeting
Tumor targeting with immunoliposomes has been studied in numerous institutes over the
past 30 years. Targeting efficiency over the non-targeted liposomes was somewhat
successful, but accelerated clearance from the blood circulation proved problematic,
particularly when whole antibodies were used as targeting moieties (Aragnol, Leserman
1986). Shorter half-life usually compromises the benefit that has been achieved by
targeting. When smaller fragments of antibodies have been used, half-lives comparable to
non-targeted liposomes have been reached (Maruyama et al. 1997). Another option to
prolong the residence time in blood circulation could be pre-targeting technology, since it
has been noticed by Harding et al. (1997) that separate injections of cetuximab and
PEGylated liposomes did not cause immune response, while immunogenicity was
potentiated by antibody-coupled liposomes.
In the current study, both direct targeting and pre-targeting (Figure 9) approaches were
used to target PEGylated liposomes to ovarian adenocarcinoma cells (SKOV-3) in vitro
and intraperitoneal xenografts in mice in vivo. Biotin-neutravidin technology was utilized
in both cases to link the endothelial growth factor receptor (EGFR) antibody (cetuximab)
to the liposomes. Direct targeting of the liposomes to SKOV-3 cells was receptor-specific
and efficient in vitro, whilst the pre-targeting was not as efficient, possibly due to
premature internalization of the cetuximab-receptor-complex. Antibody-neutravidin-
complex should remain available for interaction with biotinylated liposomes. The other
explanation could be the hindrance effect of PEG as discussed earlier. As a small
molecule, biotin could be partly covered by the PEG shield and be hindered from the
interaction with neutravidin that is bound to the receptor.
41
Figure 9 Schematic presention of pre-targeting and direct targeting approaches. In pre-
targeting method (A), antibody-linked neutravidin is administered first (step 1). Once antibody
complex has found its target, biotinylated liposomes are administered (step 2). In direct targeting
approach (B), antibodies are coupled on the surface of the liposomes and the formed
immunoliposomes are administered as a single dose.
In animal studies, accumulation of the targeted liposomes in the tumor was not higher
than that of non-targeted liposomes with either targeting approach. However,
intraperitoneally (i.p.) injected biotinylated liposomes, regardless of targeting,
accumulated faster and reached a higher concentration in tumors compared to
intravenously (i.v.) administered liposomes. This was comparable to the observation of
Lin et al. (2009), who noticed rapid tumor accumulation of PEGylated liposomes after i.p.
injection. I.p. administered liposomes can thus accomplish a fast local effect on
intraperitoneal tumors, followed by systemic drug delivery after passing into the blood
circulation. Because of the tendency of ovarian cancer to spread to the abdominal cavity,
i.p. administration of cytotoxic drugs may be beneficial to reach both the primary tumor
and the metastases.
No significantly enhanced accumulation of targeted liposomes to the tumor was
evident, when compared to non-targeted, PEGylated liposomes, even though higher
uptake was seen in cell culture. This is somewhat expected, because due to the EPR effect,
the uptake in the solid tumor tissue may not be further enhanced by targeting. Without
higher accumulation at the tumor site, enhanced cancer cell-specific internalization
(Kirpotin et al. 2006) and increased therapeutic efficacy (Mamot et al. 2005) for targeted
liposomes over non-targeted ones have been observed.
The pre-targeting method described here still requires optimization: the choice of the
antibody should be reconsidered, but also the timing between the antibody administration
and the liposome injections needs optimization. In principle, a pre-targeting approach has
potential because of separate injections of targeting agents and drug formulation; different
antigens can be targeted simultaneously with the same drug formulation. This flexibility
enables the use of the same nanoformulation in different types of cancers. When antibody
is not chemically attached to the liposome, the stability during storage could be improved,
42
and also the developmental costs may not rise as high as the costs for more complicated
directly targeted formulation.
Pre-targeting proved to be less efficient, compared to direct targeting, regarding uptake
in cancer cells in vitro. Because of fast tumor accumulation in vivo, however, the concept
of local tumor pre-targeting might be beneficial after further development.
43
12 Conclusions
1. Polymer architecture and composition had a clear effect on DNA complexation
and gene transfer efficacy. Linear structure and rubbery poly-n-butyl acrylate
block in the middle of PDMAEMA-chain improved DNA complexation and
transfection efficiency, while star-shaped and glassy polystyrene core hindered
DNA condensation as well as transfection.
2. Hydrophobin and BSA-conjugated polyamino dendrons bound to DNA with high
efficacy and were biocompatible in vitro; but transfection was observed only for
amphiphilic hydrophobin-dendron conjugates (second generation) with a low
transgene expression.
3. Lipid-coated DNA complexes with a cationic PEI/DNA core covered with
negatively charged DOPE/CHEMS mixture were successfully produced by
detergent removal method. These complexes were resistant against extracellular
glycosaminoglycans, and were able fuse with endosomal membrane at acidic pH
and mediate transfection.
4. Molecular modeling revealed that hydrophobic AETP targeting moieties were
located deep in the PEG layer of the liposomes, and thus might have been
prevented from interacting with target receptors. In addition, peptide binding to
serum proteins may further inhibit target binding. These findings may be useful in
the development of targeted nanocarriers.
5. Direct targeting with EGFR-antibody liposomes was superior to non-targeted and
pre-targeted liposomes in the cell studies, but in mice the accumulation in the
tumor remained low. Tumoral accumulation after intraperitoneal administration of
pre-targeted and non-targeted liposomes was faster and greater compared to
intravenous administration. Local tumoral pre-targeting method warrants further
studies as a potential approach in cancer therapy.
44
13 Future prospects
Nanoparticles – drugs of the future?
Nanoparticle technology in pharmaceutical research has provided great promise for more
efficient and safe drug therapy. Targeted nanomedicines in particular have garnered
growing enthusiasm among researchers reflected by a massively increasing number of
publications. Despite decades of developmental work, there are no targeted nanoparticle
formulations in clinical use and only a few are currently in clinical trials. When a drug
formulation becomes more complicated, the risks in the development phase most
definitely increase.
Somewhat surprisingly, most of the nanoparticle systems are developed for the
treatment of cancer. Development of systemically administered targeted nanoparticles
faces many technical challenges as a risk of too short half-life, poor tumoral perfusion, and
diffusion barriers at the binding site. The leakiness of the tumor vessels varies between the
tumor types, in some cases the tumor core may be not well perfused (Chauhan et al. 2011).
In addition, the expression levels of the target receptors can vary between different cancers
(Perez-Soler 2004) and the receptor density may also decrease dramatically during the
treatment (Jiang et al. 2008, Cheng et al. 2012). This brings challenges to nanoparticle
delivery to tumors. Possibly, some other disease states in closed systems, such as in the
eye, or the surgical situations, might be more beneficial delivery targets for nanomedicine
development, because of the possibility of local administration. In this case, the
nanomedicine would provide advantages, such as protection of the drug from enzymatic
degradation, prolonged activity, cell specificity, and optimization of intracellular
distribution. Non-malignant diseases also lack rapid mutation of the target cells which
makes the targeting easier.
The approval process for a nanoformulation is much more difficult than for parent
drug. This likely reflects why big pharmaceutical companies do not want to invest money
and time for a small increase in performance that might be achieved by reformulating a
currently approved drug inside a nanocarrier (Venditto, Szoka Jr. 2013). For example,
liposomal cancer therapeutics Caelyx and Ambisome are able to reduce the toxicity of the
parent drug, but improving the efficacy is still modest. Even if reduced toxicity is very
important for the patients, the minimal improvement in the efficacy may become an issue
for nanotherapeutics (Juliano 2013).
Although nanotherapeutics has not yet realised its promises, it will certainly find its
place in the pharmaceutical field. Increasing knowledge of the strengths and the
weaknesses of nanomedicines will be useful in the developmental phase. In addition,
utilization of computational modeling may be helpful in screening the best nanocandidates
before carrying out the costly preclinical and clinical studies.
45
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