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A nanoliposome-based drug delivery system for cancer By Jia Lin, MSc, MBBS Submitted in fulfilment of the requirements for the degree of Doctor of Philosophy Deakin University March, 2013
Transcript of A nanoliposome-based drug delivery system for cancer › eserv › DU:30061028 › lin-na... · A...
A nanoliposome-based drug delivery system for cancer
By Jia Lin, MSc, MBBS
Submitted in fulfilment of the requirements for the degree of
Doctor of Philosophy
Deakin University
March, 2013
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Table of Contents
List of Figures .............................................................................................................. 1
List of Tables ................................................................................................................ 4
Publications .................................................................................................................. 5
Papers submitted for publication .................................................................................. 7
Acknowledgements ...................................................................................................... 8
Abstract ...................................................................................................................... 10
List of abbreviations ................................................................................................... 13
Chapter 1 - Introduction ............................................................................................. 16
1.1 Cancer and anticancer chemotherapy .............................................................. 17
1.1.1 Cancer and anticancer chemotherapy ......................................................... 17
1.1.2 Tumour microenvironment ......................................................................... 18
1.1.2.1 Tumour vasculature ............................................................................ 18
1.1.2.2 The enhanced permeability and retention effect in tumours .............. 19
1.1.2.3 Multidrug resistance ........................................................................... 21
1.1.2.4 Cancer stem cells ................................................................................ 24
1.1.2.4.1 Breast cancer ................................................................................. 29
1.1.2.4.2 Brain cancer .................................................................................. 30
1.1.2.4.3 Colon cancer ................................................................................. 31
1.1.2.5 Targeted anticancer therapy ................................................................ 32
1.1.2.5.1 Antibody-targeted therapy ............................................................ 33
1.1.2.5.2 Small molecule inhibitors ............................................................. 33
1.1.2.5.3. Aptamer-targeted therapy ............................................................ 34
1.2. Cancer nanotechnology .................................................................................. 36
1.2.1 Nanoparticle therapeutics for cancer .......................................................... 39
1.2.1.1 Size ..................................................................................................... 40
1.2.1.2 Surface properties of nanoparticles .................................................... 41
1.2.2.2.1 Passive targeting properties .......................................................... 41
1.2.2.2.2 Active targeting ............................................................................ 43
1.3. Liposomes ....................................................................................................... 46
1.3.1 History and development of liposomes for anticancer drug delivery ........ 47
1.3.2 Characterisation of liposomes .................................................................... 50
1.3.3 Advantages of liposomes for anticancer drug delivery .............................. 52
1.3.3.1 Enhanced stability of liposomal drug formulation ............................. 53
1.3.3.2 A controlled release system of anticancer drugs ................................ 54
1.3.3.3 Reversal of MDR ................................................................................ 56
1.3.3.4 Improved PK and BD ......................................................................... 57
1.3.3.5 Aptamer and CSCs targeting .............................................................. 59
1.4. Project significance and aims ......................................................................... 60
Chapter 2 - Methodology ........................................................................................... 65
2.1 Ethics Statement .............................................................................................. 66
2.2 Materials .......................................................................................................... 66
2.2.1 Chemicals, reagents and equipment ........................................................... 66
2.2.2 Cells used in this study ............................................................................... 67
2.2.3 Animals used in this study .......................................................................... 68
2.2.4 Aptamers used in this study ........................................................................ 69
2.3 Methods ........................................................................................................... 70
2.3.1 Development of sulfatide-containing nanoliposomal doxorubicin (SCL-
DOX) and PEG-sulfatide-containing nanoliposomal doxorubicin (PEG-SCL-
DOX) ................................................................................................................... 70
2.3.1.1 Development of SCL-DOX. ............................................................... 70
2.3.1.2 Development of PEG-SCL-DOX. ...................................................... 71
2.3.1.3 Development of DOX-encapsulated, aptamer functionalised PEG-
SCL ................................................................................................................. 72
2.3.1.4 Chromatographic instrumentation and system ................................... 73
2.3.1.5 Establishing standard curves for determination of DOX loading
efficiency ........................................................................................................ 73
2.3.1.6 Determination of drug loading efficiency of SCL-DOX or PEG-SCL-
DOX ............................................................................................................... 74
2.3.1.7 Establishing standard curves for phospholipids assay ....................... 75
2.3.1.8 Determination of phospholipids in SCL-DOX or PEG-SCL-DOX ... 75
2.3.1.9 Determination of particle size and zeta potential of SCL-DOX or
PEG-SCL-DOX .............................................................................................. 76
2.3.1.10 In vitro release kinetics of SCL-DOX/PEG-SCL-DOX in PBS ...... 76
2.3.1.11 In vitro release kinetics of SCL-DOX/PEG-SCL-DOX in PBS/10%
FBS ................................................................................................................. 76
2.3.1.12 Establishing standard curves for in vitro release kinetics in PBS .... 77
2.3.1.13 Establishing standard curves for in vitro release kinetics in PBS/10%
FBS ................................................................................................................. 77
2.3.2 In vitro uptake, retention and cytotoxicity assay ........................................ 78
2.3.2.1 MTT assay .......................................................................................... 78
2.3.2.2 Confocal microscopy analysis for cellular uptake and retention of
SCL-DOX ....................................................................................................... 79
2.3.2.3 Cellular uptake and retention of aptamer-functionalised DOX-
encapsulated PEG-SCL .................................................................................. 79
2.3.2.4 Flow cytometry analysis for intracellular uptake of SCL-DOX ........ 80
2.3.3.1 Development of [3H]-tracking of liposome. ....................................... 80
2.3.3.2 Pharmacokinetics (PK) study in SD rats ............................................ 81
2.3.3.3 Biodistribution (BD) study in rats ...................................................... 81
2.3.3.4 Establishing standard curves of DOX in plasma ................................ 82
2.3.3.5 Determination of the stability of SCL-DOX in plasma ...................... 82
2.3.3.6 Establishing standard curves of DOX in tissues ................................ 82
2.3.3.7 Determination of the release of DOX from SCL-DOX in tissues ...... 83
2.3.3.8 Liquid scintillation counting assay for [3H] in plasma ....................... 83
2.3.3.9 Liquid scintillation counting assay for [3H] in tissues ....................... 84
2.3.3.10 Tumour xenograft model .................................................................. 84
2.3.3.11 Establishing standard curves of DOX in tissues and tumours in
tumour-bearing mice ...................................................................................... 85
2.3.3.12 Tumour uptake and tissue distribution of SCL-DOX in tumour-
bearing mice ................................................................................................... 85
2.3.3.13 Treatment efficacy of SCL-DOX in tumour-bearing mice ............... 85
2.3.3.14 Analysis of systemic toxicity of SCL-DOX ..................................... 86
2.3.3.15 Perfusion in healthy SD rats ............................................................. 86
2.3.3.16 Establishment of a intracranial brain tumour xenograft model ........ 87
2.3.3.17 Establishing standard curves of DOX in tissues and tumours in
intracranial brain tumour xenograft models ................................................... 88
2.3.3.18 Tumour uptake and tissue distribution in intracranial brain tumour
xenograft models ............................................................................................ 88
2.3.4 Data analysis .............................................................................................. 88
Chapter 3 - Development of sulfatide-containing nanoliposomal doxorubicin and
PEG-sulfatide-containing nanoliposomal doxorubicin .............................................. 90
3.1 Introduction ..................................................................................................... 91
3.2 Results ............................................................................................................. 93
3.2.1. Characterisation of nanoliposomes ........................................................... 93
3.2.1.1 Establishing standard curves for determination of DOX loading
efficiency ........................................................................................................ 93
3.2.1.2. Establishing standard curves for phospholipids assay ...................... 94
3.2.1.3 Characterisation of nanoliposomes .................................................... 95
3.2.2 In vitro drug retention properties. .............................................................. 96
3.2.2.1 Establishing standard curves for in vitro release kinetics in PBS and
PBS/10% FBS ................................................................................................ 96
3.2.2.2 In vitro drug retention properties ........................................................ 98
3.2.3 Flow cytometry analysis for intracellular uptake of SCL-DOX ................ 99
3.2.4. In vivo characteristics of SCL-DOX and PEG-SCL-DOX. .................... 100
3.2.4.1 Establishing standard curves for DOX concentration in plasma ...... 100
3.2.4.2 Improved pharmacokinetic properties of SCL in healthy SD rats ... 101
3.2.4.3 Establishing standard curves for tissue distribution ......................... 103
3.3.4.4 Tissue distribution advantages of SCL-DOX in SD rats .................. 105
3.2.4.5 Drug-to-lipid ratios in vivo ............................................................... 107
3.3 Discussion ...................................................................................................... 109
3.4 Conclusion ..................................................................................................... 113
Chapter 4 - Enhanced antitumour activity and reduced systemic toxicity of glioma
tumour microenvironment targeted sulfatide-containing liposomes ........................ 115
4. 1 Introduction .................................................................................................. 116
4. 2 Results .......................................................................................................... 118
4.2.1 In vitro cytotoxicity .................................................................................. 118
4.2.2 Intracellular uptake and retention of SCL-DOX in U-118MG cells ........ 119
4.2.3 Establishing standard curves for tissue distribution ................................. 123
4.2.4 Improved tumour uptake and biodistribution in U-118MG tumour
xenograft model ................................................................................................. 125
4.2.5 Enhanced therapeutic efficacy of SCL-DOX in U-118MG xenograft
tumour model .................................................................................................... 126
4.2.6 Reduced toxicity of SCL-DOX in vivo .................................................... 128
4.3 Discussion ...................................................................................................... 131
4.4 Conclusion ..................................................................................................... 136
Chapter 5 - Enhanced antitumour efficacy and reduced systemic toxicity of sulfatide-
containing nanoliposomal doxorubicin in a xenograft model of colorectal cancer . 137
5.1 Introduction ................................................................................................... 138
5.2 Results ........................................................................................................... 140
5.2.1 Intracellular uptake and retention of SCL-DOX in HT-29 cells .............. 140
5.2.2 In vitro cytotoxicity .................................................................................. 144
5.2.3 Establishing standard curves for tumour uptake ...................................... 144
5.2.4 Tissue distribution and tumour uptake Advantages of SCL-DOX ........... 145
5.2.5 Enhanced therapeutic efficacy of SCL-DOX ........................................... 147
5.2.6 Reduced systemic toxicity of SCL-DOX ................................................. 149
5.3 Discussion ...................................................................................................... 151
5.4 Conclusion ..................................................................................................... 156
Chapter 6 - Exploration of the potential of using sulfatide-containing liposomes for
drug delivery into the brain ...................................................................................... 157
6.1 Introduction ................................................................................................... 158
6.2 Results ........................................................................................................... 160
6.2.1 Enhanced brain distribution of SCL-DOX in healthy SD rats ................. 160
6.2.2. Enhanced brain distribution of SCL-DOX in perfused healthy SD rats . 161
6.2.3 Establishment of intracranial brain tumour xenograft models ................. 162
6.2.4 Establishing standard curves for tissue distribution in Wistar rats ........... 163
6.2.5 Tissue distribution and tumour accumulation of SCL-DOX in an
intracranial brain tumour xenograft model ........................................................ 166
6.3 Discussion ...................................................................................................... 167
6.4 Conclusion ..................................................................................................... 171
Chapter 7 - Development of EpCAM aptamer functionalised PEGylated sulfatide-
containing liposomal doxorubicin ............................................................................ 172
7.1 Introduction ................................................................................................... 173
7.2 Results ........................................................................................................... 175
7.2.1 Establishing a standard curve for DOX concentration in aptamer
functionalised PEG-sulfatide-containing nanoliposomal DOX (ap-PEG-SCL-
DOX) ................................................................................................................. 175
7.2.2 Characterisation of ap-PEG-SCL-DOX and PEG-SCL-DOX modified with
negative control aptamer (negative ap- PEG-SCL-DOX) ................................. 176
7.2.3 Cellular uptake of ap-PEG-SCL-DOX ..................................................... 177
7.2.4 Cellular retention of aptamer functionalised PEG-SCL-DOX ................. 180
7.3 Discussion ...................................................................................................... 185
7.4 Conclusion ..................................................................................................... 188
Chapter 8 - Summary and perspectives .................................................................... 189
8.1 Summary ........................................................................................................ 190
8.2 Perspectives ................................................................................................... 194
References ................................................................................................................ 198
Appendix .................................................................................................................. 247
1
List of Figures
Figure 1-1 The Enhanced permeability and retention (EPR) effect in tumours 21
Figure 1-2. Mechanisms of multidrug resistance (MDR) development in tumour cells
22
Figure 1-3. Stochastic versus cancer stem cell hypothesis of tumours 26
Figure 1-4. SELEX technology 36
Figure 1-5. The tumour microenvironment 43
Figure 1-6. Active cellular targeting of nanoparticles 45
Figure 1-7. Liposomes 47
Figure 1-8. Liposome classification 51
Figure 3-1. Standard curves for determination of DOX loading efficiency. 94
Figure 3-2. Standard curves for phospholipids assay. 95
Figure 3-3. Standard curves for in vitro release kinetics in PBS and PBS/10% FBS.
97
Figure 3-4. In vitro stability of SCL-DOX and PEG-SCL-DOX. 99
Figure 3-5. DOX uptake in MCF-7 and U-118MG cells. 100
Figure 3-6. Standard curve for DOX concentration in plasma. 101
Figure 3-7. Plasma clearance of different DOX formulation in healthy SD rats
102
Figure 3-8. Standard curves for tissue distribution. 104
Figure 3-9. Biodistribution of DOX encapsulated in SCL in SD rats. 106
Figure 3-10. DOX-to-lipid ratios in tissues of rats. 108
Figure 4-1. Intracellular uptake of SCL-DOX in U-118MG cells. Intracellular uptake of SCL-DOX in U-118MG cells.
120
Figure 4-2. Intracellular retention of SCL-DOX in U-118MG cells. 122
Figure 4-3 Standard curves for tissue distribution in mice. 124
Figure 4-4. BD and tumour uptake of DOX encapsulated in SCL in xenograft-bearing mice.
126
2
Figure 4-5. Improved therapeutic activity of SCL-DOX against glioma xenograft.
127
Figure 4-6. SCL-DOX enhanced survival of tumour-bearing mice. 128
Figure 4-7. Encapsulation of DOX in SCL substantially reduced the accumulation of DOX in principal sites of DOX toxicity.
130
Figure 4-8. Encapsulation of DOX in SCL significantly reduced cardiac and hepatic toxicity.
131
Figure 5-1. Intracellular uptake of SCL-DOX in HT-29 cells. 141
Figure 5-2. Intracellular retention of SCL-DOX in HT-29 cells. 143
Figure 5-3. Standard curve for DOX concentration in HT-28 tumours. 145
Figure 5-4. BD and tumour uptake of DOX encapsulated in SCL in xenograft-bearing mice.
146
Figure 5-5. SCL-DOX produced more reduction in tumour volume. 148
Figure 5-6. SC--DOX enhanced survival. 149
Figure 5-7. SCL-DOX treatment had significantly reduced myelosuppression.
150
Figure 6-1. Improved brain distribution in healthy rats. 161
Figure 6-2. Improved brain distribution in perfused healthy rats. 162
Figure 6-3. Establishment of intracranial brain tumour xenograft models. 163
Figure 6-4. Standard curves for tissue distribution in Wistar rats. 165
Figure 6-5. Tissue distribution of DOX encapsulated in SCL in C6 glioma tumour-bearing rats.
167
Figure 7-1. Standard curve for determination of DOX concentration in ap-PEG-SCL-DOX.
175
Figure 7-2. Development of ap-PEG-SCL-DOX 176
Figure 7-3. Intracellular uptake of aptamer functionalised PEG-SCL-DOX in HEK-293T cells and HT-29 cells.
179
Figure 7-4. Intracellular retention of aptamer functionalised SCL-DOX in HEK-293T cells and HT-29 cells (1 h).
181
Figure 7-5. Intracellular retention of aptamer functionalised SCL-DOX in HEK-293T cells and HT-29 cells (2 h).
182
3
Figure 7-6. Intracellular retention of aptamer functionalised SCL-DOX in HEK-293T cells and HT-29 cells (4 h).
183
Figure 7-7. Intracellular retention of aptamer functionalised SCL-DOX in HEK-293T cells and HT-29 cells (24 h).
184
4
List of Tables
Table 1-1. Summary of markers used for cancer stem cells 28
Table 3-1. Physical properties of the liposomal formulations. 96
Table 3-2. Pharmacokinetic parameters for free DOX, SCL-DOX and PEG-SCL-DOX.
102
Table 4-1. Mean IC50 values ( g/mL of DOX) for treatment with free DOX and SCL-DOX.
119
Table 4-2. Plasma levels of troponin. 131
Table 5-1. Mean IC50 values ( g/ml of doxorubicin) for treatment with free DOX and SCL-DOX
144
Table 5-2. Plasma levels of troponin 150
Table 7-1. Physical properties of the liposomal formulations. 176
5
Publications
Lin J, Yu Y, Shigdar S, Fang DZ, Du JR, Wei MQ, Danks A, Liu K and Duan W.
“Enhanced antitumor efficacy and reduced systemic toxicity of sulfatide-containing
nanoliposomal doxorubicin in a xenograft model of colorectal cancer”. PLoS ONE,
2012, 7(11): e49277. (Impact Factor = 4.092)
Sarah Shigdar, Jia Lin, Yong Li, Chaoyong James Yang, Mingqian Wei, Yimin
Zhu, Hui Liu and Wei Duan. “Cancer stem cell targeting: the next generation of
cancer therapy and molecular imaging”. Therapeutic Delivery, 2012, 3 (2) :227-44.
S Shigdar, J Lin, Y Yu, M Pastuovic, M Wei, and Wei Duan. “RNA aptamer against
a cancer stem cell marker epithelial cell adhesion molecule”. Cancer Science, 2011,
102(5):991-8. (ERA: A, Impact Factor = 3.846)
Lin J, Fang DZ, Du J, Shigdar S, Xiao LY, Zhou XD, Duan W. “ Elevated Levels of
Triglyceride and Triglyceride-Rich Lipoprotein Triglyceride Induced by a High-
Carbohydrate Diet Is Associated with polymorphisms of APOA5-1131T>C and
APOC3-482C>T in Chinese healthy young adults”. Annals of Nutrition &
Metabolism. 2011, 58(2):150-7. (Impact Factor = 2.257)
J. Du., D. Z. Fang, J. Lin, L. Y. Xiao, X. D. Zhou, S. Shigdar and Wei
Duan. “TaqIB polymorphism in the cholesterol ester transfer protein (CETP) gene
modulates the impact of HC/LF diet on the HDL profile in healthy Chinese young
adults”. The Journal of Nutritional Biochemistry. 2010, 21:1114-9. (Impact Factor
= 4.538; ERA: A*)
6
Huang X, Gong RR, Lin J, Li RH, Xiao LY, Duan W, Fang DZ (Co-corresponding
author). “Interactions of Lipoprotein lipase gene variations, high carbohydrate low
fat diet and gender on serum lipid profiles in healthy Chinese Han
youth”. Bioscience Trends, 2011, 5(5):198-204. (Impact Factor = 0.392)
7
Papers submitted for publication
Jia Lin, Yan Yu, Sarah Shigdar, Liang Qiao, Ding Zhi Fang, Andrew Danks, Ming
Q. Wei, Lingxue Kong, Lianghong Li and Wei Duan. “Enhanced antitumor activity
and reduced systemic toxicity of tumor microenvironment targeted sulfatide-
containing nanoliposome”. Submitted to Cancer Letters.
Sarah Shigdar, Christine Qian, Li Lv, Chunwen Pu, Yong Li, Lianhong Li, Manju
Marappan, Jia Lin, Lifen Wang, Wei Duan. “The use of sensitive chemical
antibodies for diagnosis: detection of low levels of EpCAM in breast cancer”.
Submitted to PLos One.
8
Acknowledgements
To my supervisor Professor Wei Duan, thank you for the great support and
encouragement you have given me not only in my PhD studies, but also in my
general life. It has been a wonderful experience to work on the projects. I have learnt
a lot from you about how to improve my observations, basic experimental skills,
scientific writing, presentation skills and response to others’ comments. Your advice
has been invaluable to me.
To Doctor Sarah Shigdar, thank you for your kindly help with my drafts, my
experimental designs, technical training, general problems in my studies and my life
when I needed assistance or advice. There is no doubt in my mind that you and your
encouragement have been a major motivator and something I will always treasure in
my life. It has been an unforgettable experience to work with you.
Moreover, I would like to thank my past and present colleagues: Yan Yu, Lei Li,
Tao Wang, Dongxi Xiang and Manju Marappan. My PhD has been full of enjoyment
because of you. Thank you for your support, help, and most of all friendship
throughout these years.
I am also grateful to all past and present staffs and students who provided a very
good research environment for me in the School of Medicine and in the Upper
Animal House.
To my family, Dad, Mum, and my husband, thank you for your endless love,
encouragement, understanding, support and encouragement over the years. You have
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always been there to motivate me and no words can describe how grateful I am to
have such a wonderful family.
10
Abstract
Cancer is still one of the leading causes of deaths in the world and chemotherapy
remains the major treatment modality for most advanced cancers. However,
traditional anticancer agents have numerous problems due to their non-specific
distribution and low therapeutic index in vivo, leading to serious side effects in
patients. Thus, there is a big challenge to develop new chemotherapeutics that can
target cancer cells effectively. In recent years, nanotechnology has been widely
developed and utilised in the treatment of cancers and the study of liposomes is a
critical direction in this area. Liposomes have become well-recognised vehicles for
drug encapsulation and have been shown to enhance the therapeutic efficacy of
several traditional antitumour drugs.
Ideally, liposomes for drug delivery have the ability to accumulate in tumour tissues
via the enhanced permeability and retention effects (the EPR effect) in tumours.
Moreover, modification with ligands on the surface of liposomes could further
enhance the targeting property of liposomes. Sulfatide, a glycosphingolipid which
can bind to several extracellular matrix proteins, has the potential to become a
targeting agent in tumours overexpressing tenascin-C. As a result, sulfatide was used
in the development of sulfatide-containing nanoliposomes (SCL) for doxorubicin
(DOX) encapsulation to overcome its dose-limiting toxicity.
This project was aimed to analyse in vitro characteristics of sulfatide-containing
liposomal DOX (SCL-DOX) and to assess its cytotoxicity in vitro, as well as
pharmacokinetics (PK) and biodistribution (BD) in healthy rats. Then, the human
glioblastoma cell line U-118MG and the human colorectal adenocarcinoma cell line
HT-29, which both have high expression of tenascin-C, were utilised as models for
11
the studies of therapeutic efficacy and systemic toxicity both in vitro and in vivo. As
there is potential for SCL-DOX to cross the blood-brain barrier (BBB), an
intracranial brain C6 tumour xenograft model in rats was established for the
evaluation of tumour uptake of SCL-DOX. To further improve the active targeting
property of SCL-DOX, an aptamer targeting the epithelial cell adhesion molecule
(EpCAM), a cell surface marker for the cancer stem cells (CSCs), was used to
modify the SCL and the cellular uptake and retention of drugs was compared
between an EpCAM-positive cell line and an EpCAM-negative cell line.
The results indicated that the SCL-DOX had a satisfactory drug loading efficiency,
a high drug-to-lipid ratio, an ideal average size as well as a remarkable in vitro
stability. It was also shown that the SCL-DOX could prolong the retention property
in both U-118MG cells and HT-29 cells in vitro when compared to the free DOX.
Besides the improved pharmacokinetic profile in healthy SD rats, in vivo data also
suggested improved tissue distributions in both healthy rats and tumour-bearing mice.
Meanwhile, the enhanced treatment efficacy and reduced systemic toxicity of SCL-
DOX were observed in U-118MG and HT-29 tumour-bearing mice compared to the
free drug. Furthermore, the findings indicated that the SCL-DOX was able to
penetrate the BBB and accumulate in the glioma. As well, modification with the
EpCAM targeting aptamer on the SCL-DOX demonstrated specific targeting to the
EpCAM-positive cell line.
In conclusion, the current project shows that the SCL could be an effective and safe
nanocarrier for the anticancer agent DOX. The SCL-DOX has the potential to target
tumours with high expression of tenascin-C. Moreover, SCL encapsulation could be
a new strategy for the treatment of diseases in the central nervous system (CNS). The
12
aptamer functionalised DOX-encapsulated PEG-SCL might be a novel nanodrug to
improve the treatment efficacy in the CSCs.
13
List of abbreviations
A
ABC ATP-binding cassette
ALDH aldehyde dehydrogenase
ALDH1 aldehyde dehydrogenase isoform 1
ANOVA analysis of variance
ap-SCL-DOX aptamer functionalised PEG-SCL-DOX
AUC Area under the plasma concentration-time curves
B
BBB blood-brain barrier
BD biodistribution
C
°C degrees celsius
CD cluster of differentiation
CHE cholesteryl hexadecyl ether
CNS central nervous system
COOH hydrophilic carboxylic acid functional group
CSCs cancer stem cells
D
DMEM Dulbecco’s modified eagles medium
DMSO dimethyl sulfoxide
DOPE 1,2-dioleoyl-snglycero-3-phosphoethanolamine
DOX doxorubicin
E
ECM extracellular matrix
EDC 1- [3-dimethylaminopropyl]-3-ethylcarbodiimide
hydrochloride
EDTA ethylenediamine tetraacetic acid
EGFR epidermal growth factor receptor
EpCAM epithelial cell adhesion molecule
EPR enhanced permeability and retention effect
F
14
FBS foetal bovine serum
FDA Food and Drug Administration
G
g G force
H
h hour
HPLC high performance liquid chromatography system
I
IF intermediate filament
i.v. intravenous
M
MCF-7 Michigan Cancer Foundation - 7
MDR multidrug resistance
MDR I multidrug-resistant I
mg milligram
min minute
mL millilitre
mAbs monoclonal antibodies
MRP MDR related protein
MTT 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyl-2H-
tetrazolium bromide
N
ng nanogram
NH2 ammine ends
NHS N-hydroxysuccinimide
NOD/SCID Nonobese diabetic/severe combined immunodeficiency
P
PBS phosphate saline buffer
PDI polydispersity index
PEG polyethylene glycol
PEG2000-DSPE 1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-
[maleimide(polyethylene glycol)-2000]
PEG-SCL-DOX PEG-sulfatide-containing nanoliposomal doxorubicin
15
PLD pegylated liposomal DOX
P-gp P-glycoprotein
PK pharmacokinetics
PPE Palmar-plantar erythrodysesthesia
PSMA prostate-specific membrane antigen
R
RES reticuloendothelial system
RT room temperature
S
SCL sulfatide-containing liposomes
SCL-DOX sulfatide-containing liposomal DOX
SD rats Sprague-Dawley rats
S.E. standard error
SELEX systematic evolution of ligands by exponential
enrichment
Sulfo-NHS sulfo-N-hydroxysuccinimide
T
t1/2 Elimination half-life
thioaptamer thioated oligonucleotide aptamer
U
g microgram
L microlitre
V
Vd volume of distribution
W
WHO World Health Organisation
16
Chapter 1 - Introduction
17
1.1 Cancer and anticancer chemotherapy
1.1.1 Cancer and anticancer chemotherapy
Cancer is still one of the most devastating diseases in the world despite vast
government funding and a heroic effort by cancer researchers to reduce mortality and
improve survival (Klein, Christoph A. 2009). Cancer treatment modalities include
surgery, radiation and chemotherapeutics (Chabner & Roberts 2005; Peer, Dan et al.
2007; Sapra, P., Tyagi & Allen 2005). However, radiation therapy, as well as surgery,
cannot eradicate metastatic tumours (Chabner & Roberts 2005). The majority of
tumour-induced fatalities are attributable to metastatic diseases, and early cancer
detection and prevention of metastases has become the centre of clinical attention
(Klein, Christoph A. 2009). Due to its spectacular success in curing certain types of
leukaemia, chemotherapy is becoming the major weapon in the management of
metastatic solid tumours (Balis 1988).
During the quest to provide a cure for cancer, cancer drug development has been
transformed from a low-budget, government-supported research effort to a high-
stakes, multi-billion dollar industry (Chabner & Roberts 2005). However, most
traditional chemotherapeutic agents lack specificity to cancer cells. Thus, they often
kill healthy cells as well as tumour cells (Alexis, Frank , Rhee, June-Wha , et al.
2008; Bawarski et al. 2008). Moreover, chemotherapy also causes severe toxicity to
patients, leading to serious side effects, low efficacy and poor quality of life (Peer,
Dan et al. 2007; Sapra, P., Tyagi & Allen 2005). In addition, most anticancer agents
have an extensive distribution in the human body after intravenous (i. v.)
administration, resulting in a low therapeutic index and a high level of toxicity to
healthy tissues. This non-specific biodistribution (BD) and the resulting side effects
18
limit the clinical application of anticancer drugs (Ganta, Srinivas et al. 2008).
Furthermore, there are many barriers to the effective delivery of anticancer drugs to
the tumour due to its pathophysiological properties. This is further exacerbated by
the fact that tumours can be located in parts of body where anticancer agents cannot
easily penetrate, such as the brain, or protected by the local environment due to
increased tissue hydrostatic pressure or altered tumour vasculature (Drummond, D. C.
et al. 1999b; Klein, Christoph A. 2009; Szakács et al. 2006). Thus, there is an urgent
need to develop new chemotherapeutics that can target tumour cells effectively.
1.1.2 Tumour microenvironment
1.1.2.1 Tumour vasculature
The structure of newly generated blood vessels in tumours are abnormal (Alexis,
Frank , Rhee, June-Wha , et al. 2008; Tang et al. 2007), with studies indicating that
the abnormal vasculature in the tumour plays an important role in tumour
permeability. Tumour blood vessels are characterised by leakage, tortuousness,
dilation and disordered interconnection (Tang et al. 2007). In some tumours, vessels
share features of arterioles, capillaries and venules. Furthermore, the walls of tumour
vessels lack normal basement membranes and perivascular smooth muscle when
compared with that of healthy tissues (Trédan et al. 2007). Moreover, the integration
of tumour cells into the vessel wall exacerbates the leakiness of the tumour
vasculature. Due to these structural abnormalities, the capillary network of tumours
is highly permeable to macromolecules when compared to their normal counterpart
(Campbell 2006; Greish 2007; Kong, Braun & Dewhirst 2000). The newly formed
tumour vessels have fenestrations ranging from 100 to 780 nm, while the pore size of
the vascular endothelium in most healthy tissues is smaller than 6 nm (Drummond, D.
19
C. et al. 1999b). Consequently, macroparticles, such as albumin, and even larger
nano-sized particles (Kratz et al. 2008), remain in the bloodstream when passing
through the normal tissues where the openings are not large enough for them to pass
in, but would extravasate through the leaky vasculature of tumours (Gullotti & Yeo
2009). In most cases, particles smaller than 200 nm would be able to cross into the
tumour microvasculature (Alexis, Frank, Rhee, June-Wha, et al. 2008).
1.1.2.2 The enhanced permeability and retention effect in tumours
The clearance route of macromolecules in the interstitial space is the lymphatic
system in both normal tissues and tumours (Maeda, H., Bharate & Daruwalla 2009).
However, tumours are characterised by an irregular lymphatic system. Due to the
absence of functional lymphatics, the clearance of macromolecules is retarded from
the tumour interstitium (Brigger, Dubernet & Couvreur 2002). Moreover, in many
solid tumours, such as rectal and breast tumours, there is an elevated interstitial fluid
pressure when compared with normal tissues (Jain 1990). Consequently, as
interstitial fluid pressure can drive the pressure gradient or fluid flux in the tumour
interstitium back to the vascular compartment, it reduces the pressure gradients from
the vascular compartment into the interstitial space (Ferretti et al. 2009; Lunt,
Chaudary & Hill 2009). As a result, limited delivery and transport of
macromolecules ensue due to the increased interstitial fluid pressure gradient from
the tumour centre to the periphery (Heldin et al. 2004). Thus, unlike normal tissues
where clearance of macromolecules proceeds rapidly and steadily via the lymphatic
system, clearance of macromolecules from the tumour lymphatic system is impaired
and they can remain in the tumour interstitium for a long period of time (Maeda,
Hiroshi 2001).
20
The characteristics of tumour vasculature enhance the permeability of
macromolecular compounds such as nanoparticles in tumour tissues, while the
impaired lymphatic system contributes to the retention of macromolecules in the
interstitial space of tumours for long periods (Maeda, Hiroshi 2001). Thus, the highly
permeable vessel structures and the lack of an intact lymphatic system in tumours
result in the phenomenon termed the enhanced permeability and retention (EPR)
effect (Figure 1-1) (Bisht & Maitra 2009). The EPR effect, discovered by Maeda and
Matsumura in 1986, is characterised by enhanced extravasation of macromolecules
from tumour blood vessels and their retention in tumour tissues (Maeda, H., Bharate
& Daruwalla 2009). The EPR effect, modulated by abnormal secretion of vascular
endothelial factors, such as bradykinin, nitric oxide, prostaglandins and matrix
metalloproteins (Alexis, Frank, Rhee, June-Wha, et al. 2008), forms the foundation
of cancer-targeting drug design as it is observed in almost all solid tumours in
humans (Maeda, H., Bharate & Daruwalla 2009; Miles & Williams 2008).
21
Figure 1-1. The Enhanced permeability and retention (EPR) effect in tumours (adapted from Cassidy and Schätzlein 2004). Normal tissue vasculatures are characterised by tight endothelial cells which thereby prevent macromolecules from passing in, whereas tumour vasculatures are leaky and hyperpermeable for the nanoparticles. Small blue dots: small molecules; Large green dots: macromolecules. (Cassidy & Schätzlein 2004)
1.1.2.3 Multidrug resistance
Multidrug resistance (MDR), the resistance of cancer cells to one or more drugs in
chemotherapy, is a significant obstacle to effective treatment of tumours by
chemotherapeutics (Peer, Dan et al. 2007). It is postulated that MDR results in lower
therapeutic effects as well as resistance to various drugs in cancer cells (Domb et al.
2007; Gottesman 2002; Gottesman, Michael M. , Fojo, Tito & Bates, Susan E. 2002;
Peer, Dan et al. 2007). The mechanisms of MDR in cancer include a decreased
uptake of water-soluble drugs, cellular changes affecting the cytotoxicity of drugs
and an increased efflux of hydrophobic cytotoxic drugs (Figure 1-2) (Jabr-Milane et
al. 2008; Szakács et al. 2006).
22
Figure 1-2. Mechanisms of multidrug resistance (MDR) development in tumour cells (adapted from Gottesman, Fojo et al. 2001). Tumour cells can develop resistance to anticancer agents by different mechanisms, such as elevated drug efflux by incresing activity of pumps or reduced drug influx. Alternatively, resistance can be induced by the activation of detoxifying proteins, disruption of apoptotic signalling patheways, or increased of DNA repair. (Gottesman, M. M., Fojo, T. & Bates, S. E. 2002)
The most widely studied mode of MDR action is drug efflux. Many clinical
anticancer drugs are derivatives of natural products, such as anthracyclines and vinca
alkaloids. These agents share properties in terms of high affinity with energy-
dependent membrane transporter proteins, leading to the efflux of drugs out of cells
(Patel & Tannock 2009). According to current wisdom, MDR occurs as a result of
overexpression of an ATP-dependent transporter protein family on the cancer cell
surface. These proteins are known as the ATP-binding cassette (ABC) transporters,
utilising ATP for exporting drugs against the concentration gradient (Dean, Rzhetsky
& Allikmets 2001; Peer, Dan et al. 2007; Spiegl-Kreinecker et al. 2002; Szakács et al.
2006). The ABC transporters are found in all cells of all species from the most
primitive microorganism to humans, and play central roles in various physiologic
systems as well as drug resistance in tumour cells (Surowiak et al. 2006).
23
P-glycoprotein (P-gp), a transmembrane glycoprotein with a molecular weight of
170 kDa, is a key protein in the ABC transporter family. It is encoded by the MDR
gene, MDR1. P-gp plays a critical role not only in regulating drug access through the
blood-brain-barrier (BBB), as it is involved in the exclusion of both neutral and
cationic organic compounds (Roya et al. 2009), but also in chemoresistance in
tumours. The role of P-gp in MDR in tumours has been confirmed by many studies
(Bradley & Ling 1994; Taheri, Mahjoubi & Omranipour 2010; Thomas & Coley
2003). It functions as a pump transporting anticancer drugs out of tumour cells in an
ATP-dependent manner, and confers cancer cells with the strongest resistance to the
widest variety of agents among the ABC transporters (Bradley & Ling 1994; Li, Y et
al. 2010; Szakács et al. 2006; Taheri, Mahjoubi & Omranipour 2010).
However, there are some resistant tumour cells overexpressing other proteins from
the MDR related protein (MRP) family instead of P-gp. Two main MRP transporters,
MRP 1 and MRP 2, contribute to drug resistance in some multidrug resistant tumour
cells. The overexpression of human MRP 1 in cancer patients has been reported by
many independent laboratories (Burger et al. 2003; Komdeur et al. 2001; Mahjoubi
et al. 2008). MRP 1 has the potential to confer drug resistance to a variety of drugs,
such as anthracyclines, vinca alkaloids, and VP-16 (Mahjoubi et al. 2008; Minko et
al. 2006), which is related to a reduced intracellular drug accumulation in tumours
(Mahjoubi et al. 2008). Another drug transporter protein in the MRP family is the
MRP 2. It has been shown that the MRP 2, a 190 kDa phosphoglycoprotein, is
expressed on the apical membrane of hepatocytes, enterocytes, and renal proximal
tubules (Surowiak et al. 2006; Zhang, Yuanyuan , Li & Vore 2007). The MRP 2 is
also involved in the transport of chemotherapeutic agents out of the tumours (Choi et
al. 2007). In vitro experiments also showed that overexpression of the human MRP 2
24
could confer resistance to platinum-containing anticancer drugs (Surowiak et al.
2006).
These transporters are associated with the efficient pumping of a drug out of the
tumour cells, and therefore, their expression level is inversely correlated with the
chemotherapeutic phenotype in various cancers (Choi et al. 2007). Conceivably, the
suppression of the P-gp expression, the human MRP 1 and/or the human MRP 2
transporters could decrease chemoresistance by increasing drug concentration inside
the cancer cells and therefore enhance the efficacy of chemotherapeutic treatments.
1.1.2.4 Cancer stem cells
There are at least two different models which can explain tumour initiation in vivo
(Figure 1-3), the stochastic model or the cancer stem cells model. In the stochastic
model, it is assuming that every cancer cells has the potential to proliferate and
regenerate a tumour (Ward & Dirks 2007). However, in the 1960s, studies had
already indicated that even if millions of cancer cells were inoculated in vivo, not all
the inplantations could develop into a palpable nodule (Brunschw, Southam & Levin
1965; Foo, Leder & Michor 2011). Thus, it is very difficult for the stochastic model
to account for this phenomenon as all cancer cells have an equal probability to
regenerate a tumour in this assumption.
Over the past several years, it has become widely accepted that the limited
efficacy of current anticancer treatments can be attributed to a small subset of cells
within tumours with the ability to proliferate and form new tumours (Al-Hajj et al.
2003; Koch, Krause & Baumann 2010). As these highly tumourigenic cancer cells
display some functional properties of normal stem cells, such as self-renewal and
25
mulipotent differentiation (Deonarain, Kousparou & Epenetos 2009; Ebben et al.
2010; Shackleton 2010), they are often referred to as cancer stem cells (CSCs). CSCs
were initially identified in leukaemias and then implicated in various solid tumours
(Alvero et al. 2009; Hanahan & Weinberg 2011; Wakimoto et al. 2009). Increasing
evidence indicates that many, perhaps all, cancers are organised hierarchically with
CSCs and non-CSCs (Heuser & Humphries 2010). CSCs have the ability to generate
a xenograft histologically similar to the parent tumour. Daughter cells generated
from CSCs are able to proliferate but unable to initiate or mantain tumours due to
their lack of intrinsic regenerative potential (O'Brien, CA, Kreso, A & Jamieson,
CHM 2010). Although CSCs only represent a very limited proportion of all tumour
cells, they are resistant to traditional chemotherapy and radiotherapy which can kill
other cells in tumours (Bao et al. 2006). Thus, CSCs have the potential to propagate
and sustain tumourigenesis (Moitra, Lou & Dean 2011; Trumpp & Wiestler 2008;
Visvader & Lindeman 2008). Therefore, CSCs are defined as a small subpopulation
of self-renewing tumour cells which have the capacity to generate heterogeneous
lineages of all cancer cells comprising a tumour (Clarke et al. 2006).
26
Figure 1-3. Stochastic versus cancer stem cell hypothesis of tumours (adapted from Ward and Dirks 2007). In the stochastic model (right), all cells within a heterogeneous human tumour have the ability to regenerate tumours. In contract, the cancer stem cell model (left) suggests that only a small subpopulation of cells within the tumour is tumourigenic.(Ward & Dirks 2007)
The CSC hypothesis is an attractive model to explain how established tumours are
able to propagate (Shackleton 2010). Based on this hypothesis, the bulk of cells in
the tumour are composed of non-CSCs which are non-tumourigenic and have limited
potential of proliferation (Koch, Krause & Baumann 2010). Compared to these
differentiated cancer cells, CSCs have slower cell cycling, lower proliferation, anti-
apoptosis genes and protective mechanisms for DNA repair (Deonarain, Kousparou
& Epenetos 2009; Phillips, TM, McBride & Pajonk 2006). Thus, CSCs are more
resistant to common chemotherapy and radiotherapy. There is increasing evidence
that CSCs share similar critical features with normal stem cells. For example, self-
renewal, the mechanism required for maintaining and expanding populations in
normal stem cells, can be found in CSCs (Hope, Jin & Dick 2004). Moreover,
differentiation, another functional capability of normal stem cells, can also be
observed in CSCs. Thus, due to these properties, the initiation of tumours and tumour
recurrences after anticancer therapies are thought to be associated with CSCs
27
(Sánchez-García, Vicente-Dueñas & Cobaleda 2007). Furthermore, Chaffer et al.
indicated that non-CSCs have the potential to convert to CSCs both in vitro and in
vivo (Chaffer et al. 2011). Therefore, studies about CSCs will lead to a better
understanding of the mechanisms driving tumour growth as well as improved
diagnosis and therapy for tumours. The development of efficient treatment strategies
requires reliable and sufficient isolation of CSCs. Measurements of specific cell
surface markers by flow cytometry have been used for characterisation of CSCs
(Table 1-1).
28
Table 1-1. Summary of markers used for cancer stem cells
Tumours Markers References Acute myeloid leukemia CD34+/CD38- (Kassem 2008)
Breast cancer
CD44+/CD24-/low
ALDH1+ CD133+
CD44+/CD24-/epithelial cell
adhesion molecule +
(EpCAM +)
(Lorico & Rappa 2011;
Stratforda et al. 2010; Velasco-
Velázqueza et al. 2012; Zheng,
Wb et al. 2012)
Brain cancer
CD133+
Nestin+
A2B5+
CD15+
(de Antonellis et al. ; Ishiwata,
Toshiyuki, Matsuda, Yoko. &
Naito, Zenya. 2011;
Tchoghandjian et al. 2010; Yu,
S-p et al. 2011)
Colon cancer
CD133+
EpCAM+
EpCAMhigh/CD44+/CD166+
CD44+/CD24+
CD24+/CD133+
CD24+/CD29+
(Kemper, Grandela & Medema
2010; LaBarge & Bissell 2008;
Levin et al. 2010; Yeung, Trevor
M. et al. 2010; Yu, X et al.
2011)
Pancreatic cancer
CD44+/CD24+/ EpCAM+
CD133+
CXCR4+
DCAMKL-1+
(May et al. 2010; Singh, S et al.
2010; Wei, H-J et al. 2011)
Prostate cancer
CD133+
CD44+/integrin 2 1+/CD133+
CD44+
FAM65Bhigh/MFI2low/LEF1low
ALDH1A1+
CD133+/ Trop-2+
(Hellsten et al. 2011; Lee, HJ et
al. 2011; Trerotola et al. 2010;
Zhang, K & Waxman 2010;
Zhang, Ye. et al. 2012)
29
1.1.2.4.1 Breast cancer
Breast cancer is a complex and heterogeneous disease (Lorico & Rappa 2011).
Amongst solid tumours, it was the first one in which CSCs were identified and
isolated (Al-Hajj et al. 2003). A subpopulation of cells with high expression of CD44
but low levels of CD24 were able to form tumours in vivo with as few as 100 cells.
On the contrary, a large number of cells without this surface marker were found to
lack of tumour-initiating capacity. Moreover, a study into the role of CD44+ breast
CSCs in mice indicated its role in metastasis (Liu, H et al. 2010). Analysis in the
clinic also found there was a relationship between CD44+/CD24-/low breast cancer
cells and metastasis. Abraham et al. indicated that breast cancer patients who had a
high percentage of CD44+/CD24-/low cells had distant metastasis (Abraham et al.
2005). However, CD44 and its isoforms can be expressed in many cancer cell types.
It can also be used as a marker for other epithelial tumours, such as colon tumours
(Cho et al. 2012). Moreover, it has been shown that the silencing of CD24 did not
contribute to tumourigenicity (Lorico & Rappa 2011). Thus, more attention has been
given to other surface markers, such as aldehyde dehydrogenase (ALDH) and CD133,
with high activity of ALDH isoform 1 (ALDH1) observed in the tumourigenic cells
(Ginestier et al. 2007). Activation of the enzymatic activity of ALDH1 in breast
epithelium has been shown to be a marker of breast CSCs. Meanwhile, breast cancer
cells with ALDH1 displayed functional properties of stem cells such as renewal and
differentiation (Kunju et al. 2011). CD133, which is also known as Prominin-1, is a
transmenmbrane glycoprotein and is one of the most commonly studied markers for
isolating CSCs. Interestingly, CD133 has a much restricted expression compared to
CD44. Breast cancer cells with CD133 expression have been reported to possess a
higher ability to initiate tumours in nonobese diabetic/severe combined
30
immunodeficiency (NOD/SCID) mice (Liu, Q et al. 2009). In other studies, EpCAM,
was used as another cell surface marker for CSCs. Fillmore et al. demonstrated that
breast cancer cells with CD44+/CD24-/EpCAM+ phenotype enriched for CSCs in
human breast cancer cell lines (Fillmore & Kuperwasser 2008).
1.1.2.4.2 Brain cancer
The cell surface marker CD133 is also used for the identification of CSCs in brain
tumours. Increasing data indicates the correlation between the CD133+ cells in
primary tumours and clinical outcome. Singh et al. demonstrated that a
subpopulation of CD133+ cells possesses stem cell properties in vitro. Meanwhile,
only the CD133+ brain tumour fraction was reported to initiate tumours in
NOD/SCID mice (Singh, SK et al. 2004). In clinical studies, an increased CD133
expression was reported to be related to a higher risk of tumour relapse in patients
with the World Health Organisation (WHO) grade 2 and grade 3 gliomas
(Zeppernick et al. 2008). Moreover, as few as 100 CD133+ cells could lead to the
initiation of tumours in NOD/SCID mice while 100-times more CD133- cells were
unable to propagate a tumour in vivo (Das, Srikanth & Kessler 2008). Nestin, an
intermediate filament (IF) protein, has been reported as another marker for brain
CSCs. The CD133+/Nestin+ cells showed self-renewal and tumour initiation
capacities (Ishiwata, Toshiyuki., Matsuda, Yoko. & Naito, Zenya. 2011). A2B5,
which is a cell surface ganglioside that marks neural precursor cells in the adult
human brain, was also demonstrated as another marker of brain CSCs (Gilbert &
Ross 2009). Tchoghandjian et al. reported that expression of A2B5 played a role in
the initiation and maintenance of brain tumours (Tchoghandjian et al. 2010).
Additionally, reports have shown that both A2B5+/CD133- and A2B5+/CD133+ cells
31
had a higher ability to form tumours than A2B5-/CD133- cells (Ogden et al. 2008).
CD15, also known as stage-specific embryonic antigen 1 (SSEA-1), has been
demonstrated as an additional CSC marker for brain tumours. Cells isolated from
glioblastomas with CD15+ phenotype displayed stem cell properties (Kim, K-J et al.
2011).
1.1.2.4.3 Colon cancer
Colon CSCs were initially identified by CD133 expression. However, the use of
CD133 as a marker for identification and isolation for colon CSCs has been debated
as it can be expressed in many other tumours, such as brain tumours, kidney tumours
and prostate tumours, and epithelial cells of normal tissues (LaBarge & Bissell 2008;
Todarolow et al. 2010). Thus, cell surface markers other than CD133 have been
studied. For instance, cells positive for EpCAM have been reported to have an
engraftment property in NOD/SCID mice (Dalerba et al. 2007). Moreover, ALDH1
has been identified as a marker to detect CSCs in colon cancer (Deng et al. 2010).
Meanwhile, CD24, CD29, CD44 and CD166 were reported to enrich for colon
cancer CSCs (Abdul Khalek, Gallicano & Mishra 2010), and
EpCAMhigh/CD44+/CD166+ was identified as a phenotype of tumourigenic cells in
colorectal cancer (Dalerba et al. 2007). Cells sorted by fluorescence-activated cell
sorting (FACS) indicated that CD44 and CD24 can also be used as cell surface
markers of CSCs in the colorectal carcinoma cell line SW1222. SW1222 cells with a
CD44+/CD24+ phenotype have the ability to initate tumours in vivo (Yeung, Trevor.
M. et al. 2010). As well, Vermeulen et al. found that cells coexpressing CD24 and
CD133 had higher clonogenic potential. Meanwhile, cells with CD24 and CD29
32
expression have been defined as the colorectal CSC population (Vermeulen et al.
2008).
In summary, as CSCs seems to be associated with mechanisms protecting tumour
cells from standard therapy, this minority subpopulation of cells could be used to
assist in the design of new and more effective anticancer therapies. For example,
disulfiram, an inhibitor of the cellular surface marker of CSCs ALDH, has been
reported to inhibit the mammosphere formation of breast cancer cells effectively in a
copper (Cu)-dependent manner (Yip et al. 2011).
1.1.2.5 Targeted anticancer therapy
Traditional anticancer therapy is based on the inhibition of cell division. Thus,
rapidly dividing cells, such as cells of the hair, gastrointestinal epithelium and bone
marrow, could be affected by drugs during treatment (Gerber 2008). In contrast,
targeted therapy is an approach which aims to block the proliferation of tumour cells
only. For example, targeted therapy can focus on the molecules which are mutant or
overexpressed in tumours. Antibodies against the cell surface markers cluster of
differentiation 20 (CD20) were utilised as one of the earlist targeted therapies in the
treatment of autoimmune disease, such as non-Hodgkin’s lymphoma (Feugier et al.
2005).
Targeted therapy includes various direct and indirect approaches. In direct
approaches, monoclonal antibodies (mAbs) or other small molecules which could
interfere with targeted proteins are used to target tumour antigens that have the
potential to alter signalling pathways; while in indirect approaches, ligands
containing a variety of effector molecules are utilised to target the antigens expressed
33
on the cell surface (Wu, H-C, Chang & Huang 2006). Thus, anticancer
chemotheraputics can target tumour cells via tumour-specific receptor-ligand binding
of mAbs or other small molecules.
1.1.2.5.1 Antibody-targeted therapy
Antibodies, also known as “targeting missiles”, play a critical role in targeted
therapy. The US Food and Drug Administration (FDA) approved muromonab-CD3
(Orthoclone OKT3) as the first monoclonal antibody to prevent acute organ rejection
after transplantation by blocking T-cell function in 1986. Since then, another 20
mAbs have been approved, and approximately half of them are utilised for anticancer
treatment (Gerber 2008). The mechanisms underlying anticancer effects of mAbs
vary, such as recruiting host immune functions to attack target cells, or carrying a
lethal payload which is toxic to the targeted cells (Adams, GP & Weiner 2005). As
their protein structures are denatured in the gastrointestinal tract, mAbs are usually
administrated via i.v. injection. To minimise the undesirable effects of the
hypersensitivity reaction to the mouse proteins in humans, more and more mAbs
used now are humanised, having an greater proportion of human components (Carter
2001).
1.1.2.5.2 Small molecule inhibitors
Extensive investigations have indicated that mutation or overexpression of protein
phosphorylation-related pathways is often found in tumour cells (Freedman et al.
2002; Lammering et al. 2001). Thus, targeted anticancer therapy could be designed
to be mono-specific to avoid side effects to normal cells. Therefore, as well as mAbs,
small molecule inhibitors which could interrupt aberrant cellular signaling by
34
interfering with key protein kinases involved, are also utilised for the targeted
anticancer therapy.
Unlike mAbs, small molecule inhibitors are chemically manufactured which is less
expensive than antibodies (Tanner 2005). However, due to the multitargeting nature
of the inhibitors, they are often less specific compared to mAbs (Imai & Takaoka
2006). Moreover, the half lives of most small molecule inhibitors are only a few
hours whereas mAbs have much longer half lives from days to weeks (Kantarjian, H
et al. 2002).
Most small molecule inhibitors are used to target plasma membrane-associated
protein tyrosine kinases (Cohen 1999). The first inhibitor was developed in the early
1980s and more than 30 agents are now in clinical trials (Dancey & Sausville 2003).
Imatinib, one of the first small molecule inhibitors, was approved in 2001 for the
treatment of chronic myeloid leukemia (Kantarjian, HM et al. 2003). Meanwhile,
small molecule inhibitors that target the epidermal growth factor receptor (EGFR)
pathway are utilised for solid cancer treatment, including non-small cell lung cancer
(Gerber 2008).
1.1.2.5.3. Aptamer-targeted therapy
Although antibodies have been widely used for active targeting in anticancer
treatments, the long clearance half-time and the incomplete penetration into target
tissues limit their use for imaging and therapy in the clinic (Hicke, Brian J. et al.
2005). Thus, new targeting agents are needed.
35
Aptamers, single-stranded or double-stranded oligonucleotides selected via an in
vitro process of systematic evolution of ligands by exponential enrichment (SELEX,
Figure 1-4), can be considered as nucleic acid analogues of antibodies (Lee, JH et al.
2010). Compared to traditional antibodies, aptamers have strong affinities and high
specificities to their targets. Moreover, aptamers can bind to various non-nucleic acid
targets (Edwards & Baeumner 2007), such as ions (Ciesiolka, Gorski & Yarus 1995),
small molecules, proteins (Latham, Johnson & Toole 1994) or bacteria. In addition,
aptamers are simple to synthesis (Ellington & Szostak 1990; Tuerk & Gold 1990)
with reduced batch-to-batch variability (Floege et al. 1999; Thiel & Giangrande 2010)
and a lack of immunogenicity compared to antibodies (Wang, P, Yang, et al. 2011a).
Furthermore, the affinity and specificity of aptamers can be further optimised via
post-selection engineering (Floege et al. 1999; Keefe & Cload 2008). Notably,
aptamers have been well-documented to be quite stable even under extreme
conditions, such as extremely high or low pHs, which means they can be denatured
and renatured multiple times without significant loss of activity (Liss et al. 2002).
Furthermore, aptamers are capable of greater specificity and affinity compared to
antibodies (Willner & Zayats 2007). Thus, aptamers have the potential to greatly
facilitate the development of novel anticancer applications. As CSCs have become
the focus of anticancer treatments nowadays, new aptamers have been developed for
targeting CSCs. For example, Shigdar et al. has developed the RNA aptamer against
a CSCs marker EpCAM and it was demonstrated to be internalised through receptor-
mediated endocytosis efficiently upon binding cancer cells (Shigdar et al. 2011).
36
Figure 1-4. SELEX technology (adapted from Esposito, Catuogno et al. 2011). The initial step of SELEX procedure is the synthesis of a single-stranded nucleic acid (RNA/DNA) library of large sequence complexity. Then, the folded oligonucleotides are selected by the capability of binding with high affinity and specificity to target molecules. Unbound oligonucleotides are removed by several stringent washing steps of the binding complexes. The target-bound oligonucleotides are eluted and subsequently amplified by the polymerase chain reaction (PCR) or reverse transcription-polymerase chain reaction (RT-PCR). The newly generated nucleic acid pool serves as a starting library for a new SELEX cycle. The number of SELEX repetitions depends on the library type used and on specific enrichment achieved per selection cycle and 6 to 20 SELEX rounds are usually needed for the selection of high affinity, target-specific aptamers. After the last round of aptamer selection, the PCR products are cloned and sequenced. (Esposito et al. 2011)
1.2. Cancer nanotechnology
Traditional anticancer chemotherapy is characterised by poor functional specificity
for tumour cells. As a result, anticancer drugs generally have relatively low
therapeutic effects as well as serious side effects, such as bone marrow suppression,
gastric erosion, hair loss, renal toxicity and cardiomyopathy, etc. (Myc et al. 2010;
Surendiran et al. 2009). Hence, one of the main goals of modern cancer
chemotherapy is to develop a safe and effective drug carrier that could act
37
preferentially at the selected targets while leaving normal tissues intact (Gullotti &
Yeo 2009). To improve the efficacy of drug delivery, nanotechnology-based drug
delivery systems are gaining considerable attention as they have the potential to
reduce side effects, minimise toxicity and improve anticancer efficacy (Suri, Fenniri
& Singh 2007).
Nanotechnology is the engineering and manufacturing of nano-sized molecules
using atomic or molecular components, which can be expected to benefit all fields of
medicine (Alexis, Frank, Rhee, June-Wha, et al. 2008). It is considered to be an
emerging technology that will have a significant impact on the advancement of
tumour therapies (Kim, KY 2007). Novel advances in drug delivery and formulations
to tumours are utilising nanotechnology that has been applied to overcome some of
the problems associated with traditional chemotherapeutic treatment of tumours
(Bawarski et al. 2008; Minko et al. 2006). Various nanotechnology platforms such as
dendrimers, liposomes, polymeric micelles and solid nanoparticles are being
developed (Bawarski et al. 2008; Surendiran et al. 2009). The utilisation of
nanoparticles as carriers of conventional chemotherapeutic agents has been shown
experimentally to improve cancer treatment efficacy (Ma et al. 2009).
Tumour tissues differ from healthy tissues in several aspects, with researchers
exploiting these differences when developing improved anticancer agents. The
defective vasculature and impaired lymphatic system leads to the EPR effect,
resulting in selective extravasation and accumulation of macromolecules in tumours
(Gullotti & Yeo 2009). The nanoparticle delivery system, such as nanoshells,
polymer micelles and liposomes, is based on the EPR effect, enhancing drug efficacy
and reducing non-specific toxicity as a reduction in systemic exposure (Ma et al.
38
2009; Maeda, Hiroshi 2001). Indeed, according to extensive studies, the
concentration of polymer drugs in the tumour could be 10 to 100 fold higher than
that of the free drug (Ganta, Srinivas et al. 2008). To take advantage of the EPR
effect, nanoparticles with passive targeting properties are utilised. Most fenestrate
openings in tumours are below 150 nm, so the nanocarriers can be made with a size
smaller than 150 nm for efficient extravasation. However, the filtration ability of the
kidney can filter particles smaller than 10 nm (about 70,000 Da), while particles
larger than 100 nm could be captured by the liver. Therefore, the ideal size of
nanocarriers for drug delivery is 10 to 100 nm as particles with such sizes minimise
extravasation in normal tissues but allow efficient accumulation at the tumour site
(Gullotti & Yeo 2009).
When nanoparticles are i.v. administered, they often bind with serum proteins and
are recognised by the scavenger receptor on the surface of macrophages, which leads
to a significant loss of nanoparticles in the circulation (Li, S-D & Huang 2008). The
proteins binding to the nanoparticles are termed “opsonins”, and the system
contributing most to the loss of injected nanoparticles is known as the
reticuloendothelial system (RES) or mononuclear phagocyte system (Drummond, D.
C. et al. 1999b). Thus, it is also critical to prevent nanoparticle uptake by the RES
(Drummond, D. C. et al. 1999b), which can destroy foreign materials by opsonisation
followed by phagocytosis (Gullotti & Yeo 2009). A widely used strategy for
shielding nanocarriers is to coat the surface of nanocarriers with a hydrophilic
polymer or glycolipid, such as polyethylene glycol (PEG) (Drummond, D. C. et al.
1999b; Gullotti & Yeo 2009; Yan et al. 2009). These flexible chained polymers or
glycolipids can occupy the sites on the surface of nanocarriers that would otherwise
bind with plasma opsonins from the RES macrophages (Drummond, D. C. et al.
39
1999b; Newton 2006; Remauta et al. 2007). Thus, the prolonged circulation time of
nanoparticles in the bloodstream increases the ability of drugs to reach their sites of
action, thereby improving the efficacy and safety as compared with conventional
preparations (Ma et al. 2009; Surendiran et al. 2009).
1.2.1 Nanoparticle therapeutics for cancer
Anticancer nanoparticle therapeutics typically utilise particles composed of
therapeutic agents that are assembled with nanomaterials, such as lipids, dendrimers,
and inorganic semiconductor nanocrystals (Davis, Chen & Shin 2008; Ganta,
Srinivas et al. 2008). The resulting nanoparticles may have functional properties that
are conferred through the assembly of individual components but not by the
individual components alone (Alexis, Frank , Rhee, June-Wha , et al. 2008). Studies
indicate that nanoparticles used in cancer chemotherapeutics possess more specific
targeting ability to tumour cells. The enhanced drug absorption and bioavailability of
anticancer agents results in an improved anticancer efficacy and decreased side
effects (Davis, Chen & Shin 2008; Li, S-D & Huang 2008). Some nanoformulations
of traditional anticancer agents have already been approved for use in the clinic. For
example, Doxil® and Abraxane® are two nanoformulations approved by the US FDA
for cancer treatment. Doxil® is a long circulating formulation of doxorubicin (DOX),
which has been well known for significant improvements over free DOX. Abraxane®
is an albumin-bound nanoparticle formulation of paclitaxel for the treatment of
metastatic breast cancer, reducing the hypersensitivity reaction associated with the
traditionally used solvent for paclitaxel (Bharali et al. 2009). Other nanocarriers,
such as co-polymer poly (lactic-co-glycolic acid), have also been approved by the US
FDA for the use of drug delivery, diagnostics and other applications of clinical and
40
basic science research, including cancer (Lü et al. 2009). The enhanced anticancer
effects and improved safety compared with their free-drug counterparts are due to the
size and surface properties of the nanoparticles.
1.2.1.1 Size
To protect the human body from invasion by foreign particles, the immune system
and other mucosal barriers function as biological barriers (Alexis, Frank , Pridgen,
Eric , et al. 2008). Anticancer nanoparticles should have the potential to bypass these
barriers to reach their target. The unique size of nanoparticles contributes to
overcoming these barriers as well as the accumulation in tumours. It is widely
accepted that the diameter of nanoparticles for cancer treatment should be in the
range of 10 to 100 nm (Davis, Chen & Shin 2008; Gullotti & Yeo 2009). The lower
limit is based on renal filtration, where particles smaller than 10 nm are rapidly
cleared by the kidneys. The upper limit is based upon pore sizes in the liver and the
spleen. The diameter of fenestrae sizes in the liver is 50-100 nm, while the size can
be up to 150 nm in the spleen (Alexis, Frank , Pridgen, Eric , et al. 2008; Campbell
2006; Drummond, D. C. et al. 1999b). Therefore, nanoparticles of 10 - 100 nm in
size which can penetrate the tumour vasculature will still be accumulated in the liver
and spleen which are two other accumulation sites of nanoparticles in addition to
tumours (Davis, Chen & Shin 2008; Drummond, D. C. et al. 1999b) and are removed
mainly by macrophages residing in these two organs. Thus, if the nanocarriers can
avoid being removed by macrophages, they are free to pass in and out of the liver
and spleen and have more opportunities to get into the tumours (Drummond, D. C. et
al. 1999b). One of the classical approaches to avoid particle clearance of
macrophages is to administer large doses of placebo carriers to impair the phagocytic
41
capacity of macrophages, thereby allowing subsequently administered drug-
encapsulated nanoparticles to remain in systemic circulation for prolonged periods or
to reach designated targets (Moghimi & Davis 1994). Other approaches, including
modification of the surface of nanocarriers, have been introduced to prolong the
circulation life (Fang, Hu & Zhang 2012).
1.2.1.2 Surface properties of nanoparticles
Nanoparticles have higher surface-to-volume ratio than larger particles. As a result,
control of nanoparticle surface properties is essential to control their behaviour
(Davis, Chen & Shin 2008). Key nanoformulation properties that are important for
drug delivery are biocompatibility and biodegradability, ensuring that unloaded
carriers are human degradable or metabolised into non-toxic components and cleared
by the body (Malam, Loizidou & Seifalian 2009). The rapid loss of nanoparticles
from the circulation before reaching their targets can affect the efficacy of the
nanoparticle therapeutics. To minimise the loss, strategies have been devised not
only with regard to optimising their size, but also surface modification of
nanoparticles. To improve tumour targeting, the surface of nanoparticles can be
modified in different ways, with both passive and active targeting strategies are
utilised to modify the surface of nanoparticles.
1.2.2.2.1 Passive targeting properties
Passive targeting strategies rely on the passive accumulation of nanoparticles due
to the properties of the delivery system and/or the pathophysiology of target tissues.
For example, retention due to the increased size of nanoparticles is one approach of
passive targeting, taking advantage of the EPR effect (Hoffman 2008). Furthermore,
42
nanoparticles that are sterically stabilised by other polymers, such as PEG, on their
surface can also increase the circulation time of nanoparticles in the bloodstream and
protect the nanoparticles from the RES, a property that is critical for nanoparticles to
escape from rapid clearance in the circulation (Drummond, D. C. et al. 1999b). In
addition, it has been well-documented that the physicochemical characteristics of
nanoparticles such as surface charge or functional groups can affect its uptake by
cells of the phagocytic system, resulting in a prolonged half-life in the bloodstream
(Gabizon, A, Shmeeda & Barenholz 2003; Ogawara, K-i et al. 2009; Yan et al. 2009).
A slightly charged surface, either positive or negative, on the nanoparticle can
protect the nanoparticles from undesired aggregation due to a reduction in self-to-self
or self-to-non-self interactions (Davis, Chen & Shin 2008; El-Aneed 2003). However,
renal filtration has the ability to filter positively charged particles (Gullotti & Yeo
2009). On the other hand, blood vessels repel negatively charged nanoparticles
(Davis, Chen & Shin 2008). Therefore, nanoparticles carrying a positively charged
surface are expected to have a high non-speci c internalisation rate and short blood
circulation half-life whereas negatively charged particles would perform better when
delivering drugs into tissues as they diffused at a faster rate (Kim, B et al.
2010). Taken together, control of surface charge will help to prevent the loss of
nanocarriers. The pH-sensitive nanoparticle is another example of passive targeting
nanocarriers. Tumour cells often proliferate rapidly. As a result, it is difficult for the
expanding tumour cell population to obtain sufficient oxygen and nutrients from the
tumour vasculature. Regions distant from vessels have slow cellular growth,
substantial cell death, and necrosis due to insufficient oxygen and nutrient supply
from vessels (Figure 1-5) (Campbell 2006). Moreover, lack of oxygen leads to
production of CO2, carbonic acid and lactic acid, resulting in low interstitial pH,
another characteristic of the tumour microenvironment (Schornack & Gillies 2003).
43
As a result, the pH around tumour tissues could be 5.5 to 6.5 (Ganta, Srinivas et al.
2008). Therefore, pH-responsive nanocarriers have significant enhanced stability in
circulation but release their payload (chemotherapy drugs) in the tumour. For
example, a 5-fluorouracil-loaded pH-responsive dendrimer nanocarrier showed a
significant faster release rate at pH 6.5 than at pH 7.4, resulting in a longer half-life
after i.v. administration and high tumour targeting in mice (Jin et al. 2011).
Figure 1-5. The tumour microenvironment (adapted from Minchinton and Tannock 2006). Tumour cells generally have accelerated proliferation which is dependent on the supply of oxygen and nutrients. Therefore, cells surrounding blood vessels proliferate actively while those away from blood vessels inevitably die. (Minchinton & Tannock 2006)
1.2.2.2.2 Active targeting
Passive targeting approaches, however, have several limitations and nanoparticle
targeting is ubiquitous and non-specific, leading to inefficient diffusion and
difficulties in the control of the drug delivery process. Moreover, certain tumours
lack the EPR effect which is a prerequisite for passive targeting (Peer, Dan et al.
2007). Functionalising the surface of nanoparticles with ligands such as antibodies,
aptamers, peptides, or small molecules that are tumour-speci c or tumour-associated
44
can promote the active binding of nanoparticles to tumours (Alexis, Frank, Rhee,
June-Wha, et al. 2008).
Recently, due to a better understanding of cancer and the revolution in medicine,
targeted therapies in cancer has been developed (Malinowsky et al. 2010). MAbs and
small molecule inhibitors are two major types of targeted therapy that can lead to the
selective inhibition of proliferation of cancer cells rather than normal rapidly
dividing cells such as hair, gastrointestinal epithelium and bone marrow (Gerber
2008). The FDA has already approved several mAbs for cancer treatment. For
example, Alemtuzumab, targeting CD52, was approved for chronic lymphocytic
leukemia successfully (Gerber 2008). Signaling pathways in tumours now become
the focus of development of targeted anticancer therapies and kinases within such
pathways are of specific intesest. For example, cetuximab, which targets the
epidermal growth factor receptor (EGFR), has been used for the treatment of
metastatic colorectal cancer in clinic (Dietel & Sers 2006). In a word, mAbs have
been extensively used for molecular targeting as its high affinity, specificity as well
as a variety of targets (Hicke, Brian J. et al. 2005; Syed & Pervaiz 2010).
Thus, combining nanoparticles with small molecules which could target cancer
cells or CSCs could improve the targeting property of nanoparticles as they can
recognise and bind to target cells via ligand-receptor interactions. Importantly, to
maximise specificity in binding, the antigen or receptor on the target cells should be
overexpressed in cancer. For example, nanoparticles with folic acid can target
tumour cells overexpressing folic receptor (Goren, D. et al. 2000). The poly[N-(2-
hydroxypropyl) methacrylamide] copolymer containing DOX, known as PK2, was
used to target the asialoglycoprotein receptor which is highly expressed on primary
45
liver cancer cells (Julyan et al. 1999). Anti-Her2-targeted immunoliposomal DOX
was designed based on the overexpression of HER2 receptors and study has
indicated that the therapeutic efficacy of the targeted liposomal formulation was
superior to its non-targeted counterpart for breast cancer (Park, JW et al. 2002).
Furthermore, binding of certain ligands to their receptors can result in receptor-
mediated internalisation, while binding of some other ligands can lead to non-
internalisation. The former approach can contribute to the release of agents inside the
cells. For example, immunoliposomes with an internalizing anti-CD19 ligand had a
significantly improved therapeutic efficacy than their counterpart with a non-
internalizing anti-CD20 ligand (Sapra, P. & Allen 2002). On the contrary, the latter
approach has the advantage of killing tumours by releasing drugs outside the tumour
cell, followed by passive diffusion into the target cells (Figure 1-6) (Allen, T. M.
1994a, 2002).
Figure 1-6. Active cellular targeting of nanoparticles (adapted from Peer, Karp et al. 2007). Binding of certain ligands on nanocarriers to their receptors can result in (i) Releasing their contents in close proximities to the target cells; (ii) Attaching to the membrane of the cell and acting as an extracellular sustained-release drug depot, or (iii) Internalising into the
cell. (Peer, Dan et al. 2007)
46
Overall, surface functionality is a critical parameter in the development of
nanoparticles. The minimisation of the loss of nanoparticles via sterical stabilisation,
control of surface charge and modification with ligands would enhance their
diagnostic and therapeutic potential (Yang, C et al. 2008).
1.3. Liposomes
Lipids are functional components of the biomembrane. Liposomes are created by
self-assembly of lipids, which enclose an aqueous space (Figure 1-7) (Bawarski et al.
2008; Hafez & Cullis 2001; Kshirsagar et al. 2005). Liposome research into drug
encapsulation and delivery has been carried out since the 1970s. Liposomes have
been used for anticancer nanoparticles since the mid-1990s (Davis, Chen & Shin
2008; Kim, KY 2007) and liposome-derived technology is established as one of the
cornerstones of nanotechnology in various fields, including medicine (Jesorka &
Orwar 2008). Several liposomal formulations, such as liposomal formulation of
DOX (Doxil®) and daunorubicin (DaunoXome®), have received clinical approval
due to their improved pharmacokinetics (PK) and pharmacodynamics of anticancer
agents. Other liposomal anticancer drugs are currently being studied in clinical trials
(Allen, T. M. & Cullis 2004).
47
Figure 1-7. Liposomes (adapted from Davis, Chen et al. 2008). Liposomes are self-assembling colloid structures composed of lipid bilayers surrounding an aqueous compartment, and can encapsulate a wide variety of therapeutic agents both in the internal aqueous space and/or in the lipid bilayer. (Davis, Chen & Shin 2008)
1.3.1 History and development of liposomes for anticancer drug delivery
Liposomes were discovered in 1965 (Bangham, Standish & Watkins 1965). As
liposomes are bilayers of lipids, they were widely used as a model of cellular
membranes initially (Metteucci & Thrall 2000). However, soon after their initial
discovery, they were recognised as novel drug carriers for diagnostic and therapeutic
agents (Phillips, WT, Goins & Bao 2009; Wiechers 2005).
Stability is a critical factor that should be taken into consideration for the design
and development of liposome-based drug delivery systems. The chemical stability of
lipids and drugs, the physical or colloidal stability of liposomes and the drug release
properties contribute to the overall stability of liposome-based formulations of drugs
(Domb et al. 2007). The classic liposomes, also known as conventional liposomes,
might be cleared by the RES quickly from circulation within minutes to hours after i.
v. administration (Domb et al. 2007; Metteucci & Thrall 2000; Oku & Namba 1994).
Therefore, due to the poor performance of conventional liposomes in vivo, they have
limited clinical applications. As a result, reduced opsonisation of liposomes is
48
expected to prolong their circulatory half-life in the bloodstream and increase the
possibility of accumulation in tumour tissue due to the EPR effect (Fenske & Cullis
2005). As opsonins and RES play important roles in the stability of nanoparticles,
minimising the binding of opsonins to nanocarriers is critical to developing long
circulating nanocarriers (Lasic, Danilo D. & Needham 1995; Ogawara, K-i et al.
2009). To avoid RES trapping of liposomes, the approach to surface modification of
liposomes to escape fast removal from the circulation has been considered (Hao et al.
2005; Oku & Namba 2005). Improvement, including the reduction in diameter and
modification of the surface of conventional liposomes had been introduced to
prolong the liposomes circulation life. One of the most important breakthroughs in
liposome-based formulations was the discovery of a PEG coating (Metteucci &
Thrall 2000). PEG is a semi-crystalline, hydrophilic polyether with a high solubility
in water and other aqueous media (Allen, T. M. et al. 1991; Senior et al. 1991). PEG
can slow down liposome recognition by opsonins via the formation of a protective
layer over the liposome surface, leading to decreased clearance of liposomes
(Gabizon, AA 2001b). As a result, PEG-coated liposomes generally have a longer
circulation half-life in the bloodstream than conventional liposomes. Numerous
studies have demonstrated that PEG-coated liposomes could accumulate in the
cancer via the EPR effect avoiding the fast clearance in the RES (Cattel, Ceruti &
Dosio 2003; Dos Santos et al. 2005; Ogawara, K et al. 2008). Thus, they are known
as sterically stabilised liposomes (Domb et al. 2007; Metteucci & Thrall 2000;
Moreira et al. 2002). The sterically stabilised liposomes also have the potential to
improve the efficacy of chemotherapy for cancer and reduce side effects (Song, CK
et al. 2009). This is best illustrated by Doxil®, a liposomal formulation of DOX.
DOX is an anthracycline anticancer drug. Although it is widely used in the clinic, the
cardiotoxicity of this drug limits its clinic utility. When a cumulative DOX dose
49
reaches 450 to 500 mg/m2, congestive heart failure and death often occur. The
advantages of Doxil® are greater efficacy and lower cardiotoxicity when compared
with free drug (Goyal et al. 2005; Hofheinz et al. 2005; Malam, Loizidou & Seifalian
2009). It is widely used to treat cancer in AIDS-related Kaposi sarcoma and multiple
myeloma in clinic (Ning et al. 2007).
Additional surface functionalities of liposomes have also been utilised.
Conjugating targeting molecules, such as antibodies, ligands, peptides, or aptamers,
to the surface of liposomes leads to active targeting liposome formulations (Gullotti
& Yeo 2009; Nallamothu et al. 2006; Newton 2006). The surface of circulating cells
such as lymphocytes, various antigens like p185HER-2, and low molecular weight
mAbs have also been the subject of research for targeting liposome formulations
(Domb et al. 2007). Tumour targeting agents can be attached to the liposome
phospholipid headgroups or to the end of a PEG derivative, and the latter approach
has proven to be more effective due to better accessibility of the antibody to its target
(Domb et al. 2007; Newton 2006). Several mAbs have been shown to deliver
liposomes to their targets (An et al. 2008; Lapalombella et al. 2008). For example,
when liposomal formulations of topotecan, vinorelbine, and DOX were attached with
the anti-CD166 single chain variable fragment, they showed efficient and targeted
uptake by prostate cancer cell lines, Du-145, PC3, and LNCaP (Roth et al. 2007).
Recent research on ligands for liposomes has focused on specific receptors
overexpressed on target cells or certain specific components of pathological cells.
For example, folic receptor is frequently overexpressed in a range of tumour cells.
The folate-targeted liposomes have already demonstrated their therapeutic potential
for murine lung carcinoma and human epidermoid KB carcinomas (Gabizon, A et al.
2003; Lu & Low 2002). Epidermal growth factor receptor-targeted liposomes were
50
shown to have improved efficacy of encapsulated DOX, epirubicin and vinorelbine
in vivo (Mamot, C. et al. 2005; Molema 2005).
1.3.2 Characterisation of liposomes
Liposomes are composed of natural components such as pure lipids or lipid
mixtures. The most commonly utilised lipids are phospholipids, in particular the
neutrally charged phosphatidylcholine and the negatively charged phosphatidic acid,
such as phosphatidylglycerol, phosphatidylserine and phosphatidylethanolamine
(Jesorka & Orwar 2008; Lammers, Hennink & G 2008). The amphiphilic
phospholipid molecule has a structure containing both a polar region and nonpolar
aliphatic chains. When placed in an appropriate environment and at specific
concentrations, these lipids self-associate into a spherical structure with an aqueous
interior core. The hydrophilic polar groups face the core, while the hydrophobic
nonpolar portions are tucked into the interior of the membrane bilayer (Bawarski et
al. 2008; Phillips, WT, Goins & Bao 2009).
Liposomes can vary in the number of lipid layers they possess and therefore be
classified into several categories according to their size and lamellarity:
multilamellar vesicles, oligolamellar vesicles, unilamellar vesicles, giant unilamellar
vesicles, large unilamellar vesicles, medium-sized unilamellar vesicles, small
unilamellar vesicles and multivesicular vesicles (Figure 1-8) (Malam, Loizidou &
Seifalian 2009), and can range from 50 nm to several micrometres (Phillips, WT,
Goins & Bao 2009). Multilamellar vesicles bigger than 0.5 m are frequently used in
pharmaceutical and cosmetic applications, whereas small unilamellar vesicles (from
20 nm to 200 nm), large unilamellar vesicles (up to 1 m) and giant unilamellar
vesicles (larger than 1 m) are the three most important groups for analytical
51
applications (Domb et al. 2007; Jesorka & Orwar 2008). Furthermore, small
unilamellar vesicles are widely used for novel drug delivery carriers. However,
considering the loading efficiency, stability in the bloodstream and the ability of
extravasation, liposomes ranging from 50 to 150 nm are most promising for drug
delivery (Metteucci & Thrall 2000; Phillips, WT, Goins & Bao 2009).
Figure 1-8. Liposome classification (adapted from Jesorka and Orwar 2008). Liposomes can be classified into different categories, such as multilamellar vesicles, giant unilamellar vesicles, large unilamellar vesicles, small unilamellar vesicles and multivesicular vesicles.
(Jesorka & Orwar 2008)
As liposomes are composed of naturally biodegradable substances, they are
metabolised and cleared while in circulation or upon reaching the target sites, making
them safe novel drug delivery carriers (Domb et al. 2007). Both water-soluble and
lipid soluble drugs can be carried by liposomes due to its amphiphilic property. The
closed bilayer vesicles have two potential compartments to encapsulate substances.
The hydrophilic substance can be entrapped by the hydrophilic aqueous core while
lipid soluble substances can be trapped by the hydrophobic interior of the membrane
(Samad, Abdus, Sultana, Y. & Aqil, M. 2007). Moreover, various methods can be
used to encapsulate substances. Depending on molecular weight, solubility and
polarity, substances can be trapped either during liposome assembly or loaded into
52
preformed liposomes by different methods. For example, drug loading into
liposomes can be achieved via liposome formation in an aqueous solution saturated
with soluble drug, via solvent exchange using organic solvents, via the use of
lipophilic drugs, and via gradient loading methods (Qiu, Jing & Jin 2008).
Amphipathic drugs such as DOX can be encapsulated into preformed liposomes with
an ammonium sulphate gradient method or a pH gradient while hydrophlic drugs can
be trapped into liposomes upon lipid hydration by passive loading (Domb et al. 2007;
Metteucci & Thrall 2000). Furthermore, the surface of liposomes can be modified by
different approaches. For example, PEGylation, the coating liposomes with PEG, or
ligand attachment are widely utilised to improve the stability and targeting of
liposomes.
1.3.3 Advantages of liposomes for anticancer drug delivery
Many current anticancer drugs have non-ideal properties, which can lead to
serious consequences, including side effects, lack of efficacy or poor quality of life
with high drug toxicity being a major barrier to the efficacy of anticancer drugs
(Allen, Theresa M. et al. 2006). Since the 1960s, liposomes as carriers for cancer
chemotherapeutic agents have been investigated in order to improve the efficacy of
anticancer drugs (Song, CK et al. 2009).
Liposomes are very versatile drug carriers. Pharmaceutical properties such as
encapsulation efficiencies of drugs and drug release rates can be manipulated. The
various sizes of liposomes contribute to a wide variety of options for utilisation
(Chrai, Murari & Ahmad 2002; Tran, Watts & Robertson 2009; Wiechers 2005;
Yang, C et al. 2008). Moreover, liposomes can carry abundant agents due to their
large payload. In addition, the outer lipid bilayer allows the incorporation of other
53
molecules, and therefore liposomes can target specific tissues by either active or
passive targeting modification (ElBayoumi, aA & Torchilin, VP 2009; Malam,
Loizidou & Seifalian 2009; Song, CK et al. 2009). Thus, different formulations of
liposomes can match different requirements in the chemotherapeutic treatment of
tumours.
As the size and surface properties of liposomes can be easily manipulated, one can
manufacture liposomes that have desired circulation times in the bloodstream. In
addition, the release kinetics of drugs in liposomes can also be modulated to match
different requirements (Davis, Chen & Shin 2008; Park, K 2007; Torchilin 2006).
Thus, the pharmacokinetic features as well as the targeting specificities of liposomes
can be controlled and modified to reduce the side effects of anticancer drugs and
enhance the efficacy.
1.3.3.1 Enhanced stability of liposomal drug formulation
Liposomes have a high therapeutic drug-to-carrier ratio relative to their diameter
(Allen, Theresa M. et al. 2006). The drug carried by the liposome (the payload),
particularly those encapsulated in the liposomes, can be substantially protected from
enzymatic degradation or inactivation in the plasma as well as unfavourable pH
which can lead to rapid breakdown of free drugs. As a result, when drugs are inside
liposomes, their half-life can be prolonged several hours or more due to enhanced
stability. Paclitaxel (Taxol), a strong antitumour agent, is a prime example. The low
aqueous solubility of paclitaxel and hypersensitivity reactions in patients to
Cremophor EL-containing paclitaxel range from mild pruritus to systemic
anaphylaxis and limit its clinical applications (Paál, Müller & Hegedûs 2001).
Studies showed that liposome formulations of Taxol enhanced the aqueous solubility
54
as well as the physicochemical stability (Yang, T, Cui, Choi, Lin, et al. 2007).
Another example is topotecan that is hydrolysed rapidly at physiological pH, limiting
its clinic applications. To improve the stability and antitumour activity of topotecan,
Drummond and colleagues encapsulated the drug in liposomal formulations with an
acidic internal pH (Drummond, D. C. et al. 2010; Hao et al. 2005).
1.3.3.2 A controlled release system of anticancer drugs
The rate of release of the payload in liposome formulations depends on drug
properties, liposome composition, the presence or absence of pH or osmotic
gradients, the liposome environment and the in vivo stability of the liposomes (Allen,
Theresa M. et al. 2006; Lamprecht , Bouligand & Benoit 2002). Arsenic trioxide can
be used for the treatment of acute promyelocytic leukemia and relapsed multiple
myeloma. However, its clinical utility has been limited by its toxicity at high doses.
Liposome encapsulated arsenic trioxide appeared to be stable in physiological
situations but released the drug in a lower pH environment, such as in tumour tissues
or in endosomes. As a result, the liposome formulation of arsenic trioxide has the
potential to decrease the toxicity of this powerful anticancer drug in healthy tissues
(Chen, H et al. 2006).
Moreover, the phospholipids at the surface of liposomes can be modulated to
provide desired drug release rates. For example, alteration of surface charge can not
only enhance drug incorporation but also influence the rate of drug release (Ranade
1989). As a result, liposomes can function as sustained release systems due to their
ability to control drug release rates and to protect assembled labile drugs. Sterically
stabilised liposomes, for example, usually have a sustained release of loaded
substances as well as prolonged circulation time, leading to the improvement of
55
bioavailability of the agents. A number of studies showed Doxil® targeted tumours
passively, and released DOX slowly in tumours once the liposomes localise in the
tumour interstitial space (Allen, Theresa M. et al. 2006; Cabanes et al. 1999;
Unezaki et al. 1994).
Furthermore, the release kinetics of agents from the liposomes can be modified by
other factors. One of the examples is pH-responsive liposomes that belong to a
specific-sensitive system (Kale & Torchilin 2010). The pH surrounding the tumour
ranging from 5.5 to 6.5, tends to be more acidic than physiological pH which is 7.4
(Trédan et al. 2007). As a result, utilising pH-responsive nanocarriers has significant
enhanced stability in circulation and accumulation in the tumour when compared
with non-pH-responsive carriers (Yu et al. 2004). Long circulation pH-responsive
liposomes can be designed by modifying sterically stabilised liposomes via insertion
of a hydrophobically modified N-isopropylacrylamide/methacrylic acid copolymer in
the lipid bilayer of sterically stabilised liposomes (Bertrand, Simard & Leroux 2010).
The pH-responsive liposomes with target-specificity are used to deliver molecules or
macromolecules into the cytoplasm (Mujumdar & Siegel 2008). By selecting the pH
and interliposomal buffer, pH-stimulated release of drugs entrapped in liposomes has
the potential to increase membrane permeability upon acidification (Ganta, Srinivas
et al. 2008). Studies regarding pH-sensitive immunoliposomes encapsulated with
cytosine arabinoside indicated that the pH-sensitive liposomal formulation could be
beneficial in the treatment of acute myeloid leukaemia (Roux et al. 2002; Simard &
Leroux 2009). Another example of the specific-sensitive systems is thermo-sensitive
liposomes that are composed of appropriate phospholipids with a phase transition
temperature around 40°C. The liposome formulation can release their contents by
selective application of hyperthermia (40-42°C). As a result, the release of drugs
56
from temperature-sensitive liposomes can be controlled via the use of the
hyperthermia. Studies on the therapeutic efficacy of DOX encapsulated temperature-
sensitive liposomes found that these liposomes could be highly efficacious carriers
for the in vivo delivery of anticancer drugs, and have potential anticancer
applications in combination with hyperthermia (Han et al. 2006; Hauck et al. 2006).
1.3.3.3 Reversal of MDR
Evidence is accumulating to indicate that liposomes have the potential to bypass
MDR, mediated by the MRP or P-gp, which is a key obstacle to cancer treatment
(Malam, Loizidou & Seifalian 2009). Targeted liposomes may overcome MDR as
they can bind to cell-surface receptors and are then endocytosed, leading to the
potential to bypass MDR conferred through cell-surface protein pump mechanisms
(Alexis, Frank, Rhee, June-Wha, et al. 2008; Malam, Loizidou & Seifalian 2009).
Previous studies have already indicated that liposomal DOX has anti-tumour effects
on both DOX-resistant and non-DOX-resistant mouse tumour C26 cancer models
(Malam, Loizidou & Seifalian 2009). Moreover, liposomes with specific antibodies
which blocked P-gp mediated efflux have been shown to increase intracellular drug
accumulation and enhanced cytotoxic effects of DOX on multidrug-resistant P388
cells (Gatouillat et al. 2007). Other liposomal anticancer drugs, such as verapamil,
had also illustrated the ability to overcome P-gp-mediated MDR phenotype in K562
leukaemia cells when conjugated to transferrin modified liposomes (Wu, J et al.
2007). Furthermore, pH-sensitive liposomes demonstrate greater antitumour activity
when compared with free drugs in a drug-resistant mouse model (Davis, Chen &
Shin 2008). Recently, liposomes conjugated to oligonucleotides and small interfering
RNA were effectively tested to overcome MDR in cancers expressing P-gp (Minko
57
et al. 2006). Finally, liposomes composed of certain phospholipids, such as
cardiolipin, have been shown to elevate the cytotoxicity of loaded anticancer agents
against resistant cells by either direct blocking of P-gp or increasing intracellular
drug concentration (Gokhale et al. 1996; Thierry et al. 1993).
1.3.3.4 Improved PK and BD
The most compelling property of liposomes is their ability to significantly alter the
PK and BD of many of their associated drugs. Encapsulation within liposomes
allows the drug to be bioavailable following their release from carriers at the tumour
site under optimal conditions. The liposomes protect drugs from metabolism and
inactivation in the circulation as well as from clearance due to the size of the free
drug (Li, S-D & Huang 2008). Furthermore, due to the size limitation of healthy
vasculature endothelium, the drug distribution in healthy tissues can be reduced, thus
drug accumulation in tumours can be increased. Investigations have demonstrated
that the accumulation of conventional liposome encapsulated DOX was higher in
tumours than that of free DOX in a mouse model (Drummond, D. C. et al. 1999b;
Ogawara, K-i et al. 2009). Moreover, drugs encapsulated in liposomes are not
bioavailable until they are released from the carriers to become free drugs (Allen,
Theresa M. et al. 2006). As long as they are released from the liposomes, the released
drugs function as free drugs. In turn, as the drugs carried by nanoparticles are located
within the particles, the amounts and types of the payload will not affect the PK and
BD of nanocarriers (Davis, Chen & Shin 2008). In addition, liposomes are high
molecular water-soluble nanoparticles. Drugs with low molecular weight
encapsulated by high molecular weight liposomes are inefficiently removed by
lymphatic drainage and therefore have more chance to accumulate in tumours
58
(Minko et al. 2006). Taken together, the liposome encapsulation approach has the
potential to improve treatment efficacy of various traditional anticancer agents.
One barrier for the early liposomal formulations is the rapid clearance in the
bloodstream. To avoid the trapping of liposomes by the RES, approaches to limiting
the size, or coating liposomes with a hydrophilic barrier to protect them from
recognition by the RES, have been considered. Sterically stabilised liposomes,
liposomes coated with PEG, have been shown to display an enhanced stability in the
circulation, a prolonged half-life and an improved pharmacokinetic profile over both
free drugs and conventional liposomes (Minko et al. 2006; Moreira et al. 2002). For
example, studies comparing different formulations of DOX illustrated that Doxil®
has an 87-fold higher area under time versus concentration, one important
pharmacokinetic parameter related to the efficacy of drugs, than free DOX seven
days post-injection. On the contrary, area under time versus concentration of
conventional liposomes with a more rapid DOX release rate was only 14-fold higher
than that of free DOX (Lasic, Danilo D. & Needham 1995; Minko et al. 2006).
PEGylated liposomal formulations of paclitaxel have also been reported to increase
the biological half-life of paclitaxel from 5.05±1.52 hours to 17.8±2.35 hours
compared to conventional liposomes in rats (Yang, T, Cui, Choi, Cho, et al. 2007).
Sterically stabilised liposomal formulation of CKD-602, a camptothecin analogue,
was shown to have more favourable pharmacokinetic parameters, such as a lower
clearance rate, both in plasma and tumour when compared with non-liposomal CKD-
602 (Zamboni et al. 2007).
The BD of anticancer agents can be improved by liposomal encapsulation. The
EPR effect is an important factor in the improvement of the BD of liposomal drug
59
formulations (Allen, Theresa M. et al. 2006). Due to their size, liposomes are thought
to extravasate more easily into the tumour vasculature due to increased gaps between
tumour endothelial cells. For example, free DOX has a wide distribution in vivo,
accumulating in most tissues to a significant extent, while studies regarding DOX-
loaded liposomes in vivo indicated a decreased drug concentration in the heart and
kidneys, but an increased concentration in tumours when compared with that of the
free DOX (Song, CK et al. 2009). This altered distribution reduced the concentration
of DOX at potential sites of toxicity, such as the heart.
1.3.3.5 Aptamer and CSCs targeting
Previous studies have conjugated aptamers with liposomes to achieve better
tumour targeting. For example, Mann et al. conjugated amino-PEG liposomes with a
thioated oligonucleotide aptamer (thioaptamer) against E-selectin (ESTA) (Mann,
Bhavane, Somasunderam, Liz Montalvo-Ortiz, et al. 2011; Mann et al. 2010). The
ESTA-liposomes showed an elevated tumour vasculature accumulation in nude mice
bearing MDA-MB-231 breast xenograft tumours. In the study using TMR-sgc8
aptamer-modified liposomes (Kang et al. 2010a), liposomes could bind specifically
to the target tumour cells in 30 min. Moreover, an RNA aptamer binding to the
prostate-specific membrane antigen (PSMA) using PLA-PEG-COOH nanoparticle
bioconjugates could target cell lines expressing PSMA protein with a 77-fold higher
cellular uptake than the control (Farokhzad et al. 2004b).
New aptamers can be developed that specifically targeting the CSCs. For example,
Shigdar et al. developed an RNA aptamer against the cancer stem cell marker
EpCAM (Shigdar et al. 2011). As the CSCs are responsible for chemoresistance and
recurrence after various treatments, the novel chemotherapeutic agents delivery
60
systems designed to target CSCs could prevent recurrence of tumours (Ambasta,
Sharma & Kumar 2011). An approach to target CSC markers is to combine
liposomes with aptamers to develop liposomes which have the potential to provide
an enhanced targeting efficiency as a selective delivery system. Thus, the
conjugation of aptamers and liposomes has the potential to improve the
chemotherapeutics efficacy loaded by liposomes.
1.4. Project significance and aims
As a novel drug delivery system for anticancer chemotherapeutics, liposomes have
been extensively studied in recent years. The liposomal formulation of anticancer
agents affords enhanced stability, improved targeting to tumours, controllable release
rate and more favourite PK and BD. However, challenges for a better drug delivery
system remain, particularly regarding targeted therapy. This project focuses on the
development and characterisation of DOX encapsulated sulfatide-containing
liposomes (SCL) that possess enhanced tumour targeting properties.
DOX is an anthracycline antibiotic widely used in oncology clinics as an
anticancer chemotherapeutic agent. It was firstly introduced in the 1970s, and has
now become one of the most commonly used anticancer drugs in the treatment for
haematological malignancies and solid tumours (Arola et al. 2000). However, when a
cumulative dose reaches 450 to 500 mg/m2, it may lead to cardiomyopathy, resulting
in congestive heart failure (Imondi et al. 1996; Olson et al. 1988). Moreover, due to
the excretion of DOX occurring predominantly by the hepatobiliary route and
partially by the kidneys (Cosan et al. 2008), recent data indicates that DOX could
induce renal damage (Bárdi et al. 2007) and decreased hepatic activity (Yokogawa et
al. 2006). As a result of ongoing efforts to develop a nanoparticle drug delivery
61
system, the liposomal formulations of DOX entering the clinic have improved
efficacy and reduced side effects (Elbayoumi, T. A. & Torchilin, V. P. 2008).
Nevertheless, the use of long-circulating formulations of liposomal DOX is limited
by its toxicity to the skin. Palmar-plantar erythrodysesthesia (PPE), also known as
hand-foot syndrome, is a relatively common dermatologic toxic reaction associated
with pegylated liposomal DOX (PLD), such as Doxil® (Farr, K. P. & Safwat, A.
2011).
The composition and structure of the extracellular matrix in tumours is different
from that in normal tissues (Liu, F et al. 2008). Certain extracellular matrix
glycoproteins are highly up-regulated in many different cancers, including breast
cancer, ovarian cancer, prostate cancer and gliomas (Fernando et al. 2008; Quemener
et al. 2007). Tenascin-C is a protein expressed at low levels in normal adult tissues
but high levels in many tumours (Mackie & Tucker 1999). The function of tenascin-
C is complex. It can influence the cell behaviour directly or indirectly. It has been
implicated in the regulation of cell migration, focal adhesion and cell proliferation
(Sage & Bornstein 1991; Trojanowska 2000). Moreover, expression of genes
modifying the extracellular matrix can be induced by tenascin-C. These genes have
the potential to promote tumour growth and invasion. Taken together, tenascin-C has
been implicated as an important target for the treatment of cancer (Adams, M et al.
2002).
A possible strategy to improve anticancer agent binding and penetration into
tumours is to modify the drug carriers with a composition that can bind to the tumour
cell surface or the extracellular matrix of tumours. Sulfatide, a lipid that is found in
humans, is involved in a variety of biological processes such as cell adhesion,
62
platelet aggregation, cell growth, protein trafficking, signal transduction, neuronal
plasticity and cell morphogenesis. Sulfatide is known to bind several extracellular
matrix glycoproteins including tenascin-C (Townson et al. 2007). Therefore,
sulfatide was chosen to be one of the components of nanoliposomes to target tumours
expressing high levels of tenascin-C in this project.
DOPE, 1,2-dioleoyl-snglycero-3-phosphoethanolamine, is a neutral lipid that is
also found in humans. It is an attractive lipid component of liposomes for
intracellular delivery of anticancer drugs. This is because DOPE is thought to be the
most superior of all the fusogenic (ability to promote membrane fusion) lipids (Hafez
& Cullis 2001). When mixed with another ionisable anionic lipid, DOPE adopts a
bilayer organisation (forming liposomes) at physiological pH due to the stabilising
effect of the helper lipids. At a lowered pH (where the pH values drops below the pK
of the helper lipid), the helper lipid is unable to stabilise the bilayer. As a
consequence, DOPE adopts a non-bilayer inverted hexagonal (HII) phase, leading to
membrane inversion, membrane fusion and the release of entrapped chemotherapy
drugs to the cytoplasm. Thus, under physiological pH, DOPE confers stability to
SCL via its inhibitory effects on liposome fusion, as the incorporation of sulfatide
into DOPE vesicles greatly enhances the stability of the liposomes formed, even in
the presence of plasma, presumably due to the hydration of the negatively charged
sulfate head-group of the glycosphingolipid (Shao, Ke et al. 2006). One of the key
aspirations in liposome research is to develop a liposome that is very stable in
physiological pH to confer a long half-life in the blood stream but quickly releases its
payload at lowered pH in the endosome once internalised. Previous research has
demonstrated that DOPE-sulfatide liposomes do enter the cells and release the
payload inside the tumour cells in vitro (Shao, Ke et al. 2006). It has been shown that
63
sulfatide was specifically required for robust uptake of liposomes by human
glioblastoma cells U-87MG. The SCL has been found to remain intact for hours after
uptake by the glioblastoma cells. As well, intracellular distribution studies have
indicated a high accumulation of DOX in the nuclei where it exerts its cytotoxic
effect after 12 h incubation with SCL at 37 °C. Meanwhile, therapeutic efficacy
studies of SCL in immunocompromised mice bearing human glioblastoma cell line
U-87MG displayed much improved therapeutic effects over the free drug (Shao, Ke
et al. 2006; Wu, X & Li 1999).
However, the in vitro and in vivo stability of the liposomes, the pharmacokinetic
behaviour and the BD pattern are yet to be established. The critical parameters of
tumour uptake, pharmacodynamic data of the sulfatide-containing liposomal DOX
(SCL-DOX) in tumour xenograft models and the safety profile remain to be
established. Moreover, due to various barriers that impede the efficient delivery of
drugs to the brain interstitium, the potential of SCL-DOX for circumventing the
blood-brain barrier (BBB) remains to be explored. Therefore, this project aims to
address these important outstanding issues in preclinical settings, especially in
human glioma and colorectal cancer models. The study also explores the active
targeting property of SCL surface functionalised by CSC targeting aptamers.
Therefore, the long term goal of this project is to develop a better drug delivery
system, and in particular, a targeted therapeutic. To achieve this goal, the current
project focused on the development and characterisation of DOX encapsulated SCL
that possesses enhanced tumour targeting properties. The specific aims of this project
are:
1) To further characterise SCL-DOX and to develop PEG-SCL-DOX.
64
2) To analyse cellular uptake and cytotoxicity of SCL-DOX in vitro.
3) To study the PK and BD of SCL-DOX in healthy Sprague Dawley (SD) rats.
4) To study tumour uptake of SCL-DOX in tumour-bearing
immunocompromised mice.
5) To study therapeutic efficacy and safety of SCL-DOX in xenograft tumour
models.
6) To study the potential of SCL-DOX in penetrating the BBB in healthy SD
rats.
7) To study tumour uptake of SCL-DOX in glioma-bearing Wistar rats.
8) To study cellular uptake and retention properties of CSCs targeting aptamer
functionalised PEG-SCL-DOX.
65
Chapter 2 - Methodology
66
2.1 Ethics Statement
The Deakin University Animal Welfare Committee has approved all animal
protocols used in this research.
2.2 Materials
2.2.1 Chemicals, reagents and equipment
The following companies supplied reagents and equipment used in this work:
Avanti Polar Lipids, In, Alabaster, Alabama
Abacus ALS, Brisbane, Australia
Auspep Pty Ltd, Tullamarine, Australia
Barwon Health Pharmacy, Geelong, Australia
BD, New Jersey, USA,
BD Medical, Scoresby, Australia
Bovogen Biologicals Pty Ltd, East Keilor, Australia
David Kopf Instruments, KOPF®, California, America
Grale Scientific Pty Ltd, Ringwood, Australia
Hitachi, Japan
HyClone, South Logan, Canada
IBA GmbH, Goettingen, Germany
Interpath Services, Melbourne, Australia
Instech Laboratories, Inc., America
Invitrogen Australia Pty Ltd, Mulgrave, Australia
67
Malvern Instruments Ltd, Worcestershire, UK
MatTek Corporation, Ashland, America
Medicago, Uppsala, Sweden
Merck, Kilsyth, Australia
Millipore, Billerica, MA, USA
NANOCS, Melbourne, Australia
Olympus, Tokyo, Japan
Pall Australia, Cheltenham, Australia
PerkinElmer, Melbourne, Australia
Promega Corporation, Madison,USA
Sigma-Aldrich Pty. Ltd, Sydney, Australia
Tecniplast S.p.A, Buguggiate, Italy
Thermo Fisher Scientific, Scoresby, Australia
ATA Scientific Pty Ltd, Taren Point, Australia
Waters, Milford Massachusetts, USA
Westlab, Madison , USA
All the organic reagents, including methanol and chloroform, used in this study
were analytical or high-performance liquid chromatography (HPLC) grade. All other
chemicals were of the highest grade available.
2.2.2 Cells used in this study
The U-118MG (human glioblastoma) cell line, HT-29 (human colorectal
adenocarcinoma) cell line, MCF-7 (human breast adenocarcinoma) cell line and C6
68
(rat glioma) cell line were purchased from American Type Culture Collection
(ATCC, Manassas, VA). HEK-293T (the human embryonic kidney 293) cell line
was a generous gift from Dr Hongjian Zhu (Melbourne University, Australia). U-
118MG, HEK-293T and C6 cells were cultured in DMEM medium (InvitrogenTM,
Australia) supplemented with 10% foetal bovine serum (FBS, Hyclone, Canada),
penicillin (50 U/mL, InvitrogenTM, Australia), and streptomycin (50 μg/mL,
InvitrogenTM, Australia) in a humidified atmosphere containing 5% CO2 at 37 °C.
HT-29 cells were cultured in McCoy's 5A medium (InvitrogenTM, Australia)
supplemented with 10% FBS, penicillin (50 U/mL), and streptomycin (50 μg/mL) in
a humidified atmosphere containing 5% CO2 at 37 °C.
For subculturing cells from monolayer, cell culture medium was removed and
cells were rinsed gently with sterilised phosphate saline buffer (PBS) twice, followed
by detachment with 0.05% trypsin-EDTA (InvitrogenTM, Australia) at 37 °C until
cells were rounded up and detached from the surface of flasks. Then, complete
medium was used to inactivate the trypsin. Cells were collected and centrifuged at
room temperature (RT), followed by resuspension with appropriate buffer or medium.
2.2.3 Animals used in this study
Animals were purchased from Deakin Upper Animal House, Monash University
and The Animal Resources Centre (Perth, Australia).
Male Sprague-Dawley (SD) rats (200 to 250 g) were housed in a temperature
controlled room (25 ± 1°C) with a 12-h light-dark cycle. Rats were fed ad libitum
with a standard diet and were fasted overnight before treatments administration.
69
Male Wistar rats (200 to 250 g) were housed in a temperature controlled room (25
± 1 °C) with a 12-h light-dark cycle. Rats were fed ad libitum with a standard diet
and were fasted overnight before treatment administration.
Six-week-old female BALB/c-Foxn1nu mice were used for U-118MG and HT-29
xenografts established. The mice were housed in the TECNIPLAST SealsafeTM
Individually Ventilated Cages which were placed in a temperature controlled room
(25 ± 1 °C) with a 12-h light-dark cycle. Mice were fed ad libitum with a standard
diet. Beddings, cages and water were autoclaved at 121 °C for 30 min while the
fodder was sterilised by ultraviolet irradiation before use.
2.2.4 Aptamers used in this study
Aptamers were synthesised by IBA GmbH.
EpCAM targeting aptamer: 5’-Amino-12 Carbon linker-G (2’-F-C) G A (2’-F-C)
(2’-F-U) G G (2’-F-U) (2’-F-U) A (2’-F-C) (2’-F-C) (2’-F-C) G G (2’-F-U) (2’-F-
C)-G-9-atom spacer- fluorescence-3’ (F: 2’-fluoro, fluorescence: FITC).
Negative control aptamer: 5’-Amino-12 Carbon linker--mG C mG A C U mG mG
U U mA C C C mG mG U C mG-9-atom spacer-fluorescence-3’ (m: 2’-O-methyl,
fluorescence: FITC).
Due to their susceptibility to degradation by serum nucleases, aptamers are
unsuitable for most therapeutic applications (Burmeister et al. 2006). Resistance to
nuclease-mediated degradation can be greatly improved by incorporating modifying
groups at the nucleotide 2’-position, and both 2’-fluoro and 2’-O-methyl
modifications have been shown to significantly increase resistance to serum
70
nucleases (Burmeister et al. 2005; Ruckman et al. 1998). Thus, the 2’-fluoro and 2’-
O-methyl modifications have been utilised for our aptamers to increase aptamer
stability both in vitro and in vivo.
However, the aptamer modified with 2’-O-methyl is an aptamer of the same
sequence as EpCAM targeting aptamer but with a different side-chain modification
which could affect the 3-dimensional structure of aptamer (Chushak & Stone 2009).
As a result, it is used as a negative control aptamer which is not able to bind to
EpCAM and does not specifically target the cancer cells with high expression of
EpCAM.
2.3 Methods
2.3.1 Development of sulfatide-containing nanoliposomal doxorubicin (SCL-
DOX) and PEG-sulfatide-containing nanoliposomal doxorubicin (PEG-SCL-
DOX)
2.3.1.1 Development of SCL-DOX.
Liposomes were prepared according to a previously published method with some
modifications (Shao, Ke et al. 2006). Briefly, DOPE unilamellar vesicles containing
30% (molar ratio) sulfatide were prepared by a hydration method followed by
polycarbonate membrane extrusion. DOPE (13.35 mM, Avanti Polar Lipids, Inc.,
Cat No: 850725P-500mg) and sulfatide (6 mM, Avanti Polar Lipids, Inc., Cat No:
131305P-100mg) were dissolved in a mixture of chloroform (Deakin University,
LES store, CH1751) and methanol (Sigma, Cat No: 366927) (2:1, v/v,), and the lipid
mixture, composed of DOPE/sulfatide (3:7, mol/mol), was transferred to glass tubes.
71
Samples were then reduced to a minimum volume under a nitrogen stream, and
stored under vacuum for 24 h at 4 °C to completely evaporate the organic solvent.
The thin lipid films were hydrated by the addition of 1 mL of 250 mM ammonium
sulfate (pH 8.5, Sigma, Cat No: A4418-500g). The samples were then placed in an
ice-water bath and sonicated under nitrogen for 2.5 min with 50% amplitude using a
sonicator (Sonics & Materials, Inc). Following sonication, the liposomes were
formed via extrusion through polycarbonate membranes (Avanti Polar Lipids, Inc.)
with consecutive pore sizes of 400 nm for 14 times, 200 nm for 14 times and 100 nm
for 19 times at room temperature. To establish a trans-bilayer ammonium sulfate
gradient, the extruded liposomes were dialyzed against a 250-fold volume of 10%
sucrose in 25 mM Trizma (Sigma, Cat No: T6066) at pH 8.5 at 4 °C for 24 h. The
external buffer was changed three times during dialysis. After dialysis of the
liposomes, DOX (Sigma, Cat No: 44583-1 mg), in 10% sucrose at a final
concentration of 5 mg/mL, was added to the liposomes at a drug-to-lipid ratio of
0.3:1 (w/w), followed by incubation in a water bath at 60 °C for 1 h. Non-
encapsulated DOX was removed by size exclusion chromatography using a
Sephadex G-50 (Sigma, Cat No: G5080-10g) column.
2.3.1.2 Development of PEG-SCL-DOX.
For development of PEG-SCL-DOX, 1,2-distearoyl-sn-glycero-3-
phosphoethanolamine-N-[maleimide(polyethylene glycol)-2000] (PEG2000-DSPE,
NANOCS, Cat No: PG2-LPCA-2K) was dissolved in chloroform (Deakin University,
LES store, CH1751) and methanol (Sigma, Cat No: 366927) (2:1, v/v). Then,
PEG2000-DSPE was added to DOPE/sulfatide (3:7, mol/mol) as a molar ratio at 3% or
72
5% of the whole lipids amount. All the following procedures were the same as those
in the Section 2.3.1.1.
2.3.1.3 Development of DOX-encapsulated, aptamer functionalised PEG-SCL
DOX-encapsulated nanoliposome (with 5% DSPE-PEG-COOH) was prepared as
previously described (Section 2.3.1.1). The coupling of aptamer to liposomes was
performed through an amide bond between the carboxylic groups of the linker lipid,
DSPE-PEG-COOH, and amino groups of the aptamer in the presence of a water-
soluble carbodiimide, 1- [3-dimethylaminopropyl]-3-ethylcarbodiimide
hydrochloride (EDC) (Otsubo et al. 1998). The reaction of carboxyl group in the
presence of EDC forms an amine-intermediate, which after reacting with sulfo-N-
hydroxysuccinimide (Sulfo-NHS) will form a semi-stable amine reactive NHS ester
that afterwards will react with amino groups of aptamer leading to a stable amide
bond. For conjugation, EDC (Thermo, Cat No: 22980) and NHS (Thermo, Cat No:
24520) were added into the DOX-encapsulated nanoliposome in PBS (pH= 6.0,
Medicago, Cat No: 09-9400-100) to a final concentration of 2 mM and 5 mM,
respectively. The mixture was then incubated for 0.5 h at room temperature with
gentle stirring (Dhar et al. 2008). The pH value of the mixture was raised to 7.2-7.5
with phosphate buffer (Medicago, Cat No: 09-9400-100) (or other non-amine buffer)
immediately before the reaction to the amine-containing molecule. When conjugated
with PEG-SCL, aptamer from a stock solution in RNA-free H2O was added into the
NHS-activated nanoliposomes at 1:10 molar ratio (aptamer-to-DSPE-PEG-COOH in
nanoliposome ratio) and gently stirred for 2 h at room temperature. Unreacted EDC,
Sulfo-NHS and free aptamer were removed by gel filtration using a Sephdex-G50
column (Estevez et al. 2010b; Karathanasis, Bhavane & Annapragada 2007). After
73
purification, samples were collected for spectrophotometry, fluorescence and particle
size analysis. All the analytical procedures were the same as described in Section
2.3.1.
2.3.1.4 Chromatographic instrumentation and system
Chromatographic instrumentation was used based on a previously published
method with some modifications (Alvarez-Cedron, Sayalero & Lanao 1999). Briefly,
the High Performance Liquid Chromatography (HPLC) system (Milford, MA, USA)
used in this study consists of a Waters e2695 Separation Module and a Waters 2475
Multi Fluorescence Detector. The excitation and emission wavelengths were set at
the 470 nm and 585 nm, respectively. Chromatographic separation was performed
using a Nova-Pak® C18 column (3.9 × 150 mm i.d., 4 m, Waters, USA) with a
Nova-Pak® C18 guard column (3.9 × 20 mm i.d., 4 m, Waters, USA). A mixture
(55:45, v/v) of methanol and 10 mM phosphate buffer (pH = 3.0, Fluka, Cat No:
71500) was used as the mobile phase. The flow-rate used in the assay was 1 mL/min
and the column was maintained at 40 ± 5 °C throughout the chromatographic process.
All solvents for HPLC procedures were prepared freshly and filtered with 0.22 m
membrane before using.
2.3.1.5 Establishing standard curves for determination of DOX loading
efficiency
DOX in 10% sucrose was loaded into the liposomes as described above without
passing through the Sephadex G-50 (Sigma, Cat No: G5080-10g) column. The
solution was then diluted with a methanol and phosphate buffer mixture (55:45, v/v)
to a final DOX concentration of 20 g/mL. The mixture was diluted with the same
74
solution (methanol and phosphate buffer, 55:45, v/v) to different concentrations,
including 50 ng/mL, 200 ng/mL, 400 ng/mL, 800 ng/mL, 1,000 ng/mL, 4,000 ng/mL,
8,000 ng/mL, 10,000 ng/mL and 20,000 ng/mL. The samples were centrifuged at
21,000 × g for 5 min. The supernatant layers were collected into HPLC vials, which
were measured via HPLC as described in the Section 2.3.1.4. Calibration curves
were constructed by plotting peak area of fluorescence derived from DOX vs. DOX
concentrations. A linear regression was used for quantitation. The standard formulas
were determined by linear regression as y = mx + b, where y is the peak area of DOX
and x is the DOX concentration.
2.3.1.6 Determination of drug loading efficiency of SCL-DOX or PEG-SCL-
DOX
The amount of DOX encapsulated in SCL-DOX or PEG-SCL-DOX was
determined by disrupting the liposomes with methanol, followed by quantification of
DOX using a fluorescence detector in the HPLC system. Sample concentrations were
determined by linear regression using the standard formula y = mx + b obtained from
Section 2.3.1.5, where y is the peak area and x is the concentration of the standard in
ng/mL. Briefly, 10 L aliquot of the SCL-DOX eluted from a Sephadex G-50 column
was diluted in 100-fold phosphate buffer/methanol (45:55, v/v), and the mixture was
centrifuged at 21,000 × g for 5 min. Then, the supernatant was measured via using
HPLC. Encapsulation efficiency was calculated by the following equation:
Encapsulation efficiency (%) = DOX encapsulated in liposomes ×100% DOX added to liposomes
75
2.3.1.7 Establishing standard curves for phospholipids assay
The concentration of phospholipids (DOPE, Avanti Polar Lipids, Inc., Cat No:
850725P-500mg) in the liposomes was determined as previously described (Stewart
1980). For the standard curve of phospholipids assay, 10 mg DOPE was dissolved in
10 mL of chloroform as a stock. Then, different concentrations (0, 100, 300, 500, 700,
900 and 1,000 g/mL) of DOPE were prepared with dilution in chloroform.
Following this, 1 mL chloroform (Deakin University, LES store, CH1751) and 0.5
mL ferri-thiocyanate reagent (containing 1M ferric chloride hexahydrate, MPBio,
Cat No:153500-100g, and 0.4 M ammonium thiocyanate, Westlab, Cat No: AA010-
500G) were added to the sample, followed by vortexing for 1 min, and centrifugation
at 12,000 × g for 5 min. Following the removal of supernatant, the absorbance of
samples was measured at 488 nm against the chloroform blank using
spectrophotometer (U-1800, HITACHI, Japan). Absorbance were plotted against the
concentration to obtain the analytical curve followed by linear regression analysis.
The method was considered linear when R2>0.99. The calibration curve was obtained
by plotting the absorbance (y) vs. the concentrations (x).
2.3.1.8 Determination of phospholipids in SCL-DOX or PEG-SCL-DOX
The concentration of phospholipids (DOPE) in liposomes was determined as
previously described (Stewart 1980). Briefly, 1 mL chloroform and 0.5 mL ferri-
thiocyanate reagent (containing 1M ferric chloride hexahydrate and 0.4 M
ammonium thiocyanate) were added to a 100 L aliquot of SCL-DOX. The
following steps were the same as described in Section 2.3.1.6. The DOPE
concentration in the samples was calculated according to a standard curve of its
absorbance intensity vs. DOPE concentration.
76
2.3.1.9 Determination of particle size and zeta potential of SCL-DOX or PEG-
SCL-DOX
Following the size exclusion chromatography assay, a 10 L aliquot of SCL-DOX
was diluted in 990 L PBS and mixed gently. The vesicle size and zeta potential of
SCL were measured using Zetasizer Nano ZS Particle Characterisation System from
Malvern Instruments (Malvern, UK).
2.3.1.10 In vitro release kinetics of SCL-DOX/PEG-SCL-DOX in PBS
The in vitro leakage of DOX from SCL-DOX/PEG-SCL-DOX was measured by a
dialysis method (Rai et al. 2008; Song, H et al. 2006a). Briefly, 2.5 mL SCL-
DOX/PEG-SCL-DOX was added into a Slide-A-Lyzer Dialysis Cassette (Pierce,
molecular weight cut-off of 2 kDa). The dialysis cassette was placed into a beaker
containing a 250-fold excess of PBS. The SCL-DOX/PEG-SCL-DOX was dialysed
with stirring for 72 h at 37°C. At various time points (0 h, 0.5 h, 1 h, 4 h, 8 h, 24 h,
48 h and 72 h), a 500 L aliquot was withdrawn from the external buffer for release
kinetics analysis, and replaced with the same volume of fresh external buffer. For
HPLC measurement, the aliquots were mixed with 500 L methanol, followed by
centrifuged at 21,000 × g for 5 min. Supernatants were then collected for
measurement by HPLC. The drug concentration in the external buffer was calculated
according to a standard curve of DOX concentration vs. its fluorescence intensity.
2.3.1.11 In vitro release kinetics of SCL-DOX/PEG-SCL-DOX in PBS/10% FBS
The dialysis procedure was the same as that in the Section 2.3.1.9 except the
dialysis buffer consisted of PBS containing 10% FBS, penicillin (50 U/mL), and
77
streptomycin (50 μg/mL). In addition, 1 mL methanol was added to a 500 L aliquot
of external buffer for protein precipitation. After centrifuged at 21,000 × g for 5 min,
supernatants were collected for measurement of the DOX concentration by HPLC.
The drug concentration in the external buffer was calculated according to a standard
curve of DOX concentration vs. its fluorescence intensity as described in Section
2.3.1.11.
2.3.1.12 Establishing standard curves for in vitro release kinetics in PBS
DOX stock was dissolved in methanol and diluted with PBS/methanol (1:1, v/v) at
the final concentration of 20 g/mL. The stock was then diluted by PBS and
methanol mixture (1:1, v/v) to cover three concentration ranges (0.1-100 ng/mL, 50-
1000 ng/mL and 800-20000 ng/mL), each of which contained five or eight standard
solution (0.1-100 ng/mL: 0.1, 0.5, 1, 5, 10, 25, 50 and 100 ng/mL; 50-1,000 ng/mL:
50, 200, 400, 800 and 1,000 ng/mL; 800-20000 ng/mL: 800, 4,000, 8,000, 10,000,
20,000 ng/mL). The samples were centrifuged at 21,000 × g for 5 min. The
supernatant layers were collected into the vials, followed by measurement under the
conditions as described in Section 2.3.1.4. Calibration curves were constructed as
described in Section 2.3.1.5.
2.3.1.13 Establishing standard curves for in vitro release kinetics in PBS/10%
FBS
DOX stock was dissolved in methanol and diluted by methanol/PBS (2:1, v/v) at a
final concentration of 20 g/mL. The stock was then diluted with PBS and methanol
mixture (2:1, v/v) to the three concentration ranges (0.1-100 ng/mL, 50-1,000 ng/mL
and 800-20000 ng/mL), each of which contained five or eight standard solution (0.1-
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100 ng/mL: 0.1, 0.5, 1, 5, 10, 25, 50 and 100 ng/mL; 50-1000ng/mL: 50, 200, 400,
800 and 1,000 ng/mL; 800-20,000 ng/mL: 800, 4,000, 8,000, 10,000, 20,000 ng/mL).
The samples were centrifuged at 21,000 × g for 5 min. The supernatant layers were
collected into the vials, followed by measuring under the conditions as described in
Section 2.3.1.4. Calibration curves were constructed as described in Section 2.3.1.5.
2.3.2 In vitro uptake, retention and cytotoxicity assay
2.3.2.1 MTT assay
The viabilities of treated and untreated cells was determined by the (3-(4,5-
dimethylthiazol-2-yl)-2,5-diphenyl-2H-tetrazolium bromide) MTT assay which
measures the mitochondrial conversion of MTT to formazan as detected by the
change of optical density at 570 nm (Jung et al. 2009b; Zhang, FY et al. 2009).
Briefly, the MTT (Sigma, Cat No: M5655) stock was dissolved with Milli Q water to
a final concentration of 5 mg/mL. This solution was filtered through a 0.22 m filter
and stored at -20°C in the dark before use. One T-75 flask of U-118MG cells or HT-
29 cells was trypsinised and 5 mL of complete media was added to the trypsinised
cells. The cell suspension was centrifuged in a sterile 15 mL Falcon tube at 500 × g
in the swinging bucket rotor for 5 min. Then, U-118MG cells were plated at a density
of 4.0 × 103 cells per well in 100 l DMEM medium and HT-29 were plated at a
density of 3.0 × 103 cells per well in 100 L McCoy's 5A medium in 96-well plates
and allowed to grow for 24 h. The cells were then exposed to a series of
concentrations (0.1 g/mL, 0.5 g/mL, 1 g/mL, 5 g/mL, 10 g/mL, 50 g/mL and
100 g/mL) of free DOX, SCL-DOX or blank SCL for 48 h at 37° C under 5% CO2.
Ten microlitres of MTT reagent was added into each well and incubated for 4 h. The
reaction was terminated by removing MTT prior to the addition of 150 L/well
79
solubilisation reagent (Dimethyl Sulfoxide, DMSO, Merck, Cat No: 1.02952.1000).
The absorbance of the wells, including the blanks, was measured at 570 nm using a
VICTOR TM X5 Multilabel HTS Reader (PerkinElmer Life and Analytical
Sciences). DOX concentration leading to 50% cell-killing (IC50) was calculated using
the statistical software package SPSS 13.0 (Cheng et al. 2009). All experiments were
performed in triplicate and repeated thrice.
2.3.2.2 Confocal microscopy analysis for cellular uptake and retention of SCL-
DOX
U118MG cells (1.5 × 105 cells/well) or HT-29 cells (1 × 105 cells/well) were
seeded in 35-mm glass bottom dishes and incubated at 37°C in 5% CO2 for 24 h. The
medium was then replaced with full culture medium containing 2 g/mL free DOX
or SCL-DOX. Twenty-four hours later, cells were washed twice with phosphate
buffered saline (PBS) and imaged for cellular uptake studies. For retentions studies,
cells were first exposed to 2 g/mL free DOX or SCL-DOX in full cell culture
medium for 24 h and then washed twice with PBS. Cells were then incubated with
fresh cell culture medium and serially imaged at 1 h, 2 h, 4 h, and 24 h using a
Fluoview FV10i fluorescence laser scanning confocal microscopy (Olympus, Japan).
2.3.2.3 Cellular uptake and retention of aptamer-functionalised DOX-
encapsulated PEG-SCL
HEK-293T cells (1.0 × 105 cells/well) or HT-29 cells (1.0 × 105 cells/well) were
seeded in 35-mm glass bottom dishes and incubated at 37°C in 5% CO2 for 24 h. The
medium was then replaced with full culture medium containing 2 g/mL free DOX,
SCL-DOX, aptamer-functionalised DOX-encapsulated SCL (ap-PEG-SCL-DOX)
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and SCL-DOX conjugated with negative control aptamer (negative ap-PEG-SCL-
DOX). Cellular uptake and retention was evaluated as described in Section 2.3.2.2.
2.3.2.4 Flow cytometry analysis for intracellular uptake of SCL-DOX
To measure the intracellular uptake of free DOX and SCL-DOX, U-118MG
glioblastoma cells or MFC-7 breast cancer cells were cultured in a 12-well plate to
achieve approximately 80% confluence. DOX was dissolved in PBS at a
concentration of 1 mg/mL, and diluted in DMEM to a final concentration of 4 g/mL.
SCL-DOX was prepared fresh before each experiment. One millilitre of 4 g/mL
DOX or SCL-DOX in full culture medium was added into each well. Following
incubation at 37 °C for different periods (0 h, 1 h, 2 h, 4 h, 6 h and 24 h), the culture
medium was discarded. Cells were then washed three times with ice-cold PBS,
detached by trypsinisation, centrifuged at 500 × g for 5 min, resuspended in 0.1 mL
ice-cold PBS, and then analysed using a BD FACSCantoTM II flow cytometer
(Becton-Dickinson) using excitation wavelength of 488 nm and emission between
543 and 627 nm. Fluorescent histograms were recorded by BD FACSCantoTM II flow
cytometer and analysed using BD FACSDiva software (v6.0). A minimum of 10,000
events were analysed for each sample from three independent experiments (Gariboldi
et al. 2007; Meschini et al. 2002; Tang et al. 2007; Zheng, C et al. 2009). 2.3.3 In
vivo methods
2.3.3.1 Development of [3H]-tracking of liposome.
[3H] Cholesteryl hexadecyl ether (CHE, Perkin Elmer, Cat No: NET859001MC)
was added at a molar ratio of 1:24000 (CHE/DOPE) as a nonmetabolised,
nonexchangeable lipid tracer into the lipid mixture (Hussain et al. 2007). The
81
preparation of SCL-DOX or PEG-SCL-DOX was carried out as described above
(Section 2.3.1.1 and 2.3.1.2).
2.3.3.2 Pharmacokinetics (PK) study in SD rats
To investigate the PK properties of SCL-DOX and PEG-SCL-DOX in vivo,
healthy male SD rats were randomly divided into 3 experimental groups (five to six
rats per group) and were injected i.v. with free DOX, SCL-DOX or PEG-SCL-DOX
via the tail vein with a single dose of 5 mg DOX/kg (administration rate 0.4 mL/min).
After injection, blood was serially sampled from the tail of the same animal at 2 min,
30 min, 2 h, 6 h, 24 h and 48 h. Blood samples (500 L) were collected in
heparinised tubes, and then centrifuged at 3,000 × g at 4 °C for 10 min to separate
the plasma and stored at -20 °C until assayed for DOX and [3H] (Ahmed et al. 2009;
Rahman et al. 1986b).
2.3.3.3 Biodistribution (BD) study in rats
For BD study, healthy male SD rats were randomly divided into 3 experimental
groups (five to six rats per group). Free DOX (5 mg/kg) or SCL-DOX (5 mg/kg in
DOX) was administrated via the tail vein. Rats were then sacrificed by Lethabarb R
(100 mg/kg) at 0.5 h, 2 h, 4 h and 24 h after the injection and tissues (heart, liver,
spleen, lung, kidney and brain) were collected. Tissues were then lightly washed in
cold physiological saline to remove any excess blood, blot-dried using filter paper
and weighed (Xiong, Huang, Lu, Zhang, Zhang, Nagai, et al. 2005). Tissues were
then stored at -20 °C until assayed for DOX and [3H].
82
2.3.3.4 Establishing standard curves of DOX in plasma
Free DOX was dissolved in methanol to a stock concentration of 1 mg/mL. For
each concentration (2, 5, 10, 50, 100, 200, 500, 1000, 2000, 5000 and 10000 ng/mL),
stock was then diluted with 100 L blank plasma and methanol/phosphate buffer
mixture (55:45, v/v) to the final volume of 1 mL. Samples were vortexed for 1 min,
and centrifuged at 21,000 × g for 10 min at 4 C. The supernatant was transferred to
another tube followed by the addition of 2 L of freshly prepared perchloric acid
(35%, v/v,). Then, samples were vortexed for another 1 min and centrifuged at
21,000 ×g for 10 min at 4 °C. The supernatant was collected into vials, followed by
measuring of DOX via HPLC under the conditions as described in Section 2.3.1.4.
Calibration curves were constructed as described in Section 2.3.1.5.
2.3.3.5 Determination of the stability of SCL-DOX in plasma
To determine DOX levels in plasma, 495 L of methanol and 405 L of phosphate
buffer was added to 100 μL plasma. Then, samples were vortexed and treated as the
procedure above (Section 2.3.3.4). The supernatant (100 L) was injected into the
HPLC system (Alvarez-Cedrón, Sayalero & Lanao 1999). The chromatographic
instrumentation and system were used as described in Section 2.3.1.4. The drug
concentration was calculated according to a standard curve of DOX concentration vs.
its fluorescence intensity.
2.3.3.6 Establishing standard curves of DOX in tissues
To establish a standard curve of DOX in tissues, 100 g of tissue was placed in a
tightly sealed 2-mL screw-capped tube. Then, 1 mL methanol/phosphate buffer
83
mixture (55:45, v/v) with different free DOX concentrations (50, 100, 500, 1000 and
5000 ng/mL) was added into the tubes. Samples were homogenised via the
FastPrep®-24 tissue and cell homogeniser (MP Biomedicals, US). The tissue
homogenate was centrifuged at 21,000 × g for 10 min at 4 °C. The supernatant was
transferred to another tube and extracted with the addition of 2 L freshly prepared
perchloric acid (35%, v/v), vortexed for 1 min, and centrifuged at 21,000 × g for 10
min at 4 °C. The supernatant was collected into vials, followed by measurement
under the conditions as described in Section 2.3.1.4. Calibration curves were
constructed as described in Section 2.3.1.5.
2.3.3.7 Determination of the release of DOX from SCL-DOX in tissues
To determine DOX levels in tissues, 100 g of tissue was added to 495 L of
methanol and 405 L of phosphate buffer in a tightly sealed 2-mL screw-capped tube
followed by homogenisation using the FastPrep®-24 tissue and cell homogeniser
(MP Biomedicals, US). The preparation procedure was the same as that in Section
2.3.3.6. The chromatographic instrumentation and system were used as described in
Section 2.3.1.3. The drug concentrations in the tissues were calculated according to a
standard curve of DOX prepared in the respective tissue homogenate.
2.3.3.8 Liquid scintillation counting assay for [3H] in plasma
To determine lipid levels in plasma, 1 mL of Solvable (PerkinElmer, Cat No:
6NE9100) was added to 100 μL plasma in a screw-cap air-tight glass tube and
digested for 1 h at 50 °C. After cooling to room temperature, 200 μL of 100 mmol/L
ethylenediamine tetraacetic acid (EDTA, Scharlau chemie SA, Cat No: AC0965) was
added, followed by the addition of 200 μL of 30% (v/v) H2O2. Samples were allowed
84
to stand for 15 min at room temperature followed by incubation at 55 °C for 1 h and
cooling to room temperature. After the addition of 10 mL of scintillation cocktail
(ULTMA GOLDTM, PerkinElmer, Cat No: 6013329), the samples were analysed in
a Tri-Carb 2910TR liquid scintillation analyser (PerkinElmer, USA) after the
adaption for light and temperature for 1 h in the counter prior to counting.
2.3.3.9 Liquid scintillation counting assay for [3H] in tissues
For determination of [3H] in tissues, 2 mL of Solvable was added to 100 mg
tissues. Tissues were then digested for 3 h at 50 °C. The following steps and
analytical procedures were the same as described in Section 2.3.3.8.
2.3.3.10 Tumour xenograft model
Six-week-old female BALB/c-Foxn1nu mice were allowed a 7 day acclimatisation
period after arrival. The animal weight was recorded upon arrival. For uptake and
therapeutic studies, single cell suspensions were harvested by trypsinisation. The
cells were washed in PBS, and resuspended in PBS at a seeding concentration of
5×107 cells/mL for U-118MG or 3×107/mL for HT-29. Cell suspensions (5 × 106
cells in 0.1 mL PBS for U-118MG, 3 × 106 in 0.1 mL PBS for HT-29) were
inoculated subcutaneously (s. c.) into the right flank of the mice with a 1 mL syringe
and 26 gauge needle. Tumour size was assessed using a digital calliper every other
day after implantation and approximate tumour burden (mm3) was calculated as
length × width2/2 (V=lw2/2), where length and width are the longest and shortest axis
in millimetres (Elbayoumi, T. A. & Torchilin, V. P. 2008).
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2.3.3.11 Establishing standard curves of DOX in tissues and tumours in tumour-
bearing mice
The establishment of standard curves of DOX in tissues and tumours in tumour-
bearing mice were the same as described in Section 2.3.3.6.
2.3.3.12 Tumour uptake and tissue distribution of SCL-DOX in tumour-bearing
mice
Once the tumour reached a volume of 150 mm3, the nude mice were randomly
divided into 2 experimental groups (5 to 6 mice per group) for injection.
Formulations of either free or SCL-DOX at doses of 5 mg/kg DOX or equivalent
were administered via the tail vein at a rate of 0.4 mL/min. After injection, mice
were sacrificed by injection of Lethabarb R (100 mg/kg) at 24 h post-injection.
Tumours and tissues were collected and then processed as described in Section
2.3.3.7 (Cheng et al. 2009; ElBayoumi, aA & Torchilin, VP 2009; Li, X et al. 2009).
2.3.3.13 Treatment efficacy of SCL-DOX in tumour-bearing mice
For therapeutic experiments, the nude mice were randomly divided into 4
experimental groups (8 to 10 mice per group) for injection when the xenograft
tumours reached 150 mm3 (for U-118MG) or 35 mm3 (for HT-29). Mice received
injection of saline, free DOX (5 mg/kg), SCL-DOX (5 mg/kg in DOX) or blank SCL
via tail vein once a week for 6 weeks (for U-118MG) or twice a week for 3 weeks
(for HT-29). Tumour growth was monitored by measuring tumour diameter every
other day with a calliper and animal weights were monitored at the same time (Lopes
de Menezes et al. 2000; Tang et al. 2007).
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2.3.3.14 Analysis of systemic toxicity of SCL-DOX
To evaluate the general toxicity of free DOX, SCL-DOX and blank SCL, blood
was collected when the mice for therapeutic experiments were sacrificed. Plasma
biochemistry, blood cells counts and troponin were analysed by a veterinary
pathology laboratory (Gribbles Veterinary Pathology, Clayton, Victoria, Australia).
Blood smears were obtained for each animal to obtain a relative white cell count
adapted from the Fonio method for platelet counting (Adams, E 1948; Oliveira et al.
2003), with minor modifications. Slides were stained with Giesma and an area of the
blood smear was chosen where the red cells abutted each other but did not overlap,
with consecutive fields chosen to eliminate bias. The total number of white cells per
1500 red cells were counted (n=3 for each slide) and compared for each group.
2.3.3.15 Perfusion in healthy SD rats
The perfusion in healthy SD rats was performed based on a previously published
method (Balasubramanian et al. 2010). Briefly, the heavily anesthetised rat was
placed on its back in a dissection tray. A cut along the sternum (~8 cm long) was
made to expose the sternum's end. Sharp scissors were used to cut diaphragm
laterally on both sides and cut toward the head across ribs and parallel to lungs. The
heart was exposed and held steady with forceps, and then a large bore blunt needle
(18 gauge) was inserted into left ventricle to extend straight up about 5 mm. Rats
were perfused through the left cardiac ventricle with saline (>150 mL) until the
venous return was clear. The cortex was collected and then processed as described in
Section 2.3.3.7
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2.3.3.16 Establishment of a intracranial brain tumour xenograft model
The establishment of an intracranial brain tumour xenograft model was based on
previously published methods with some modifications (Madhankumar et al. 2009;
Yamashita et al. 2007). Briefly, seven-week-old male Wistar rats were allowed a 7
days acclimatisation period after arrival. The animal weight was recorded upon
arrival. Immediately before surgery, single cell suspensions of C6 were harvested by
trypsinisation. The cells were washed in PBS, and resuspended with serum free
DMEM at a seeding concentration of 1×107 cells/mL. Rats were deeply anesthetised
by i.p. injection of Ketamine-Xylazine (75 mg/kg - 10 mg/kg body weight) and then
placed in a stereotactic device (Model 900 Small Animal Stereotaxic Instrument,
KOPF®, USA). The head was mounted in a head holder. Hair was shaved and the
head was cleaned with Betadine Surgical Scrab solution, followed by 80% ethanol
twice. The skin was incised in the midline and a small hole was made with a drill at a
point approximately 3 mm from mid-line, 2 mm posterior to the coronal suture in the
anterior part of the right skull. One million C6 cells, suspended in a total volume of
10 L DMEM, was injected using the Harvard Apparatus Pump 11 Elite Syringe
Pumps (USA) with the 26 gauge needle (Harvard apparatus, USA) at the rate of 1
L/min over a period of 10 min into the right brain hemisphere, 6 mm beneath the
bone surface. After injection, the needle was left in the brain for 10 min to minimise
movement of the cells away from the site and to improve reproducibility. Then, the
needle was withdrawn slowly, with 5 min suspension for each millimetre. The needle
hole on the scalp was sealed with bone wax, and the incision was closed using
sutures. Rats were monitored every day. Eighteen days after the surgery, free DOX
or SCL-DOX (at the dosage of 5 mg/kg of DOX) was given via i. v. administration
through the tail vein. Rats were sacrificed 24 h after injection by injection of
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Lethabarb R (100 mg/kg). Tumours and tissues were collected for DOX
concentration analysis. All the measurement procedures were the same as described
in Section 2.3.3.7.
2.3.3.17 Establishing standard curves of DOX in tissues and tumours in
intracranial brain tumour xenograft models
The establishment of standard curves of DOX in tissues and C6 gliomas were the
same as described in Section 2.3.3.6.
2.3.3.18 Tumour uptake and tissue distribution in intracranial brain tumour
xenograft models
The glioma-bearing rats were randomly divided into 2 experimental groups (3 mice
per group) for injection at day 18 after tumour inoculation. Formulations of either
free or SCL-DOX at doses of 5 mg/kg DOX or equivalent were administered via the
tail vein at a rate of 0.4 mL/min. After injection, rats were sacrificed by injection of
Lethabarb R (100 mg/kg) at 24 h time point. Tumours and tissues were collected and
then processed as described in Section 2.3.3.7.
2.3.4 Data analysis
The pharmacokinetic parameters were calculated from the average plasma
concentrations using the pharmacokinetic software DAS 2.0 software (Mathematical
Pharmacology Professional Committee of China, Shanghai, China). Data and results
were reported as means and standard error (means ± S.E.) unless otherwise stated.
The differences in the mean values among different groups were analysed using a
89
one-way analysis of variance (ANOVA) using SPSS 13.0. Significance was
considered at values of p<0.05.
90
Chapter 3 - Development of sulfatide-
containing nanoliposomal doxorubicin and
PEG-sulfatide-containing nanoliposomal
doxorubicin
91
3.1 Introduction
Indiscriminate exposure of all cells in the body to a systemically administered
chemotherapy agent kills healthy cells as well as tumour cells (Alexis, F. et al. 2008;
Bawarski et al. 2008), causing severe toxicity to the patients and leading to serious
side effects, low efficacy and poor quality of life (Peer, Dan. et al. 2007; Sapra, P.,
Tyagi & Allen 2005). This non-specific tissue distribution and the resulting side-
effects limit the clinical application of anticancer drugs (Ganta, S. et al. 2008). Thus,
there is an urgent need to develop new chemotherapeutics that can target tumour
cells effectively. Over the past few decades, nanoscale therapeutic systems have
emerged as novel therapeutic modalities for combating cancer (Segal & Saltz 2009).
The nanoparticle formulations of traditional free antitumour drugs may have
improved pharmacokinetics (PK) and biodistribution (BD) profiles, enhanced
antitumour efficacy, as well as reduced toxicity to healthy tissues.
Liposomal drugs were the first approved and widely used such nanomedicine for
the treatment of cancers (Barenholz 2012; Krown et al. 2004; Sawant & Torchilin
2012; Wang, AZ, Langer & Farokhzad 2012). Liposomes are microscopic
phospholipid vesicles with a bilayered membrane structure. Preclinical and clinical
studies have shown that the pharmacokinetic profiles, as well as the targeting
specificities, of liposomes can be controlled and modified to reduce the side effects
of encapsulated drugs and enhance their efficacy (Davis, Chen & Shin 2008; Park, K
2007; Torchilin 2006).
The targeted delivery of anticancer agents to the tumour microenvironment is a
promising avenue for the therapy of metastatic tumours. Tenascin-C, a large
extracellular matrix (ECM) hexabrachion glycoprotein, is highly expressed in the
92
microenvironment of most solid tumours, but is absent or greatly reduced in most
adult tissues (Brellier & Chiquet-Ehrismann 2012). As sulfatide, a lipid that is found
in humans, is known to bind several ECM glycoproteins including tenascin-C
(Kappler et al. 2002), it has the potential to be one of the components of
nanoliposomes to target tumours expressing high levels of tenascin-C. We had
recently developed a novel liposomal carrier system that was composed of two lipids
found in humans, sulfatide and DOPE (He et al. 2010; Townson et al. 2007). Under
physiological pH, DOPE confers stability to sulfatide-containing liposomes (SCL)
via its inhibitory effects on liposome fusion, as the incorporation of sulfatide into
DOPE vesicles greatly enhances the stability of the liposomes formed, even in the
presence of plasma, presumably due to the hydration of the negatively charged
sulfate head-group of the glycosphingolipid (Shao, Ke et al. 2006). Previous study
indicated that the SCL could significantly improve the cellular uptake in cells with
high level of tenascin-C expression compared to the nanoliposomes without sulfatide
(Shao, Ke et al. 2006). We had also demonstrated that the interaction between
sulfatide and tenascin-C mediates the binding of SCL to the ECM and endocytic
uptake of the liposomes by tumour cells at least in vitro (Shao, Ke et al. 2006; Shao,
K. et al. 2007). As a result, the sulfatide-containing liposomal carrier system
represents a new class of natural lipid-guided intracellular delivery system targeting
the tumour microenvironment.
In exploring such a new direction for the development of more effective anticancer
chemotherapeutics, it is important to understand the in vivo behaviour of the
nanocarrier, as the unique distribution governed by the properties of the nanocarrier
can change the therapeutic efficacy as well as alter the toxicity profile of the
encapsulated drug. However, the in vitro and in vivo stability of the drug
93
encapsulated SCL, as well as its pharmacokinetic behaviour and the BD pattern are
yet to be established. Thus, this chapter explores the physiochemical properties of the
sulfatide-containing liposomal doxorubicin (SCL-DOX) in vitro, as well as the
pharmacokinetics (PK) and BD in vivo.
The work presented here was performed in collaboration with Yan Yu.
Parts of this Chapter have been published (Lin J, Yu Y, Shigdar S, Fang DZ, Du
JR, Wei MQ, Danks A, Liu K and Duan W. “Enhanced antitumor efficacy and
reduced systemic toxicity of sulfatide-containing nanoliposomal doxorubicin in a
xenograft model of colorectal cancer”. PLoS ONE, 2012, 7(11): e49277.)
3.2 Results
3.2.1. Characterisation of nanoliposomes
3.2.1.1 Establishing standard curves for determination of DOX loading
efficiency
To evaluate DOX loading efficiency in SCL via HPLC, standard curves were
constructed by plotting peak area of fluorescence derived from DOX vs. DOX
concentrations. To mimic samples to be determined, SCL-DOX or PEG-SCL-DOX
without purification using size-exclusion chromatography were used as standards and
diluted to series of DOX concentrations from 50 ng/mL to 20,000 ng/mL, followed
by the HPLC measurement. A linear regression analysis was used for quantitation.
As shown in Figure 3-1, a good linearity between peak areas and concentrations was
obtained within the tested concentrations with a correlation coefficient of R2=0.9980
(for SCL-DOX and 3% PEG-SCL-DOX) or R2=0.9991 (for 5 % PEG-SCL-DOX).
94
Figure 3-1. Standard curves for determination of DOX loading efficiency. The standard formulas were determined by linear regression as y = mx + b, where y is the peak area of DOX and x is the DOX concentration. (a) Standard curve for determination of DOX loading efficiency in SCL. (b) Standard curve for determination of DOX loading efficiency in SCL with 3% PEG2000-DSPE. (c) Standard curve for determination of DOX loading efficiency in SCL with 5% PEG2000-DSPE. Liposomal formulations were diluted with a methanol and phosphate buffer mixture (55:45, v/v) to different concentrations and measured using HPLC.
3.2.1.2. Establishing standard curves for phospholipids assay
To determine the DOPE concentration in the SCL-DOX, DOPE standards with
different concentrations were prepared and measured using a spectrophotometer for
construction of the calibration curve by plotting absorbance (y) vs. concentrations (x).
As shown in Figure 3-2, the correlation coefficients (R2) of the standard curve was
0.9996, suggesting a good linear regression of the absorbance vs. the concentration
95
within the concentration range of 0 – 1,000 g/mL in UV-spectrophotometry for
DOPE.
Figure 3-2. Standard curves for phospholipids assay. One mg/mL DOPE stock in chloroform were diluted with chloroform to different concentrations and measured by UV-spectrometry. The standard formulas were determined by linear regression as y = mx + b, where y indicated the absorbance while x was the concentration of DOPE. Data are shown as
means S.E. (n=3).
3.2.1.3 Characterisation of nanoliposomes
As the composition and preparation method of liposomes could directly affect
their physicochemical characteristics (Fatouros & Antimisiaris 2001), the loading
efficiency, size, drug-to-lipid ratio and surface charge were analysed for drug-
incorporating SCL and PEG-SCL for comparison. The physicochemical
characteristics of SCL and PEG-SCL are presented in Table 3-1. The mean vesicle
size of SCL incorporating DOX was 92.32 ± 1.31 nm with a polydispersity index
(PDI) of 0.15 ± 0.01. At an initial weight ratio of DOX-to-DOPE of 0.3:1, the
efficiency of DOX loading to SCL using an ammonium sulfate gradient was 94.11%
± 2.27%, consistent with the previous report (Fritze et al. 2006). The zeta potential
value of SCL was -26.38 ± 2.20 mV. The DOX to DOPE weight ratio after DOX
encapsulation into SCL was 0.50:1. When compared within different formulations,
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SCL showed the highest loading efficiency, as well as the best DOX to DOPE ratio,
while the SCL with 3% PEG2000-DSPE indicated the smallest particle size (75.23 ±
6.73 nm), the lowest DOX encapsulation efficiency (81.47% ± 1.89%) and the
lowest drug-to-lipid ratio (0.23:1). The loading efficiency of SCL with 5% PEG2000-
DSPE was 89.54% ± 1.36%. The size and PDI were found to be 75.86 ± 5.62 nm and
0.11 ± 0.1, respectively. The average value of zeta potential was -11.37 ± 0.26 mV.
Table 3-1. Physical properties of the liposomal formulations.
Composition DOX loading efficiency (%)
Particle size (nm)
DOX-to-DOPE ratio (w/w)
Polydispersity index (PDI)
Zeta potential (mV)
SCL-DOX 94.11 ± 2.27 92.32 ± 1.31 0.50:1 0.15 ± 0.01 -26.38 ± 2.20 PEG-SCL (3 % PEG) 81.47% ± 1.89 75.23 ± 6.73 0.23:1 0.14 ± 0.01 -19.44 ± 1.82PEG-SCL (5 % PEG) 89.54 ± 1.36 75.86 ± 5.62 0.24:1 0.11 ± 0.01 -11.37 ± 0.26
Data are shown as means ± S.E. of at least three independent experiments.
3.2.2 In vitro drug retention properties.
3.2.2.1 Establishing standard curves for in vitro release kinetics in PBS and
PBS/10% FBS
To analyse the stability of SCL-DOX in vitro, standard curves for the
measurement of DOX in PBS or in PBS/10%FBS were established using 16 standard
DOX solution concentrations covering a range from 0.1 ng/mL to 20,000 ng/mL and
measured using HPLC. As shown in Figure 3-3, peak area of fluorescence derived
from DOX vs. analytic concentration showed a good linear relationship with a
correlation coefficient of R2 > 0.999.
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Figure 3-3. Standard curves for in vitro release kinetics in PBS and PBS/10% FBS. The standard formulas were determined by linear regression as y = mx + b, where y is the peak area of fluorescence derived from DOX and x is the DOX concentration. (a) Standard curves of SCL-DOX for in vitro release kinetics in PBS. (b) Standard curves of SCL-DOX for in
vitro release kinetics in PBS/10% FBS. (c) Standard curves of SCL-DOX with 3% PEG2000-DSPE for in vitro release kinetics in PBS. (d) Standard curves of SCL-DOX with 3% PEG2000-DSPE for in vitro release kinetics in PBS/10% FBS. (e) Standard curves of SCL-DOX with 5% PEG2000-DSPE for in vitro release kinetics in PBS. (f) Standard curves of SCL-DOX with 5% PEG2000-DSPE for in vitro release kinetics in PBS/10% FBS. DOX stock in methanol was diluted with methanol/PBS (v/v=1:1, for standard curve in PBS, or v/v=2:1, for standard curve in PBS/10% FBS) to different concentrations and quantified using HPLC.
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3.2.2.2 In vitro drug retention properties
Having established that SCL showed high loading efficiency, ideal size for in vivo
delivery and negative surface charge, study was proceeded to evaluate the stability of
drug encapsulation as it is essential for efficient drug delivery to the target site. In
vitro DOX release from SCL or PEG-SCL was determined by dialysing SCL-DOX
or PEG-SCL-DOX against PBS or PBS with 10% FBS at 37 °C and measuring the
DOX concentration over time from the fluid within the dialysis container. As shown
in Figure 3-4, there was minimal DOX leakage from the SCL during the first 48 h
dialysis period, with more than 99% of DOX retained in the SCL after 48 h under
both PBS and PBS/serum dialysis conditions. The release of DOX was found to
increase following 48 h incubation. The percentage of DOX retained in the SCL after
72 h was 84.06% ± 8.63% in PBS and 91.91% ± 1.36% in PBS with 10% serum,
respectively. The percentage of DOX retained in the PEG-SCL with 3% PEG after
72 h was 99.96% ± 1.21% in PBS and 99.98% ± 0.95% in PBS with 10% FBS
respectively. Similarly, there were more the 99% of DOX retained in the SCL with 5%
PEG2000-DSPE in both PBS and PBS/10% FBS dialysis condition (99.97 ± 0.74%
and 99.98% ± 0.82%, respectively). Thus, all liposome formulations in the study
were very stable for 48 h 37 °C and retained substantial physical stability after 72 h
in vitro.
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Figure 3-4. In vitro stability of SCL-DOX and PEG-SCL-DOX. (a) In vitro stability of SCL-DOX. (b) In vitro stability of SCL-DOX with 3% PEG2000-DSPE. (c) In vitro stability of SCL-DOX with 5% PEG2000-DSPE. The stability of SCL-DOX or PEG-SCL-DOX was studied by analysing the release of DOX via dialysis assay in PBS or PBS with 10% FBS at 37 °C. Aliquots of dialysis buffer were collected at designated time points (0, 0.5, 1, 4, 8, 24, 48 and 72 h). DOX released into the dialysis buffer was quantified using HPLC. Data are
shown as means S.E. of at least three independent experiments.
3.2.3 Flow cytometry analysis for intracellular uptake of SCL-DOX
To evaluate the accumulation of SCL-DOX in vitro, the intracellular uptake of
DOX and SCL-DOX was studied using flow cytometry utilising the natural
fluorescent property of DOX. Human glioblastoma cells U-118MG and human
adenocarcinoma cells MCF-7 were incubated with 4 g/mL DOX or equivalent
amounts of SCL-DOX for 1 h, 2 h, 4 h, 6 h or 24 h followed by washing with PBS
three times. Cells were then trypsinised and the fluorescence of cells was measured.
As shown in Figure 3-5, the intracellular fluorescence intensity of SCL-DOX was
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lower than that of free DOX in both MCF-7 and U-118MG cells at all time points
during the time period assayed, suggesting that free DOX was taken up more rapidly
by the breast cancers and glioma cells than SCL-DOX in vitro.
Figure 3-5. DOX uptake in MCF-7 and U-118MG cells. Cells were treated with 4 g/mL of either free DOX or SC-DOX for the time indicated. The intracellular mean intensity of DOX fluorescence was analysed using flow cytometry (excitation wavelength: 552 nm, emission wavelength: 578 nm). Data are expressed as means ± S.E. of triplicates.
3.2.4. In vivo characteristics of SCL-DOX and PEG-SCL-DOX.
3.2.4.1 Establishing standard curves for DOX concentration in plasma
Before the quantitative determination of DOX in plasma, plasma spiked with a
serial dilution of DOX concentrations were prepared to obtain the calibration line for
the following PK study. A linear regression was used for quantitation. The
correlation coefficient (R2) for the calibration curve was 0.9998 (Figure 3-6),
indicating a good linearity between peak areas and DOX concentrations was obtained
within the tested concentrations ranging of 2 ng/mL to 10,000 ng/mL.
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Figure 3-6. Standard curve for DOX concentration in plasma. The standard curve was constructed by measurements of plasma spiked with a serial dilution of DOX solutions. The calibration curve was generated by plotting peak area of fluorescence derived from DOX against the DOX concentrations in solution. Sample concentrations were determined by linear regression using the formula y = mx + b, where y is the peak area of DOX ( V × sec) and X is the DOX concentration (ng/mL).
3.2.4.2 Improved pharmacokinetic properties of SCL in healthy SD rats
Pharmacokinetic parameters were evaluated after the single i.v. administration of 5
mg/kg free DOX, SCL-DOX or PEG-SCL-DOX. Due to better encapsulation
efficiency, higher drug-to-lipid ratio and similar stability in vitro compared to SCL
with 3% PEG2000-DSPE (Table 3-1 and Figure 3-4), SCL with 5% PEG2000-DSPE
were chosen as the formulation for PK studies. The pharmacokinetic properties in
healthy male SD rats were determined from plasma concentration-time profiles of
free DOX, SCL-DOX and PEG-SCL-DOX. The main pharmacokinetic parameters
are shown in Table 3-2. In the study, the terminal half-life with free DOX in plasma
was 20.65 ± 1.34 h.. However, the terminal half-life was 41.89 ± 3.58 h with SCL-
DOX, showing a 2.03-fold increase in the terminal half-life. Indeed, the steady state
Vd of free DOX (18.36 L/kg) was 25.15-fold higher than SCL-DOX (0.73 L/kg). The
clearance rate of DOX encapsulated by SCL-DOX (1.39 L/h/kg) was significantly
lower than that of DOX solution (2.68 L/h/kg, p < 0.01). The area under the plasma
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concentration-time curves (AUC0- ) of DOX delivered through SCL was 2.06-fold
higher than free DOX (3597.03 ± 99.36 g/Lh and 1746.87 ± 69.94 g/Lh,
respectively). PK parameters given in Table 3-2 also indicated that PEG-SCL
formulation had the highest AUC0- (73730.54 ± 6161.04 g/Lh), the slowest rate of
clearance (0.62 ± 0.02 L/h/kg) and smallest Vd (0.05 ± 0.01 L/kg) among the tested
formulations. However, DOX in the serum of rats injected with PEG-SCL-DOX
decreased significantly after 2 h (Figure 3-7), which resulted in a decreased
elimination half-life (T1/2) when compared with that of free DOX and SCL-DOX.
Table 3-2. Pharmacokinetic parameters for free DOX, SCL-DOX and PEG-SCL-DOX. Formulations AUC0- ( g/Lh) t1/2 (h) CL (L/h/kg) Vd (L/kg) Free DOX 1746.87 ± 69.94 20.65 ± 1.34 2.68 ± 0.22 18.36 ± 0.80 SCL-DOX 3597.03 ± 99.36** 41.89 ± 3.58** 1.39 ± 0.04** 0.73 ± 0.01**
PEG-SCL-DOX 73730.54 ± 6161.04*** ### 0.06 ± 0.01*** ### 0.62 ± 0.02*** # 0.05 ± 0.01***##
Data are shown as means ± S.E. of at least three independent experiments. AUC: Area under the plasma concentration-time curves. t1/2, Elimination half-life. Vd: Volume of distribution. **, p<0.01 compared to free DOX. ***, p<0.001 compared to free DOX. #, p<0.05 compared to SCL-DOX. ##, p<0.01 compared to SCL-DOX. ###, p<0.001 compared to SCL-DOX.
Figure 3-7. Plasma clearance of different DOX formulation in healthy SD rats. Free DOX or DOX containing nanoliposomes were injected i. v. via tail vein with a single dose of 5 mg DOX/kg. At selected time points after injection (0.033, 0.5, 2, 6, 24 and 48 h), rats
were euthanised and the DOX was extracted and quantified. Data are shown as means S.E. for g DOX per g of tissue (n=4-5).
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3.2.4.3 Establishing standard curves for tissue distribution
Before the quantitative determination of DOX in tissues, tissues (including heart,
liver, spleen, lung and kidney) spiked with a serial dilution of DOX solutions were
prepared to obtain calibration line which was used for assessing the concentration
corresponding to the different peaks in the chromatograms. The relationship between
peak areas and DOX concentrations was assessed over a range of 50-5,000 ng/mL.
The correlation coefficients (R2) for calibration curves was better than 0.996 (Figure
3-8), suggesting a good linear regression within the tested ranges.
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Figure 3-8. Standard curves for tissue distribution. The standard curve was constructed by measurements of tissues spiked with a serial dilution of DOX solutions. The calibration curve was generated by plotting the peak area of fluorescence derived from DOX against the DOX concentration in solution. Standard formulas were determined by linear regression as y = mx + b where y is the peak area of DOX ( V × sec) and x is the DOX concentration (ng/mL).
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3.3.4.4 Tissue distribution advantages of SCL-DOX in SD rats
Next, the BD characteristics were evaluated of SCL-DOX after i. v. injection of a
single dose of 5 mg/kg free DOX or SCL-DOX in healthy SD rats. As shown in
Figure 3-9, treating the rats with SCL-DOX led to a significantly decreased DOX
accumulation in the heart, lung and kidneys. In the heart and lung, the DOX
concentration was significantly lower in the animals receiving SCL-DOX than those
receiving free DOX at all time points, with as much as a 9-fold and 7-fold lower
concentration in the SCL-DOX group at 2 h for the heart and lung, respectively
(Figures 3-8(a) and (d)). Moreover, distribution of DOX at 0.5 h, 2 h and 4 h time
points in the kidney was significantly reduced by i. v. administered SCL-DOX when
compared to free DOX (Figure 3-9(e)). Consistent with the reported enhanced
sequestration of nanoparticles, including liposomes, by organs of the
reticuloendothelial (RES) system (Rahman et al. 1986a; Sugiyama & Sadzuka 2011),
the DOX concentration was significantly higher in the liver of animals receiving
SCL-DOX than those received free DOX (Figure 3-9(b)). In the spleen, the DOX
concentration in the SCL group was significantly higher than the free DOX group
only during the first 2 h following administration. There was no statistically
significant difference in the DOX concentration at 4 h and 24 h points in the spleen
between different treatment groups (Figure 3-9(c)).
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Figure 3-9. Biodistribution of DOX encapsulated in SCL in SD rats. Biodistribution of DOX encapsulated in SCL in SD rats. Healthy rats were injected with a single dose of 5 mg/kg free DOX or SCL-DOX i. v. Rats were euthanised at 0.5 h, 2 h, 4 h and 24 h after the administration. Organs were harvested, washed, weighed, and the DOX was extracted and
quantified. Data are shown as means S.E. for g DOX per g of tissue (n=5-6). ***p<0.001 compared to free DOX.
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3.2.4.5 Drug-to-lipid ratios in vivo
In order to evaluate the in vivo stability of SCL-DOX in various tissues, DOX-to-
lipid ratios in tissues were measured after i. v. injection of a single dose of 5 mg/kg
free DOX or SCL-DOX in healthy SD rats. [3H] cholesteryl hexadecyl ether (CHE)
was added to DOPE as a nonmetabolised, nonexchangeable lipid tracer (Hussain et al.
2007). The [3H] labelled lipids in tissues were analysed using a liquid scintillation
analyser and DOX concentrations were measured using HPLC. The DOPE levels in
tissues were determined by scintillation counting of [3H] followed by converting [3H]
to DOPE using the calculating DOPE-to-[3H] CHE ratio of 1:24000. As shown in
Table 3-1, the DOX-to-DOPE ratio before administration (set as the baseline for
intact SCL-DOX) was 0.5:1. The DOX-to-lipid ratios in different tissues and in
plasma of healthy rats were determined and shown in Figure 3-10. As 3H-CHE is a
nonmetabolised, nonexchangeable lipid tracer, DOX-to-lipid ratios increased when
DOX accumulated or liposome contents were lost. As shown in Figure 3-10, there
were significant increased DOX-to-DOPE ratios in tissues except in the liver and the
spleen which are two major organs for the clearance of liposome formulations
(Drummond, D. C. et al. 1999a) after the i.v. administration, suggesting the breaking
down of DOX encapsulated liposomes by the macrophages of the reticuloendothelial
system (RES) (De Jong & Borm 2008). Meanwhile, the elevated DOX-to-DOPE in
the kidney might be related to the renal clearance of small molecules (Haas et al.
2002). On the other hand, the peak ratio occurred at 2 h point in the spleen and lung
(Figure 3-10 (a)), while in the heart, the highest ratio was at 4 h point following
injection. Conversely, the peak ratio was reached at 0.5 h after SCL-DOX injection
in the liver and kidney (Figure 3-10 (b) and 10 (c)). The early time to peak ratio in
the liver may reflect the sequestration of liposomes by the RES. Moreover, the ratio
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decreased significantly in all tissues after the peak point. Notably, unlike other
organs which displayed an elevated ratio at 24 h compared to 4 h, there was a
persistent decrease of drug-to-lipid ratio in the heart. The early accumulation but
rapid clearance of the drug in the heart (Figure 3-9 (a) and Figure 3-10 (a)) indicated
the potential of reduced cardiotoxicity of DOX encapsulated by SCL. Furthermore,
the significantly increased ratio data in the liver at 24 h suggests that drug-depleted
SCL continued to accumulate in this organ. Meanwhile, there was a significant
increase of drug-to-lipid ratio in the bloodstream at 0.5 h. Then, the ratio decreased
significantly after the peak (Figure 3-10 (d)), indicating that there might be a DOX
accumulation in the plasma in a short period after i.v. administration.
Figure 3-10. DOX-to-lipid ratios in tissues and plasma of rats. Healthy rats were injected with a single dose of with 5 mg/kg [3H]-tracking SCL-DOX i. v. For tissue distribution study, rats were euthanised at 0.5 h, 2 h, 4 h and 24 h after the administration. For plasma study, rats were euthanised at 0.033 h, 0.5 h, 2 h, 6 h, 24 h and 48 h after the administration. Organs
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were harvested, washed, weighed, and [3H] were quantified using liquid scintillation
counting. Data are shown as means S.E. (n=5-6).
3.3 Discussion
In the present study, the physicochemical characteristics and in vitro stability of
different liposomal formulations were investigated. In addition, in vitro cellular
uptake of DOX and SCL-DOX in U118MG and MCF-7 were compared. Meanwhile,
following the instrumental conditions reported in the Methodology Chapter, the
correlation coefficient (R2) for all calibration curves were equal to 0.990 or higher,
presenting a good linear relationship between instrument response and analytical
concentration within the analytical concentration ranges. Furthermore, this is the first
study investigating the PK properties and BD profiles of SCL-DOX in healthy SD
rats.
DOX was successfully encapsulated in the SCL or PEG-SCL efficiently. The PDI
of different formulations were less than 0.2 for all formulations, which indicated that
liposomes had a homogeneous size distribution. Zeta potential is another important
index for the stability of liposomal formulations. High absolute values of zeta
potential indicates high electric charge on the surface of the drug loaded liposomes,
which can prevent the aggregation of liposomes in buffer due to strong repellent
forces (Kim, JY et al. 2009). It was shown that formulations in the study have
negative surface charge. Meanwhile, the zeta potential of SCL and PEG-SCL in this
study was within -30 mV (Table 3-1). The decreased absolute value of zeta potential
in PEG-SCL with 5% PEG2000-DSPE and with 3% PEG2000-DSPE compared to SCL
might be attributed to the inclusion of PEG (Sezgin, Yuksel & Baykara 2006). One
of the possible mechanisms for the PEG effect on surface charge is that the presence
of PEG on the liposome may lead to the slipping plane away from liposome surface
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and hence reduce the zeta potential. Another possible mechanism is the reduced
mobility of liposomes caused by the presence of the PEG chains on the liposome
surface, which could also lead to the decreased zeta potential (Kim, JY et al. 2009).
In the literature, most of the reported DOX-to-lipid ratio after loading was between
0.2:1 and 0.3:1 (w/w) (Hernández, Martí & Estelrich 1991; Hussain et al. 2007;
Mayer et al. 1989) while in this study, a DOX-to-DOPE ratio was 0.50:1 (w/w) for
SCL after encapsulation (Table 3-1). Therefore, the SCL might have the potential to
encapsulate more drugs than many other published formulations of nanoliposomes.
However, the loading efficiency and drug-to-lipid ratio was lower in the PEG-coated
SCL. In a previous study, Wang et al. had indicated that encapsulation efficiency of
liposomes would be decreased significantly when PEG was added (Wang, CH &
Huang 2003). Reduced entrapping efficiency of PEGylated liposome might be a
result of the 3-dimensional steric hindrance around the liposomes inducing by
inserting PEG into lipids bilayers, which would keep drugs from entering the inside
of the vesicles (Wang, CH & Huang 2003).
The stability of the encapsulated drug in nanocarriers, which is essential for
efficient drug delivery to the target site, is another important concept to consider
when interpreting these data. The in vitro stability study suggested that SCL were
very stable for 48 h 37 °C and retain substantial physical stability at 72 h in vitro
(Figure 3-4). Interestingly, the percentage of liposomal DOX leakage after 24 h
incubation in other studies was generally more than 5% (Huang et al. 2010; Lim et al.
1997; Sadzuka et al. 2003; Song, H et al. 2006b), in contrast to minimal payload
leakage of SCL-DOX at 48 h in this study. Moreover, PEG-SCL formulations
displayed better DOX retention properties during dialysis as expected due to the
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sterically stabilised function of PEG (Meyer et al. 1998). Therefore, both SCL and
PEG-SCL displayed superior DOX retention properties in vitro under the
experimental conditions used.
In the cellular uptake study evaluated by flow cytometry, the mean intensity of
fluorescence of DOX in cells treated with free drugs was higher than those treated
with SCL-DOX (Figure 3-5), suggesting a lower cellular uptake of SCL-DOX in
culture dishes. Similar results can be found in previous studies. For example, PEG-
complexed cationic liposome had been shown to have a lower uptake than free DOX
in murine melanoma cells (Jung et al. 2009a). Meanwhile, in human cancer cell lines,
including A375P (human melanoma), MCF-7 (human breast adenocarcinoma) and
HepG2 (human hepatoma), free DOX displayed a higher level of uptake when
compared with their liposomal formulation as well (Song, CK et al. 2009). As the
lower intracellular uptake might be associated with the lower cytotoxicity, the
liposomal formulation might be less toxicity to tumour cells in vitro. However, cells
cultured in monolayer are exposed to a constant drug concentration throughout the
entire assay period, and often take up free DOX more rapidly than liposomal
formulations (Accardo et al. 2012). Therefore, uptake in vitro may not provide a
reliable prediction of therapeutic efficacy in vivo (Xu & Anchordoquy 2011).
The comparison of pharmacokinetic properties between free DOX and SCL-DOX
in healthy SD rats displayed a significantly decreased clearance rate of SCL-DOX
compared to free DOX (p<0.01). The pharmacokinetic studies of SCL-DOX in
healthy SD rats also revealed the elimination half-life of SCL-DOX to be
significantly longer than that of free DOX (p<0.01). The same improvements were
found in the area under the plasma concentration-time curve from time 0 to infinity
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and apparent Vd (Table 3-2), indicating significantly prolonged circulation time and
enhanced bioavailability of SCL-DOX compared to free DOX. In the study, the short
terminal half-life with free DOX in plasma was consistent with data from previous
studies (Rahman et al. 1986a; Wei, G et al. 2008). The terminal half-life showed a
2.03-fold increase in the terminal half-life, suggesting a significantly increased
circulation time of SCL-DOX (p<0.01) as well as a decrease of volume of
distribution (Vd) compared to free drug. Moreover, the significant decreased
clearance rate of DOX encapsulated by SCL-DOX suggested a different rate of
clearance of SCL-DOX compared to free drug. The increased AUC0- of DOX
delivered through SCL (2.06-fold higher than free DOX), suggesting a decreased in
non-specific binding as well as a selective sequestration of the drug to tissues when
administered entrapped in SCL, leading to enhanced bioavailability. The decreased
half-life of PEG-SCL-DOX compared with that of free DOX and SCL-DOX might
be related to the significant clearance from the bloodstream 2 h after
injection.( Figure 3-7), which resulted in a decreased elimination half-life (T1/2).
Unexpectedly, liposomes modified with PEG showed a very short circulation half-
life in vivo. Therefore, although PEG coating has been widely used to prolong the
half-life of liposomal formulations and most studies have indicated that PEG-coated
liposomes generally have a longer circulation life in the bloodstream than
conventional liposomes (Allen, T. M. et al. 1991; Immordino, Dosio & Cattel 2006),
the PEG-SCL-DOX in this study showed no advantage to the circulation life.
Another study has also indicated a shorter half-life of PEG-coated liposomal
doxorubicin compared to free DOX (Wei, G et al. 2008). The possible reasons for the
reduced half-life of our PEG-SCL might be the rapidly exchange out of the lipid
bilayer in vivo (Gregoriadis, G (ed.) 2006). Due to the decreased particle size of
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PEG-SCL-DOX compared with SCL-DOX, the smaller particles might result to the
reduced half-life in the blood stream (Desai, N 2012).
In addition, there was a significant improvement of BD profile of SCL-DOX. As a
widely used and efficient antitumour drug, the risk of severe cardiotoxicity is a well-
known barrier to the use of free DOX in clinical therapy (Tokarska-Schlattner et al.
2005). Encapsulation of DOX into SCL resulted in an approximate 4-fold lower
DOX concentration in the heart of rats receiving SCL-DOX (Figure 3-9), which
represented a significant improvement over those reported by others showing an
approximately 1.5 times lower DOX accumulation in the heart of other liposomal
DOX formulation than free DOX (Xiong, Huang, Lu, Zhang, Zhang & Zhang 2005;
Xiong, Huang, Lu, Zhang, Zhang, Nagai, et al. 2005). As a result, the significant
reduction in accumulation of DOX in the heart indicated the potential of SCL-DOX
in reducing the cardiotoxicity of DOX. This was consistent with observations from
DOX-to-lipid ratio in tissues. The data demonstrated that although there was an early
accumulation of SCL-DOX in the heart, DOX was eliminate rapidly from the tissue
(Figure 3-10 (a)), suggesting an improved safety profile of SCL-DOX in vivo.
3.4 Conclusion
In conclusion, a series of studies had been designed and carried out to
analyse SCL for DOX encapsulation. The simple two-lipid liposomes demonstrated a
satisfactory loading efficiency of DOX and a high stability in vitro. Moreover, this
nanodrug had favourable pharmacokinetic attributes in terms of prolonged
circulation time, decreased clearance rate, reduced volume of distribution and
enhanced bioavailability in healthy rats. As a result of the improved biodistribution
and the targeting property of sulfatide to tenascin-C, SCL might have the potential to
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be an effective and safer nano-carrier for DOX for the treatment of tumours with
elevated expression of tenascin-C.
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Chapter 4 - Enhanced antitumour activity and
reduced systemic toxicity of glioma tumour
microenvironment targeted sulfatide-containing
liposomes
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4. 1 Introduction
Glioblastoma is very heterogeneous and the most common form of malignant
primary brain tumours (Ohgaki 2009). It presents with an unfavourable prognosis
due to its aggressiveness and likelihood of recurrence (Veenman et al. 2010). Thus,
the development of novel anti-glioblastoma treatments is mandatory. Doxorubicin
(DOX) is an anthracycline antibiotic widely used in oncology clinics as an anticancer
chemotherapeutic agents. It was first introduced in the 1970s, and has now become
one of the most commonly used anticancer drugs in the treatment of haematological
malignancies and solid tumours (Arola et al. 2000). However, when a cumulative
dose reaches 450 to 500 mg/m2, it may lead to cardiomyopathy, resulting in
congestive heart failure (Imondi et al. 1996; Olson et al. 1988), which has become
one of the well-known barriers of DOX in clinic utility. Moreover, serious toxicity to
normal tissues, such as myelosuppression, is also leading to a poorer quality of life
for patients (Charrois & Allen 2003). As a result of ongoing efforts in developing
nanoparticle drug delivery systems, the liposomal formulations of doxorubicin, such
as Doxil®, have entered the clinic with improved efficacy and reduced side effects
(Elbayoumi, T. A. & Torchilin, V. P. 2008).
One promising strategy to improve anticancer agents binding and penetrating into
tumours is to develop drug carriers with a component that binds to the tumour cell
surface or the extracellular matrix of tumours. Sulfatide, a lipid that is found in
humans, is involved in a variety of biological processes, such as cell adhesion,
platelet aggregation, cell growth, protein trafficking, signal transduction, neuronal
plasticity and cell morphogenesis. It is known to bind several extracellular matrix
glycoproteins including tenascin-C (Townson et al. 2007) which is overexpressed in
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the microenvironment of most solid cancers, including malignant brain tumours
(Orend & Chiquet-Ehrismann 2006). Studies had recently shown that sulfatide was
specifically required for robust uptake of sulfatide-containing nanoliposomes by
human glioblastoma cells U-87MG (Shao, Ke et al. 2006; Wu, X & Li 1999). The
unique feature of this nanoliposome is that it is comprised of two natural lipids found
in human cells, namely sulfatide and DOPE. Thus, this nanoliposome is totally
human compatible and degradable. The sulfatide-containing liposomal (SCL-DOX)
had been found to remain intact for hours after uptake by the glioblastoma cells
(Shao, Ke et al. 2006; Wu, X & Li 1999). Intracellular distribution studies had
indicated a high accumulation of DOX in the nuclei where it exerted its cytotoxic
effect after 12 h incubation with SCL-DOX at 37 °C (Shao, Ke et al. 2006; Wu, X &
Li 1999).
Recognising the potential of the use of a tumour environment targeting ligand as
one of the main structural constituents of the nanocarriers capable of both passive
and active targeting, a series of studies had been designed and carried out to
determine the in vitro stability, the pharmacokinetic behaviour and the
biodistribution pattern of the SCL-DOX (Chapter 3). The data have indicated that the
sulfatide-containing liposomal carrier system represents a new class of natural lipid-
guided intracellular delivery systems which have the potential to increase treatment
efficacy and reduce systematic toxicity of the drug encapsulated.
Therefore, in the current study, cellular uptake and retention of SCL-DOX in U-
118MG glioblastoma cells were examined. The potential clinical utilities of this
novel class of natural lipid-guided and tumour microenvironment targeting
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nanoliposome were demonstrated via studies on its biodistribution in both U-118MG
tumour-bearing mice, as well as its antitumour efficacy and toxicity profiles.
The work presented here was performed in collaboration with Yan Yu.
Parts of this work have been submitted to Nanomedicine: Nanotechnology,
Biology and Medicine for publication.
4. 2 Results
4.2.1 In vitro cytotoxicity
In order to assess the in vitro cytotoxicity of SCL-DOX to glioma cells, U-118MG
glioblastoma cells were incubated with different formulations of DOX at various
concentrations (0-100 g/mL) for 48 h and the cytotoxicity of free DOX and SCL-
DOX to cells was compared using the MTT assay which measures the mitochondrial
conversion of 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyl-2H-tetrazolium bromide
(MTT) to formazan as detected by the change of optical density at 570 nm (Jung et al.
2009a). The IC50 values for DOX and SCL-DOX are shown in Table 4-1. The IC50
for free DOX (2.51 ± 0.33 g/mL) in U-118MG cells was lower than that observed
in U-118MG cells treated with SCL-DOX (19.55 ± 0.68 g/mL). These results
suggested that U-118MG cells were relatively more sensitive to free DOX than to
SCL-DOX following exposure to a constant concentration of the agent. There was no
significant toxicity of blank SCL after 48 h incubation (data not show). The lower
toxicity resulting from SCL-DOX to U-118MG cells in vitro is in good agreement
with the literature for other liposomes loaded with free DOX (Chen, Z et al. 2012;
Shmeeda et al. 2010) and is consistent with the previous stability study (Section in
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Chapter 3.2.2.2) as the in vitro cytotoxicity is related to the release rate of DOX from
nanocarriers (Song, H et al. 2006a).
Table 4-1. Mean IC50 values ( g/mL of DOX) for treatment with free DOX and SCL-
DOX. Cell line Free Dox ( g/ml) SCL-DOX ( g/ml) U-118MG 2.51 ± 0.33 19.55 ±0.68
Data are shown as means ± S.E. of three independent experiments performed in triplicate.
4.2.2 Intracellular uptake and retention of SCL-DOX in U-118MG cells
The accumulation and retention of DOX and SCL-DOX in U-118MG cells were
studied using laser scanning confocal microscopy utilising the natural fluorescent
property of DOX. Cellular uptake of free DOX or SCL-DOX by U-118MG was
examined following 24 h incubation with 2 g/mL DOX or equivalent amounts of
SCL-DOX. As shown in Figure 4-1, both free DOX and SCL-DOX accumulated in
the glioblastoma cells. The overlay of Hoechst staining (nucleus) and red
fluorescence (DOX) indicated that SCL-DOX was not adhered to the cell surface but
actually penetrated into the nucleus. However, there was a slightly stronger DOX
fluorescence (red) in cells treated with free DOX when compared to those treated
with SCL-DOX after 24 h incubation. Interestingly, SCL-DOX was better retained
by the U-118MG glioma cells. As shown in Figure 4-2 (a) and (c), there was a
significant decrease of DOX fluorescence in cells treated with free DOX only 2 h
after washing with PBS. In contrast, DOX fluorescence could be found in glioma
cells treated with SCL-DOX even 24 h after washing (Figure 4-2 (b) and (d)),
suggesting more sustained retention of DOX in U-118MG cells when delivered via
SCL-DOX than by free DOX. The improved retention of DOX encapsulated with
SCL in vitro implies the potential of better treatment efficacy of SCL-DOX in vivo.
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Figure 4-1. Intracellular uptake of SCL-DOX in U-118MG cells. Intracellular uptake of
SCL-DOX in U-118MG cells. U-118MG cells were incubated with 2 g/mL free DOX or equivalent SCL-DOX for 24 h. Cells were washed twice with PBS and imaged with a confocal fluorescence microscope. (a) Cells treated with free DOX (low magnification). (b) Cells treated with SCL-DOX (low magnification). (c) Cells treated with free DOX (high magnification). (d) Cells treated with SCL-DOX (high magnification). Red: fluorescence from DOX; blue: nuclei stained with Hoechst 33342. Data are representative of three independent experiments. Scale bars: 200 m (for (a) and (b)) or 20 m (for (c) and (d)).
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Figure 4-2. Intracellular retention of SCL-DOX in U-118MG cells. U-118MG cells were
first incubated with 2 g/mL free DOX or equivalent SCL-DOX for 24 h. Cells were then
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washed with PBS twice and cultured in fresh full culture medium. The same wells of cells were imaged serially at 1 h, 2 h, 4 h and 24 h after washing using fluorescence confocal microscopy. (a) Cells treated with free DOX (low magnification). (b) Cells treated with SCL-DOX (low magnification). (c) Cells treated with free DOX (high magnification). (d) Cells treated with SCL-DOX (high magnification). Data are typical of three independent experiments. Red: fluorescence from DOX; blue: nuclei stained with Hoechst 33342. Scale bars: 200 m (for (a) and (b)) or 20 m (for (c) and (d)).
4.2.3 Establishing standard curves for tissue distribution
Before the quantitative determination of DOX in tissues, the linearity of the
calibration curves correlating the fluorescence intensity measured by HPLC and the
amounts of DOX in different tissues (including heart, liver, spleen, lung, kidney and
tumour) was verified over the range of 10-500 ng/mL using tissues spiked with a
serial dilution of DOX solutions. As shown in Figure 4-3, the correlation coefficients
(R2) for calibration curves were better than 0.996. Thus, under the instrumental
conditions reported in Chapter 2 (Section 2.3.1.3), a good linearity between peak
areas and concentrations was obtained over the analytical concentration ranges.
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Figure 4-3 Standard curves for tissue distribution in mice. Standard curve was constructed by measurements of tissues or tumours spiked with a serial dilution of DOX solutions. The calibration curve was generated by plotting the peak area of DOX against the DOX concentration in solution. Standard formulas were determined by linear regression as y
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= mx + b where y is the peak area of DOX ( V × sec) and X is the DOX concentration (ng/mL).
4.2.4 Improved tumour uptake and biodistribution in U-118MG tumour
xenograft model
Next, the ability of SCL-DOX to enhance the delivery of therapeutic agents to the
tumour in vivo was investigated using a mice tumour xenograft model. The uptake of
DOX in various organs and the tumour was determined 24 h after a single dose of 5
mg/kg DOX or SCL-DOX by i. v. administration in U-118MG glioblastoma-bearing
nude mice. Consistent with the results from the study in healthy SD rats in the
preceding chapter (Chapter 3), the DOX concentration in organs responsible for
dose-limiting toxicity in clinics, i.e. the heart, in the tumour-bearing mice treated
with SCL-DOX was significantly lower than those treated with free DOX (33.07 ±
1.34 g/g versus 49.07 ± 2.35 g/g) (Figure 4-4 (a)). The DOX concentration in the
other known major DOX toxicity organ, the skin, was also statistically significantly
lower in the SCL-DOX group compared to the free DOX animals (0.97 ± 0.06 versus
1.61 ± 0.11 g/g), as well as in the kidneys (403.94 ± 9.99 versus 472.35 ± 20.08
g/g). On the other hand, the measured concentration of DOX in other organs
indicated that there were significantly higher levels of DOX in the liver, spleen and
lung in the group treated with SCL-DOX when compared to the free DOX group
(1.40-fold, 3.78-fold and 1.71-fold, respectively) (Figures 4-4(b) and 4-4(e)). As for
the xenograft glioma (Figures 4-4(d)), there was a statistically significant elevation
of DOX level in tumour tissue in the SCL-DOX group compared to that of free DOX
(1.33-fold), indicating the enhanced intratumoural DOX delivery by SCL-DOX in
vivo.
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Figure 4-4. BD and tumour uptake of DOX encapsulated in SCL in xenograft-bearing
mice. Nude mice bearing human glioblastoma U-118MG xenografts were treated with a single dose of 5 mg/kg free DOX or SCL-DOX i.v. Mice were euthanised 24 h later. Organs and tissues were harvested, washed, weighed, and the DOX content in tissues, expressed as
g DOX per g tissue, was determined. Data are shown as means S.E. (n=5-6). *, p<0.05 compared to free DOX; **, p<0.01 compared to free DOX; ***, p<0.001 compared to free DOX.
4.2.5 Enhanced therapeutic efficacy of SCL-DOX in U-118MG xenograft
tumour model
Given that the SCL-DOX is able to deliver more therapeutic agents to the
xenograft tumour (Figure 4-4d), the antitumour activity of SCL-DOX in vivo was
proceeded to determine. Mice bearing U-118MG tumours were injected with saline,
blank SCL, DOX in solution or encapsulated within SCL once a week for 6 weeks
when subcutaneous tumours reached a volume of 150 mm3. Figure 4-5 shows tumour
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growth inhibition. There was no significant difference in tumour sizes between mice
treated with saline control and blank SCL during the study period. Importantly, at a
dose of 5 mg/kg, all the DOX formulations were effective in preventing tumour
growth compared to saline and blank liposome control after the 2nd injection. The
final mean tumour load was 97.29 10.71 mm3 in the SCL-DOX treatment group
while in free DOX group it was 154.76 12.53 mm3. Thus, SCL-DOX formulations
displayed stronger tumour growth suppression than free DOX.
Figure 4-5. Improved therapeutic activity of SCL-DOX against glioma xenograft. Mice bearing U-118MG xenografts were injected i.v. with saline, 5 mg/kg of free DOX, SCL-DOX or empty SCL once a week for 6 weeks when tumour volume reached approximately
150 mm3. Data shown are means S.E. (n=5-6). *, p<0.05 compared to saline; **, p<0.01 compared to saline; #, p<0.05 compared to free DOX; ***, p<0.001 compared to free DOX; ##, p<0.05 compared to free DOX; &, p<0.01 compared to blank SCL; &&, p<0.01 compared to blank SCL; &&&, p<0.001 compared to blank SCL.
To further confirm the antitumour efficacy, the survival rates of tumour-bearing
mice were compared following different treatment regimens. As shown in Figure 4-6,
the median survival days for saline, free DOX and SCL-DOX group were 45, 61 and
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93 days, respectively. Therefore, the medium life-span was increased in mice treated
with SCL-DOX by 2.07-fold and 1.52-fold compared to those treated with saline or
free DOX, respectively. Taken together, i. v. administration of SCL-DOX 6 times
over a 6-week period displayed not only a better tumour inhibitory effect but also an
improved survival of U-118MG xenograft-bearing mice.
Figure 4-6. SCL-DOX enhanced survival of tumour-bearing mice. The Kaplan-Meier survival curve shows improvement of life span of U-118MG xenograft-bearing mice treated with SCL-DOX (n= 8-10 per group). Mice were treated as indicated in the legend for Fig. 6 and were sacrificed when tumour burden reached 800 mm3.
4.2.6 Reduced toxicity of SCL-DOX in vivo
Clinically, the efficacy of DOX is limited by dose-limiting toxicities. One
objective of delivery of chemotherapy agents in a nano-formulation is to reduce
systemic toxicities. For this aim, the toxicity of SCL-DOX and free DOX was
evaluated after repeated injection in tumour-bearing nude mice. The tissue
concentration of DOX was measured 24 h after the 6th administration of either SCL-
DOX or free DOX (Figure 4-7). Consistent with the findings in healthy rats, the
concentrations of DOX in the heart and skin were significantly lower, 0.56-fold and
0.08-fold, respectively, in the tumour-bearing mice treated with SCL-DOX than
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those treated with free DOX following repeated administrations. On the contrary,
compared with the free DOX group, 1.28-fold more DOX accumulation was
observed in the tumour in SCL-DOX group, clearly confirming the enhanced
intratumoural DOX delivery by SCL-DOX in vivo. As for plasma biochemistry
analysis carried out 72 days after the last injection (Figure 4-8), DOX treatment
alone induced a significant increase in serum creatine kinase (CK) concentration
(1048.00 ± 100.95 U/L), indicative of heart damage, compared to saline control
(588.50 ± 167.37 U/L), blank SCL (543.40 ± 86.47) and SCL-DOX groups (430.60 ±
82.94 U/L). Furthermore, compared with the free DOX group, the plasma
concentrations of aspartate transaminase (AST), a sensitive indicator of liver damage,
in the SCL-DOX group (p<0.01) and blank SCL group were significantly lower
(p<0.05). Moreover, as a marker of DOX-damaged myocytes, troponin was
measured for the evaluation of DOX-induced cardiomyopathy in vivo (Herman et al.
1999; Singal & Iliskovic 1998). The method used in the present study has a cut-off
threshold of <0.01 g/L for normal subjects (Koh, Nakamura & Takahashi 2004). As
shown in Table 4-2, there was a 1.8-fold higher level of troponin in the mice treated
with free DOX than those with saline control while in both SCL-DOX and blank
SCL group, there were no significant elevation compared to the control. Thus, the
results indicated that the SCL-DOX has the potential to minimise the cardiotoxicity
of free DOX.
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Figure 4-7. Encapsulation of DOX in SCL substantially reduced the accumulation of
DOX in principal sites of DOX toxicity. U-118MG xenograft-bearing mice were treated as indicated in the legend for Figure 4-5. Blood was collected immediately after the mice were
sacrificed upon reaching the end point. Data shown are means S.E. (n=5-6). *, p<0.05 compared to saline; **, p<0.01 compared to saline; ***, p<0.001 compared to free DOX.
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Figure 4-8. Encapsulation of DOX in SCL significantly reduced cardiac and hepatic
toxicity. U-118MG xenograft-bearing mice were treated as indicated in the legend for Figure 4-5. Blood was collected immediately after the mice were sacrificed upon reaching the end point. Serum enzymes indicative of cardiac and hepatic toxicity were analysed. Data shown
are means S.E. (n=3-5). *, p<0.05 compared to saline; #, p<0.05 compared to free DOX; ##, p<0.01 compared to free DOX.
Table 4-2. Plasma levels of troponin.
Treatment Number of nude mice Troponin ( g/L) Saline 4 <0.01 Free DOX 5 0.018 ± 0.003 SCL-DOX Blank SCL
5 5
<0.01 <0.01
Data are presented as means ± S.E. (n = 4-5).
4.3 Discussion
As liposomes are composed of naturally biodegradable substances, they are
metabolised and cleared while in circulation or upon reaching the target sites, making
them safe novel drug delivery carriers (Domb et al. 2007). The current study
provided the first report on antitumour efficacy of tenascin-C-targeted SCL-DOX in
a xenograft model of glioma that expresses high levels of tenascin-C (Reardon,
Zalutsky & Bigner 2007).
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As for in vitro cytotoxicity between different liposomes and cancer cell lines,
published studies have reported contradictory results. For example, in MLLB2 cells
(rat prostate cancer) (Wang, J et al. 2005) and MCF-7/ADR cells (human breast
carcinomas) (Li, B et al. 2012), the IC50 of the liposomal formulation was
significantly lower than that of free DOX, which indicated a higher cytotoxicity of
liposomal DOX in vitro. On the contrary, in HepG2 cells (human hepatoma) (Li, X et
al. 2009) and U-87 cells (human glioblastoma) (Shao, Ke et al. 2006), free DOX
seemed to have higher intracellular uptake with associated higher cytotoxicity than
that of liposomal DOX. Obviously, the kinetic properties are different between
liposomal DOX and free DOX. The half-life of liposomal DOX can be up to several
days while free DOX can be eliminated in a few minutes in vivo (Allen, Theresa M.
et al. 2006; Cabanes et al. 1999; Unezaki et al. 1994). Moreover, the MTT assay used
to derive the IC50 is carried out with monolayers in culture dishes, which are very
different when compared to the 3-dimentional tissue architecture in vivo (Xu &
Anchordoquy 2011). Thus, comparison of IC50 in vitro, which is relevant to the
cytotoxicity under the same exposure time and constant concentration, is not a
reliable predictor of the therapeutic efficacy in vivo (Wu, J et al. 2007). Furthermore,
SCL-DOX had been found to remain in the nuclei for several hours even after
washing (Figure 4-2 (b) and (d)). This is in agreement with another study using the
same DOX delivery system in a different cell line (Shao, Ke et al. 2006). Notably,
despite the higher IC50 value of SCL-DOX in vitro, it had a much better anti-tumour
efficacy over the free DOX in U-118MG tumour-bearing nude mice (Figure 4-5).
As a widely used and efficient antitumour drug, the risk of severe cardiotoxicity is
a well-known barrier of free DOX in clinical therapy (Tokarska-Schlattner et al.
2005). Encapsulation of DOX into the SCL resulted in an approximate 4-fold lower
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DOX concentration in the heart of healthy SD rats (Figure 3-8 in Chapter 3).
Reduced accumulation of DOX derived from SCL in the heart was also found in U-
118MG tumour-bearing nude mice after single injection and repeated injection
(Figures 4-4 and 4-7). Therefore, the significant reduction in accumulation of DOX
in the heart indicated the potential of SCL-DOX in reducing the cardiotoxicity of
DOX. This had been reinforced by the biochemical studies of the serum creatine
kinase activity, which is a toxicological indicator of severe cardiotoxicity (Bagchi et
al. 2003) as well as cardiac troponin, another biomarker for the detection and
prevention of cardiotoxicity at an earlier phase (Mercuro et al. 2007). Previous
studies by others revealed an increase in serum troponin levels from week 10 after
the first administration of DOX in Wistar rats treated with liposomal DOX (Koh,
Nakamura & Takahashi 2004). This present study revealed no discernable increase in
serum troponin levels in mice treated with SCL-DOX even 14 weeks after the onset
of treatment, suggesting a remarkable reduction in cardiotoxicity of DOX delivered
via SCL. Moreover, the study also found a significant reduction in the amounts of
serum AST in mice treated with SCL-DOX (Fig. 9). This is important as DOX is
excreted predominantly through the hepatobiliary route (Cosan et al. 2008) and there
is a good negative correlation between serum AST activity and hepatic activity
(Yokogawa et al. 2006).
Although encapsulation with liposomes has been successful in overcoming
cardiotoxicity and myelosuppression (for free DOX), the toxicity of liposomal DOX
has shifted to cutaneous toxicity (Charrois & Allen 2003). Palmar-plantar
erythrodysesthesia (PPE), also called hand-foot syndrome, is a toxic reaction
associated with high accumulation of cytotoxic chemotherapeutics, including
pegylated liposomal doxorubicin formulation in the skin (Farr, Katherina
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Podlekareva. & Safwat, Akmal. 2011). High liposome localisation in the skin had
been reported previously in both nude mice and human (Goren, D. et al. 1996). In
this study, a significantly reduced concentration of DOX in the skin was observed in
mice treated with SCL-DOX repeatedly when compared to the free DOX group
(Figure 4-7), suggesting that the SCL formulation of DOX could help to reduce the
dose-limiting cutaneous toxicity displayed by other liposomal formulations of DOX.
In contrast, the uptake of DOX in liver and spleen, which are tissues rich in cells
of the reticuloendothelial system (RES), was higher for SCL-DOX when compared
to free DOX (Figure 4-4), in agreement with previous studies (Elbayoumi, TA &
Torchilin, VP 2009; Xiong, Huang, Lu, Zhang, Zhang & Zhang 2005). The increased
accumulation of SCL-DOX in the organs of RES might be related to the particle size
of the SCL (Campbell 2006; Drummond, D. C. et al. 1999b). Therefore, further work
is needed to reduce the uptake of SCL-DOX by RES.
As well as a benefit of reduced toxicity by SCL-DOX formulations, better
treatment efficacy from animals treated with SCL-DOX over free DOX was found in
the tumour-bearing mice (Figures 4-5 and 4-6). Significantly higher concentrations
of DOX in tumour from SCL-DOX treated mice were found both after single
injection (Figure 4-4) and repeated injections (Figure 4-7). Administration of SCL-
DOX had superior tumour inhibitory effects compared to that of free DOX (Figure 4-
5), manifested as both the inhibition of tumour growth and the increased life span
(Figure 4-6). Importantly, tumours grew much slower in the mice receiving SCL-
DOX when compared with those receiving free DOX. As the repeated injection of
liposomal formulation could lead to the change of pharmacokinetics of drugs (Cui et
al. 2008), the difference of tissue concentrations between after single injection and
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after repeated injections might be resulted from the separated pharmacokinetic
property of SCL-DOX.
Nano-carriers can be used to improve the treatment efficacy and reduce the side
effects of drugs they encapsulate. Due to its high surface-to-volume ratio,
functionalising the surface of nanoparticles with ligands such as antibodies, aptamers,
peptides, or small molecules that are tumour-speci c or tumour-associated can
promote the active binding of nanoparticles to tumours (Alexis, F. et al. 2008). The
composition and structure of the extracellular matrix in tumours are different from
that in the normal tissues (Liu, F et al. 2008). Certain extracellular matrix
glycoproteins are highly up-regulated in many different cancers, including gliomas,
breast cancer and ovarian cancer (Fernando et al. 2008; Quemener et al. 2007).
Tenascin-C is a protein expressed at low levels in normal adult tissues but high levels
in many tumours (Mackie & Tucker 1999), including gliomas. Therefore, tenascin-C
has been implicated as an important target for the treatment of cancer (Adams, M et
al. 2002). In a previous study, the key component of SCL, sulfatide, was
demonstrated to mediate the binding and endocytic uptake of SCL in tumour cells via
the interaction with tenascin-C (Alvarez-Cedron, Sayalero & Lanao 1999). It was
evident from the current results that SCL conferred sufficient therapeutic activity, as
well as reduced side effects of its encapsulated drug, in the tenascin-C expressing
tumour model (U-118MG) used here. Thus, modification with other ligands, which
could target the cancer stem cells, on the surface of SCLs will likely aid the
development of novel therapeutic strategies for treating tenascin-C positive tumours.
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4.4 Conclusion
In summary, the tenascin-C targeting SCL displayed increased therapeutic effects
in addition to the reduced toxicity in vivo. These results taken together, therefore,
demonstrate that SCL has the potential to be an effective and safer nano-carrier for
the treatment of tumours with elevated expression of tenascin-C. Further
improvement in therapeutic index can be afforded by the functionalisation of SCL
with other cancer-targeting ligands, such as antibodies or aptamers.
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Chapter 5 - Enhanced antitumour efficacy and
reduced systemic toxicity of sulfatide-
containing nanoliposomal doxorubicin in a
xenograft model of colorectal cancer
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5.1 Introduction
Colorectal cancer is the third most common cause of cancer-related deaths
worldwide (Sala-Vila & Calder 2011; Szepeshazi et al. 2002), with up to 25% of
patients presenting with metastatic disease. Despite surgery and chemotherapy, many
of these patients eventually succumb to metastatic diseases. Adjuvant therapies,
including radiotherapy and chemotherapy, are designed to target residual tumour
cells. In stage III patients with colorectal cancer, chemotherapy remains the main
treatment strategy (Bhattacharya et al. 2009). However, the success of these therapies
is limited by the emergence of therapy-resistant cancer cells as well as dose-limiting
toxicities (Segal & Saltz 2009). Liposomes are well-recognised vehicles for drug
delivery in recent years. They have been indicated to enhance the therapeutic
efficacy of several traditional antitumour drugs (Lasic, D. D. 1998). Ideal designed
liposomes for drug delivery have the ability to passively accumulate in tumour
tissues via the EPR effects in tumours. The passive accumulation can be further
enhanced by actively targeting liposomes modified by coupling ligands on the
vesicle surface (Forssen, Coulter & Proffitt 1992).
The tumour microenvironment has been shown to be one of the key factors
contributing to the development of cancer (Hanahan & Weinberg 2011). Tenascin-C
is a glycoprotein found in the extracellular matrix and it is an important target for
targeted-drug delivery in the treatment of cancer (Hicke, B. J. et al. 2006). The novel
strategy employed in this study is to use a lipid component of the liposome to target
the tumour microenvironment. Due to the targeting property of sulfatide, a
glycosphingolipid known to interact with several extracellular matrix proteins, to
tenascin-C, sulfatide-containing liposomes (SCL) have the potential to be an
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effective and safer nano-carrier for the treatment of tumours with elevated expression
of tenascin-C.
In view of the limited success of chemotherapy in colorectal cancer and high
toxicity of DOX, a SCL encapsulation approach was taken to overcome these
barriers. In Chapter 4, the efficacy of SCL against glioma suggested an enhanced
treatment efficacy of sulfatide-containing liposomal DOX (SCL-DOX) and a
reduction in the accumulation of DOX in the drug’s principal toxicity organs
achieved by SCL-DOX led to a diminished systemic toxicity as evident from the
plasma biochemical analyses. The SCL-DOX had also demonstrated a good stability
in vitro, as well as improved pharmacokinetic behaviour and enhanced BD profiles
(Chapter 3). Now, the study was extended to colorectal cancer.
In this study, a mouse xenograft model of human colorectal adenocarcinoma (HT-
29) that is known to express tenascin-C (De Santis et al. 2006; Mukaratirwa et al.
2005) was utilised to study the biodistribution (BD), antitumour efficacy and toxicity
of sulfatide-containing liposomal carrier system.
The work presented here was performed in collaboration with Yan Yu.
Parts of this Chapter have been published (Lin J, Yu Y, Shigdar S, Fang DZ, Du
JR, Wei MQ, Danks A, Liu K and Duan W. “Enhanced antitumor efficacy and
reduced systemic toxicity of sulfatide-containing nanoliposomal doxorubicin in a
xenograft model of colorectal cancer”. PLoS ONE, 2012, 7(11): e49277.)
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5.2 Results
5.2.1 Intracellular uptake and retention of SCL-DOX in HT-29 cells
Taking advantage of the natural fluorescent property of DOX, the cellular uptake
and retention of free DOX or SCL-DOX was studied using laser scanning confocal
microscopy. HT-29 cells were incubated with 2 g/ml DOX or SCL-DOX for 24 h.
Following washing, the cellular uptake of different formulations of DOX was
examined. As shown in Figure 5-1, both free DOX and SCL-DOX were taken up by
the colorectal adenocarcinoma cells and there was accumulation of DOX in the
nucleus in both groups (Figure 5-1), albeit cells treated with free DOX showed
slightly stronger red fluorescence (DOX) than those treated with SCL-DOX after 24
h incubation. Interestingly, the retention of SCL-DOX in HT-29 cells was better than
that of free DOX. As shown in Figure 5-2, following washing with PBS and
incubation in fresh media for 4 h, the DOX fluorescence in the free DOX group
decreased significantly. Moreover, cells treated with free DOX showed diminished
red fluorescence 24 h after washing. Conversely, the red fluorescence for SCL-DOX
was more stable compared to that of the free DOX group. Even 24 h after washing,
DOX fluorescence could be readily observed in the nuclei of cells treated with SCL-
DOX (Figures 5-2(b) and (c)). The enhanced retention of SCL-DOX in vitro suggests
that the SCL formulation of DOX might exhibit better treatment efficacy in vivo.
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Figure 5-1. Intracellular uptake of SCL-DOX in HT-29 cells. HT-29 cells were incubated
with 2 g/mL free DOX or equivalent SCL-DOX for 24 h. Following two washes with PBS, cells were imaged with a confocal fluorescence microscope. (a) and (c) Cells treated with free DOX. (b) and (d) Cells treated with SCL-DOX. Red: fluorescence from DOX; blue: nuclei stained with Hoechst 33342. Scale bars: 200 m (for (a) and (b)) or 10 m (for (c) and (d)).
142
143
Figure 5-2. Intracellular retention of SCL-DOX in HT-29 cells. HT-29 cells were first
incubated with 2 g/mL free DOX or equivalent SCL-DOX for 24 h. After two washes with PBS to remove the drugs, cells were cultured in fresh full culture medium followed by imaging serially at 1 h, 2 h, 4 h and 24 h using fluorescence confocal microscopy. (a) and (c) Cells treated with free DOX. (b) and (d) Cells treated with SCL-DOX. Red: fluorescence
144
from DOX; blue: nuclei stained with Hoechst 33342. Scale bars: 200 m (for (a) and (b)) or 10 m (for (c) and (d)).
5.2.2 In vitro cytotoxicity
To study the in vitro cytotoxicity, HT-29 cells were exposed to various
concentrations of free DOX or SCL-DOX for 48 h, and the cell viability was
measured using the MTT assay which measures the mitochondrial conversion of
MTT (3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyl-2H-tetrazolium bromide) to
formazan as detected by the change of optical density at 570 nm (Li, G et al. 2012).
As shown in Table 5-1, the IC50 of DOX for HT-29 cells was 1.74 ± 0.10 g/mL
while the IC50 of SCL-DOX was 2.77 ± 0.06. Thus, under in vitro conditions where
cells were exposed to a constant concentration of the agents throughout the entire
assay period, free DOX was more toxic than SCL-DOX. Empty SCL did not show
any effects on cell survival (data not shown).
Table 5-1. Mean IC50 values ( g/ml of doxorubicin) for treatment with free DOX and SCL-DOX
Cell line Free dox ( g/ml) SCL-dox ( g/ml) HT-29 1.74 ± 0.10 2.77 ±0.06
Data are shown as means ± S.E. of triplicate in three independent experiments.
5.2.3 Establishing standard curves for tumour uptake
Before the quantitative determination of DOX in tissues, HT-29 tumours spiked
with a serial dilution of DOX concentrations were prepared to obtain the calibration
curve which was used for determination of the concentration of DOX from the
measured fluorescence intensity in chromatograms.. As shown in Figure 5-3, the
correlation coefficient (R2) for the calibration curve was 0.996. Thus, under the
instrumental conditions described in Chapter 2, a good linearity between peak areas
145
and concentrations was obtained over the analytical concentration ranges within the
range of 1-100 ng/mL.
Figure 5-3. Standard curve for DOX concentration in HT-28 tumours. A standard curve was constructed by measuring fluorescence in tissues or tumours spiked with a serial dilution of DOX solutions. The calibration curve was generated by plotting the peak area of fluorescence derived from DOX against the DOX concentration in solution. Standard formulas were determined by linear regression as y = mx + b where y is the peak area of DOX ( V × sec) and x is the DOX concentration (ng/mL).
5.2.4 Tissue distribution and tumour uptake Advantages of SCL-DOX
Studies comparing the accumulation of free DOX or SCL-DOX in tumours and
organs were performed in a BALB/c nude mice HT-29 tumour xenograft model.
Animals were injected i.v. with a single dose of free DOX or SCL-DOX (5 mg/kg)
and there was no statistically significant difference in DOX concentration in the
kidneys between the two treatment groups 24 h after administration. The lungs and
liver showed higher DOX accumulation (4.74-fold and 12.94-fold, respectively) with
SCL-DOX treatment (Figure 5-4(a)), and the spleen, a major organ of the
reticuloendothelial system, showed a 17-fold higher DOX accumulation with SCL-
DOX treatment. However, SCL-DOX treatment in the two principal organs that
display dose-limiting toxicities of DOX clinically, namely the skin and heart,
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decreased the DOX accumulation to 59.0% (0.039 0.001 g/g versus 0.066 0.003
g/g) and 77.4% (0.956 0.073 g/g versus 1.235 0.083 g/g) compared to the
free DOX, respectively (Figure 5-4(b) and (c)). Moreover, SCL encapsulation
significantly enhanced DOX accumulation (1.3-fold; 0.060 0.005 g/g versus
0.047 0.003 g/g) in the xenograft tumour compared to free DOX (Figure 5-4(d)),
clearly confirming the enhanced intratumoural DOX delivery by SCL-DOX in vivo.
Figure 5-4. BD and tumour uptake of DOX encapsulated in SCL in xenograft-bearing
mice. Nude mice bearing human colorectal cancer HT-29 xenografts were treated with 5 mg/kg free DOX or SCL-DOX i.v. Mice were euthanised 24 h later. Organs and tissues were harvested, washed, weighed, and the DOX was extracted and quantified. Data are
shown as means S.E. (n=5~6). *, p<0.05 compared to free DOX; **, p<0.01 compared to free DOX.
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5.2.5 Enhanced therapeutic efficacy of SCL-DOX
Having determined the tissue distribution of SCL-DOX, the antitumour activity of
SCL-DOX was evaluated using the BALB/c nude mice HT-29 tumour xenograft
model. Once the tumour had grown to approximately 35 mm3, the animals were
divided randomly into four groups (n=5~10) in order to minimise difference in
weight and tumour size among the groups. The following regimens were
administered i.v. twice a week for 3 weeks: (i) saline; (ii) empty SCL; (iii) free DOX
(5 mg/kg) and (iv) SCL-DOX (5 mg/kg). The body weight of the animals and the
tumour size were then monitored until the size of the tumour in the control animals
reached the end point of the study. As presented in Figure 5-5, for the control groups
of mice receiving saline or empty SCL, the treatment did not show any efficacy, and
the mean tumour sizes at the end of the study were 1129.03 55.06 mm3, and
1188.63 137.54 mm3, respectively (mean S.E.; n=5~6). The SCL-DOX treatment
group demonstrated superior efficacy, with a final mean tumour load of 586.52
29.63 mm3, compared to 809.13 43.75 mm3 in the free DOX group. Thus,
compared to saline or free DOX treatment, the efficacy of SCL-DOX to suppress
tumour growth at the dose of 5 mg/kg was significantly improved by ~1.9-fold and
~1.4-fold, respectively.
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Figure 5-5. SCL-DOX produced more reduction in tumour volume. Mice bearing HT-29 xenografts were injected i.v. with saline, 5 mg/kg of free DOX, SCL-DOX or empty SCL twice a week for 3 weeks as indicated, starting on the day when tumour volume reached ~35
mm3. Data shown are means S.E. (n=5~6). *, p<0.05 compared to saline; **, p<0.01 compared to saline; #, p<0.05 compared to free DOX; &&, p<0.01 compared to blank SCL; &&&, p<0.001 compared to blank SCL.
Next, the survival rates of tumour-bearing mice following the four different
treatment regimens were compared. As shown in Figure 5-6, median survival times
for the four different groups were 26 days (saline), 33 days (free DOX), 36 days
(SCL-DOX) and 32 days (blank SCL), respectively. Thus, SCL-DOX treatment
increased medium life-span by 38.5% compared to the saline control group, by 12.5%
compared to the blank SCL group and by 9.1% compared to the free DOX group.
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Figure 5-6. SCL-DOX enhanced survival. The Kaplan-Meier survival curve shows improvement of life span of xenograft-bearing mice treated with SCL-DOX (n=9~10 per group). Mice were treated as indicated in Figure 5-3 and were sacrificed throughout the study period upon reaching the study end point.
5.2.6 Reduced systemic toxicity of SCL-DOX
DOX-induced cardiomyopathy is one of the key dose-limiting toxicities of the
drug (Rahman et al. 1986b). Cardiac troponin is released from DOX-damaged
myocytes (Wei, G et al. 2008), therefore, measurement of serum levels of this protein
provides a sensitive assessment of early cardiotoxicity of DOX. The method used in
the present study has a cut-off threshold of <0.01 g/L for normal subjects
(Sugiyama & Sadzuka 2011). As shown in Table 5-2, the free DOX treatment
resulted in a 75-fold higher troponin serum level compared to the controls,
confirming the known cardiotoxicity of the free drug. However, no elevation of
serum troponin was observed for the SCL-DOX treatment group, which remained
below cut-off levels, as with the control groups, indicating that treating xenograft-
bearing mice with 6 doses of SCL-DOX over the period of 4 weeks had minimal
cardiotoxicity.
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Table 5-2. Plasma levels of troponin Treatment Number of nude mice Troponin (ug/L) Saline 4 <0.01Free DOX 3 0.750±0.234 SCL-DOX Blank SCL
53
<0.01<0.01
Data are presented as means ± S.E. (n = 3-5).
To investigate whether encapsulation of DOX into SCL had any impact on the
severity of bone marrow suppression (myelosuppression), the most common adverse
effect of DOX chemotherapy (Singal & Iliskovic 1998), the changes in peripheral
white blood cells count were studied. As shown in Figure 5-7, there was no
statistically significant difference in the total number of white blood cells between
the saline treated or SCL-DOX-treated groups. Furthermore, compared with mice
treated with free DOX, those treated with SCL-DOX had a 2.0-fold and a 3.3-fold
higher count for lymphocytes and monocytes, respectively.
Figure 5-7. SCL-DOX treatment had significantly reduced myelosuppression. HT-29 xenograft-bearing mice were treated as indicated in Figure 5-3. Blood was collected immediately after the mice were sacrificed upon reaching the end point. Data shown are
means S.E. (n=3~5). *, p<0.05 compared to saline; ***, p<0.001 compared to saline; #, p<0.05 compared to free DOX; ##, p<0.01 compared to free DOX.
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5.3 Discussion
This study is the first to evaluate the tissue distribution, in vivo anticancer
activities and toxicity profile of SCL-DOX in a xenograft mouse model of human
colorectal adenocarcinoma. The current study have shown several important points:
(a) SCL-DOX was readily taken up by colon cancer cells and displayed prolonged
retention; (b) encapsulation of DOX in SCL resulted in a decreased distribution of
the drug to the principal sites of acute and chronic toxicity of free DOX, namely the
heart and the skin, as well as markedly lowered cardiotoxicity and myelosuppression;
(c) sulfatide-containing liposomal drug displayed an enhanced therapeutic efficacy in
a mouse model of human colorectal adenocarcinoma.
The intracellular uptake of SCL in glioma cells was shown to be a result of
endocytic uptake of the liposomes in previous work (Shao, K. et al. 2007). In this
study, the results confirmed the intracellular uptake of SCL-DOX by human
colorectal adenocarcinoma cells, HT-29, using confocal microscopy. Importantly, it
has been demonstrated that the encapsulated chemotherapy drug was delivered
intracellularly to the site of action, the nuclei, of the HT-29 cells (Figures 5-1 and 5-
2). Furthermore, the delivered liposomal drug was retained by the tumour cells even
24 h after washing. The in vitro cytotoxicity study compared the viability of
colorectal cancer cells treated with SCL-DOX and free DOX using the MTT assay
and showed that both forms possess overt cytotoxicity. Interestingly, the SCL-DOX
had an IC50 59% higher than that of free DOX (Table 5-1). This is consistent with
observations from others that the IC50 of the free and liposomal drug, when assayed
in vitro, varies dependent on the cell lines used and the nature of the liposomes. For
example, Wang et al. found in the rat prostate cancer cell line MLLB2, the IC50 of
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the liposomal formulation was significantly lower than free DOX (Wang, J et al.
2005). In a study of resistant MCF-7/ADR cells (human breast adenocarcinoma), the
liposomal DOX showed a 30-fold lower IC50 compared to free DOX (Li, B et al.
2012). However, in other studies free DOX seemed to have higher intracellular
uptake and displays higher cytotoxicity than that of liposomal DOX. For example, in
the hepatocellular carcinoma cell line, HepG2, free DOX had been indicated to
possess a higher cytotoxicity compared to that of DOX-loaded stealth liposomes (Li,
X et al. 2009). Furthermore, polyethylene glycol (PEG) coated-liposomal DOX had
been shown to have less toxicity than free DOX in the glioma cell line, U-87 cells
(Shao, Ke et al. 2006). It is important to appreciate that the in vivo pharmacokinetics
are very different between liposomal DOX and free DOX. The half-life of liposomal
DOX can be several days while free DOX can be eliminated in a few minutes in vivo
(Allen, Theresa M. et al. 2006; Cabanes et al. 1999; Unezaki et al. 1994). In cell
culture dishes, cells are exposed to a constant drug concentration throughout the
entire assay period, and often take up free DOX more rapidly than liposomal
formulation (Accardo et al. 2012). Furthermore, cells for MTT assays are mostly
cultured in monolayers, which have a spatial organisation drastically different from
the in vivo 3-dimensional tissue architecture (Xu & Anchordoquy 2011). Thus,
comparison of IC50 between a free drug and a nanoparticle-formulation of the drug in
vitro provides a measurement of the cytotoxicity under a constant concentration
during a chosen assay period, and therefore it may not be able to provide a reliable
prediction of the therapeutic efficacy in vivo (Wu, J et al. 2007). In the present study,
although the IC50 of SCL-DOX was higher than that of free DOX in HT-29 cells in
vitro, the SCL formulation was shown to display a much improved tumour inhibitory
effect over the free DOX in HT-29 tumour-bearing nude mice (Figures 5-4 and 5-5).
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Gastrointestinal tumours are known to be relatively resistant to chemotherapeutics.
In 80% of untreated colon tumours, there is an elevated level of the multidrug-
resistant I (MDR I) gene (Oudard et al. 1991). Notably, significantly higher
concentrations of DOX in tumours from SCL-DOX treated mice were found after
single administration (p<0.05). Moreover, it is interesting to note that SCL
encapsulation not only resulted in a significant increase in tumour uptake in vivo, but
also enhanced anti-tumour efficacy of DOX in HT-29 tumour-bearing nude after
repeated administration. As the statistically significant difference in tumour growth
between free DOX and SCL-DOX groups appeared after the 5th administration
(Figure 5-4), the therapeutic efficacy may be attributed to the preferential
accumulation of SCL-DOX in the tumour via the enhanced permeability and
retention effect (EPR effect) over a period of multiple administrations. Moreover,
after intracellular delivery, the SCL-DOX was retained in the tumour cells much
better than free DOX (Figure 5-2). Thus, the increased DOX concentration in
tumours treated with SCL-DOX was consistent with a better treatment efficacy in
vivo.
In most cases, cancer chemotherapy is limited by the low therapeutic index of the
anticancer drugs due to serious toxicity to normal tissues. Indeed, the therapy-
limiting toxicity of DOX is cardiomyopathy, which may lead to congestive heart
failure and death (Singal & Iliskovic 1998). Encapsulation of free DOX using a
cyanoacrylate nanoparticle shifted toxicity from the heart to the kidney (Manil,
Couvreur & Mahieu 1995), while the PEGylated liposomal DOX (Doxil)
accumulates in skin, resulting in a shift in the toxicity profile from cardiotoxicity and
myelosuppression to cutaneous toxicity, known as palmar-plantar erythrodysesthesia
(Charrois & Allen 2003). The current study aimed to develop a nanoliposome that
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improves the biodistribution characteristics and changes the toxicologic properties of
the encapsulated drug. In current preclinical toxicology experiments, the intended
clinical administration route, i.e. intravenous (i. v.) injection, was utilised throughout
the studies with the HT-29 tumour-bearing mice models. The data (Figure 5-4)
demonstrated that the accumulation of liposomal DOX in the skin and heart were
significantly lower than that of free DOX. Since tissue levels of DOX were reduced
by SCL delivery, the decreased concentration of DOX at these sites was likely to
result in a reduced risk of the development of side effects. Indeed, the biochemical
and haematological analyses (Figure 5-7 and Table 5-3) demonstrated that the
improved therapeutic efficacy of SCL-DOX was obtained without an increase in
toxicity to the heart or to the bone marrow. It is noteworthy that the troponin level
was significantly lower in the SCL-DOX group than that of the free DOX group. As
for indicators of myelosuppression, the higher level of total white cell numbers,
lymphocytes and monocytes in the SCL-DOX group indicated that formulations of
SCL decreased the toxic myelosuppression associated with free drug. The increased
accumulation of SCL-DOX in the liver, spleen and lung compared to free DOX
might be related to the particle size of SCL, as nanoparticles with sizes of 100-200
nm preferentially accumulate in organs of the reticuloendothelial system (Alexis,
Frank , Pridgen, Eric , et al. 2008; Campbell 2006; Drummond, D. C. et al. 1999b) .
By encapsulating drugs in nanoparticles, nanomedicine reduces drug concentration
in normal host tissues and increases the concentration of active drug within the
tumour. To further improve the selectivity and specificity, it is desirable to actively
target the nanodrugs to the site of the tumour. Targeting nanodrugs using antibodies
that recognise the tumour-associated antigens is a widely adopted approach.
However, its application might be limited by altered expression or low percentages
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of tumour cells that express any given antigen. Therefore, the development of
alternative targeted drug delivery systems may have important implications for future
novel anticancer therapeutics. It had been previously demonstrated that sulfatide
mediates the binding and endocytic uptake of SCL in tumour cells via the interaction
with tenascin-C (Shao, K. et al. 2007). The extracellular matrix is a major constituent
of the tumour microenvironment that is very different from that of their normal
counterparts. Tenascin-C is a large extracellular matrix hexabrachion glycoprotein
and is absent or greatly reduced in most normal adult tissues. Tenascin-C is highly
expressed in the majority of malignant solid tumours, including gliomas, and cancer
of the breast, uterus, ovaries, prostate, colon, stomach, pancreas, lung, liver, skin and
kidney (Brellier & Chiquet-Ehrismann 2012). Furthermore, high tenascin-C
expression correlates with a low survival prognosis in cancers such as glioma, breast,
colon and lung carcinoma. In addition to its roles in promoting tumour cell
proliferation and angiogenesis, tenascin-C had recently been shown to provide breast
cancer cells with a key metastatic niche to colonise the lungs by promoting tumour
cell dissemination and survival during the early steps of metastasis and enhancing the
fitness of the disseminated cancer cells at the site of colonisation (Oskarsson et al.
2011). Thus, the production of tenascin-C enhances the ability of micrometastatic
colonies to survive and expand. Given that up to 20% of new cases of colorectal
cancer present with metastatic disease, and of the patients who present with localised
disease, about 20% will subsequently relapse with distant metastases (Segal & Saltz
2009), targeting tenascin-C would constitute a promising strategy in the effort to
combat micrometastasis. Sulfatide had been shown to bind to several extracellular
matrix proteins, including tenascin-C (Miura et al. 1999; Pesheva et al. 1997). In this
study, the in vivo utilities of the SCL in enhanced efficacy in colorectal cancer
xenografts known to express tenascin-C as well as reduced concentration and toxicity
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in tissues that are susceptible to the main side effects of the encapsulated drug, i.e.
the heart, the bone marrow and the skin were further demonstrated. Future research
in combining active targeting ligands, such as antibodies and aptamers, which target
cancer stem cell surface makers with natural lipid-guided intracellular delivery, may
open a new direction for the development of more effective anticancer
nanotherapeutics.
5.4 Conclusion
In conclusion, the current study had demonstrated that SCL-DOX showed an
improved toxicity profile and enhanced efficacy in a human colorectal cancer
xenograft model. It may therefore provide a potent and safe nanomedicine platform
for treatment of colorectal cancer that has overexpression of sulfatide-binding
proteins, especially tenascin-C. Functionalising SCL with antibodies or aptamers
may further enhance the clinical utilities of this natural lipid-guided liposomal
formulation of anticancer drugs for both primary tumour and micrometastasis via
effective targeting the tumour microenvironment.
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Chapter 6 - Exploration of the potential of
using sulfatide-containing liposomes for drug
delivery into the brain
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6.1 Introduction
Gliomas are the most frequently diagnosed primary brain malignancy in the
central nervous system (CNS) (Liu, Y, Zhou & Zhu 2012). Although plenty of
efforts have been done in diagnostic procedures, surgical techniques, radiation
therapy and chemotherapy, the majority of patients succumb to the disease within 1
year after diagnosis (Aoki et al. 2004). One of the reasons for high mortality of
gliomas is that the tumours have the ability to invade diffusely into surrounding
healthy brain tissue, thereby precluding successful surgical removal (Skog et al.
2008). Thus, chemotherapeutics play a critical role in the treatment of brain tumours.
However, the major challenge for chemotherapeutics in the management of CNS
tumours is the limited penetration of many drugs through the blood–brain barrier
(BBB) (Ambruosi et al. 2006). The endothelial cells of the blood vessels in the brain
form cerebral capillaries which are characterised by tight continuous circumferential
junctions, restricting the penetration of most polar solutes across the cerebral
endothelium (Begley 1996; Marin-Padilla 2012).
Traditional anticancer agents suffer from various limitations ranging from poor
treatment efficacy to serious side effects (Allen, T. M. 1994b). Breakdown of the
BBB is present in several CNS disorders, including brain tumours (Gerstner & Fine
2007; Nagaraja et al. 2011). However, the BBB could remain intact, especially in the
early stage of tumour growth and at the tumour invasion edge of the brain tissue
(Bulnes et al. 2012; Huynh, Deen & Szoka 2006). Thus, most current anticancer
chemotherapeutics have poor BBB penetration, which prevent them from targeting
tumours in brain. Therefore, it is critical to develop new approaches which are safe
and effective against brain tumours. Nowadays, nanomaterials such as liposomes or
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polymer nanoparticles, offer several therapeutic strategies for treatments of brain
tumours (Auffinger et al. 2012; Zhou et al. 2012).
Liposomal drugs are promising candidates for local delivery within the
CNS. Increasing evidence has indicated the advantages of nanoliposome technology
in the treatments of gliomas (Gong et al. 2011; Saito et al. 2004). Since they can be
combined with a variety of drugs that are not able to cross the BBB, as a new form of
targeted tissue specific delivery, drug encapsulated nanoliposomes can greatly
improve drug distribution and therefore, its therapeutic effect in brain tumours was
enhanced (Mamot, C. et al. 2004). For example, in in vivo studies, liposomes showed
high accumulation in gliomas (Shibata, Ochi & Mori 1990; Siegal, Horowitz &
Gabizon 1995), indicating the potential of liposomal formulations of anticancer
agents to overcome the BBB. However, it had been identified that liposomes do not
undergo significant transport through the BBB in the absence of ligand-mediated
drug delivery (Misra et al. 2003). In addition, various studies investigating liposome
penetration into the BBB were carried out in in vitro models (Lohmann, Huwel &
Galla 2002; Visser et al. 2005) or in healthy animals which is not a reliable predictor
of the therapeutic efficacy in the tumour-bearing model (Soni, Kohli & Jain 2005;
Xie et al. 2005).
The sulfatide-containing liposomal doxorubicin (SCL-DOX) used in the current
study had shown reduced toxicity, extended longevity, and enhanced tumour specific
targeting property in U-118MG and HT-29 subcutaneous xenograft models (Chapter
4 and Chapter 5). The question arises as to whether the SCL-DOX can penetrate the
BBB to deliver systemically administered therapeutic agents intracranially. Therefore,
in the present study, DOX encapsulated in SCL was evaluated for its capacity of
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delivering chemotherapeutic agents into the brain. The extent of brain delivery was
first studied in healthy rats to examine the potential of SCL-DOX to cross the BBB.
Furthermore, tumour uptake, as well as tissue distribution, in intracranial C6 brain
tumour xenograft rats was studied after a single i.v. administration of SCL-DOX.
The work presented here was performed in collaboration with Yan Yu, Tao Wang
and Dongxi Xiang.
6.2 Results
6.2.1 Enhanced brain distribution of SCL-DOX in healthy SD rats
This study initially evaluated the accumulation of SCL-DOX after i. v. injection in
healthy SD rats with an intact BBB. Therefore, a single dose of 5 mg/kg free DOX or
SCL-DOX were injected via the tail vein and the brain cortexes of healthy SD rats
were collected at 4 predetermined time points following injection. As shown in
Figure 6-1, free DOX showed limited ability to cross the BBB. Conversely, SCL-
DOX led to a significantly elevated DOX accumulation in the brain compared to rats
treated with free DOX at all time points (p<0.001).
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Figure 6-1. Improved brain distribution in healthy rats. Healthy rats were injected with a single dose of 5 mg/kg free DOX or SCL-DOX i. v. Rats were euthanised at 0.5 h, 2 h, 4 h and 24 h after injection. Brain cortex was harvested, washed, weighed, and the DOX was
extracted and quantified. Data are shown as means S.E. for g DOX per g of tissue (n=5-6). ***p<0.001 compared to free DOX.
6.2.2. Enhanced brain distribution of SCL-DOX in perfused healthy SD rats
As brain tissue collected from unperfused animal might contain up to 2% blood, it
is desirable to study the drug delivery in perfused animal to minimise the possibility
of overestimation of the extent of drug delivery due to the contamination of the blood.
To verify the results obtained in Section 6.2.1, rats were perfused with saline 24 h
after injection and the brain cortex was collected followed by measurement using
HPLC. As shown in Figure 6-2, although there was a 3.61-fold decreased (3.57 g/g-
tissue vs. 12.89 g/g-tissue) SCL-DOX accumulation compared to that in the rats
without perfusion (Figure 6-1), the distribution of DOX in the cortex was
significantly increased by i. v. administered SCL-DOX when compared to free DOX
after perfusion.
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Figure 6-2. Improved brain distribution in perfused healthy rats. Healthy rats were injected with a single dose of 5 mg/kg free DOX or SCL-DOX i. v. Rats were perfused with saline 24 h later followed by euthanasia. Brain cortex was harvested, washed, weighed, and
the DOX was extracted and quantified. Data are shown as means S.E. for g DOX per g of tissue (n=5-6). ***p<0.001 compared to free DOX.
6.2.3 Establishment of intracranial brain tumour xenograft models
Having established that SCL-DOX could penetrate the BBB in healthy rats and
mice, the study proceeded to investigate whether the SCL system could penetrate the
BBB to deliver DOX to brain tumours. Taking the volume of the implantation into
account, 1×106 C6 glioma cells in 10 L volume were implanted into the
intracerebral region of healthy Wistar rats using a Harvard Apparatus Pump 11 Elite
Syringe Pumps and stereotaxic device (Figure 6-3 (a)). On day 18 after implantation,
rats were sacrificed. Tissues and tumours were collected and DOX distribution was
measured via HPLC. As shown in Figure 6-3 (b) and (c), C6 glioma cells formed a
tumour in the right hemisphere section, suggesting the successful establishment of
intracranial brain tumour xenograft models in Wistar rats.
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Figure 6-3. Establishment of intracranial brain tumour xenograft models. C6 Glioma cells were implanted into the intracerebral region of rats using syringe pump and stereotaxic device. Rats were euthanised 18 days after implantation. (a) The instrumental set-up used for the implantation of C6 glioma cells in the brain of rats. (b) The exterior view of a representative brain from a tumour-bearing rat. (c) The vertical view of a dissected right hemisphere with the implanted C6 glioma. Arrows: the location of the implanted brain tumour.
6.2.4 Establishing standard curves for tissue distribution in Wistar rats
Before the quantitative determination of DOX in tissues and tumours, tissues and
tumours spiked with a serial dilution of DOX solutions were prepared to obtain a
calibration curve which was used for assessing the DOX concentration
corresponding to different peaks in the chromatograms. Peak areas were plotted
against DOX concentrations to obtain the analytical curve followed by linear
164
regression analysis. The linearity of the calibration curves for distribution of DOX in
different tissues as well as in xenograft C6 gliomas were evaluated over the range of
50-5,000 ng/mL. As shown in Figure 6-3, the correlation coefficients (R2) for
calibration curves were better than 0.990, suggesting a good linear regression within
the analytical concentration ranges.
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Figure 6-4. Standard curves for tissue distribution in Wistar rats. The standard curve was constructed by measurements of tissues spiked with a serial dilution of DOX solutions. The calibration curve was generated by plotting the peak area of fluorescence derived from DOX against the DOX concentration in solution. Standard formulas were determined by
166
linear regression as y = mx + b where y is the peak area of DOX ( V × sec) and x is the DOX concentration (ng/mL).
6.2.5 Tissue distribution and tumour accumulation of SCL-DOX in an
intracranial brain tumour xenograft model
The ability of SCL to improve the delivery of DOX into the intracranial tumour was
evaluated using a rat glioma xenograft model. The tissue distribution and tumour
uptake was determined 24 h after a single dose of 5 mg/kg DOX or SCL-DOX i. v.
administration in C6 glioma-bearing Wistar rats. Consistent with the results from the
study in healthy SD rats in Chapter 3, there was a significantly reduced accumulation
in organs responsible for dose-limiting toxicity in patients, i.e. the heart, in the
tumour-bearing rats treated with SCL-DOX compared to those injected with free
DOX (0.41 ± 0.05 g/g versus 0.64 ± 0.02 g/g) (Figure 4-4 (a)). In the kidney, one
of the other known major DOX toxicity organs (Burke et al. 1977; Giri et al. 2004),
DOX concentration was also statistically significantly lower in the SCL-DOX group
compared to that of the free DOX animals (0.96 ± 0.13 versus 3.98 ± 0.36 g/g). On
the other hand, in organs which are at least partly responsible for clearance of
liposomes from the circulation, such as the liver and spleen, the measured
concentration of DOX indicated that there were significantly higher levels of DOX in
the group treated with SCL-DOX when compared to the free DOX group (89.17-fold
and 3.05-fold, respectively) (Figures 6-5(d)). Notably, there was a statistically
significant elevation of DOX level in the brain and tumour tissue in the SCL-DOX
group compared to that of free DOX (p<0.01 and p<0.05, respectively) (Figures 6-
4(b) and 6-4(c)).
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Figure 6-5. Tissue distribution of DOX encapsulated in SCL in C6 glioma tumour-
bearing rats. Tumour-bearing rats were injected with a single dose of 5 mg/kg free DOX or SCL-DOX i. v. Rats were euthanised 24 h later. Tumours and organs were harvested,
washed, weighed, and the DOX was extracted and quantified. Data are shown as means S.E. for g DOX per g of tissue (n=3-4). *p<0.05 compared to free DOX. **p<0.01 compared to free DOX. ***p<0.001 compared to free DOX.
6.3 Discussion
Malignant gliomas are heterogeneous, diffuse and highly infiltrating by nature
(Linz 2010). Despite wide surgical resection and improvements in radiotherapy and
chemotherapy, the prognosis of patients with gliomas remains extremely poor, with
the majority of patients having a survival rate of less than 10% at five years (Stupp et
al. 2009). As the most vital organ in the body, the brain is isolated and protected
from the intrusion of potentially toxic substances in the blood stream (Pathan et al.
2009). The critical brain supporting system restricts and ensures selective access of
168
nutrients and other essential chemicals across the BBB (Kabanov & Batrakova 2004).
However, the efficacy of anticancer therapeutic agents in targeting brain tumours has
been limited due to poor penetration of the BBB (Chen, Y, Dalwadi & Benson 2004).
The BBB is thus considered a major obstacle for chemotherapeutic drugs to
efficiently reach tumours in the brain (Koukourakis et al. 2000). Various strategies
had been developed to overcome the BBB, including its temporary disruption by
hyperosmotic solutions (Hall et al. 2006) or vasoactive agents such as bradykinin
(Matsukado, Sugita & Black 1998). However, disruption of the BBB has the
potential to result in non-localised drug delivery, as well as an elevated cellular
uptake of plasma albumin and other protein components in the blood which were
toxic to the brain (Vykhodtseva, McDannold & Hynynen 2008). Therefore, new
alternative approaches, such as nanocarriers which include liposomes, had been
being developed for drug transportation across the BBB for drug delivery (Huwyler,
Wu & Pardridge 1996).
In the new drug delivery strategy, the carried drug selectively accumulated in
tissues with increased vasculature permeability, such as tumour tissue in the brain,
(Golub et al. 1994). In vivo studies had shown that liposomes were indeed highly
accumulated in the brain (Yuan et al. 2010) or in the experimental gliomas (Shibata,
Ochi & Mori 1990; Siegal, Horowitz & Gabizon 1995), suggesting that liposomal
drugs might overcome the BBB effectively. Therefore, we used SCL as nanocarriers
for DOX and evaluated its potential in tumour targeting in intracranial brain tumour
xenograft rats in the current study. Consistent with the previous studies regarding
other liposomal formulations of DOX (Saito et al. 2004; Tiwari & Amiji 2006), SCL
encapsulation has the potential to overcome the limitations of free therapeutic drug
molecules in crossing the BBB (Figure 6-1 and Figure 6-2).
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Previous studies had indicated that the C6 rat glioma model simulates the growth
and invasion of human glioblastoma and could be used as a suitable animal model
(Grobben, De Deyn & Slegers 2002). For example, Inoue at al. (Inoue et al. 1987)
studied the therapeutic effects of hyperosmotic BBB disruption using this in situ
glioma model in SD rats, while Kaye et al. (Kaye, Morstyn & Ashcroft 1985) had
studied glioma growth in the C6 tumour-bearing Wistar rats. Thus, unlike
subcutaneous tumour models, the intracranial brain tumour xenograft model
performed in the present study mimics the natural behaviour of brain tumours.
Although both Wistar and SD rats were utilised for the study of C6 tumours, the
C6/Wistar model is more similar to the human Glioblastoma multiforme (GBM)
characteristics, such as a high mitotic rate, cells forming palisades, and
neovascularisation (Miura et al. 2008). As a result, Wistar rats were used as the C6
tumour xenograft model in our study.
The comparison of brain distribution in healthy SD rats treated with different
DOX formulations revealed a significantly increased accumulation of SCL-DOX
compared to the free DOX group (Figure 6-1). As a widely used and efficient
antitumour drug, DOX delivery to the brain is often restricted because of the poor
penetration of this therapeutic molecule through the BBB (Rousselle et al. 2001).
The statistically significant increase of DOX in the brain of perfused healthy SD rats
injected with SCL-DOX showed that encapsulating DOX by SCL had the potential
to cross the BBB. Moreover, compared with the free group, SCL encapsulation
resulted in an elevated DOX accumulation in the cortex of perfused healthy SD rats
(Figure 6-2), indicating a high accumulation of SCL-DOX in brain cells. As a result,
the present nanodrug system might be capable of increasing therapeutic efficacy of
drugs that are unable to penetrate in the BBB in the treatment of brain diseases.
170
In the intracranial C6 glioma model, the BBB would have been, at least partly,
disrupted by surgery and tumour growth. Nonetheless, there was a significant
improvement of tumour uptake in tumour-bearing rats (Figure 6-5(c)). DOX
encapsulation within SCL resulted in 1.67-fold higher uptake by glioma compared to
free DOX in the current study. Encapsulation of DOX into SCL resulted in a
significantly higher DOX accumulation in the brain of rats receiving liposomal
formulation than those in previous studies (Arnold et al. 2005; Yang, FY et al. 2012).
Moreover, the tumour uptake in the present study was also higher than that in the
previous study which demonstrated an enhanced therapeutic effect in vivo (1.79
g/g-tissue in this study vs. 0.16 g/g-tissue reported by others) (Arnold et al. 2005).
Therefore, the SCL-DOX in the present study indicated its potential to be a better
nanodrug delivery system for brain tumours or other brain diseases. However, due to
time constraint, the experiments with C6 glioma model using perfused animals were
unable to be repeated. Further studies in perfused glioma-bearing rats should be
performed to confirm the observations presented in this Chapter.
The increased accumulation of free DOX in the brain of tumour-bearing rats
compared with the brain uptake in healthy rats was most likely due to the disruption
of BBB in the brain tumour model (Chekhonin et al. 2007; Wang, P et al. 2010). The
glioma tumours are characterised by the invasive property, therefore, the capillary
endothelium could be destroyed by tumour cells when the tumour expand beyond 2.0
mm3 (Erwin 2009). As a result, the penetration of free DOX into the brain could
increase, leading to the increased uptake by tumour cells as well as the brain in
general. Future studies, such as the histopathological analysis (Mohamed, Karam &
Amer 2011) or studies about apoptosis pathways (Tangpong et al. 2011), need to be
performed to confirm the toxicity of DOX to normal brain cells. In addition, the use
171
of SCL for the treatment of C6 glioma was also advantageous in view of other tissues.
The biodistribution (BD) data in this C6 glioma model demonstrated that
accumulation of liposomal DOX in the heart and kidney were significantly lower
than those of free DOX (Figure 6-5(a) and (d)). Therefore, the decreased
concentration of DOX is likely to result in a reduced systemic toxicity in tumour-
baring rats as the reduced tissue levels of DOX afforded by SCL delivery.
6.4 Conclusion
In summary, the present proof-of-principle study had demonstrated that SCL-DOX
was able to penetrate into the BBB and accumulate in the C6 glioma in vivo.
Moreover, the improved tissue distribution was confirmed in the glioma-bearing rats.
Therefore, such nanodrug delivery systems may represent a new strategy for the
development of effective anticancer chemotherapeutics targeting the brain tumour.
Further investigations are planned to study the efficacy in inhibition of glioma
growth, the survival benefit, as well as systemic toxicity.
172
Chapter 7 - Development of EpCAM aptamer
functionalised PEGylated sulfatide-containing
liposomal doxorubicin
173
7.1 Introduction
Many DNA intercalators have been shown to have antitumour and antibiotic
activity (Lee, CJ et al. 2004). Doxorubicin (DOX), which is well documented to
function via blocking the progression of topoisomerase II and thus preventing DNA
replication and cell division, is widely used in the treatment of many malignant
diseases due to its broad antitumour activity (Kaplan et al. 2010). Cardiotoxicity is a
major limiting factor in anticancer therapy of DOX (Yeh et al. 2004). Cardiotoxicity
induced by DOX may present as either acute or chronic cardiomyopathy. While
chronic cardiotoxicity is dose dependent, due to the limitation of the cumulative dose
of DOX to <450 mg/m2, acute cardiotoxicity, occurring within 2 to 3 days of its high
dose administration, is now rare (Zhang, YW et al. 2009). Other preventive strategies
include modifying DOX with targeting molecular agents (Bagalkot et al. 2006) or
establishing new drug delivery systems (Nagykalnai 2010), which have the potential
to improve its distribution in vivo.
Aptamers, rapidly generated through the Systematic Evolution of Ligands by
EXponential enrichment (SELEX), are single-stranded RNA or DNA
oligonucleotides (Yang, X, Li & Gorenstein 2011). The interactions of aptamers with
their target molecules mainly rely on hydrogen bonding, electrostatic, and
hydrophobic interactions rather than Watson–Crick base pairing (Hermann & Patel
2000). Aptamers possess several advantages when compared to antibodies. These
include increased stability, smaller size, high target affinity and specificity,
reproducibility via inexpensive and consistent chemical synthesis, easy modification
for both labelling and selective changes in sequence, and the possibility against toxic
or poorly immunogenic targets (Pendergrast et al. 2005; Wang, P, Hatcher, et al.
174
2011; Wang, P, Yang, et al. 2011b). Therefore, nucleic-acid aptamers have become
increasingly important molecular tools with wide applications in modern medicine,
including diagnostics and anticancer therapeutics (Rerole et al. 2011).
Due to their advantages over other ligands used in drug delivery, aptamers had
been utilised for the targeted delivery of DOX (Zhang, H 2004) or nanocarriers (Herr
et al. 2006). A study by Zhang successfully conjugated DOX to an aptamer and
showed that this aptamer-DOX conjugate was capable of specifically targeting
prostate cancer cells via the prostate-specific membrane antigen (PSMA) and
delivering DOX inside the cell (Zhang, H 2004). As well, a Sgc8-PEG-liposome for
targeting CCRF-CEM (T-cell Acute Lymphoblastic Leukemia) was reported by
Kang et al (Kang et al. 2010b).
Our lab recently developed a RNA aptamer against a cancer stem cell marker,
epithelial cell adhesion molecule (EpCAM), and this aptamer efficiently internalises
upon binding to target cells (Shigdar et al. 2011). EpCAM was firstly described in
1979 as a 40 KDa cell surface cancer-associated antigen, which is a type I single
span transmembrane glycoprotein with extracellular epidermal growth factor-like
(EGF-like) and thyroglobulin (TY) motifs as well as an intracellular domain
containing an internalization motif and several -actinin binding sites (Balzaret, M al.
1999). Previous studies had shown that the expression of EpCAM in various cancers
is inversely related to the prognosis of the patients (Baeuerle & Gires 2007). Thus,
aptamers against EpCAM may have the ability to target cancer cells, especially those
with high expression of EpCAM. Therefore, preliminary in vitro studies were carried
out to explore the potential of aptamers in enhancing cellular uptake of sulfatide-
containing nanoliposomal DOX (SCL-DOX) in cells overexpressing EpCAM.
175
7.2 Results
7.2.1 Establishing a standard curve for DOX concentration in aptamer
functionalised PEG-sulfatide-containing nanoliposomal DOX (ap-PEG-SCL-
DOX)
To evaluate DOX concentration in aptamer functionalised PEG-sulfatide-
containing nanoliposome (ap-PEG-SCL) via HPLC, standard curves were
constructed by plotting peak area of fluorescence derived from DOX vs. DOX
concentrations. To mimic samples to be determined, ap-PEG-SCL-DOX without
unconjugated aptamer purification using size-exclusion chromatography were used
as standards and diluted to series of DOX concentrations from 50 ng/mL to 800
ng/mL, followed by HPLC measurement. A linear regression analysis was used for
quantitation. As shown in Figure 7-1, a good linearity between peak areas and
concentrations was obtained within the tested concentrations with a correlation
coefficient of R2=0.9984.
Figure 7-1. Standard curve for determination of DOX concentration in ap-PEG-SCL-
DOX. The standard formulas were determined by linear regression as y = mx + b, where y is the peak area of DOX and x is the DOX concentration.
7.2.2 C
with ne
The
SCL w
as that
charge
PEG-S
incorpo
6.30 nm
respect
30.37 ±
Table 7Compositap-PEG-Snegative aData are
Figure
The cochemicaPEG us
Characteris
egative con
bioconjuga
with 5% PEG
in Table 3
were analy
CL-DOX.
orating DOX
m with a P
tively. Mean
± 0.77 mV,
7-1. Physicaltion SCL-DOX ap-PEG-SCL-De shown as m
7-2. Develo
njugation bal reaction beed for SCL a
ation of ap
ntrol aptam
ation in the
G (Figure 7
3-1 (Chapte
ysed for com
As shown
X and nega
PDI of 0.16
nwhile, the
respectively
l properties Particl130.60
DOX 131.30means ± S.E.
pment of ap
etween EpCetween the hand the free a
p-PEG-SCL
mer (negativ
current stu
7-2). Thus, t
er 3). The s
mparison be
in Table 7
ative ap-PE
6 ± 0.01 an
zeta potenti
y.
of the liposole size (nm) 0 ± 6.30 0 ± 4.42 of at least th
p-PEG-SCL
CAM aptamhydrophilic camine group
176
L-DOX and
ve ap- PEG
udy occurre
the average
size, polydi
etween ap-P
7-1, the me
EG-SCL inc
nd 131.30
ial value of
omal formulPolydispersity 0.16 ± 0.04 0.15 ± 0.02
hree indepen
L-DOX (ada
mer and PEGcarboxylic ac (-NH2) mod
d PEG-SCL
G-SCL-DOX
ed after DO
e loading ef
spersity ind
PEG-SCL-D
ean vesicle
corporating
± 4.42 nm
f SCL was -3
lations. index (PDI)
ndent experim
apted from F
G-SCL-DPXcid functionadified on the
L-DOX mo
X)
OX encapsul
fficiency wa
dex (PDI) a
DOX and n
size of ap
DOX wer
m with a PD
32.00 ± 1.2
Zeta potentia-32.00 ± 1.22-30.37 ± 0.77
ments.
Farokhzad e
X was obtainal group (-COaptamers.
dified
lated in the
as the same
and surface
negative ap-
p-PEG-SCL
e 130.60 ±
DI of 0.15,
2 mV and -
al (mV)27
et al. 2004).ned via theOOH) on the
e
e
e
-
L
±
,
-
. e e
177
7.2.3 Cellular uptake of ap-PEG-SCL-DOX
To evaluate the binding specificity of the ap-PEG-SCL-DOX, two different
aptamers were utilised, including the FITC-labelled EpCAM targeting aptamer and
the FITC-labelled negative control aptamer which is an aptamer of the same
sequence but with a different side-chain modification which could affect the 3-
dimensional structure of aptamer (Chushak & Stone 2009). Therefore, the negative
control aptamer does not bind to EpCAM and does not specifically target the cancer
cells with high expression of EpCAM. The accumulation of free DOX, PEG-SCL-
DOX, ap-PEG-SCL-DOX and negative ap-PEG-SCL-DOX in the EpCAM-negative
cell line (HEK-293T) and the EpCAM-positive cell line (HT-29) were studied using
laser scanning confocal microscopy utilising the natural fluorescent property of DOX
and the FITC label on the aptamer.
After 24 h incubation with 2 g/mL DOX, PEG-SCL-DOX, ap-PEG-SCL-DOX or
negative ap-PEG-SCL-DOX followed by washing with PBS, cellular uptake of free
DOX or liposomal formulations by HEK-293T cells and HT-29 cells were examined.
As shown in Figure 7-3(a), there was no significant difference between HEK-293T
cells treated with ap-PEG-SCL-DOX and those with negative ap-PEG-SCL-DOX.
The overlay of Hoechst staining (nucleus) and red fluorescence (DOX) in HT-29
cells indicated that ap-PEG-SCL-DOX was not adhered to the cell surface but
actually penetrated into the nucleus (Figure 7-3(b)). Notably, confocal microscopy of
HT-29 cells exposed to ap-PEG-SCL-DOX demonstrated increased DOX
fluorescence (red) compared to other groups after 24 h incubation. Moreover, the
strong green fluorescence (from FITC that was used to label the aptamer) in HT-29
cells treated with ap-SCL-DOX indicated the accumulation of EpCAM aptamer in
178
the cells. The overlay of nucleus staining and FITC fluorescence in the cells
suggested that the EpCAM RNA aptamer functionalised PEG-SCL-DOX is
efficiently internalised after the binding to EpCAM on cell surface. As a control,
fluorescence of ap-PEG-SCL-DOX was determined in the EpCAM-negative cell line
HEK-293T. Results indicated that there was cellular uptake of ap-PEG-SCL-DOX in
EpCAM-negative cell line (Figure 7-3(a)). Moreover, negative ap-PEG-SCL-DOX
was also taken up by both two cell lines (Figure 7-3), which might have resulted
from non-specific binding following the long incubation.
Figure
293T ce
or aptam
imaged fluorescfluorescbars: 20
7-3. Intrac
ells and HT-
mer function
with a confocence from cence from a0 m.
ellular upta
-29 cells. Cenalised PEG-
ocal fluorescDOX; Blu
aptamer. Dat
ake of aptam
ells were incuSCL-DOX f
cence microsue: nuclei ta are represe
179
mer functio
ubated with 2for 24 h. Cel
cope. (a) HEstained wit
entative of th
onalised PE
2 g/mL freells were was
EK-293T celth Hoechst hree indepen
G-SCL-DO
e DOX, PEGhed twice w
ls. (b) HT-233342; G
ndent experim
X in HEK-
G-SCL-DOXwith PBS and
9 cells. Red:Green: FITC
ments. Scale
-
X d
: C e
180
7.2.4 Cellular retention of aptamer functionalised PEG-SCL-DOX
For the study of cellular retention of ap-PEG-SCL-DOX, cells were first incubated
with different formulations of DOX for 24 h at the dose of 2 g/mL. Cells were then
washed with PBS twice and cultured in fresh culture medium. The same wells of
cells were imaged serially at 1 h, 2 h, 4 h and 24 h after washing using fluorescence
confocal microscopy. The DOX fluorescence in both cell lines treated with liposomal
formulations decreased slower when compared to cells treated with free DOX,
suggesting better retention of our liposomal formulation in vitro. In contrast, the
DOX fluorescence in free DOX and negative ap-PEG-SCL-DOX group decreased
rapidly after washing with PBS (Figure 7-4). There was a significant reduction of
DOX fluorescence in free DOX group only 2 h after washing with PBS (Figure 7-5).
Interestingly, compared to the EpCAM-positive HT-29 cells, the DOX fluorescence
(red), as well as aptamer fluorescence (green), diminished significantly in the HEK-
293T cells which are EpCAM-negative. HEK-293T cells treated with ap-SCL-DOX
or negative ap-PEG-SCL-DOX both showed significant decreased red fluorescence
and green fluorescence 4 h after washing (Figure 7-6(a)). Conversely, the DOX
fluorescence and aptamer fluorescence for ap-PEG-SCL-DOX was more stable in
HT-29 cells compared to those treated with other DOX formulations and those in
HEK-293T cells. Even 24 h after washing, DOX fluorescence and FITC fluorescence
(aptamer) could be readily observed in the HT-29 cells treated with ap-PEG-SCL-
DOX (Figure 7-7 (b)).
Figure
cells an
SCL-DOmedium
microscnuclei srepresen
7-4. Intrace
nd HT-29 ce
OX for 24 h.m. The same
copy. (a) HEstained withntative of thr
ellular reten
lls (1 h). Cel Cells were twells of cel
EK-293T ceh Hoechst 3ree independ
ntion of apt
lls were firstthen washedlls were ima
ells. (b) HT-33342; Greendent experime
181
amer functi
t incubated w with PBS tw
aged at 1 h w
-29 cells. Rn: FITC fluents. Scale b
ionalised SC
with 2 g/mLwice and cultwashing usin
Red: fluorescuorescence fars: 20 m.
CL-DOX in
L free DOX otured in freshng fluorescen
cence from Dfrom aptame
HEK-293T
or equivalenth full culturence confocal
DOX; Blue:er. Data are
T
t e l
: e
Figure
cells an
SCL-DOmedium
microscnuclei srepresen
7-5. Intrace
nd HT-29 ce
OX for 24 h.m. The same
copy. (a) HEstained withntative of thr
ellular reten
lls (2 h). Cel Cells were twells of cel
EK-293T ceh Hoechst 3ree independ
ntion of apt
lls were firstthen washedlls were ima
ells. (b) HT-33342; Greendent experime
182
amer functi
t incubated w with PBS tw
aged at 2 h w
-29 cells. Rn: FITC fluents. Scale b
ionalised SC
with 2 g/mLwice and cultwashing usin
Red: fluorescuorescence fars: 20 m.
CL-DOX in
L free DOX otured in freshng fluorescen
cence from Dfrom aptame
HEK-293T
or equivalenth full culturence confocal
DOX; Blue:er. Data are
T
t e l
: e
Figure
cells an
SCL-DOmedium
microscnuclei srepresen
7-6. Intrace
nd HT-29 ce
OX for 24 h.m. The same
copy. (a) HEstained withntative of thr
ellular reten
lls (4 h). Cel Cells were twells of cel
EK-293T ceh Hoechst 3ree independ
ntion of apt
lls were firstthen washedlls were ima
ells. (b) HT-33342; Greendent experime
183
amer functi
t incubated w with PBS tw
aged at 4 h w
-29 cells. Rn: FITC fluents. Scale b
ionalised SC
with 2 g/mLwice and cultwashing usin
Red: fluorescuorescence fars: 20 m.
CL-DOX in
L free DOX otured in freshng fluorescen
cence from Dfrom aptame
HEK-293T
or equivalenth full culturence confocal
DOX; Blue:er. Data are
T
t e l
: e
Figure
cells an
equivalefull cul
fluorescfrom Daptamer
7-7. Intrace
nd HT-29 c
ent SCL-DOlture medium
cence confocDOX; Blue: r. Data are re
ellular reten
cells (24 h)
OX for 24 h. Cm. The sam
cal microscopnuclei stain
epresentative
ntion of apt
). Cells werCells were thme wells of
py. (a) HEKned with Hoe of three ind
184
amer functi
re first incubhen washed wf cells were
K-293T cellsoechst 3334dependent exp
ionalised SC
bated with with PBS twe imaged a
. (b) HT-29 2. Green: Fperiments. S
CL-DOX in
2 g/mL frwice and cultuat 24 h wa
cells. Red: fFITC fluoresScale bars: 20
HEK-293T
ree DOX orured in fresh
ashing using
fluorescencescence from0 m.
T
r h g
e m
185
7.3 Discussion
Aptamers are highly selective nucleic acids which have been developed for
numerous therapeutic applications (Yang, X et al. 2011). In the current study, we
developed a novel specialised aptamer-liposome bioconjugate for targeted delivery
of anticancer agents into EpCAM expressing cancer cells. SCL with 5% PEG used in
our study possesses a negative surface charge (Table 3-1, Chapter 3), which was a
desirable characteristic as positively charged nanoparticles might lead to nonspecific
interaction with the negatively charged aptamers and diminish the binding property
(Farokhzad et al. 2004a). The negative charge on the SCL could repel the similarly
charged EpCAM aptamers and thus minimise the charge interaction between
aptamers and SCL surface. As a result, the conjugation in the current study was
obtained via the chemical reaction between the hydrophilic carboxylic acid
functional group (-COOH) on the PEG used for SCL and the free amine group (-NH2)
modified on the aptamers (Figure 7-2). In the presence of sulfo-N-
hydroxysuccinimide (sulfo-NHS) and 1- [3-dimethylaminopropyl]-3-
ethylcarbodiimide hydrochloride (EDC), the carboxylic acid group on the SCL
surface would easily convert to the NHS ester and link with the NH2-modified
aptamers (Estevez et al. 2010a; Healy et al. 2004).
RNA aptamers used in the current study were previously selected against the
extracellular region of EpCAM (Shigdar et al. 2011). By conjugating the PEG-SCL
and the EpCAM aptamer, aptamer-liposome bioconjugates were generated for
assessment in this study. The elevated particle size and increased absolute value of
zeta potential in ap-PEG-SCL-DOX and negative ap-PEG-SCL-DOX compared to
186
SCL (Table 3-1, Chapter 3) might be attributed to the insertion of the aptamer
(Wilner et al. 2012).
Because of previously established EpCAM aptamer binding speci city, EpCAM
targeting aptamers and negative control aptamers which did not target the EpCAM,
as well as two cell lines, were used to analyse the cellular uptake in vitro. The human
colorectal adenocarcinoma cell line HT-29 was known to express EpCAM, which is
a 40-kDa transmembrane glycoprotein frequently expressed in solid tumours
(Herrmann et al. 2010; Winkler et al. 2009). Conversely, HEK-293T, human
embryonic kidney cells, was reported to be the EpCAM-negative (Shigdar et al.
2011). Comparison of cellular uptake between these two cell lines would evaluate the
specific interaction between EpCAM targeting aptamer functionalised PEG-SCL-
DOX and EpCAM-positive cells.
The un-functionalised SCL-DOX displayed an enhanced accumulation of DOX in
HT-29 cells compared to free DOX (Chapter 5). Through current cellular uptake
studies, confocal microscopy also revealed differences among uptakes of different
DOX formulations. The data demonstrated that the binding of ap-PEG-SCL-DOX
bioconjugates to HT-29 cells, which are EpCAM-positive, was significantly
enhanced when compared with the negative ap-PEG-SCL-DOX complex and the
SCL-DOX formulation (Figure 7-3 (b)). The overlay of nucleus staining, DOX and
FITC fluorescence from the aptamer indicated intracellular delivery of our aptamer
functionalised SCL-DOX to the site of DOX action, the nucleus. In contrast, the
results in HEK-293T cells, which do not express EpCAM, showed no significant
difference in cellular uptake between the conjugates and the control groups (Figure
7-3 (a)). Moreover, the non-specific binding and internalisation of negative ap-PEG-
187
SCL-DOX in both HEK-293T and HT-29 cells might have resulted from the long
time incubation (24 h incubation). As shown in previous study, the binding of this
EpCAM targeting aptamer and cells was observed 0.5 h after incubation (Shigdar et
al. 2011). To confirm if this was a transient increase observed in the EpCAM
negative cell line, studies were then proceeded to evaluate the retention of the ap-
PEG-SCL-DOX to both EpCAM-positive and EpCAM-negative cell lines. The
retention experiment showed that this non-specific binding was transient. The
fluorescence of aptamer (green), as well as the DOX (red), diminished very fast in
HEK-293T cells (Figure 7-4 (a), Figure 7-5 (a), Figure 7-6 (a) and Figure 7-7 (a)).
The rapid clearance of fluorescence in HEK-293T cells suggested the interaction of
the ap-PEG-SCL-DOX and EpCAM-negative cells might be non-specific. As a result,
although there was non-specific uptake in the EpCAM negative cell line, both the
drug and aptamer fluorescence diminished rapidly, indicating the drug solutions had
also been removed from the cells rapidly. Moreover, in the monolayer model, cells
cultured in the monolayer are exposed to a constant drug concentration throughout
the entire assay period (Accardo et al. 2012), which is not a reliable predictor of the
therapeutic efficacy in vivo where the cells wouldn’t be subjected to a constant
concentration for so long. Therefore, the fluorescence in the cells in the uptake study
is not truly representative of what would be seen in in vivo studies.
On the other hand, there were significantly different retention patterns observed in
the EpCAM-positive cell line compared to that in the EpCAM-negative cell line
during the period of observation (24 h). As expected, unlike in the HEK-293T cells
(EpCAM-negative), in the EpCAM-positive cells, HT-29, the red fluorescence for
DOX and the green fluorescence for aptamer were more stable compared to cells
treated with other DOX formulation. Even 24 h after washing, DOX fluorescence
188
and aptamer fluorescence could be readily observed in the nuclei of cells in HT-29
cells treated with ap-PEG-SCL-DOX compared to those in other groups (Figure 7-7
(b)). Thus, the prolonged retention of ap-PEG-SCL-DOX in the nuclei for several
hours even after washing suggested the potential of a better anti-tumour efficacy of
our bioconjugates.
7.4 Conclusion
In this preliminary study, an EpCAM aptamer functionalised SCL delivery system
for DOX had been developed. The enhanced cellular uptake and prolonged retention
of the aptamer-SCL-DOX conjugates studies suggested the potential for efficient
targeted delivery of anticancer chemotherapeutics to cells overexpressing EpCAM.
The non-specific binding of the EpCAM aptamer functionalised SCL delivery
system was transient and was diminished from the EpCAM negative cells within a
short period of time, indicating that the EpCAM aptamer functionalised SCL
delivery drug was specifically retained by EpCAM positive cells. More studies, such
as a semiquantitative cellular uptake study (flow cytometry assay), quantitative
cellular uptake study (HPLC assay) and in vitro toxicity study (MTT assay) should
be utilised to confirm the improved drug delivery property of the EpCAM aptamer
functionalised SCL delivery system in the EpCAM positive cell lines. Moreover, as
EpCAM is a cell surface marker of cancer stem cells (CSCs), our EpCAM aptamer
functionalised SCL delivery system has the potential to be an effective nano-carrier
for the specific targeting of CSCs. Further studies, including the tissue distribution
and treatment efficacy of these vehicles in vivo, are needed to evaluate the specificity
to target the CSCs.
189
Chapter 8 - Summary and perspectives
190
8.1 Summary
Cancer is still the leading cause of death in worldwide despite advances having
been made in reducing morality and improving survival (Klein, C. A. 2009). Current
cancer treatments mainly include surgical intervention, radiation and
chemotherapeutic drugs (Miller et al. 2012). However, most deaths from cancer
result from distant metastatic disease, and surgery is not generally able to
treat metastatic cancer (Ruiterkamp & Ernst 2011). Thus, effective management of
metastases has moved to the centre of clinical attention and chemotherapy has been
widely used for the treatment of metastases (Wolf et al. 2012). Nevertheless, in the
anticancer area, one of the biggest challenges that still remain is the low therapeutic
index of the majority of anticancer agents (Wang, Y, Newell & Irudayaraj 2012). As
traditional anticancer chemotherapeutics often kill healthy cells as well as tumour
cells and cause toxicity to the patient, it would therefore be desirable to develop
chemotherapeutics that could either passively or actively target cancerous cells
effectively (Urbinati, Marsaud & Renoir 2012). Present applications of
nanotechnologies for the treatment of cancers represent a new direction for future
cancer therapeutics and liposomes play a critical role in the formulation of potent
drugs to improve the therapeutic index (Samad, A., Sultana, Y. & Aqil, M. 2007).
The utility of liposomes as a nanodrug delivery system has been demonstrated by its
ability to efficiently deliver cytotoxic agents to tumour tissue while minimizing the
undesired negative side effects associated with these drugs (Cukierman & Khan
2010). Moreover, new formulations of liposomes are targeted to the reduced toxicity
and increase accumulation at the target site via passive or active targeting
modification (Elbayoumi, T. A. & Torchilin, V. P. 2008; Gabizon, AA 2001a).
However, significant challenges remain.
191
As a glycosphingolipid which can bind to several extracellular matrix proteins,
sulfatide has the potential to become an effective targeting agent for tumours
overexpressing tenascin-C in their microenvironment. To overcome the dose-limiting
toxicity of the anticancer agent DOX, a SCL encapsulation approach was employed
to improve treatment efficacy and reduce side effects of free DOX. This study
analysed in vitro characteristics of SCL-DOX and assessed its cytotoxicity in vitro,
as well as pharmacokinetics (PK), biodistribution (BD), therapeutic efficacy and
systemic toxicity in vivo. Due to the overexpression of tenascin-C on cell surface, a
human glioblastoma U-118MG xenograft model and a human colon carcinoma HT-
29 cell line were chosen as models in the study for the targeting property of SCL.
Firstly, the characteristics of SCL-DOX in vitro and in healthy rats were analysed
(Chapter 3). The SCL-DOX was shown to achieve the highest drug to lipid ratio of
0.5:1 among all such systems reported so far, a high encapsulation of DOX with an
ideal average size less than 100 nm and a remarkable in vitro stability. The in vitro
uptake and retention studies suggested that DOX encapsulated in SCL was delivered
into the nuclei and displayed prolonged retention over free DOX in both U-118MG
cells and HT-29 cells in vitro. Although there were no advantages of cytotoxicity in
culture dishes in vitro compared to free DOX, this simple two-lipid SCL-DOX
nanodrug has favourable pharmacokinetic attributes in terms of prolonged circulation
time, reduced volume of distribution, slow clearance rate and enhanced
bioavailability in healthy rats. The improved tissue distribution in SD rats revealed in
the present study, indicated that SCL-DOX significantly reduced the DOX
accumulation in the heart and kidney compared to free DOX, suggesting the potential
of decreased toxicity to these organs of the SCL-DOX in vivo.
192
The prolonged cellular retention in vitro, as well as the improved pharmacokinetic
and tissue distribution profiles in vivo, suggested the potential of enhanced treatment
efficacy and reduced systemic toxicity of SCL-DOX. Therefore, the in vivo BD and
antitumour treatments efficacy of SCL-DOX was evaluated in the nude mice bearing
U-118MG glioblastoma or HT-29 colorectal cancer (Chapter 4 and Chapter 5). As
expected, in addition to the improved PK and BD profiles in healthy SD rats, an
enhanced treatment efficacy of SCL-DOX was found in tumour-bearing mice
compared to the free drug. The use of this nanodrug delivery system to deliver DOX
for treatment of tumour-bearing mice produced a much improved therapeutic
efficacy in terms of suppression of tumour growth and extended survival in contrast
to the free drug. Furthermore, a reduction in the accumulation of DOX in the drug’s
principal toxicity organs, the heart and the skin, in U-118MG and HT-29 xenograft
mice models was observed. Also, SCL-DOX led to the diminished systemic toxicity
as evident from the plasma biochemical analyses. Moreover, reduced
myelosuppression was observed in HT-29 tumour-bearing mice treated with SCL-
DOX. Thus, SCL has the potential to be an effective and safer nano-carrier for
targeted delivery of therapeutic agents to tumours with elevated expression of
tenascin-C in their microenvironment. Such natural lipid-guided nanodrug delivery
systems may represent a new strategy for the development of effective anticancer
chemotherapeutics targeting the tumour microenvironment for both primary tumours
and micrometastases.
The ability of a drug to penetrate the blood-brain barrier (BBB) is crucial for its
utility in the pharmaceutical treatment of the central nervous system (CNS) diseases,
including gliomas (Lohmann, Huwel & Galla 2002). In the BD study of healthy SD
rats, a significantly improved DOX accumulation induced by SCL encapsulation was
193
observed in normal brain and in the perfused brain, indicating the SCL-DOX was
able to penetrate the BBB and remain in the brain cells (Chapter 6). To evaluate the
ability of SCL-DOX to cross the BBB and the treatment efficacy of DOX
encapsulated by the SCL in gliomas, an intracranial brain tumour xenograft model in
Wistar rats was established. SCL-DOX was administrated via i. v. injection and
DOX concentration in tumour and brain were analysed. As shown in Chapter 6, the
significantly improved tumour uptake and brain accumulation, as well as improved
tissue distribution, indicated that SCL-DOX could effectively increase the
penetration of DOX across the BBB and reduce its systemic toxicity. Therefore, SCL
has the potential to be utilised as an anticancer agent nanocarrier for the delivery of
drugs across to the BBB.
The SCL-DOX has been identified as a liposome formulation which could target
the tumour with high expression of tenascin-C effectively (Chapter 3 and Chapter 4).
However, previous studies have indicated that conjugating targeting molecules, such
as antibodies, ligands, peptides, or aptamers, to the surface of liposomes could lead
to an improved targeting liposome formulation (Herrmann et al. 2010). As a result,
the aptamer functionalised PEG-SCL-DOX (ap-PEG-SCL-DOX) was developed to
further improve the targeting property of SCL-DOX (Chapter 7). The treatment of
HT-29 tumour, which has been reported to have a high expression of epithelial cell
adhesion molecule (EpCAM), was showed to be effectively improved by SCL-DOX
in the tumour xenograft mice. However, DOX has been reported to not be able to
eradicate cancer stem cells (CSCs) which are related to the recurrence of tumours
(Zheng, X et al. 2010). Thus, PEG-SCL-DOX modified with aptamer targeting
EpCAM, one of the cell surface markers for CSCs, has the potential to target CSCs.
In the Chapter 7, the aptamer-facilitated cellular uptake and retention of DOX
194
encapsulated PEG-SCL showed the accumulation, as well as retention time, of DOX
had been enhanced via the highly specific RNA aptamer modified PEG-SCL in HT-
29 cells. Therefore, the strategy of delivering DOX via the EpCAM targeted aptamer
functionalised PEG-SCL might be suitable for systemic targeted therapy of EpCAM-
positive cancers, or even CSCs.
8.2 Perspectives
As shown in Chapter 6, it is likely that SCL could improve the delivery efficiency
of DOX into the brain. However, due to time constraints, the assays for evaluation of
SCL for brain delivery mainly focused on preliminary studies of tissue distribution
and tumour uptake in vivo. More studies need to be designed and carried out to study
the therapy efficacy and safety of SCL-DOX in the treatment of gliomas. Thus, the
next step following current studies would investigate the outcome of tumour
inhibition, survival and systemic toxicity of SCL-DOX in intracranial brain tumour
xenograft rats with treatment of free DOX as a control. Moreover, histopathological
analysis of tissue and tumour sections would be useful to investigate DOX toxicity to
the normal brain cells as well as DOX accumulation in the xenograft glioblastoma in
vivo. Furthermore, the study about tissue distributions in perfused glioma-bearing
rats would be utilised to confirm the enhanced tumour uptake of SCL-DOX in vivo.
Another essential future research direction is to study the targeting property of
aptamer functionalised PEG-SCL to CSCs. The concept of CSCs is that only a
minority of cells within a tumour is able to generate a new tumour (Lin et al. 2011).
Most of current antitumour therapies can only treat rapidly dividing tumour cells
effectively (Bitarte et al. 2011). However, tumour proliferation may be driven by
CSCs which divide slowly and are relatively resistant to cytotoxic drugs (Koch,
195
Krause & Baumann 2010). Thus, many tumours may progress because CSCs are not
sensitive to the treatment. Therefore, targeting CSCs for the prevention of tumour
progression and treatment could become a novel strategy for better treatment.
EpCAM, a glycosylated, 30- to 40-kDa type I membrane protein, which is expressed
in a variety of human epithelial tissues, cancers, and progenitor and stem cells, is a
cell surface marker for CSCs in numerous types of cancers, including hepatocellular
carcinoma and colon carcinomas (Munz, Baeuerle & Gires 2009). Recently, our lab
developed a RNA aptamer targeting EpCAM overexpressed on the cancer cells
(Shigdar et al. 2011). Current results of the study indicated the effective binding and
internalisation of EpCAM targeting aptamer functionalised PEG-SCL-DOX in the
EpCAM-positive cell line. However, confocal microscopy is more a qualitative study.
To further confirm the specific targeting properties of CSCs of our ap-PEG-SCL-
DOX, more specific studies need to be designed. For example, a semiquantitative
cellular uptake study (flow cytometry assay), quantitative cellular uptake study
(HPLC assay) and in vitro toxicity study (MTT assay) should be utilised to confirm
the improved drug delivery property of the EpCAM aptamer functionalised SCL
delivery system in the EpCAM-positive cell lines and compared with that in the
EpCAM-negative cell lines. Furthermore, the newly developed CD133 aptamers in
our laboratory could be utilised as another targeting agent against glioblastoma stem
cells.
As the in vitro monolayers in culture dishes are very different when compared to
the 3-dimentional tissue architecture in vivo (Chen, SF et al. 2012; Xu &
Anchordoquy 2011), the 3-dimentional tumoursphere (gliosphere and colonosphere)
should be utilised for the evaluation of cellular uptake, penetration and retention of
ap-PEG-SCL-DOX in vitro, which would be a more reliable predictor of the
196
therapeutic efficacy in vivo. Moreover, the side-population or the aldehyde
dehydrogenase activity assay of cell fraction can be assessed by flow cytometry,
which are assays utilised to identify CSCs after treatment in vitro (Peeters et al.
2006). The ap-PEG-SCL-DOX is expected to reduce the percentage of CSCs when
compared to other DOX formulation as it targets CSCs as well as the normal bulk
cancer cells. In vitro tumoursphere formation assays can be utilised to evaluate the
ability of cells treated with ap-PEG-SCL-DOX to grow as flowing spheres in
nonadherent conditions in serum-free environment, which reflects the capacity of
CSC to self-renew in vitro (Azzi et al. 2011). In order to verify the reduction of
CSCs resulted from the ap-PEG-SCL-DOX treatment, in vivo limiting dilution assays
(LDA) and xenotransplantation shall be employed in the future, which is the gold
standard to demonstrate the reduction of CSC after treatment (O'Brien, CA, Kreso, A
& Jamieson, CH 2010; Zhang, M, Atkinson & Rosen 2010).
In conclusion, an effective and safe nanocarrier, SCL, using two human
compatible and degradable lipids for the anticancer agent DOX was characterised in
this project. Due to the SCL encapsulation, DOX is better delivered into the target
cells with a prolonged retention. Furthermore, the favourable pharmacokinetic
properties and improved tissue distributions induced by SCL encapsulation translated
into enhanced antitumour activity, which result in significant growth inhibition,
prolonged life-span and reduced systemic toxicity compared to free DOX. Moreover,
results from preliminary data suggest that SCL have a great potential to overcome the
BBB. Aptamer-directed PEG-SCL was then developed by conjugating the EpCAM
targeting aptamer to the surface of PEG-SCL loaded with DOX. The aptamer
functionalised DOX-encapsulated PEG-SCL showed improved uptake and retention
in EpCAM-positive tumour cells in vitro. With the successful development and
197
characterisation of a novel and safe nanocarrier for DOX, future studies will
investigate the potential of SCL in improving the therapeutic index of the traditional
anticancer drugs, in overcoming the BBB and in targeting CSCs.
198
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