dendrimer functionalized magnetic nanoparticles as …...(saumya nigam) (iitb: 09411414) (mu:...
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Dendrimer functionalized magnetic nanoparticles
as promising platforms for cancer theranostics
Submitted in the partial fulfilment of the requirements of the degree of
Doctor of Philosophy
of the
Indian Institute of Technology Bombay, India
and
Monash University, Australia
by
Saumya Nigam
Supervisors:
Prof. D. Bahadur (IIT Bombay)
Dr. X. Chen (Monash University)
The course of study for this award was developed jointly by
the Indian Institute of Technology, Bombay and Monash University, Australia
and given academic recognition by each of them.
The programme was administered by The IITB-Monash Research Academy
(Year 2016)
Dedicated to...
My Grandfather
Declaration
I declare that this written submission represents my ideas in my own words and where others’
ideas or words have been included, I have adequately cited and referenced the original
sources. I also declare that I have adhered to all principles of academic honesty and integrity
and have not misrepresented or fabricated or falsified any idea/data/fact/source in my
submission. I understand that any violation of the above will be cause for disciplinary action
by the Institute and can also evoke penal action from the sources which have thus not been
properly cited or from whom proper permission has not been taken when needed.
Notice 1
Under the Copyright Act 1968, this thesis must be used only under the normal conditions of
scholarly fair dealing. In particular no results or conclusions should be extracted from it, nor
should it be copied or closely paraphrased in whole or in part without the written consent of
the author. Proper written acknowledgement should be made for any assistance obtained
from this thesis.
Notice 2
I certify that I have made all reasonable efforts to secure copyright permissions for third-
party content included in this thesis and have not knowingly added copyright content to my
work without the owner’s permission.
(Saumya Nigam)
(IITB: 09411414)
(MU: 22667563)
Date: 24.10.2016
i
Abstract
Cancer therapeutics deals with development of preclinical therapeutic drugs for
cancer which are developed through innovative research and technologies. This itself is a
very challenging task and it involves various approaches to prevent, diagnose and treat
cancer. Various functional nanomaterials have been developed for improving the efficiency
of therapeutic drugs against cancer. For use in biomedical applications, the nanomaterials
must exhibit unique properties like reduced size, aqueous stability, biocompatibility, and
interactive functional groups. The ‘active’ surfaces of nanoparticles could be modified by
organic or inorganic materials, such as macromolecules, bio-molecules, drugs, etc. Amongst
various functional materials, magnetic nanomaterials have emerged as versatile nanosystems
promising for the detection, diagnosis and treatment of cancers. Superparamagnetic iron
oxide (Fe3O4) nanoparticles have been thoroughly investigated as drug delivery vectors,
magnetic drug targeting agents, contrast agents in magnetic resonance imaging (MRI) and
hyperthermia treatment of cancer. Dendrimers are another emerging class of functional
nanomaterials which are hyper branched, mostly symmetrical polymers with repetitive
branching units. The presence of multiple functional groups makes them ideal candidates for
anchoring guest molecules and therefore, they are assessed as delivery vectors, MR imaging
agents, stabilizers of molecules, catalysis, sensing etc. Combining these two nanomaterials
would contribute towards the development of ‘smart’ and versatile nanosystems with desired
properties.
Citric acid functionalized Fe3O4 (CA–Fe3O4) and glutamic acid functionalized Fe3O4
(Glu–Fe3O4) aqueous colloidal magnetic nanoparticles were synthesized. The surface
engineering made the Fe3O4 hydrophilic and facilitated its aqueous suspensions. Their
successful synthesis was confirmed by X-Ray diffraction (XRD) studies while infrared
spectroscopy (FTIR) was used to confirm the surface modification. The electron
micrographs showed that the nanoparticles were spherically-shaped, evenly dispersed and
magnetometry confirmed their superparamagnetic nature (MS = 57 emu/g at 20 kOe). Time-
dependent calorimetric measurements determined the specific absorption rate (SAR) of CA–
Fe3O4 and Glu–Fe3O4 which is an important parameter in evaluating their heating efficacy
in the presence of alternating magnetic field (ACMF). The SAR of CA–Fe3O4 was found to
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be 49.24 W/g at an applied field of 10 kA/m. On the other hand, SAR of Glu–Fe3O4
nanoparticles was ~134 W/g. These values strongly suggested that these nanoparticles could
also act as effective heating source for magnetic hyperthermia. Doxorubicin hydrochloride
(DOX) was used as a model drug to evaluate their performance for drug delivery. The DOX
molecules were released in substantial amounts in the mild acidic environments. These
nanoparticles showed no loss of cell proliferation activity with the mouse fibroblast (L929),
human cervical carcinoma (HeLa), oral carcinoma (KB), prostate cancer (PC-3) and human
breast cancer (MCF-7) cells.
A peptide dendrimer of second generation was synthesized by divergent method
following the condensation reaction of L-lysine and L-arginine as monomeric units. Its
structural characterization was carried out by various sophisticated techniques such as
nuclear magnetic resonance (NMR), FTIR and X-ray photoelectron spectroscopy (XPS).
The Glu–Fe3O4 were functionalized by various generations of polyamidoamine
(PAMAM) and as-prepared 2nd generation peptide dendrimer and their performances as
delivery vectors for cationic (doxorubicin) and anionic (epigallocatechin gallate) drugs were
assessed. It was seen that the drug loading and release efficiency increased with an increase
in the dendrimer generation. Further, the peptide dendrimers exhibited similar drug loading
efficiency as the PAMAM dendrimers, however, their drug release capacity was
significantly improved. These nanosystems were also found to be potentially effective in
magnetic hyperthermia and were thus used towards in vitro combinatorial chemo-thermo
therapy of cancer using human cervical cancer (HeLa) cells as exemplary model. For the
magnetic hyperthermia treatment, the exposure of these cells to ACMF for 10 min was
successful in reducing the viable cell population by 50% (LD50). While exploring the
combinatorial therapy, it was seen that DOX in synergism with magnetic hyperthermia,
enhanced therapeutic effects and successfully reduced the viable cell population to ~2%.
The in vivo studies investigated the performance of dendrimer functionalized Fe3O4
nanoparticles in a subcutaneous syngeneic murine melanoma model. The systemic exposure
of these nanoparticles (15 mg/kg body weight) caused changes in various blood and serum
parameters. This assisted with the information about their toxicity and non-specific uptake
by various organs. All the vital organs (heart, kidneys, lungs, liver, spleen, brain, stomach
and thigh muscles) showed no loss of activity. The biochemical parameters also remained
iii
unaltered in the treated mice in comparison to the control mice population confirming a
healthy liver and renal activity. The atomic emission spectroscopy (ICP-AES) of the organs
showed that the magnetic nanoparticles were mainly accumulated in the liver, lung and
spleen of the mice with a meagre amount also seen in heart and kidneys. The evaluation of
efficient magnetic drug targeting (MDT) revealed ~6-fold accumulation of iron in the
tumour region when compared to the tumour of mice which were not exposed to the
magnetic field. This high localisation pattern led to high concentrations of DOX in the
tumour and thus was effective in arresting the tumour growth significantly. It was seen that
lower number of doses are sufficient to suppress the tumour growth in combination with
magnetic field than that required without the magnetic field. By the end of 14th day, the
average tumour volume was 55 ± 8.3 mm3 as compared to the control animals in which the
tumour volume was seen to be 4794 ± 844 mm3 (~88-fold decrease).
Furthermore, the dendrimer functionalized Fe3O4 nanoparticles were also seen to be
MR active and showed higher relaxivities when compared to the commercial magnetic
contrast agent. Various physico-chemical parameters potentiate and affect the MR contrast
properties. The relaxivities of the nanoparticles were evaluated under the varying parameters
of iron concentration, buffer environments (ultrapure water, buffered saline and simulated
body fluid) and temperatures (25, 37 and 45 °C). It was seen that under the conditions of
simulated body fluid environment at 37 °C, the peptide dendrimer functionalized Fe3O4
nanoparticles show higher r2 (spin-spin) relaxivity of 220 mM-1s-1. In vitro T2 weighted MR
imaging of dendrimer functionalized Fe3O4 treated (for various treatment times) HeLa cells
showed increased contrast when compared to the untreated cells. Owing to the high
magnetisation, high specific absorption rate and shorter transverse relaxation time, the
dendrimer functionalized Fe3O4 nanoparticles could be a suitable platform for MR imaging
and multimodal cancer theranostics.
Keywords: Cancer, Chemotherapy, Contrast agent, Dendrimers, Drug delivery, Functional
nanomaterials, Hyperthermia, Iron oxide, Magnetic drug targeting, Magnetic
nanoparticles, Magnetic resonance imaging, Superparamagnetism, Theranostics.
iv
Table of Contents
Abstract ................................................................................................................................. i
List of Figures ................................................................................................................... viii
List of Tables .................................................................................................................... xiv
Abbreviations and Nomenclatures ................................................................................... xv
Chapter 1 .............................................................................................................................. 1
Functional Nanomaterials and their Applications in Cancer Theranostics ................... 1
1.1 Superparamagnetic Iron Oxide Nanoparticles ............................................................. 1
1.2 Dendrimers ................................................................................................................... 2
1.3 Cancer ........................................................................................................................... 3
1.4 Cancer Theranostics ..................................................................................................... 5
1.4.1 Chemotherapy ........................................................................................................ 7
1.4.2 Magnetic Hyperthermia and Chemo-thermotherapy ........................................... 13
1.4.3 Diagnostic Bio-imaging ....................................................................................... 16
1.5 Objectives of the Thesis ............................................................................................. 20
1.6 Outline of the Thesis .................................................................................................. 21
1.7 References .................................................................................................................. 21
Chapter 2 ............................................................................................................................ 28
Citrate-stabilised Fe3O4 Nanoparticles: Conjugation and Release of Doxorubicin for
Cancer Therapeutics .......................................................................................................... 28
2.1 Introduction ................................................................................................................ 28
2.2 Experimental and Characterisation Techniques ......................................................... 30
2.2.1 Synthesis of Citrate-stabilised Fe3O4 Nanoparticles ............................................ 30
2.2.2 Drug Loading and Release ................................................................................... 30
2.2.3 Calorimetric Measurements ................................................................................. 31
2.2.4 In vitro Evaluation ............................................................................................... 32
2.2.5 Characterisation Techniques ................................................................................ 32
2.3 Results and Discussions ............................................................................................. 33
2.3.1 Nanoparticle Characterisation.............................................................................. 33
2.3.2 Drug Loading and Release ................................................................................... 36
2.3.3 Calorimetric Measurements ................................................................................. 38
2.3.4 In vitro Evaluation ............................................................................................... 39
v
2.4 Summary .................................................................................................................... 40
2.5 References .................................................................................................................. 40
Chapter 3 ............................................................................................................................ 44
PEG-Modified PAMAM-Fe3O4-drug Triads with the Potential for Improved
Therapeutic Efficacy .......................................................................................................... 44
3.1 Introduction ................................................................................................................ 44
3.2 Experimental Techniques ........................................................................................... 47
3.2.1 Synthesis and surface modification of Fe3O4 nanoparticles ................................ 47
3.2.2 Analysis of Drug Loading and Release ............................................................... 48
3.2.3 Evaluation of Carriers’ Biocompatibility and Therapeutic Efficacy ................... 49
3.3 Results and Discussions ............................................................................................. 49
3.3.1 Characterisation of Fe3O4 Nanoparticles ............................................................. 49
3.3.2 DOX Loading and Release .................................................................................. 53
3.3.3 In vitro Evaluation of Cellular Toxicity and the Therapeutic Effect ................... 57
3.4 Summary .................................................................................................................... 59
3.5 References .................................................................................................................. 60
Chapter 4 ............................................................................................................................ 65
Section A ............................................................................................................................. 65
4.1 Synthesis of Peptide Dendrimer and its in vitro Characteristics ................................ 65
4.1.1 Introduction .......................................................................................................... 65
4.1.2 Experimental and Characterisation Techniques .................................................. 67
4.1.3 Results and Discussions ....................................................................................... 69
4.1.4 Summary .............................................................................................................. 72
4.1.5 References ............................................................................................................ 73
Chapter 4 ............................................................................................................................ 75
Section B .............................................................................................................................. 75
4.2 Dendritic Fe3O4 Nanoparticles for Combinatorial Therapy: Peptide Dendrimers with
Enhanced Efficiency as Alternative Platforms for PAMAM Dendrimers ....................... 75
4.2.1 Introduction .......................................................................................................... 75
4.2.2 Experimental Techniques .................................................................................... 78
4.2.3 Results and Discussions ....................................................................................... 81
4.2.4 Summary .............................................................................................................. 95
vi
4.2.5 References ............................................................................................................ 95
Chapter 5 ............................................................................................................................ 98
Assessment of Doxorubicin-loaded Dendritic Fe3O4 Nanoparticles for Magnetic Drug
Targeting in Murine Melanoma Model............................................................................ 98
5.1 Introduction ................................................................................................................ 98
5.2 Experimental and Characterisation Techniques ....................................................... 100
5.2.1 In vitro Evaluation ............................................................................................. 100
5.2.2 In vivo Therapeutic Efficacy Studies ................................................................. 101
5.3 Results and Discussion ............................................................................................. 106
5.3.1 In vitro Evaluation ............................................................................................. 106
5.3.2 In vivo Therapeutic Efficacy Studies ................................................................. 109
5.4 Summary .................................................................................................................. 115
5.5 References ................................................................................................................ 116
Chapter 6 .......................................................................................................................... 118
MR Contrast Properties of Dendritic Fe3O4 Nanoparticles ......................................... 118
6.1 Introduction .............................................................................................................. 118
6.2 Experimental Techniques ......................................................................................... 120
6.3 Results and Discussions ........................................................................................... 120
6.3.1 Relaxivity studies ............................................................................................... 120
6.3.2 In vitro T2 weighted imaging ............................................................................. 123
6.4 Summary .................................................................................................................. 124
6.5 References ................................................................................................................ 124
Chapter 7 .......................................................................................................................... 126
Conclusions and future scope .......................................................................................... 126
7.1 Conclusions .............................................................................................................. 126
7.2 Future scope ............................................................................................................. 128
Appendix I ........................................................................................................................ 130
Thermally-activated delivery of curcumin using magnetic liposomes ........................ 130
AI.1 Introduction ............................................................................................................ 130
AI.2 Experimental Techniques ....................................................................................... 133
AI.2.1 Synthesis of Fe3O4 Nanoparticles and Magnetic Liposomes .......................... 133
AI.2.2 Analysis of Drug Loading and Release ........................................................... 134
vii
AI.2.4 Evaluation of Biocompatibility and Therapeutics ........................................... 135
AI.3 Results and Discussion ........................................................................................... 136
AI.3.1 Characterisation of the Synthesised Nanoparticles ......................................... 136
AI.3.2 Encapsulation (Loading) of Curcumin ............................................................ 142
AI.3.3 Release of Curcumin at Elevated Temperatures .............................................. 143
AI.3.4 Biocompatibility of Magnetic Liposomes and Anticancer Therapeutic Effect of
Curcumin-Loaded Magnetic Liposomes .................................................................... 144
AI.4 Summary ................................................................................................................ 146
AI.5 References .............................................................................................................. 146
List of Publications ........................................................................................................... 151
Acknowledgments ............................................................................................................ 154
viii
List of Figures
Figure 1.1 Crystal Structure of Magnetite (© University of Minnesota) .............................. 2
Figure 1.2 Example of Generation 2 PAMAM dendrimer (© Dendritech Inc.) ................... 3
Figure 1.3 A schematic representation of the growth of a tumour (© National Cancer
Institute) ................................................................................................................................. 4
Figure 1.4 Schematic representation of the surface modification process of SPIONs53 ....... 9
Figure 1.5 Drug release (%) in the absence of an AC magnetic field (at room temperature)
and presence of an AC magnetic field (at 42° C) in an acetate buffer (pH 5). Inset:
Fluorescence spectra of the released drug from the DMNCs in the absence and presence of
an AC magnetic field54 ......................................................................................................... 10
Figure 1.6 Intracellular uptake of G3- (a) and G3-FA- (c) stabilised Fe3O4 NPs. (b) and (d)
panels show the binding of G3- and G3-FA-stabilised Fe3O4 NPs with the pre-incubation of
free FA, respectively. Panel (e) shows the morphology of control cells without treatment56
.............................................................................................................................................. 11
Figure 1.7 Schematic representation of the MNP-g-MPS-MTX58 ...................................... 12
Figure 1.8 (a) Confocal laser scanning microscopy image of GL261 mouse glioma cells
following a 2 h incubation with nanoparticle and (b) MRI signal intensity of unlabelled cells
and cells labelled with nanoparticle, measured at 3T using a spin echo sequence91 ........... 19
Figure 1.9 T2-weighted fast spin echo images after injection of 2.5 mg/m1 conjugates per
mouse in 1 h (a) FA-PEG-G3.5 (b) PEG-G3.5@IONPs (c) FA-PEG-G3.5@IONPs. The
arrows denote allograft tumours, which are marked by circles52 ......................................... 20
Figure 2.1 (a) XRD pattern and (b) TEM micrograph of CA–Fe3O4 (Inset of (b) shows the
selected area electron diffraction pattern of CA–Fe3O4)...................................................... 33
Figure 2.2 FTIR spectra of pure CA and CA–Fe3O4 .......................................................... 34
Figure 2.3 (A) TGA-DTA plots of CA–Fe3O4 (B) Zeta-potential of CA–Fe3O4 at different
pH values. Inset ‘a’ shows the hydrodynamic diameter of CA–Fe3O4 obtained from DLS
measurements; inset ‘b’ shows the possible schematic representation of CA–Fe3O4 (CA
coating on the surface of the Fe3O4) .................................................................................... 35
Figure 2.4 Field dependence of magnetisation (M vs. H) plot of CA–Fe3O4 at 300 K. Inset
shows the temperature dependence of magnetisation (M vs. T) measurement .................... 36
ix
Figure 2.5 (a) Normalised fluorescence spectra of (a) 10 mg/ml of DOX (1 ml) reacted with
different amounts (0, 20, 40, 60, 80 and 100 mg) of CA–Fe3O4 for 15 min. Inset shows the
loading efficiency (binding isotherm of DOX with CA–Fe3O4) obtained from quenching of
fluorescence intensities; (b) 10 mg/ml of pure DOX (1 ml) at different time intervals; and
(c) Drug release profile of DOX-loaded CA–Fe3O4 in a cell mimicking environment
(reservoir: pH 5 and sink: pH 7.3 at 37 °C) ......................................................................... 38
Figure 2.6 (a) Time-dependent calorimetric measurements of CA–Fe3O4; and (b) Viabilities
of HeLa cells incubated with medium that contains CA–Fe3O4 .......................................... 39
Figure 3.1 (a) XRD pattern of Glu–Fe3O4 nanoparticles; TEM micrographs showing (b)
Glu–Fe3O4 nanoparticles (c) particle size distribution (d) selected area electron diffraction
(SAED) pattern, and (e) high-resolution TEM image of Glu–Fe3O4 nanoparticles ............ 50
Figure 3.2 FTIR spectra of (a) glutamic acid (Glu) and glutamic acid-coated Fe3O4 (Glu–
Fe3O4) nanoparticles, (b) PEG-PAMAM of generations 3, 5, 6, and (c) Fe3O4–DGX
nanoparticles ........................................................................................................................ 51
Figure 3.3 TGA-DTA of glutamic acid-coated Fe3O4 (Glu–Fe3O4) nanoparticles ............. 52
Figure 3.4 (a) Room-temperature field-dependent magnetisation of Glu–Fe3O4 and Fe3O4
nanoparticles coated with dendrimers of generations 3, 5, and 6 (Fe3O4–DGX) at 20 kOe (b)
Time-dependent calorimetric measurements of Glu–Fe3O4 in different ACMFs. The SAR
was calculated to be ∼134 W/g from the initial slope of the time-temperature curve. ........ 52
Figure 3.5 (a,b,c) Fluorescence intensity of DOX-loaded Fe3O4–DGX nanoparticles. (d)
Drug loading efficiency versus dendrimer generation. The difference between DG3 and DG6
was significant (p < 0.05). There were no significant differences between DG3 and DG5, and
DG5 and DG6 (p > 0.05) ....................................................................................................... 54
Figure 3.6 (a) Loading efficiencies of EGCG on to Fe3O4–DGX nanoparticles (b) Loading
efficiencies of three groups of Fe3O4–DGX (x = 3, 5 and 6) nanoparticles against dendrimer
generation. The difference between DG3 and DG6 was significant (p < 0.05). ................... 55
Figure 3.7 Drug-release profile of (a) DOX-loaded Fe3O4–DGX nanoparticles at pH 5.0 and
7.4 and (b) EGCG-loaded Fe3O4–DGX nanoparticles at pH 7.4 ......................................... 57
Figure 3.8 Percentage relative cell viability versus the concentration of Fe3O4–DGX
nanoparticles. A decreasing trend of biocompatibility of these nanoparticles is seen with
increase in dendrimer generation and concentration. ........................................................... 58
x
Figure 3.9 Confocal laser scanning images of (a) control HeLa cells and (b) treated HeLa
cells after 5 h of incubation (scale−20 μm) (c) Viability of HeLa cells in a medium that
contained either Fe3O4–DG5 or DOX–Fe3O4–DG5 nanoparticles in various amounts. The
difference between any pair of data at each concentration was insignificant (p > 0.05). .... 59
Figure 4.1 Schematic representation of the synthesis of the EDA-KR2 dendrimer ............ 68
Figure 4.2 High-resolution X-ray Photoelectron spectra showing C1s, N1s and O1s core
levels of the EDA-KR2 dendrimer ....................................................................................... 70
Figure 4.3 (a) FTIR spectra (b) thermal degradation profile of the EDA-KR2 dendrimer . 71
Figure 4.4 Cell viabilities incubated with peptide dendrimer. The dendrimer is seen to be
biocompatible for a variety of cell lines, even at concentrations of as high as 20 mg/ml. .. 72
Figure 4.5 (a) XRD pattern; (b) FTIR spectra; (c) thermal degradation profiles and (d, e, f)
electron micrographs of Glu–Fe3O4, PAMAM–IO, and KR2–IO nanoparticles, respectively
(σ ≤ 15%) ............................................................................................................................. 82
Figure 4.6 High resolution XPS spectra for C1s, Fe2p, O1s and N1s of the Glu–Fe3O4,
PAMAM–IO and KR2–IO nanoparticles ............................................................................. 84
Figure 4.7 Drug loading profiles of (a) PAMAM–IO and (b) KR2–IO nanoparticles;
Gaussian profiles of (c) DOX-PAMAM–IO and (d) DOX-KR2–IO; and Stern–Volmer plots
of DOX with (e) PAMAM–IO (R2=0.973) and (f) KR2–IO (R2=0.995) .............................. 85
Figure 4.8 Drug release profile of (a) PAMAM–IO and (b) KR2–IO nanoparticles under the
stimulus of pH 5.0 and 7.3 ................................................................................................... 87
Figure 4.9 Time-dependent calorimetric measurements of (A) PAMAM–IO and (B) KR2–
IO nanoparticles at different ACMF and concentrations ..................................................... 89
Figure 4.10 Dependence of SAR on the strength of applied ACMF for (a) PAMAM–IO and
(b) KR2–IO ........................................................................................................................... 90
Figure 4.11 (a) Synergistic effects of DOX-loaded nanoparticles and hyperthermia on HeLa
cells (b) Viable HeLa cell population in the presence of ACMF for varying treatment times
for both PAMAM–IO and KR2–IO nanoparticles ............................................................... 92
Figure 4.12 Synergistic effects of DOX and magnetic hyperthermia (MHT) on viable HeLa
cell population after exposure to both, DOX-loaded dendritic nanoparticles and ACMF... 92
xi
Figure 4.13 Laser Scanning Confocal microscopy images of treated HeLa cells with FITC-
nanoparticles (control), DOX-PAMAM–IO and MHT, and DOX-KR2–IO and MHT. The
scale bar of the images is 50 μm. ......................................................................................... 94
Figure 5.1 Melanoma tumour bearing C57BL/6 mouse being administered intravenously
with DOX-loaded KR2–IO nanoparticles. The skin around the tumour region was shaved
and neodymium magnet was stuck using a medical tape for 6 h. ...................................... 105
Figure 5.2 (a, b) Biocompatibility profile of PAMAM–IO and KR2–IO with B16F10
melanoma cells after 24 and 48 h respectively (c) Dose-dependent cell viability profile of
DOX-loaded dendritic Fe3O4 nanoparticles after 24 and 48 h. Their IC50 values showed
significant difference for 24 h and 48 h (p < 0.05) while there was no significance between
the DOX-PAMAM–IO and DOX-KR2–IO at same time point (p > 0.05).. ...................... 106
Figure 5.3 (a) Schematic representation of magnetically-guided cellular internalization of
dendritic Fe3O4 nanoparticles, (b) (A, C) Confocal micrographs and (B, D) fluorescence
profiles of PAMAM–IO and KR2–IO nanoparticles internalized by B16F10 melanoma cells
after 24 h (c) Amounts of iron internalized by the melanoma cells estimated using atomic
emission spectroscopy……………………………………………………………………108
Figure 5.4 Histograms depicting variations in blood parameters for (a) red blood cells, (b)
haemoglobin, (c) Platelets, (d) white blood cells, (e) lymphocytes and (f) monocytes. The
study showed variations across 14 days of study after intravenous administration of dendritic
Fe3O4 nanoparticles. The results are expressed as mean ± s.d. (n=3) with statistical
significance *p < 0.05 and **p < 0.01 with respect to control animals. ........................ 11010
Figure 5.5 Histograms showing (a) mean corpuscular volume, (b) mean corpuscular
haemoglobin and (c) mean corpuscular haemoglobin. The results are expressed as mean ±
s.d. (n=3) with statistical significance *p < 0.05 and **p < 0.01 with respect to control
animals. .......................................................................................................................... 11010
Figure 5.6 Histograms depicting variations in serum biochemical parameters for (a) SGPT,
(b) SGOT, (c) ALP, (d) Creatinine and (e) blood urea nitrogen (over different time points
during the 14-day study period). The results are expressed as mean ± s.d. (n=3) with
statistical significance *p < 0.05 and **p < 0.01 with respect to control animals. .......... 1111
Figure 5.7 Biodistribution of (a) PAMAM–IO and (b) KR2–IO nanoparticles through
quantification of iron accumulated in various vital organs represented at different time points
xii
post intravenous administration. The quantified iron is represented as mean ± s.d. after
control values were deducted from the plot. (c) Body weight profiles of mice across 14 days
of study. The changes in weight of animals are non-significant (p > 0.05) and does not show
any critical changes. ........................................................................................................... 113
Figure 5.8 (a) Body weight profiles of mice across 35 days of study. The changes in weight
of control animals are due to the uncontrolled tumour size. The weight of animals of group
II, III and IV are significantly decreased (*p < 0.05) in comparison to the control animals,
(b) tumour volume profiles of untreated control animals against the treated groups II, III and
IV. Tumour growth was significantly inhibited with DOX-loaded KR2–IO nanoparticles as
against animals treated with free DOX. (c) Quantification of iron accumulated in various
vital organs. In absence of magnetic field, the thigh muscles (tumour) tend to accumulate
small amount of iron which is significantly elevated with application of magnetic field. 114
Figure 5.9 Histological analyses of the excised organs after 24 h. Optical micrographs of
the treated tumor exhibited zones with compromised cellular cohesion and structural
integrity with significantly different iron content in absence and presence of magnetic field
(MF). Other vital organs showed minimal accumulation of iron causing little or no toxicity
to normal tissues and bodily functions……………………………………………………115
Figure 6.1 Plots of transverse relaxivity (r2) values of (a) PAMAM–IO and (b) KR2–IO in
ultrapure water ................................................................................................................... 121
Figure 6.2 Plots of transverse relaxivity (r2) values of (a) PAMAM-IO and (b) KR2-IO in
phosphate buffered saline (pH 7.3) (c) phantom images showing substantial reduction in
transverse relaxation times of both the nanoparticles ........................................................ 122
Figure 6.3 Plots of transverse relaxivity (r2) values of (a) PAMAM-IO and (b) KR2-IO in
simulated body fluid (pH 7.4) (c) phantom images showing substantial reduction in
transverse relaxation times of both the nanoparticles ........................................................ 123
Figure 6.4 T2-weighted MR images of HeLa cells with PAMAM-IO (a,c) and KR2-IO (b,d)
in the absence (a,b) and presence (c,d) of magnetic field .................................................. 124
Figure AI.1 (a) XRD pattern of Dx–Fe3O4, TEM micrographs of (b) Dx–Fe3O4 (inset shows
the selected area diffraction pattern of Dx–Fe3O4 and (c) magnetic liposomes (inset shows
magnified image of MLs)................................................................................................... 136
xiii
Figure AI.2 FTIR spectra of (a) dextrin and Dx–Fe3O4 nanoparticles (b) MLs and Cur-MLs
............................................................................................................................................ 138
Figure AI.3 Thermal degradation profiles of (a) Dx–Fe3O4 and (b) magnetic liposomes 140
Figure AI.4 Zeta potential of Dx–Fe3O4 and MLs as a function of pH. Inset shows the
hydrodynamic diameter of the MLs obtained from DLS measurements (1.2 µm) ............ 141
Figure AI.5 (a) Field-dependent magnetisation (M vs. H) plot of Dx–Fe3O4 and MLs at room
temperature (b) Time-dependent specific absorption measurements of Dx–Fe3O4
nanoparticles (Inset depicts the dependence of heat generation on the applied ACMF) ... 142
Figure AI.6 (a) Absorbance spectra of curcumin-loaded MLs against pure curcumin (b)
Drug release profiles of curcumin-MLs at physiological and hyperthermic temperatures 143
Figure AI.7 (a) Percentage of the cell viability of MLs incubated with mouse fibroblasts and
cervical cancer cells for 24 h (b) Dose-dependent evaluation of Curcumin-MLs for
determination of IC50 with HeLa cells ............................................................................... 145
xiv
List of Tables
Table 4.1 The area under the peaks of DOX-loaded PAMAM–IO and KR2–IO nanoparticles
(σ≤10%) ............................................................................................................................... 86
Table 4.2 IC50 values of (a) DOX-PAMAM–IO and (b) DOX-KR2–IO nanoparticles with
various cancer cells. The values are represented in µg/ml of formulations with ≤ 15% (IC50
of pure DOX was 1.1-1.5μM with different cell lines) ........................................................ 88
Table 5.1 Animal denomination, dosage of dendritic nanoparticles and time points of sample
collection for biocompatibility and biodistribution studies (h=hour; d=day) .................... 102
Table 5.2 Animal denomination, dosage of dendritic nanoparticles and time points of sample
collection for biocompatibility and biodistribution studies (h=hour; MF=magnetic field)
............................................................................................................................................ 104
Table 5.3 IC50 values of DOX-PAMAM–IO and DOX-KR2–IO nanoparticles with B16F10
melanoma cells. The values are represented in mg/ml of formulations with ≤ 15% (IC50 of
pure DOX was 0.11 – 0.17 μM25) ...................................................................................... 107
xv
Abbreviations and Nomenclatures
ACMF Alternating Current Magnetic Field
ALP Alkaline Phosphatase
BBB Blood Brain Barrier
BUN Blood Urea Nitrogen
CA Citric Acid
CBC Complete Blood Count
DAPI 4′, 6-Diamidino-2-phenylindole
DCM Dicholoromethane
DOX Doxorubicin
DLS Dynamic Light Scattering
DMSO Dimethyl sulfoxide
DTA Differential Thermal Analysis
EC Epicatechin
ECG Epicatechin-3-gallate
EGC Epigallocatechin
EGCG Epigallocatechin-3-gallate
EPR Enhanced Permeation and Retention
FBS Foetal Bovine Serum
FITC Fluorescein isothiocyanate
FTIR Fourier Transform Infrared Spectrometer
Glu Glutamic Acid
GRAN Granulocytes
Hb Haemoglobin
HCT Haematocrit
HRTEM High-Resolution Transmission Electron Microscope
ICP-AES Inductive Coupled Plasma - Atomic Emission Spectroscopy
LSCM Laser Scanning Confocal Microscope
LYM Lymphocytes
MCH Mean Corpuscular Haemoglobin
xvi
MCHC Mean Corpuscular Haemoglobin Concentration
MCV Mean Corpuscular Volume
MDT Magnetic Drug Targeting
MF Magnetic Field
MHT Magnetic Hyperthermia
MNP Magnetic Nanoparticles
MON Monocytes
MRI Magnetic Resonance Imaging
NMR Nuclear Magnetic Resonance
PAMAM Polyamidoamine
PBS Phosphate Buffered Saline
PEG Polyethylene Glycol
PEI Polyethylene Imine
PI Propidium Iodide
PLT Platelets
PPI Polypropylene Imine
RBC Red Blood Cells
RES Reticulo-endothelial System
SAED Selected Area Electron Diffraction
SAR Specific Absorption Rate
SBF Simulated Body Fluid
SGOT Serum Glutamic Oxaloacetic Transaminase
SGPT Serum Glutamic Pyruvic Transaminase
SPION Superparamagnetic Iron Oxide Nanoparticles
SRB Sulforhodamine-B
TC Curie Temperature
TE Echo Time
TEM Transmission Electron Microscope
TGA Thermo-gravimetric Analysis
TR Repetition Time
VSM Vibrating Sample Magnetometer
xvii
WBC White Blood Cells
XRD X-Ray Diffraction
XPS X-ray Photoelectron Spectroscopy
Parts of this chapter have been published in J. Biomed. Nanotech, 2014, 10, 32-49.
Chapter 1
Functional Nanomaterials and their Applications in
Cancer Theranostics
The fabrication of materials along with a scaling down of their size to nanometres imparts
certain distinctive physical and chemical properties to them due to size and surface effects.
These enhanced spectrum of properties of functional materials show promise of improvements
in diverse applications. At nanoscale dimensions, the physico-chemical, optical, mechanical,
catalytic, electrical, electronic and magnetic behaviour of materials is no longer similar or
dependent on their bulk behaviour. This has led to an increase in the number of materials, both
inorganic and organic. These materials are tailored and applied to various applications. This
chapter introduces two promising functional nanomaterials briefly and elaborately discusses
the various aspects of their use in cancer therapeutics and imaging diagnostics.
1.1 Superparamagnetic Iron Oxide Nanoparticles
Iron oxide nanoparticles primarily exist in two forms: magnetite (Fe3O4) and
maghemite (γ-Fe2O3). Fe3O4 nanoparticles have an inverse spinel crystal structure, with oxygen
forming a face-centred cubic system in which the tetrahedral sites are occupied by Fe3+ and the
octahedral sites are filled with both Fe2+ and Fe3+ ions (Figure 1.1). At nanoscale sizes,
magnetic nanoparticles have a single magnetic domain. The strongly coupled magnetic spins
on each atom combine to produce a particle with a single ‘giant’ spin. This effect is known as
2
superparamagnetism. Magnetic nanoparticles (MNPs) show superparamagnetic behaviour at
room temperature, which is attributed to the size reduction in the material; that is, these
particles show no hysteresis loop in nanoscale dimensions. In recent years, researchers have
optimised the properties of these nanoparticles for the improvement of their multimodal
performance. MNPs have undergone thorough investigation and evaluation in a variety of
biomedical applications such as hyperthermia treatment of cancer, contrast agents for magnetic
resonance imaging (MRI), magnetic separation and sorting of cells and proteins (bio-
recognition), and controlled and targeted drug delivery1-3. This has led to the rise of an entire
field of nanobiomedicine.
Figure 1.1 Crystal Structure of Magnetite (© University of Minnesota)
1.2 Dendrimers
Dendrimers are monodisperse, branched macromolecules with three-dimensional
spatial conformations that have a defined radial symmetry4, 5. This structural uniqueness
imparts a wide range of physical and chemical properties to dendritic molecules. They also
characteristically set them apart from the classical polymeric molecules. Dendrimers consist of
three architectural denominations: the central core, the branching units and the terminal
functional groups. The number of repeated branching molecules assigns generation to the
dendrimers and is also responsible for their globular shape. It is the hyperbranching of the
molecule from the centre of the dendrimer towards the periphery that results in homostructural
layers between the focal points (branching points). The number of focal points from the core
towards the outer surface is the generation number. Thus, generation refers to the number of
repeated branching cycles performed during the synthesis. The core of the dendrimer is denoted
as generation ‘zero’ (G0). The branches of lower generation dendrimers tend to radiate out
3
towards the periphery and exist in open conformation. On the other hand, since the number of
the generation increases, the branches tend to retract and adopt globular conformations in three-
dimensional space, with intramolecular hydrogen bonding governing the structures. The
generation-dependent conformational changes that are confirmed by X-ray analysis
demonstrated that the higher generations are more spherical when compared to the lower linear
generations6. Due to their structural and functional uniqueness, dendrimers have found
successful biomedical applications7 in the fields of MRI8, drug delivery9, nucleotide delivery10,
11 and the like.
Figure 1.2 Example of Generation 2 PAMAM dendrimer (© Dendritech Inc.)
1.3 Cancer
The body is made up of many types of cells that grow and divide in a controlled fashion,
according to the body’s requirement and to keep it healthy. When cells become old or damaged,
they are replaced with new cells, which maintains the cycle of life. When this functional order
process goes out of control, the cells bypass their death and new cells are continuously formed
even though the body does not need them. The cells that have disrupted cellular machinery
divide uncontrollably and become autonomous, which gives rise to a diseased state and is
termed as cancer. Cancerous cells and tissues have abnormal growth rates, shapes, sizes, and
functioning. Cancer cells do not grow faster than normal cells; rather their growth is just
uncontrolled. These extra, unwanted and anomalous cells form a solid mass that is known as a
tumour. These cells also have the tendency to spread to and accumulate in other parts of the
4
body and are responsible for tumours at sites that are at a distance from the origin (metastasis).
This uncontrolled cell proliferation disrupts the balance between cell growth rates and cell
mortality rates, giving rise to dysfunctional cell masses that have irregular shapes and sizes
(Figure 1.3). This disorder can be a result of (a) uncontrolled cell growth and (b) loss of a cell’s
ability to undergo apoptosis, which is programmed cell death. The immune system of the body
is unable to recognise and remove these cells because they are erroneously identified as ‘self’
and not foreign, which poses a major challenge in the treatment of cancer.
Figure 1.3 A schematic representation of the growth of a tumour (© National Cancer
Institute)
Tumours are further classified into following types:
(a) Benign tumours: Are not cancerous. They can often be removed, and, in most cases,
they do not come back. Cells in benign tumours do not spread to other parts of the body.
(b) Malignant tumours: Are cancerous. Cells in these tumours can invade nearby tissues
and spread to other parts of the body.
Cancer types can be grouped into broader categories. The main categories of cancer include:
(a) Carcinoma: Cancer that begins in the skin or in tissues that line or cover internal organs.
(b) Sarcoma: Cancer that begins in bone, cartilage, fat, muscle, blood vessels, or other
connective or supportive tissue.
(c) Leukemia: Cancer that starts in blood-forming tissue such as the bone marrow and
causes large numbers of abnormal blood cells to be produced and enter the blood.
(d) Lymphoma and myeloma: Cancers that begin in the cells of the immune system.
5
(e) Central nervous system cancers: Cancers that begin in the tissues of the brain and spinal
cord.
The cancer therapies that are currently under investigation consist primarily of the drugs
or other molecules that block the growth and spread of cancer. These molecules act by
interfering with biochemical signal cascades that are involved in tumour proliferation and
growth. Thus, these target molecules are called as “molecular targets,” targeted therapies are
referred to as “molecularly targeted therapies”. The molecular and cellular changes that the cell
undergoes are specific to the type of cancer. Therefore, these targeted therapies may be more
effective than classical therapies (including chemotherapy and radiotherapy) due to their
specificity and selectivity. Another approach focuses on the concept of starvation as the rapidly
dividing tumor cells require nutrients proportionately. This requirement leads to rapid
proliferation of blood vessels in the region, angiogenesis, due to increased levels of angiogenic
growth factors locally. Blocking of these local signal molecules, in theory, should lead to death
of cancer cells due to starvation. The general approach in overcoming the current hindrances
is to attach the therapeutic molecules to a nanoscale carrier that will release them at the target
site over an extended period of time or when specifically triggered to do so12. In addition, the
surfaces of these nanoscale carriers may be engineered to seek out and become localised at a
disease site, for example, by identifying and attaching to cancer sites.
1.4 Cancer Theranostics
In the recent past, significant advances have been made in the design and development
of drug delivery systems through the use of multidisciplinary teams that utilise biological,
chemical, physical, and engineering sciences. The underpinning science aims to achieve a
greater understanding of the physico-chemical properties of drugs, their interaction with the
delivery systems, their biological fate and the targeting of drugs at the molecular, membrane
and cellular levels. The rise of modern pharmaceutical and nanobiotechnology promotes the
development of novel nanomaterials-based delivery systems that can overcome biological
barriers, particularly in cancer therapy13-16. The ultimate requirement is to formulate a system
that can circulate through the blood stream, detect molecular changes due to cancer, image the
targeted cancer, deliver the therapeutic agent to the desired location and then monitor the
effectiveness of the therapy. Keeping this in mind, researchers are now inclined towards the
tailoring of nanomaterials with systematic variation and properties, which leads to unique and
varied applications that are well beyond the traditional ones. Among the long list of nanoscale
6
materials that are currently being explored for their prospective use in nanomedicine17-20,
superparamagnetic Fe3O4 nanoparticles are widely favoured because of their biocompatibility
profile21, 22 and reactive surface that can be modified with a variety of molecules23-26. This
flexibility of nanostructure design has led to the use of these nanoparticles in diverse
applications that range from magnetic separation27, biosensors28, 29, medical imaging30-34, drug
delivery33, 35-38 and tissue repair39 to hyperthermia40-42.
MNPs are chosen as drug carriers due to their specific nanostructure fabrication,
tailored drug release properties, and lower toxicity and immunogenicity, which results in
improved treatment efficacy and reduced side effects. These drug loaded MNPs can be guided
to the desired target area using an external magnetic field while simultaneous tracking of the
biodistribution of the particles takes place. Besides this, these nanoparticles respond
significantly and generate heat when exposed to an alternating current magnetic field (ACMF),
which can provide promising therapeutic solutions in the issue of hyperthermia42, 43. However,
for use in therapeutic applications, the colloidal stability of these nanoparticles in aqueous and
physiological mediums, which is usually obtained by either charging the surface or by creating
steric hindrance using surfactant molecules, is crucial. In order to control the surface properties
of these nanoparticles, the surface may be modified with different types of molecules, which
prevents large agglomerates, changes in the structural configuration, biodegradation, and
anchors various cargo molecules44. The main hindrance faced by these magnetic nanoparticles
in delivering the intended cargo to the desired location is their non-specific internalisation by
the cells45. Dendrimers have received significant attention for their ability to functionalise and
stabilise nanoparticles. Their peculiar architecture and flexibility of modification in different
ways have provided them with unique properties46 and could possibly be used in biomedical
applications. They have an edge over other macromolecules and polymers due to their
capability to provide varied functional groups. A second advantage is that they have an internal
cavity for entrapment of guest molecules such as drugs, imaging molecules and targeting
ligands. The presence of multiple functional groups with symmetric perfection and nanometre-
scale internal cavities enables the use of dendritic stabilised nanoparticles in remarkable
biomedical applications such as drug delivery, imaging and sensing44. Lately, researchers have
started to attempt the integration of the unique features of dendrimer chemistry with the
versatile magnetic nanoparticles in a single nanovehicle, in order to enhance their properties;
this broadens the scope of their biomedical applications.
7
Dendrimer-functionalised magnetic nanoparticles can be coupled with target
molecules, drugs and imaging moiety to provide a multifunctional platform for real-time
monitoring and targeted drug delivery. Such delivery, controls the release of the therapeutic
agent and enables its release only at the target site without healthy tissue being exposed to the
drug, thereby minimising side effects. This chapter reviews the recent progress in the design
and fabrication of dendrimer-based magnetic nanoparticles for biomedical applications. An
extensive literature survey shows that there is only a small number of works available on this
particular subject, indicating that these nanomaterials are still in the nascent stage. It also
indicates that and plentiful work, from fundamental biological scientific studies to
commercially viable nanobiotechnology, can be carried out using these exceptionally attractive
materials.
1.4.1 Chemotherapy
As the name suggests, chemotherapy involves the use of various chemicals for the
treatment of diseases. These chemicals interfere with the internal machinery of the cells to
rectify the diseased state to a normal state. Chemotherapy is one of the classical therapies in
the treatment of cancer. A wide spectra of molecules have been explored and used as
therapeutic agents in a variety of cancers. These chemicals damage the internal biochemical
cascades (that is, DNA, RNA and proteins) to initiate apoptosis. However, the main challenge
that is faced by chemotherapy is its unspecific distribution and action in the body as these
molecules are as detrimental to normal cells as to cancer cells. This demands the development
of carriers and delivery vectors for these therapeutic agents, which can identify and selectively
deliver these agents only to the cancer sites47. The development of drug delivery vectors is
currently dominated by liposomes48 and polymeric systems22, which is marked by some critical
disadvantages. Therefore, researchers are in constant search of efficient drug carriers.
Normally, a drug delivery system comprises three components: a therapeutic agent that
is, the drug, a targeting moiety and a carrier. The drug is either incorporated by passive
absorption or chemical conjugation to the carrier. The choice of the carrier molecule is crucial
because it significantly affects the pharmacokinetics and pharmacodynamics of the drugs. A
wide range of materials such as natural or synthetic polymers, lipids, surfactants and
dendrimers have been employed as drug carriers49. One of the major limitations of a drug
delivery system is the challenge of releasing the desired amount of drug to its actual site of
action at a specific rate within a stipulated time. In the current scenario of drug delivery, the
8
bottleneck is caused due to poor solubility and high toxicity of drugs, requirement of high
dosages, non-specific delivery, in vivo degradation and short circulating of the half-lives of the
delivery system50. Therefore, advances in biotechnology and nanomedicine aim to design a
targeted drug delivery system in which the drugs can be effectively and conveniently conveyed
to the desired site, and to improve the pharmacokinetics of easily degradable materials, increase
patient compliance and reduce healthcare costs.
1.4.1.1 Non-targeted drug delivery
Dendrimer-modified magnetic nanoparticles have been explored recently for their
potential as drug delivery vehicles. Wu et al.51 investigated the preparation of dendritic-linear
block copolymer-modified superparamagnetic iron oxide nanoparticles (SPIONs) that have a
core of Fe3O4 nanoparticles and a dendritic-linear block copolymer, polyamidoamine-type
dendron-b-poly(2-dimethylaminoethylmethacrylate)-b-poly(N-isopropylacrylamide)
(PAMAM-b-PDMAEMA-b-PNIPAM) as a shell in order to stabilise the nanoparticles in an
aqueous medium. In the buffer solution, the release amounts of doxorubicin (DOX) from
uncross-linked block copolymer-modified SPIONs were 28.8 and 15.5% at 25 °C and 37 °C,
respectively, while the release amounts from the cross-linked block copolymer-modified
SPIONs were 26.8 and 13.7% at 25 °C and 37 °C, respectively. These results indicate that both,
the uncross-linked and cross-linked polymer-modified SPIONs show thermo-sensitive drug
release behaviours (due to the presence of the thermo-sensitive PNIPAM component). Chang
et al.52 developed a pH-responsive drug release system by conjugating PAMAM dendrimers
with DOX (PAMAM–DOX) and superparamagnetic Fe3O4 nanoparticles. PAMAM
(generation 2.5) was modified by PEG and further modified by hydrazine in order to anchor
DOX molecules. The formulations that were synthesised were found to be sensitive towards
the change in proton concentration in the surrounding medium (pH sensitivity). An interesting
feature of this result was that, after 15 h, around 80% of the DOX was released in pH 5, while
less than 8% of the DOX was released in pH 7.4. The amount of the drug released was governed
by the acidic hydrolysis of these hydrazone bonds.
He et al.53 described dendritic−linear−brush-like triblock copolymer polyamidoamine-
b-poly(2-(dimethylamino)-ethyl methacrylate)-b-poly(poly(ethylene glycol) methyl ether
methacrylate) (PAMAM-b-PDMAEMA-b-PPEGMA)-grafted SPIONs and studied their drug
release properties. PDMAEMA and PPEGMA were added step-by-step onto the surface of
SPIONs using a Cu-mediated radical polymerisation method (Figure 2.1). The release amounts
9
of DOX from the copolymer-modified SPIONs for a period of 48 h at pH 4.7, 7.4 and 11.0
were 83.1, 64.7 and 8.3%, respectively. The results showed that free DOX was able to inhibit
the cell proliferation of NIH 3T3 and HeLa cells more efficiently when compared to the DOX
released from the formulations; this indicated a delay in the release of the drug due to a lower
diffusion rate from the polymeric shell, which showed promise of lower toxicity to the normal
cells. Therefore, on the basis of the biocompatibility and drug release effects, the modified
SPIONs could provide a unique opportunity to design excellent drug delivery systems for
therapeutic applications.
Figure 1.4 Schematic representation of the surface modification process of SPIONs53
Chandra et al.54 synthesised dendrimer-functionalised Fe3O4 nanoparticles for delivery
of DOX. The nanoparticles were synthesised and modified by 3-aminopropyltrimethoxysilane
(APTS). The dendritic structures were grown on these silane-coated Fe3O4 nanoparticles by
using methylacrylate and arginine as the monomeric units. The drug release revealed a
sustained release profile and achieved a plateau after 8 h for a maximum release of 54% of the
drug. Chandra et al. also evaluated the efficacy of the drug release under ACMF, marking the
efficiency of these platforms in combination therapy of cancer for hyperthermia. In the absence
of any external magnetic field, the percentage of drug release was approximately 35% for the
initial 45 mins, which shot up to ~80% on application of ACMF for the next 15 min. Under the
applied field conditions, acid hydrolysis and the increase in temperature contribute
synergistically towards the drug release profile (Figure 1.5).
10
Figure 1.5 Drug release (%) in the absence of an AC magnetic field (at room temperature)
and presence of an AC magnetic field (at 42° C) in an acetate buffer (pH 5). Inset:
Fluorescence spectra of the released drug from the DMNCs in the absence and presence of an
AC magnetic field54
Another important parameter related to drug delivery is the biodistribution and toxicity
of the nanocarriers. It has been found in dendrimer-based MNPs that smaller generation
dendrimers show rapid renal elimination while higher generation dendrimers show
accumulation in the liver55. The binding of the ethylene glycol units renders positive charges
in the dendrimers, which enhances permeation and retention (EPR) and increases circulation
time in the blood. The encapsulation of the drugs within the ethylene glycol arms of the
dendrimers leads to reduced toxicity and lesser accumulation in body organs. Therefore, it is
expected that if the dendrimers are covered with ethylene glycol entity, it can make the
dendrimers attractive drug carriers in vivo.
1.4.1.2 Targeted drug delivery
1.4.1.2.1 Ligand-targeted Drug Delivery
The specificity of the target location still remains a challenge in cancer therapy. Non-
targeted systems tend to deliver the therapeutic molecules to the cancer as well as the
neighbouring normal cells, which leads to side effects. To overcome this barrier, delivery
systems are modified by biomolecules that identify the target cell and, thus, enhance the
specificity of the chemotherapy. Along these lines, Shi et al.56 reported dendrimer-modified
11
Fe3O4 nanoparticles that utilise folic acid as the targeting ligand. Folic acid (FA) was
conjugated with the PAMAM (generation 3) using amide bonding and then incubated with
Fe3O4 nanoparticles. The cell uptake studies reveal that both G3- and G3-FA-stabilised Fe3O4
nanoparticles bind to the KB cells that overexpress FA receptors (Figure 1.6). The presence of
free FA molecules in the growth media also affected (lowered) by the binding of nanoparticles
to the receptors because free FA is the preferred ligand for the receptors. Thus, the results of
Shi et al. suggest the potential use of dendrimer-functionalised Fe3O4 nanoparticles in targeted
delivery of various therapeutic agents.
Figure 1.6 Intracellular uptake of G3- (a) and G3-FA- (c) stabilised Fe3O4 NPs. (b) and (d)
panels show the binding of G3- and G3-FA-stabilised Fe3O4 NPs with the pre-incubation of
free FA, respectively. Panel (e) shows the morphology of control cells without treatment56
Chang et al.57 described a PAMAM-modified Fe3O4 nanocarrier system for targeted
delivery of doxorubicin to MCF-7 cells. The drug release profiles were investigated in a buffer
of pH 5.0 and pH 7.4 at 37°C. After 15 h, about 75% of the DOX was released in low pH while
less than 8% of the DOX release was observed in a physiological pH (7.4), thus establishing
the pH-sensitive drug release behaviour. On similar lines, Li et al.58 fabricated polyglycerol-
conjugated Fe3O4 nanoparticles for targeted delivery of methotrexate (MTX) to cancer cells.
12
An esterification reaction was carried out to form conjugates of 3-(trimethoxysilyl)
propylmethacrylate-coated MNPs (MNP-g-MPS) and MTX, which involved the reaction
between the hydroxyl groups of hyperbranched polyglycerol (HPG) and the carboxylic acid
groups of MTX (Figure 1.7). The drug release profiles were evaluated in a phosphate buffer of
pH 7.4.
Figure 1.7 Schematic representation of the MNP-g-MPS-MTX58
The formulations were evaluated for their cytotoxicity in RAW mouse macrophages, 3T3
fibroblasts and KB cells. The uptake of iron from MNP-g-HPG nanoparticles was very low and
showed the evasion of endocytosis by the cells (0.4-0.5 pg/cell). The conjugation of MTX, on
the other hand, significantly increased the endocytosed nanoparticles for KB cells, (~1.2
pg/cell) while for 3T3 fibroblasts and RAW macrophages, it was less than 0.7 pg/cell. The
formulation of MNP-g-HPG-MTX2 (12.5 µg/mg of nanoparticle) showed a 20-fold increase
in the uptake by the KB cells (12.1 pg/cell) when compared to MNP-g-HPG nanoparticles,
while for 3T3 fibroblasts and RAW macrophages, the uptake was 11 and 8 times lower,
respectively. Depending on the amounts of MTX internalised, the nanoparticles were capable
of killing ~50% of the KB cells while exhibiting low cytotoxicity towards 3T3 fibroblasts and
RAW macrophages.
1.4.1.2.2 Magnetically-guided (Magnetic Drug Targeting) Drug Delivery
Magnetic drug targeting (MDT) in cancer therapy refers to the attachment of
therapeutic molecules to magnetic delivery vectors, after which the molecules are concentrated
13
and released at the cancer site under the guidance of an external homogenous magnetic field59,
60. Ideally, the drugs should be chemically bound to the magnetic nanocarriers, before
administering them into the systemic circulation, magnetically guiding them along with the
blood flow and trapping them around the tumour region. The magnetic field gradient interacts
with these magnetic nanocarriers in the circulation and leads them to accumulate in the tumour
region. For all in vivo applications, it is desirable that MNPs retain sufficient hydrophilicity
and, do not exceed 100 nm so that rapid clearance by the reticulo-endothelial system (RES)
can be avoided. This can be achieved by stabilising the magnetic nanoparticles with dendrimers
through which a smaller, more neutral and hydrophilic surface, which has a longer plasma half-
life61, can be obtained. It has also been shown that a strong magnetic field at the site of the
tumour induces accumulation of drug-loaded magnetic nanoparticles. However, as a proof of
concept, Alexiou and co-workers found quite interesting and encouraging results when they
employed the MDT in the local accumulation of mitoxantrone in the squamous cell carcinoma
implanted in New Zealand White rabbits62. They administered their nanoparticles in the
femoral artery while an external magnetic field was focussed on the tumour. They found a
significant (p < 0.05), complete, and permanent remission of the squamous cell carcinoma as
compared to the control group of animals. They successfully combined intratumoural
nanoparticle accumulation and locoregional cancer treatment without any kind of systemic
toxicity. The principles of magnetic guidance of MNP-conjugated drugs have reached clinical
trials for cancer therapy. MDT has been known to result in desired therapeutics in superficial
tumours, but its ability to effectively reach the deep-seated tissues/cancer sites remains
inadequate63. Despite the advantages of the dendrimer-stabilised magnetic nanoparticles, we
could hardly find any work utilising these nanoparticles for magnetic drug targeting, which
suggested that this area is still unexplored and has promising prospects. Most of the recent
research has centred on drug delivery and, therefore, it is time to devote studies to be carried
out beyond ‘proof of concept’.
1.4.2 Magnetic Hyperthermia and Chemo-thermotherapy
The classical routes of cancer treatment include chemotherapy, radiotherapy and
surgical removal of the tumour mass. These methods of combating cancer are non-selective
and cannot differentiate between cancer and normal cells. A solid tumour has very poor
vasculature, which makes the cells compete for nutrients and oxygen for their high energy
demands. The intercellular interactions are disrupted and the cells become autonomous. The
14
continued, uncontrolled cell division results in a solid tumour mass that turns hypoxic (low on
oxygen) and is responsible for a large number of chemical and metabolic differences between
cancerous and normal cells64. Although this hypoxic nature makes these cells insensitive to
radiations and chemotherapeutic agents, which also hampers its clinical treatment, it also
makes these cancer cells sensitive towards a rise in temperatures65. Due to their compact three-
dimensional structure, it is difficult for these cells to dissipate the applied heat, which in turn
proves lethal to the cells, resulting in apoptosis as a direct response. The heating causes the
coagulation of proteins, denaturing them and obstructing the cellular machinery, leading to cell
death. Even if the cells do not die, this rise in temperatures in combination with the hypoxic
conditions renders these cells vulnerable to chemotherapeutic agents or to being radio-
sensitised. To overcome the shortcomings of classical cancer therapies and to improve
therapeutic effectiveness, hyperthermia has been explored as an alternative therapy in cancer
treatment.
Hyperthermia literally translates to ‘elevated temperatures’. This heat generation, when
obtained through power dissipation of magnetisation reversal processes of nanoparticles, is
termed as magnetic hyperthermia. When a colloidal suspension of magnetic nanoparticles is
exposed to an alternating current magnetic field (ACMF), the reversal of magnetisation occurs
through hysteresis loss and Neel and Brownian relaxation losses66. During the undertaking of
these processes, substantial amounts of energy is dissipated in the microenvironment as heat
and has found applications in hyperthermia treatment of cancer67-70. In ferromagnetic materials,
the heat is generated due to susceptibility and hysteresis losses. Therefore, the power generated
in the case of ferromagnetic materials is proportional to the product of the frequency of ACMF
and the area of the hysteresis loop. There are no Brownian losses (whole particles rotation),
and the vector sum of the rotational losses of the magnetic moments within the particle (Néel
relaxation) tend to cancel each other because the magnetic domains are randomly aligned in a
multi-domain particle system. On the other hand, in superparamagnetic materials this heat is
a direct result of power dissipation due to Brownian and rotational relaxations of single domain
particles. This heating power is governed by various factors and is primarily dependent on the
particle size, magnetic properties of the material, strength (H) and frequency (ω) of the applied
ACMF and solvent properties71-74. Despite being more proficient in their heat-generating
properties, ferromagnetic materials are not chosen over superparamagnetic nanoparticles for
magnetic hyperthermia applications. Recently, various researchers have explored the
possibility of combining magnetic hyperthermia with the chemotherapy of cancer48, 75. It is
15
seen that they act synergistically, which yields more promising results than when they function
in separate therapy options. Since hyperthermia causes severe changes in cellular
macromolecules, affecting all cellular functions, there are two primary concerns in its clinical
use. First, similar to classical therapies, hyperthermia also misses selectivity and affects both,
cancer and normal cells. Second, the foremost challenge in applying hyperthermia to cancer
treatment is the maintaining of local temperatures below 45 °C, so that the neighbouring normal
cells suffer minimally. Also, the combining of chemotherapy with thermotherapy is expected
to result in improved therapeutic performances and in the regression of tumours.
Guardia et al.76 studied the heating behaviour of iron oxide nanocrystals as a function
of size. Between the sizes of 13 and 40 nm, the nanoparticles with an average diameter of 19±3
nm showed high clinically significant SAR values and were able to check cell proliferation in
oral carcinoma cells in vitro efficiently. To address the challenge of localised heating,
multifunctional magnetic nanoparticles have been explored for their heat-generating property
that can be also be modulated according to the required clinical settings. Baba et al.77 undertook
a comparative study to evaluate the uptake and hyperthermic efficiency of superparamagnetic
and ferromagnetic nanoparticles with human breast cancer cells (MCF-7). Their results
demonstrated that the cell death due to ferromagnetic particles was significantly higher than
the cell death due to superparamagnetic nanoparticles that reached temperatures above 45 °C.
Sadhukha and co-workers synthesised 12±3 nm-sized aqueous-stable Fe3O4 nanoparticles with
high magnetisation and SAR properties78. They evaluated the hyperthermic efficiency of these
nanoparticles with human lung adenocarcinoma (A549) and human mammary adenocarcinoma
(MDA-MB-231) cell lines and found magnetic hyperthermia to be much more effective than
conventional hyperthermia (water bath). They concluded that magnetic hyperthermia treatment
resulted in pleiotropic effects with induction of apoptosis and generation of reactive oxygen
species as the mechanistic causes. In one of their other works, Sadhukha and co-workers
developed epidermal growth factor receptor (EGFR)-targeted inhalable Fe3O4 nanoparticles to
target non-small cell lung cancer in order to selectively target the epithelial linings of tumours
in lungs79. Female Fox Chase SCID® Beige mice were chosen to generate an orthotopic lung
tumour model and were subjected to magnetic hyperthermia after successful inhalation and
deposition of these nanoparticles. Exposure of ACMF in these animals was seen to cause an
insignificant decrease in tumour size (p>0.05). This decrease in tumour size was relative to the
control animals that received the same dose but were not exposed to ACMF.
16
It has been seen that the synergistic action of chemotherapy and magnetic hyperthermia
yield much more promising results than as separate therapy options. Qu et al.80 evaluated
camptothecin in combination with hyperthermia against ovarian and liver cancer cells, while
Li et al.81 used 5-fluorouracil, which specifically targeted cancer cells by the anti-human
epidermal growth factor receptor 2 (anti-HER2). These nanoparticles were therapeutically
evaluated with murine bladder cancer cells (MBT-2) and were able to reduce the viable cell
population by 70% after an ACMF exposure of 15 min. In male C3H/HeN mice, these
nanoparticles also showed some promising local hyperthermia effects when administered both
intratumourally as well as intravenously, with significant reduction in the tumour volume in
small tumours (<50 mm3). On similar lines, Kossatz and co-workers found that hyperthermia
along with doxorubicin successfully reduced the tumour volume by 40% in subcutaneous
MDA-MB-231 tumour-bearing female athymic nude mice82. Various formulations have been
fabricated and studied for applications in hyperthermia and combinatorial therapy, but
dendrimer functionalised magnetic nanoparticles have not yet been explored much.
1.4.3 Diagnostic Bio-imaging
An accurate diagnosis and detection of cancer at an early stage is one of the major
limitations in existent cancer therapy. Typically, the tumours that grow in the deep-seated
tissues/organs go undetected, leading to inadequate treatment measures that have high
morbidity, which leads to lower chances of the survival of the patient. Tumour detection, thus,
would require a non-invasive diagnostic technique with high spatial resolution, which would
investigate the anatomy and physiology of the tumour even in the early stages. New and
customised nanoparticulate formulations exploit the properties of multiple imaging modalities
in order to improve cancer diagnosis and to monitor the response to chemotherapy. Two such
imaging techniques are optical imaging and magnetic resonance imaging (MRI). Optical
imaging primarily utilises the fluorophores or the tagging molecules that provide sensitivity
and spatial and temporal resolution to the nanoparticles; however, their imaging through
penetration is limited to a few millimetres below the tissue.
MRI is a high-resolution medical imaging technique that is widely used in clinical
settings for the physiology of both healthy and diseased bodies. It uses the magnetic relaxation
of the proton nuclei of the water in the body and generates three-dimensional tomographic
images. Proton relaxation is a very slow process that reduces the resolution of the images,
resulting in compromised quality and unclear information. In order to improve the contrast of
17
the images obtained, additional mediator materials are required. These are known as contrast
agents. Gadolinium-based paramagnetic compounds have been used to aid the quickening of
proton relaxation, which results in adequate MR signals that generate high-quality, brighter
images83, 84. Another class of contrast agents are superparamagnetic nanoparticles. Magnetic
nanoparticles have not only shown potential in selective delivery of anticancer drugs but have
also exhibited highly desirable sensitivity and specific imaging capabilities. The higher
magnetic moment of these nanoparticles gives rise to alterations in the surrounding magnetic
fields of the MRI machine. This alteration in the field gradient results in the shortening of the
relaxation time of the proton nuclei. As a result, an image that has high contrast characteristics
is generated along with sufficient information about the region in question, especially in the
case of tumours85, 86. MRI results provide improved image resolution and tissue contrast
through the use of nanoparticles and contribute towards the revelation of detailed tissue
morphology and anatomy for whole body imaging of animals and humans.
1.4.3.1 Optical Imaging
Tumour detection by using fluorescent probes may be quite advantageous in terms of
improved biocompatibility when compared to other types of contrast agents, but they suffer
from poor penetration of light through tissues. To provide a minimally invasive solution to this
barrier, nanotechnology has enabled the use of water-soluble, functionalised MNPs that are
highly stable and, when tagged with fluorescent molecules, can manifest extensive fluorescent
properties for in vivo imaging of live cells. Although dendrimers have not been explored much
in the context of MNPs, they have been used as matrices for the synthesis of lanthanide-
magnetic nanohybrids for imaging. Luwang et al.87 reported the use of a poly (amido amine)
(PAMAM) dendrimer for the first time as a cross-linking agent for the synthesis of a
luminescent lanthanide-based multifunctional nanohybrids (YPO4:Tb3+@Fe3O4) with strong
luminescence properties that show their ability to label cervical cancer cells for bio-imaging.
In this respect, Wate et al.88 envisaged the immense biological scope of dendrimers and Fe3O4
nanoparticles. They realised that, if used concurrently with graphene oxide (GO), the
dendrimer-modified Fe3O4 nanoparticles would provide physico-chemical advantages such as
a solution/dispersion state and a superparamagnetic property, especially with regard to their
biological implications and in reducing the cellular toxicity of GO. Hence, a multicomponent
GO nanostructured system that consists of Fe3O4 nanoparticles, the PAMAM-G4-NH2
dendrimer and Cy5 has been fabricated, which exhibited high dispersion in an aqueous
18
medium, strong NIR optical absorbance and magnetically responsive properties. In vitro
experiments provide a clear indication of successful uptake of the nanoparticles by MCF-7
breast cancer cells, and this multicomponent GO nanostructured system is seen to behave as a
bright and stable fluorescent marker.
1.4.3.2 Magnetic Resonance imaging (MRI)
MRI is a non-invasive imaging technique with high spatial resolution and tomographic
capabilities. The technique utilises the magnetic relaxation of water protons in the body to
generate images. However, the signals generated from the water protons are not enough for
accurate image construction, which might lead to improper diagnostics. This problem is
addressed by the use of contrast agents. In simplistic terms, the contrast agents create a local
magnetic field, which in turn affects the relaxation of water protons. These enhanced signals
are, thus, used to provide better images with improved contrast. The sensitivity in biological
targets depends on the specific and selective accumulation of the contrast agents at the target
site. Since quite some time, targeted MNPs have been used to enhance the MR (magnetic
resonance) signal sensitivity for in vivo tumour detection. In this direction, dendrimers have
been seen to have a significant effect on the corresponding MR relaxivities and physiological
properties of MNPs89. Thus, dendrimer-MNPs agents are expected to provide sharper images
with physiologically relevant contrast, enhanced blood-pool retention time and specific organ
uptake. The fine tuning of the size and end-group functionalities of the dendrimers provide
added advantage in this respect90. Superparamagnetic Fe3O4 nanoparticles have been shown to
be effective contrast agents in labelling cells in order to provide high sensitivity in MRI, but
this sensitivity depends on their ability to label cells with sufficient quantities of the
nanoparticles, which is otherwise challenging.
To address this issue, a cell-penetrating polyester dendron with peripheral guanidines
was developed and conjugated to the nanoparticles surface by Martin and his co-workers91. In
GL261 mouse glioma cells, the dendritic guanidine exhibited remarkable cell-penetrating
capabilities to the HIV-Tat 47-57 peptide for the transport of fluorescein, and when conjugated
to the nanoparticles, it exhibited significantly enhanced uptake, in comparison to nanoparticles
that have no dendron or dendrons with hydroxyl or amine peripheries. T2 values of the cell
19
pellets, using a spin echo sequence on a 3T MRI scanner, were found to be 65 and 6.1 ms for
unlabelled and labelled cell pellets, respectively (Figure 1.8). The substantial decrease
observed in the transverse relaxation time (T2) of labelled cells that are relative to control cells
illustrates the potential utility of these nanoparticles in labelling cells for detection by MRI.
Figure 1.8 (a) Confocal laser scanning microscopy image of GL261 mouse glioma cells
following a 2 h incubation with nanoparticle and (b) MRI signal intensity of unlabelled cells
and cells labelled with nanoparticle, measured at 3T using a spin echo sequence91
Duanmu et al.92 reported water-soluble Fe3O4 nanoparticles that were coated with three
different generations (G1, G2 and G3) of melamine dendrons, and investigated the MRI
contrast-enhancement potential of dendron-functionalised nanoparticles. The R2 relaxivities for
the G2 and G3-modified nanoparticles were found to be significantly larger than those with G1
(with the G2, R2 value being the largest at ~260 mM-1s-1), which may be likely due to the
physico-chemical nature of the dendron. Though a detailed understanding of these generation-
specific effects requires further study, the strong transverse relaxivities with generation-
specific values and tunable physico-chemical properties that were obtained made them an ideal
choice as contrast agents in MRI applications. Acknowledging the status mentioned above,
Chang et al.52, 57 presented a breakthrough in the development of new synthetic dendrimer-
modified magnetic Fe3O4 conjugates in the enhancement of anatomical MR contrast. The
system is based on the conjugates of FA), poly (ethylene glycol) (PEG)-modified dendrimers
(PAMAM) with DOX and superparamagnetic Fe3O4 (FA-PEG-PAMAM-DOX@IONPs). For
MR imaging, C57BL/6 female mice (6–7 weeks) were implanted with B16F10 melanoma
cancer cells. Tumours are seen as hyper-intense areas in T2-weighted MR images, and for all
mice injected with IONPs, darkening of T2-weighted MR images from the tumour areas at 1 h
20
post-injection relative to pre-injection indicated that the IONPs accumulated and made the
tumours clearly detectable (Figure 1.9).
Figure 1.9 T2-weighted fast spin echo images after injection of 2.5 mg/m1 conjugates per
mouse in 1 h (a) FA-PEG-G3.5 (b) PEG-G3.5@IONPs (c) FA-PEG-G3.5@IONPs. The
arrows denote allograft tumours, which are marked by circles52
Another encouraging attempt was made to achieve dual objectives of demonstrating
magnetic resonance and fluorescence imaging simultaneously by grafting small-sized dendrons
on the surface of Fe3O4 nanoparticles93. From the T1 and T2 values measured by relaxometry at
1.5T, mean relaxivity values of 272 mM-1s-1for r2 and with a 26.4 r2/r1 ratio were obtained for
the nanoparticles bearing carboxylate functions at their periphery. The authors attributed the
improved relaxivity values to the optimal design of the small-sized dendritic organic shell,
attractive magnetic properties of the inorganic core and the dendrons anchored through
phosphonate functionalities. This study confirmed the superiority of the dendritic approach to
develop new, smart and multimodal contrast agents.
1.5 Objectives of the Thesis
(a) Synthesis and characterisation of magnetite nanoparticles by the soft chemical route;
(b) Synthesis and characterisation of the peptide dendrimer;
(c) Development of multifunctional dendritic magnetic nanoparticles and optimisation of their
properties relevant for various cancer theranostic applications;
(d) Thorough in vitro assessment of these formulations for cancer therapeutics, magnetic
resonance imaging and in vivo assessment of their compatibility, biodistribution and magnetic
drug targeting; and
21
(e) Establishing the comparative potential of as-prepared peptide dendrimers against the widely
used commercial polyamidoamine (PAMAM) dendrimers.
1.6 Outline of the Thesis
This thesis is organised into seven chapters and an appendix. The first chapter
introduces the functional nanomaterials, Fe3O4 nanoparticles and dendrimers, and thoroughly
reviews their current and probable use as theranostic agents in combating cancer. In the chapter
two, the synthesis and characterisation of aqueous colloidal suspension of citric acid-coated
Fe3O4 is reported. The chapter evaluates the biocompatibility, heating ability and therapeutic
efficacy of the drug-loaded nanoparticles against cervical cancer cells. Chapter three describes
the functionalisation of Fe3O4 with polyamidoamine dendrimer nanosystems to develop a
conjugate drug delivery platform. A comparative therapeutic evaluation of this system is dealt
with on the basis of the generation of the dendrimer. Section one of the fourth chapter, describes
the elaborate synthesis and characterisation of the peptide dendrimer. The properties of this
dendrimer were optimised so as to enhance the biocompatibility while minimally
compromising on their physico-chemical properties. Section two of the fourth chapter
describes the synthesis and characterisation of dendrimer-functionalised Fe3O4 nanoparticles
(both PAMAM and as-prepared peptide). An elaborate evaluation of their biocompatibility,
drug delivery efficiency, therapeutic efficacy and their potential use in in vitro combinatorial
chemo-thermotherapy is evaluated. Taking a step forward with these formulations, chapter five
reports the comparative in vivo assessment of biocompatibility, bio-distribution and therapeutic
efficacy of these nanoparticles in C57BL/6 black mice. The efficacy of DOX-loaded dendritic
Fe3O4 nanoparticles were evaluated for tumor regression via intravenous administration, based
on the concept of MDT. Chapter six evaluates the MR relaxivity properties of these
nanoparticles and its dependence on various parameters, mainly elevated temperatures. In
chapter seven, the conclusions of the present work and their intended prospective investigation
are discussed. The appendix one explores the combination of magnetic nanoparticles with lipid
vesicles that form magnetic liposomes for therapeutic efficacy of a hydrophobic drug.
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Chemistry of Materials, 2001, 13, 2201-2209.
90. J. W. Bulte, T. Douglas, B. Witwer, S. C. Zhang, E. Strable, B. K. Lewis, H. Zywicke,
B. Miller, P. van Gelderen, B. M. Moskowitz, I. D. Duncan and J. A. Frank, Nature
biotechnology, 2001, 19, 1141-1147.
91. A. L. Martin, L. M. Bernas, B. K. Rutt, P. J. Foster and E. R. Gillies, Bioconjugate
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27
92. C. Duanmu, I. Saha, Y. Zheng, B. M. Goodson and Y. Gao, Chemistry of Materials,
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93. G. Lamanna, M. Kueny-Stotz, H. Mamlouk-Chaouachi, C. Ghobril, B. Basly, A.
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This chapter has been published in J. Magn. Magn. Mater., 2011, 323, 237–243.
Chapter 2
Citrate-stabilised Fe3O4 Nanoparticles: Conjugation and
Release of Doxorubicin for Cancer Therapeutics
2.1 Introduction
Magnetic nanoparticles have received a great deal of attention due to their potential in
biomedical applications such as hyperthermia treatment of cancer, as a contrast agent for
magnetic resonance imaging, magnetic separation and sorting of cells and proteins (bio
recognition), and controlled and targeted drug delivery1-6. In the recent past, among the vast
varieties of magnetic nanoparticles, superparamagnetic Fe3O4 nanoparticles have emerged as
an excellent candidate for biomedical applications due to their better chemical stability and
biocompatibility in comparison to other metallic magnetic nanoparticles7, 8. Many methods
have been developed to prepare superparamagnetic Fe3O4 nanoparticles9. The thermal
decomposition of organometallic precursors in an organic solvent (that has a high boiling point)
at elevated temperatures in the presence of surfactants has been successfully used for the
synthesis of monodisperse Fe3O4 nanoparticles3, 10. However, the Fe3O4 nanoparticles prepared
by these methods are highly hydrophobic, which hampers their biomedical applications,
especially, in drug delivery, hyperthermia treatment of cancer and magnetic resonance
imaging. Although many ligand-exchange processes have been established to offer them a
hydrophilic surface characteristic for aqueous stability, their magnetic field responsiveness has
29
not been effectively improved3, 9. Therefore, much effort has been focused on the fabrication
of biocompatible aqueous stable superparamagnetic Fe3O4 nanoparticles and engineering them
with the desired functionality, which offer good magnetic responses, through soft-chemical
routes2, 11-16.
The strategies used for surface functionalisation comprise grafting of, or coating with
organic species, such as surfactants or polymers, or coating with an inorganic layer such as
silica or gold13-22. Further, the presence of these biocompatible layers on the surface not only
stabilises the Fe3O4 nanoparticles, but also provides an accessible surface for routine
conjugation of biomolecules through the well-developed bioconjugation chemistry. Thus, the
stability of the bonding between functional molecules and Fe3O4 nanoparticles is crucial from
the point of view of the application. The small-molecule targeting groups are attractive due to
their ease of preparation and simple conjugation chemistry18, 23, 24. Further, multiple grafting or
coating of small molecules can provide multivalent systems that exhibit significantly enhanced
efficacy in drugs and biomolecules25. On the other hand, some binding affinity may be lost
through steric hindrances by large surfactant molecules or long polymer chains, which could
be easily overcome by the use of small molecules that have multiple functional groups such as
carboxyl (COOH), amine (NH2), thiol (SH) and the like. Furthermore, the presence of a large
number of uncoordinated functional groups on the surface of the magnetic nanoparticle can be
used for the linking of various biomolecules and drugs.
Among various small molecules, citrate moiety has been extensively used in the
preparation of aqueous stable Fe3O4 nanoparticles and in their conjugation to biomolecules and
drugs by exploiting the uncoordinated carboxyl groups present on the surface of nanoparticles.
Liu et al.26 fabricated highly water-dispersible Fe3O4 particles that have a uniform size by
solvothermal reaction at 200 °C through the reduction of FeCl3 in the presence of trisodium
citrate as a stabiliser. They found that these Fe3O4 nanoparticles with surface citrate groups can
effectively enrich peptides at a trace level. Munnier et al.12 developed a new method for
reversible association of drug (DOX) to citrate-stabilised Fe3O4 nanoparticles that were
prepared by agitating bare nanoparticles in a citric acid solution. Khosroshahi and Ghazanfari17
fabricated citrate-modified Fe3O4 nanoparticles by stirring bare Fe3O4 nanoparticles in
trisodium citrate solution as an intermediate to obtain silica-coated Fe3O4 core-shell
nanoparticles. However, most of these works on the fabrication of aqueous stabilised Fe3O4
nanoparticles are achieved at either elevated temperatures2, 3, 26 or involve multiple synthesis
steps12, 13, 15. The goal of this chapter was to develop a facile single-step process for the
preparation of highly biocompatible citric acid functionalised (citrate-stabilised) Fe3O4
30
aqueous colloidal magnetic nanoparticles (CA–Fe3O4), with optimal magnetisation at low
temperatures, as potential drug carriers, which can also be used as effective heating source for
the hyperthermic treatment of cancer.
2.2 Experimental and Characterisation Techniques
2.2.1 Synthesis of Citrate-stabilised Fe3O4 Nanoparticles
In a typical synthesis, 4.44 g of FeCl3 and 1.732 g of FeCl2 were dissolved in 80 ml of
water in a round-bottomed flask and the temperature was slowly increased to 70 °C in a
refluxing condition under a nitrogen atmosphere with constant mechanical stirring at 1000 rpm.
The temperature was maintained at 70 °C for 30 min. Further, 20 ml of ammonia solution was
added promptly to the reaction mixture and kept at the same temperature for another 30 min.
Then 4 ml of an aqueous solution of citric acid (0.5 g/ml) was added to the reaction mixture
mentioned above and the reaction temperature was slowly raised to 90 °C under reflux and
reacted for 60 min with continuous stirring. Black precipitates were obtained by cooling the
reaction mixture to room temperature, followed by thorough rinsing with water. During each
rinsing step, samples were separated from the supernatant by using a permanent magnet.
2.2.2 Drug Loading and Release
The anticancer agent, doxorubicin hydrochloride (DOX) was used as a model drug to
estimate the drug release behaviour of the CA–Fe3O4. In order to investigate the interaction of
the drug molecules with CA–Fe3O4, we have studied the fluorescence spectra of pure DOX and
DOX loaded CA–Fe3O4 in addition to zeta-potential measurements. The aqueous dispersion of
different amounts of CA–Fe3O4 (0, 20, 40, 60, 80 and 100 mg from a stock suspension of 2
mg/ml) was added to 1 ml of the DOX solution (10 µg/ml) and mixed thoroughly by shaking
at room temperature for 15 min. The fluorescence spectra of the supernatant (obtained after
magnetic sedimentation of the drug loaded CA–Fe3O4) were then recorded. The fluorescence
spectra of 1 ml of pure DOX (10 mg/ml) were also taken at different time intervals for
comparative studies. The fluorescence intensities of supernatants (washed drug molecules were
also taken into consideration for calculations) against those of the pure DOX solution were
used to determine the loading efficiency (binding isotherm of DOX with CA–Fe3O4). The
loading efficiency (w/w%) was calculated using the following relation:
31
% 𝐿𝑜𝑎𝑑𝑖𝑛𝑔 𝑒𝑓𝑓𝑖𝑐𝑖𝑒𝑛𝑐𝑦 = 𝐼𝐷𝑂𝑋− 𝐼𝑆− 𝐼𝑊
𝐼𝐷𝑂𝑋 × 100 (Eqn. 2.1)
where, IDOX is the fluorescence intensity of the pure DOX solution; IS, the fluorescence
intensity of supernatant; and IW, the fluorescence intensity of washed DOX (physically
adsorbed DOX molecules).
The drug release study was carried out under a reservoir-sink conditions. For the release
study, we have quantified the amount of DOX-loaded CA–Fe3O4 according to the binding
isotherm. The loading was carried out at an increased scale by incubating 2 ml of the aqueous
solution of DOX (2 mg/ml) with 1 ml of the aqueous suspension of CA–Fe3O4 (10 mg/ml) for
1 h in the dark (however, no decrease in fluorescence intensities was observed after 15 min of
incubation). The drug loaded CA–Fe3O4 (10 mg) was immersed in 5 ml of an acetate buffer
(pH 5), and then put into a dialysis bag. The dialysis was performed against 200 ml of
phosphate buffered saline (PBS) with a pH of 7.3, under continuous stirring at 37 °C to mimic
the environment of the cell. 1 ml of the external medium was withdrawn and replaced with
fresh PBS at fixed times in order to maintain the sink conditions. The amount of doxorubicin
released was determined by measuring the fluorescence intensity at λex = 490 nm and λem =
535±35 nm against the standard plot that was prepared under similar conditions. Each
experiment was performed in triplicates; the standard deviation is given in the plot.
2.2.3 Calorimetric Measurements
Time-dependent calorimetric measurements were done to evaluate the heating
capabilities of these CA–Fe3O4. Towards this end, 1 ml (10 mg/ml of Fe) of Fe3O4 colloidal
suspension was taken in a glass sample holder with suitable arrangements to minimise the heat
loss. An alternating current magnetic field (ACMF) of 7.64, 8.2 and 10.0 kA/m, and a fixed
frequency of 425 kHz were used to evaluate the specific absorption rate (SAR). The SAR was
calculated using the following equation3, 16, 27:
𝑆𝐴𝑅 = 𝐶 × ∆𝑇
∆𝑡 ×
1
𝑚𝐹𝑒 (Eqn. 2.2)
where, C is the specific heat of the solvent (C=Cwater=4.18 J/g °C); ΔT/Δt is the initial slope of
the time-dependent temperature curve; and mFe is the mass fraction of Fe in the sample.
32
2.2.4 In vitro Evaluation
The Sulforhodamine-B (SRB) assay was performed to evaluate the cytocompatibility
of the CA–Fe3O4 with HeLa cells. The cells were seeded into 96-well plates at densities of
1×104 cells per well for 24 h. Further, different concentrations of the CA–Fe3O4 colloidal
suspension (0, 31.250, 15.625, 7.813, 3.906, 1.953, 0.977, and 0.488 mg/ml of Fe) were added
to the cells and incubated for 24 h at 37 °C and 5% CO2. Thereafter, the cells were washed
thrice with PBS and processed for the SRB assay to determine the cell viability. For this, cells
were fixed with a solution of 10% trichloroacetic acid and stained with 0.4% SRB that was
dissolved in 1% acetic acid. The cell-bound dye was extracted with 10 mM unbuffered Tris
buffer solution (pH 10.5), after which the absorbance was measured at 560 nm by using a
multiwell plate reader. The cell viability was calculated using the following formula:
% 𝐶𝑒𝑙𝑙 𝑉𝑖𝑎𝑏𝑖𝑙𝑖𝑡𝑦 = 𝐴𝑏𝑠𝑜𝑟𝑏𝑎𝑛𝑐𝑒 𝑜𝑓 𝑡𝑟𝑒𝑎𝑡𝑒𝑑 𝑐𝑒𝑙𝑙𝑠
𝐴𝑏𝑠𝑜𝑟𝑏𝑎𝑛𝑐𝑒 𝑜𝑓 𝑐𝑜𝑛𝑡𝑟𝑜𝑙 𝑐𝑒𝑙𝑙𝑠 × 100 (Eqn. 2.3)
2.2.5 Characterisation Techniques
The X-ray diffraction (XRD) pattern was recorded on a Philips powder diffractometer
PW3040/60 with Cu Kα radiation. The surface modification of the nanoparticles was analysed
using a Fourier transform infrared spectrometer (FTIR) (Magna 550, Nicolet Instruments
Corporation, USA) in the range of 4000–400 cm-1. The particle size was determined by using
a high-resolution transmission electron microscope (HRTEM), JEOL JAM 2100F, which
operated at 200 kV. The thermal analyses were performed by a TA Instruments SDT Q600
analyser under N2 atmosphere from room temperature to 800 °C, with a heating rate of 10
°C/min. The elemental analysis was carried out by the FLASH EA 1112 series CHNS (O)
analyser (Thermo Fennigan, Italy). The hydrodynamic diameter and zeta-potential were
determined by dynamic light scattering (DLS) and Zeta PALS, respectively, (BI-200
Brookhaven Instruments Corp). The magnetic measurements of dried and powdered samples
were carried out using a vibrating sample magnetometer (VSM, LakeShore, Model-7410). The
Curie temperature was measured in an applied field of 100 Oe. In order to evaluate the SAR,
the amount of iron in the nanoparticle suspension was determined by UV (Cecil, Model No.
CE3021) by using the established phenanthroline spectrophotometric method28. The standard
curve was prepared from a stock iron solution under similar conditions (R2=0.998). The heat-
generating capability of the nanoparticles was evaluated by their SAR under the ACMF by a
radio frequency generator (EASY HEAT, EZLI5060) that operated at a fixed frequency of 247
33
kHz. All the fluorescence spectra were recorded on the Hitachi F 2500 fluorescence
spectrophotometer. The fluorescence intensities of aliquots for determination of the amounts
of the released drug were recorded using a Perkin Elmer 1420 multilabel counter.
2.3 Results and Discussions
2.3.1 Nanoparticle Characterisation
Figure 2.1 shows (a) XRD pattern and (b) TEM micrograph of CA–Fe3O4. The XRD
pattern reveals the formation of a single-phase Fe3O4 inverse spinel structure with a lattice
constant, a = ~8.378 Å, which is very close to the reported value of magnetite (JCPDS Card
No. 88-0315, a = 8.375 Å). The presence of sharp and intense peaks confirmed the formation
of highly crystalline nanoparticles. The crystallite size of CA–Fe3O4 is estimated to be
approximately 8 nm from X-ray line broadening using the Scherrer formula. From the TEM
micrograph, it is clearly observed that Fe3O4 nanoparticles are almost spherical and are around
8–10 nm (σ≤10%) in diameter. The selected area electron diffraction pattern (SAED) of this
sample (inset of Figure 2.1b) can be indexed to the reflections of the inverse spinel Fe3O4
structure and shows only the diffraction intensity that is associated with highly crystalline
Fe3O4, which is consistent with the XRD result.
Figure 2.1 (a) XRD pattern and (b) TEM micrograph of CA–Fe3O4 (Inset of (b) shows the
selected area electron diffraction pattern of CA–Fe3O4)
30 40 50 60 70
Inte
ns
ity
(a.u
.)
2 (Degree)
(22
0)
(31
1)
(40
0)
(42
2)
(51
1)
(44
0)
(a)
(b)
34
Figure 2.2 shows the FTIR spectra of pure CA and CA–Fe3O4. The absorption bands
for the pure CA are well resolved, but those of the CA–Fe3O4 are rather broad and few. The
1710 cm-1 peak assignable to the C=O vibration (asymmetric stretching) from the COOH group
of CA shifts to an intense band at about 1600 cm-1 for CA–Fe3O4, revealing the binding of a
CA radical to surface of Fe3O4 nanoparticles by chemisorption of carboxylate (citrate) ions 18,
29. Carboxylate groups of CA form complexes with Fe atoms on the surface of Fe3O4, imparting
a partial single bond character to the C=O bond. This results in the weakening of the C=O
bond, which shifts the stretching frequency to a lower value. Furthermore, the vibrational
modes that appear at 1400, 1250 and 1065 cm-1 in CA–Fe3O4 corresponds to the symmetric
stretching of COO–, symmetric stretching of C–O, and the OH group of CA30. The strong IR
band observed at around 575 cm-1 in CA–Fe3O4 can be ascribed to the Fe–O stretching
vibrational mode of Fe3O4.
Figure 2.2 FTIR spectra of pure CA and CA–Fe3O4
Figure 2.3A shows the TGA-DTA plots of CA–Fe3O4. A weight loss of ~7.5% with a
sharp endothermic peak at ~60 °C can be ascribed to the removal of physically absorbed water
and CA molecules on the Fe3O4 nanoparticles. The weight loss of about 13.5% with a broad
exothermic peak at ~260 °C can be associated with the removal of chemically attached CA
molecules from the surface of Fe3O4 nanoparticles. The weight loss of ~3.0% beyond 400 °C,
with a sharp exothermic peak at ~500 °C is associated with the phase transformation of Fe3O4–
Fe2O3. Furthermore, elemental analysis shows the presence of organic components (carbon and
35
hydrogen) on CA–Fe3O4. Thus, the FTIR, TGA and CHNS(O) results confirmed that Fe3O4
nanoparticles have been functionalised with citric acid during the course of the synthesis.
Figure 2.3 (A) TGA-DTA plots of CA–Fe3O4 (B) Zeta-potential of CA–Fe3O4 at different
pH values. Inset ‘a’ shows the hydrodynamic diameter of CA–Fe3O4 obtained from DLS
measurements; inset ‘b’ shows the possible schematic representation of CA–Fe3O4 (CA
coating on the surface of the Fe3O4)
Figure 2.3B shows the zeta-potential of CA–Fe3O4 at different pH values. From zeta-
potential measurements, it has been observed that adsorption of CA onto the surface of Fe3O4
nanoparticles results in a highly negative surface charge; also, the isoelectric point is not
observed in the measured pH range of 3–6 (isoelectric point of bare Fe3O4 nanoparticles is 6.7).
These highly negative values of the zeta-potential for the CA–Fe3O4 further confirmed the
presence of negatively charged carboxylate groups on the surface of the Fe3O4 nanoparticles.
Furthermore, DLS measurements (Inset ‘a’ of Figure 2.3B) indicate that these samples result
in an aqueous colloidal suspension with mean hydrodynamic diameters (almost constant with
invariable change in the polydispersity index) of about 25 nm (σ≤10%) due to the presence of
associated and hydrated organic layers3, 31. Specifically, some of the carboxylate groups of
citric acid strongly coordinate to iron cations on the Fe3O4 surface to form a robust coating (a
possible schematic representation of CA–Fe3O4 is shown in the inset ‘b’ of Figure 2.3B), while
uncoordinated carboxylate groups extend into the water medium, conferring a high degree of
water stability to the Fe3O4 nanoparticles. Further, the electrostatic repulsive forces originate
among the highly negatively charged CA–Fe3O4 (–25.5 mV at pH 6) in aqueous suspensions,
which also play an important role in their water stabilisation.
36
Figure 2.4 shows the field-dependent magnetisation (M vs. H) plot of CA–Fe3O4 at 300
K. The CA–Fe3O4 exhibits superparamagnetic behaviour without magnetic hysteresis and
remanence. The maximum magnetisation of CA–Fe3O4 was found to be 57 emu/g at 20 kOe.
The observed magnetisation is comparable to that of neat Fe3O4 nanoparticles (60.5 emu/g),
which is obtained by the co-precipitation method32 and is higher than the aqueous stable Fe3O4
nanoparticles (43.2 emu/g) that are obtained by the high-temperature thermal decomposition
method followed by subsequent surface functionalisation through ligand-exchange strategy3.
Thus, these aqueous stable Fe3O4 nanoparticles that have a high magnetic response can be
exploited for magnetic drug targeting, hyperthermia treatment and magnetic resonance
imaging. Further, the temperature dependence of magnetisation (M vs. T) measurement (inset
of Figure 2.4) shows that the Curie temperature (TC) of CA–Fe3O4 is around 580 °C, which is
in agreement with that reported for Fe3O43, 33, whereas, the TC of γ-Fe2O3 is around 645 °C34.
These results confirm that the phase formed is Fe3O4 rather than γ-Fe2O3. The slight hump
observed in the M vs. T curve at around 300 °C may be assigned to an increase in inter-particle
interaction near the surface due to the removal of organic moiety, (citric acid) as is suggested
by the TGA-DTA analysis.
Figure 2.4 Field dependence of magnetisation (M vs. H) plot of CA–Fe3O4 at 300 K. Inset
shows the temperature dependence of magnetisation (M vs. T) measurement
2.3.2 Drug Loading and Release
We have used the fluorescence spectra and zeta-potential analyses to investigate the
interactions of drug molecules with CA–Fe3O4, and their loading efficiency. At a low pH (4–
37
6), the protonated primary amine present on the drug induces a positive charge in the
doxorubicin molecule (cationic DOX)35. While at low pH, the carboxylic moiety of citric acid
(pK1=3.13, pK2=4.76 and pK3=6.40) is deprotonated and carries a negative charge36. The
surface charge of the CA–Fe3O4 at pH 6 (the pH at which the drug loading experiment was
carried out) was found to be negative (–25.5 mV) from the zeta-potential measurement. An
increase in the surface charge, from –25.5 to –10.5 mV, is observed in the CA–Fe3O4 after drug
loading (100 µg DOX reacted with 200 µg CA–Fe3O4 from their stock solutions of 2 mg/ml)
through the zeta-potential measurement. This result strongly suggested that positively charged
drug molecules are bound to negatively charged CA–Fe3O4 through electrostatic interactions.
The impressive affinity of doxorubicin for negatively charged molecules such as oleate ions
and phospholipids has been the subject of numerous investigations37-39. The interaction of DOX
molecules with CA–Fe3O4 was also evident from the predominant quenching of DOX
fluorescence in the presence of CA–Fe3O4 (Figure 2.5a), whereas, self-quenching (DOX–DOX
interaction due to π–π stacking) of pure DOX is not observed (Figure 2.5b). Furthermore, the
fluorescence intensity of DOX decreases (till saturation loading is achieved) on increasing the
concentration of CA–Fe3O4, which is obvious due to the increase in the loading efficiency of
DOX into CA–Fe3O4. The fluorescence spectroscopy (the fluorescence intensity is highly
dependent on the state of the molecule, that is, free or in the attached form) has been
successfully used to study interactions between DOX and its surrounding in previous literature,
for instance, when the drug intercalates DNA or penetrates within membrane models or
liposomal drug carriers40-43. The loading efficiency (binding isotherm of DOX with CA–Fe3O4)
obtained from quenching of fluorescence intensities is shown in the inset of Figure 2.5a. From
the inset of Figure 2.5a, it has been observed that loading efficiency is strongly dependent on
the ratio of the particles to DOX in the reaction solution and a maximum of around 90% (σ≤5%)
drug loading efficiency (w/w) could be achieved by electrostatic interactions. The obtained
loading efficiency is much higher than that reported (14%) by Munnier et al.12. They have
stated that the drug molecules (DOX–Fe2+ complex) were attached to the surface –OH groups
of the citric acid treated Fe3O4 nanoparticles through Fe2+ ions. The presence of citric acid
moieties on the surface of Fe3O4 nanoparticles may provide steric hindrance to the attachment
of the DOX–Fe2+ complex with the surface –OH groups, thereby reducing the loading of the
drug onto the nanoparticles. However, the drug loading is attributed primarily to the
electrostatic interactions between positively charged DOX molecules and negatively charged
carboxyl moieties that are present on the surface of Fe3O4 nanoparticles (as is suggested from
zeta-potential measurements), showing comparatively higher drug loading.
38
Figure 2.5 (a) Normalised fluorescence spectra of (a) 10 mg/ml of DOX (1 ml) reacted with
different amounts (0, 20, 40, 60, 80 and 100 mg) of CA–Fe3O4 for 15 min. Inset shows the
loading efficiency (binding isotherm of DOX with CA–Fe3O4) obtained from quenching of
fluorescence intensities; (b) 10 mg/ml of pure DOX (1 ml) at different time intervals; and (c)
Drug release profile of DOX-loaded CA–Fe3O4 in a cell mimicking environment (reservoir:
pH 5 and sink: pH 7.3 at 37 °C)
Figure 2.5c shows the drug release profile of DOX-loaded CA–Fe3O4 in a cell mimicking
environment (reservoir: pH 5 and sink: pH 7.3 at 37 °C). It has been observed that drug
molecules release slowly over a period of 50 h; the shape of the release profile suggests that
the complete release of drug was not attained. The drug loaded CA–Fe3O4 released around 60%
of the loaded drug in acetate buffer (pH 5) against the PBS (pH 7.3) after 50 h. The release of
DOX could be attributed to the weakening of the electrostatic interactions between the drug
and the partially neutralised carboxyl groups on the nanoparticle surface, which is due to an
increase in the protons in the colloidal solution. Munnier et al.12 discussed that release of DOX
molecules (released gradually over a period of 1 h and, thereby, attaining a plateau) from loaded
particles (achieved by chelation of the DOX–Fe2+ complex with –OH groups on the surface of
citric acid treated Fe3O4 nanoparticles) is primarily due to stimulated hydrolysis of drug
molecules. However, in this study, it was seen that weakening of the electrostatic interactions
is a slower process, which leads to the sustained release of drug molecules over a period of 50
h. These results indicate that the bound drug molecules will be released in appreciable amounts
in the mild acidic environments of the tumours.
2.3.3 Calorimetric Measurements
Time-dependent calorimetric measurements were performed on a suspension of CA–
Fe3O4 in order to investigate their heating efficacy (Figure 2.6a) as an additional functionality.
The SAR of CA–Fe3O4 was found to be 32.26, 38.63 and 49.24 W/g of Fe with an applied field
c
39
(H) of 7.64, 8.82 and 10.0 kA/m, respectively (at a fixed frequency of 425 kHz). It has been
observed that the time required to reach 43 °C (hyperthermia temperature) decreases with an
increase in the field (Figure 2.6a), which is obvious as the heat generation is proportional to
the square of the applied ACMF. These SAR values should not be viewed in terms of
performances, but only as the demonstration of the fact that these nanoparticles are effective
heating sources for hyperthermia treatment of cancer.
Figure 2.6 (a) Time-dependent calorimetric measurements of CA–Fe3O4; and (b) Viabilities
of HeLa cells incubated with medium that contains CA–Fe3O4
2.3.4 In vitro Evaluation
Figure 2.6b shows the viabilities of HeLa cells incubated with medium that contains
CA–Fe3O4. The SRB assay results indicate that the viability of the HeLa cells is not affected
by the mere presence of CA–Fe3O4, which results in the registering of normal growth in the
cells, suggesting that nanoparticles are reasonably biocompatible, and do not have a toxic
effect, and may be used under further in vivo settings. The percentage of cell viability is slightly
above 100% at lower concentration, which may be due to the presence of iron (Fe3O4
nanoparticles) that sometimes facilitates cell growth3, 31.
This study discussed the formation of aqueous-stable, highly crystalline, biocompatible
citric acid functionalised Fe3O4 nanoparticles that have optimal magnetisation, higher drug
loading efficiencies and a good SAR, which could find promising applications in drug delivery
and hyperthermia treatment of cancer.
40
2.4 Summary
A simple facile approach for the preparation of citrate stabilized Fe3O4 aqueous
colloidal nanoparticles of 8–10 nm by using a soft chemical approach is described in this
chapter. The detailed structural analyses by FTIR, TGA-DTA, CHNS(O) and zeta-potential
confirmed the functionalisation of Fe3O4 nanoparticles with citric acid. These nanoparticles
exhibited good colloidal stability, optimal magnetisation, in vitro biocompatibility and good
specific absorption rate (under an external ACMF). It was evident that the positively charged
drug molecules such as DOX could easily bound to the negatively charged CA–Fe3O4 through
electrostatic interactions. More specifically, a drug loading efficiency of about 90% (w/w) was
achieved by electrostatic interactions of drug molecules (DOX) with CA–Fe3O4. The drug
release profile in a cell mimicking environment indicated that the bound drug molecules will
be released in appreciable amounts in the mildly acidic environments of the tumours. Thus,
CA–Fe3O4 can be used as a potential carrier for effective magnetic drug targeting and
hyperthermia treatment of cancer.
However, the drug loaded on to CA–Fe3O4 is attached to the surface functional groups
and, thus, is very vulnerable to pH change in its immediate microenvironment. In order to avoid
this undesired release of DOX, various macromolecules could be further used to bind on the
nanoparticles surface. These macromolecules could effectively provide an accessible surface
for routine conjugation of biomolecules through the well-developed bioconjugation chemistry.
Therefore, dendrimers were used to conjugate to the surface groups of Fe3O4 nanoparticles in
the subsequent work. This also enhanced the aqueous stability and physico-chemical properties
of Fe3O4, and developed them into an improved platform as drug delivery vectors.
2.5 References
1. J. Cheon and J.-H. Lee, Accounts of Chemical Research, 2008, 41, 1630-1640.
2. K. C. Barick, M. Aslam, P. V. Prasad, V. P. Dravid and D. Bahadur, Journal of
Magnetism and Magnetic Materials, 2009, 321, 1529-1532.
3. K. C. Barick, M. Aslam, Y.-P. Lin, D. Bahadur, P. V. Prasad and V. P. Dravid, Journal
of Materials Chemistry, 2009, 19, 7023-7029.
4. H. Gu, K. Xu, C. Xu and B. Xu, Chemical Communications, 2006, DOI:
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41
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Parts of this chapter has been published in Langmuir, 2014, 30, 1004-1011.
Parts of this chapter has been “Just Accepted” in IEEE Trans. Magn. DOI: 10.1109/TMAG.2016.2517602.
Chapter 3
PEG-Modified PAMAM-Fe3O4-drug Triads with the
Potential for Improved Therapeutic Efficacy
3.1 Introduction
Chemotherapy uses chemicals to damage the internal machinery (that is, DNA, RNA
and proteins) of cells in order to trigger cell cycle arrest or apoptosis, however, such agents
generally induce apoptosis in both cancer and normal cells1. Doxorubicin (DOX), for example,
is an anthracycline molecule, which is a widely used cationic anticancer drug that interferes
with the replication process by intercalating DNA, resulting in cell apoptosis. DOX is
selectively destructive to rapidly proliferating cells, but is also toxic to cardiac tissue, which
might result in life-threatening heart damage2, 3. These side effects of DOX could be reduced
by the use of a delivery system that can safely and specifically deliver it to the targeted tissue.
Catechins, which belong to one of the sub-classes of polyphenols, are the major alkaloids found
in extracts of green tea (Camellia sinensis). Catechins comprise of epigallocatechin-3-gallate
(EGCG), epigallocatechin (EGC), epicatechin-3-gallate (ECG), epicatechin (EC) and
theaflavins4, 5. EGCG is an anionic polyphenol and has been explored as a chemopreventive
and chemotherapeutic drug for cancer6-8. It was found to inhibit the production of the vascular
endothelial growth factor (VEGF) in some types of cancer cells9-11 and to activate VEGF
receptors in leukemia12 and colon13 cancer cells. Catechins are highly unstable in an aqueous
45
solution and degrade through oxidative processes14. Their rapid degradation in blood and other
body fluids hampers their efficient delivery by traditional routes of administration, such as
orally, through enteric absorption, or by intravenous injection to the systemic circulation. The
absorption and bioavailability of catechins, across intestinal membranes, have been reported to
vary largely based on their molecular weight and lipophilicity15. Indeed, the aqueous instability
and uncontrollable absorption of EGCG limit their practical use as cancer therapeutics. The
challenges in the delivery of both DOX and EGCG, their chemical stability, their
bioavailability and bioactivity have necessitated a delivery system. This delivery system should
be able to deliver these drugs to the targeted tissue and selectively release them by responding
to local microenvironment conditions such as temperature and pH values.
In the last chapter, we evaluated CA–Fe3O4 nanoparticles for their potential as a
delivery agent for cationic DOX. The findings concluded that these nanoparticles have
appropriate properties and performances for use as drug delivery vehicles and can therefore,
be customised for better performances. The desired tailoring in the physico-chemical properties
and performances can be achieved by the surface modification of the Fe3O4 nanoparticles using
a variety of molecules. Without surface modification, bare Fe3O4 nanoparticles agglomerate
due to van der Waals forces and magnetic dipole-dipole interactions because of their high
surface energy and large surface area-to-volume ratio. When modified, the surface energy of
Fe3O4 is reduced and the nanoparticles form a stable, uniform dispersion in aqueous
suspensions16, which is a desirable performance for the administration of drugs. The two
primary strategies to engineer the surface of nanoparticles are (i) embedding the nanoparticles
in matrices17 or (ii) encapsulating them within macromolecules18, 19. Based on these
approaches, the development of surface-modified drug delivery systems has been dominated
by liposomal and polymeric systems20, 21. Both, however, have critical disadvantages.
Liposome capsules, although possessing a uniform size, generally have a poor physiological
stability in blood, which hampers efficient delivery22. The nanocarrier systems prepared from
polymers, on the other hand, have a tuneable stability but high water/pH degradability and
uncontrolled absorption after administration; these systems are also highly polydispersed and
have a large size distribution23. The drawbacks mentioned above, however, could be obviated
by the use of branched macromolecules (for example, dendrimer)24, 25. The dendrimers not only
reduce particle agglomeration, but also provide high number of functional groups. In fact, over
the past two decades, dendrimers have been intensively investigated as drug delivery systems
because of their unique structures and properties, monodispersity, water solubility,
multivalency and encapsulation ability to entrap hydrophobic drugs26. Among a number of
46
dendrimers, polyamidoamine (PAMAM) dendrimers are the most studied as drug delivery
vehicles27-29. The PAMAM dendrimer-drug system is also often loaded with magnetic
nanoparticles (for example, Fe3O4) as an agent that performs additional functions, such as
magnetic resonance imaging or hyperthermia treatment30, 31. In addition, dendrimers can be
modified with poly (ethylene glycol) (PEG) to decrease the toxicity of the dendrimers and to
improve their systemic circulation time, resulting in enhanced retention effects32, 33.
There are two strategies in the loading of a drug onto dendrimers: encapsulation and
conjugation. In the process of encapsulation, drug molecules are internalised by a carrier.
Encapsulation involves either hydrophobic interactions between the relatively non-polar
cavities of dendrimers and a hydrophobic drug, or hydrogen bonding between dendrimeric
cavities and drug molecules. Conjugation is a process of loading drug molecules onto the
surface of a carrier. It can be achieved either by covalent or by electrostatic interactions
between the dendrimer and drug. While EGCG is still in its preliminary stages of evaluation as
an anticancer drug, DOX has been successfully delivered using dendritic nanosystems.
PAMAM-based dendritic-linear block copolymers (that is, dendrons) have been investigated
to encapsulate both DOX and Fe3O4 particles34, 35. These works showed that while the PAMAM
dendrimers showed improved encapsulation efficiency, their drug release performance was
unsatisfactory. Their findings demonstrated that the releasing efficiency of dendritic systems
was acceptably high at a temperature that was much lower (25 °C) than 37 °C, which would
make clinical administration difficult34. Also, while the system could also be pH-responsive,
the high releasing efficiency (> 50 %) at normal physiological conditions, that is, pH 7.434, 35,
predicted that much of the loaded DOX, if administrated systemically in vivo, would be
released immediately and uncontrollably into the bloodstream, resulting in a severe side effect.
The release of conjugated drugs from a dendrimer system has been investigated under
various external stimuli such as temperature and, especially, pH36-41 based on the fact that the
microenvironments around a tumour and inside cells are typically acidic42. In general, the
release efficiency of covalently conjugated drugs is lower than that of electrostatically bound
drugs. For example, when DOX molecules were covalently bound to PEG-modified PAMAM
by thiolated transferrin molecules by Michael addition, the release of DOX was only 30% at
pH 4.543. Recent work on electrostatic conjugation of DOX with various generations (for
example, G3, G5 and G6) of dendrimers demonstrated that the release efficiency of DOX was
40% in a PBS buffer at pH 5.218, 44, but was increased to 80 % when pH was reduced to 4.218.
Note that the acidic pH around 5 is found neither extracellularly (normally) nor within most of
the cell cytoplasm; this is, however, the pH of the lysosome, which is where many internalised
47
molecules end up and are broken down by the acid hydrolases and other acid-active enzymes.
Further, the PEGylation of PAMAM dendrimers increases the drug loading by enhancing
hydration around the dendrimer periphery. It also induces the release of the drug in a controlled
fashion over a prolonged time. Cationic PAMAM interacts with negatively charged plasma
membranes, thereby promoting cell death by interfering with membrane fluidity. In contrast,
anionic PEG-PAMAM maintains the integrity of the plasma membrane, thus, reducing the
cytotoxicity 45, 46.
The primary objective of this chapter, therefore, was to establish optimal fabrication
procedures to develop multifunctional nanocarriers that are capable of carrying and delivering
both cationic and anionic drugs. To this end, different generations (G3, G5 and G6) of PEG-
PAMAM were used to modify the surface of Fe3O4, and DOX and EGCG were electrostatically
conjugated on the surface of the PAMAM-Fe3O4 nanocarriers. The PAMAM-Fe3O4-drug triads
were thoroughly characterised and explored in vitro as drug delivery systems for cancer
chemotherapeutics.
3.2 Experimental Techniques
3.2.1 Synthesis and surface modification of Fe3O4 nanoparticles
Iron oxide (Fe3O4) nanoparticles were synthesised using a soft chemical route. In this
process, 4.44 g of FeCl3 and 1.732 g of FeCl2 were dissolved in 80 ml of water in a round-
bottomed flask, and the temperature was slowly increased to 60 °C with refluxing under a
nitrogen atmosphere and mechanical stirring at 1000 rpm. After 30 min, 25 ml of ammonia
solution was added to the reaction mixture for precipitation of magnetite. An aqueous solution
of 5 ml of glutamic acid (0.5 g/ml) was then added to the reaction mixture mentioned above,
and the temperature was raised to 95 °C under reflux and maintained for a further 90 min with
continuous stirring. A black precipitate of glutamic acid-coated iron oxide nanoparticles (Glu–
Fe3O4) was obtained and then was thoroughly washed with ultrapure water. During each
washing step, the supernatant was decanted using a permanent magnet.
The Glu–Fe3O4 were then modified with PEG-PAMAM dendrimers to anchor the drug
molecules. In this process, 500 µl of Glu-Fe3O4 (at 1 mg/ml) was mixed with 200 µl of PEG-
PAMAM (0.1 mg/ml; generation 3, 5, or 6, marked as GX, where x=3, 5, or 6), and the volume
was made up to 1 ml by ultrapure water. The mixture was incubated at room temperature
overnight under continuous shaking. The dendrimer modified-Fe3O4 nanoparticles
48
(abbreviated as Fe3O4–DGX, where x=3, 5, or 6) were then collected by a permanent magnet
and rinsed with ultrapure water 3–4 times.
3.2.2 Analysis of Drug Loading and Release
In order to evaluate the interactions of the drug molecules with the Fe3O4–DGX
nanoparticles, absorption and fluorescence spectroscopy was followed for EGCG and DOX,
respectively. The host-guest molecular complexation affects the characteristic electronic
absorption spectra of the reaction system (Benesi–Hildebrand method) and, therefore, was
employed to study the binding of EGCG to the Fe3O4–DGX nanoparticles. The fluorescence
intensity is highly dependent on the free or attached state of the molecule and, thus, has been
extensively used to study interactions between fluorophores and quencher47-50. The aqueous
suspension of varying amounts of Fe3O4–DGX nanoparticles (200, 400, 600, 800, 1000 and
1200 µg from a stock solution of 1 mg/ml) were added to 1 ml of each of the drug solutions
(10 µg/ml). The mixture was incubated at room temperature for 15 min. The absorbance and
fluorescence spectra of the supernatant after magnetic sedimentation were recorded for EGCG
and DOX, respectively. The intensities of the supernatant against the pure drug solution were
used to determine the loading efficiency of the Fe3O4–DGX nanoparticles. The loading
efficiency (w/w %) was calculated using the following relation;
% 𝐿𝑜𝑎𝑑𝑖𝑛𝑔 𝐸𝑓𝑓𝑖𝑐𝑖𝑒𝑛𝑐𝑦 = (𝐼𝐷𝑟𝑢𝑔− 𝐼𝑆)
𝐼𝐷𝑟𝑢𝑔 × 100 (Eqn. 3.1)
where IDrug is the intensity of the pure drug, and Is is the intensity of the supernatant.
For drug release experiments, the amounts of the drug and the Fe3O4–DGX
nanoparticles were quantified on the basis of loading efficiency. Along the lines of typical drug
loading experiment, 500 µg of Fe3O4–DGX (x = 3, 5 & 6) nanoparticles were incubated with
the 500 µg of the drug (1 mg/ml) in the dark for 3 h. The drug-loaded nanoparticles were
collected with a permanent magnet and were utilised for drug release experiments. For a typical
drug release study, drug-loaded Fe3O4–DGX nanoparticles were immersed in 5 ml of a sodium
acetate buffer (pH 5) and put into a dialysis bag, which was then suspended in 200 ml PBS (pH
7.3) to form the reservoir-sink settings. The dialysis was performed under continuous stirring
at 37 °C. At different time intervals, 1 ml aliquots from the sink (PBS) were withdrawn and
replaced with equal amounts of fresh PBS in order to maintain the interfacial concentration
gradient across the semi-permeable membrane. The amount of the drug released was
determined by the measurement of absorbance (λmax = 274 nm) and fluorescence (λex = 490 nm
49
and λem = 560 nm) of the aliquots respectively, against standard plots that were prepared under
similar conditions (EGCG: R2= 0.996; DOX: R2= 0.999).
3.2.3 Evaluation of Carriers’ Biocompatibility and Therapeutic Efficacy
The biocompatibility of the Fe3O4–DGX (x = 3, 5 & 6) was assessed with the human
cervical cancer cell line (HeLa). In order to establish the potential of drug-loaded Fe3O4–DGX
in releasing the respective drugs, a dose-dependent study was undertaken to evaluate its 50%
inhibitory concentration (IC50) values over 24 h. The HeLa cells were grown in a supplemented
essential medium at 37 °C at 5% CO2 in a humidified atmosphere. Once they reached 90%
confluency, they were treated with a trypsin-EDTA solution in order to initiate detachment
from the substrate surface. The obtained cell suspension was centrifuged and the supernatant
was discarded. The cells were then re-suspended in a fresh supplemented medium and were
counted by the trypan blue exclusion method51. The cells were then re-seeded in 96-well plates
at a cell density of 2×104 cells per well and grown for a further 24 h. After 24 h, the spent
growth medium was discarded and replaced by a fresh supplemented medium that contained
different concentrations of the Fe3O4–DGX and drug-loaded Fe3O4–DGX (individual, serially
diluted test samples); this was followed by an additional 24 h of incubation. The cells were
then carefully washed with sterile PBS and the viable cell population was determined by the
sulforhodamine B (SRB) colorimetric assay52 as described in section 2.2.4. The IC50 value was
determined from dose-responsive sigmoidal curves that were generated from the spectroscopic
data obtained from Origin 8.0 software.
3.3 Results and Discussions
3.3.1 Characterisation of Fe3O4 Nanoparticles
The characterisation of the prepared nanomaterials was undertaken as described in
section 2.2.5. The XRD pattern revealed the formation of single-phase Fe3O4 nanoparticles
with an inverse spinel structure that had a lattice constant, a = ~8.38 Å, which was consistent
with the reported value of magnetite (JCPDS Card No. 88-0315, a = 8.375 Å). A high degree
of crystallinity in nanoparticles was indicated by the presence of sharp and intense peaks
(Figure 3.1a). The Scherrer formula was used to calculate the crystallite size of Glu–Fe3O4
(diffraction line broadening), which was ~8 nm53.
50
Figure 3.1 (a) XRD pattern of Glu–Fe3O4 nanoparticles; TEM micrographs showing (b) Glu–
Fe3O4 nanoparticles (c) particle size distribution (d) selected area electron diffraction (SAED)
pattern, and (e) high-resolution TEM image of Glu–Fe3O4 nanoparticles
The microstructure examination in TEM showed that the particles were spherical with
regular morphology (Figure 3.1b), and their size ranged between 10 and 15 nm (Figure 3.1c).
The SAED rings in Figure 3.1d were indexed to be (220), (311), (400), (511) and (440)
diffraction planes of the inverse spinel Fe3O4, which was consistent with the XRD pattern.
From the high-resolution micrograph, the lattice spacing of the crystallite (Figure 3.1e) was
measured to be 2.5 Å, which corresponds to the (311) plane of Fe3O4.
The FTIR spectra of the synthesised nanoparticles showed successful grafting of
glutamic acid onto the surface of Fe3O4 nanoparticles. Figure 3.2a shows that the absorption
bands for the pure glutamic acid are well resolved when compared to the Glu–Fe3O4, which
are broadened and few. The strong IR band observed around 580 cm-1 in Glu–Fe3O4 could be
ascribed to the Fe–O stretching vibration mode of the magnetite. The spectral band at ~1650
cm-1 (C=O vibration of –COOH groups for glutamic acid) shifted to 1620 cm-1, indicating the
binding of glutamic acid to the surface of Fe3O4 nanoparticles by chemical absorption of
carboxylate (glutamate) ions. The carboxylate groups of the glutamic acid forms complexes
with Fe atoms that are present on the surface of Fe3O4, rendering a partial single bond character
to the C=O bond54. This is expected to weaken the C=O bond, which shifts the stretching
frequency to a lower value. The frequencies that appear at the 1540, 1420 and 1260 cm-1 in
Glu–Fe3O4 correspond to the symmetric stretching of COO–, C-O and OH groups, respectively,
confirming the successful grafting of glutamic acid on the surface of Fe3O4 nanoparticles. The
FTIR spectra of the pure PEG-PAMAM (Figure 3.2b) and PEG-PAMAM-modified Fe3O4
(Figure 3.2c) nanoparticles show well-resolved characteristic vibrations of PAMAM
51
dendrimers, which broadened after conjugation with the Glu–Fe3O4 nanoparticles. Fe3O4–DGX
samples revealed a broad N–H stretching vibrations at 3400 cm-1, C–H stretching vibrations at
2880 cm-1 (2900 cm-1), COO- asymmetric stretching at 1625 cm-1, C=O stretching at 1550 cm-
1, N-O stretching at 1325 cm-1 (1350 cm-1), C–O–C stretching vibrations at 1250 cm-1, and
interfering C–O stretching with C–N stretches at 1100 cm-1. The values in parenthesis represent
the vibrational peaks of pure PEG-PAMAM. Due to the intense fingerprint region of PEG-
PAMAM dendrimers, the metal oxygen vibrations of iron oxide have been masked.
Figure 3.2 FTIR spectra of (a) glutamic acid (Glu) and glutamic acid-coated Fe3O4 (Glu–
Fe3O4) nanoparticles, (b) PEG-PAMAM of generations 3, 5, 6, and (c) Fe3O4–DGX
nanoparticles
Thermo-gravimetric analysis (TGA) and differential thermal analysis (DTA)
degradation profiles (Figure 3.3) show a weight loss of 9.6% with a sharp DTA endothermic
peak at ~80 °C, which could be attributed to the removal of physically absorbed water
molecules on the Fe3O4 nanoparticles. The weight loss of 5.0% with a broad DTA exothermic
peak at ~290 °C could be associated with the removal of chemically attached glutamic acid
molecules from the surface of Fe3O4 nanoparticles during the degradation of Glu–Fe3O4
nanoparticles. The weight loss of ~2.8% beyond the temperature of ~400 °C, indicated by a
sharp DTA exothermic peak at ~565 °C, is likely associated with the phase transformation of
magnetite to maghemite (Fe2O3)55.
52
Figure 3.3 TGA-DTA of glutamic acid-coated Fe3O4 (Glu–Fe3O4) nanoparticles
In order to investigate whether the magnetic performance of Fe3O4 nanoparticles is
compromised by coatings, the field-dependent magnetisation (M vs. H) of Glu–Fe3O4 and
Fe3O4–DGX nanoparticles was measured at 300 K at a field of 20 kOe (Figure 3.4a). The Glu–
Fe3O4 nanoparticles exhibited superparamagnetic behaviour that was characterised by zero
coercivity and remenance, and had a maximum magnetisation of 57 emu/g at 20 kOe. The
magnetisations of the Fe3O4–DGX nanoparticles were calculated to be approximately 48, 37
and 33 emu/g for generations 3, 5 and 6, respectively. The presence of PEG-PAMAM on the
surface of these nanoparticles resulted in a decrease in the saturation magnetisation value due
to the increase in the non-magnetic components. This decrease is seen to be in proportion to
the generation of the PEG-PAMAM.
Figure 3.4 (a) Room-temperature field-dependent magnetisation of Glu–Fe3O4 and Fe3O4
nanoparticles coated with dendrimers of generations 3, 5, and 6 (Fe3O4–DGX) at 20 kOe (b)
Time-dependent calorimetric measurements of Glu–Fe3O4 in different ACMFs. The SAR was
calculated to be ∼134 W/g from the initial slope of the time-temperature curve.
53
The ability of the superparamagnetic nanoparticles to generate localised heat in the
presence of ACMF makes them potentially useful in killing tumour cells. To investigate their
potential in the hyperthermic treatment of cancer, time-dependent calorimetric measurements
were carried out with Glu–Fe3O4 nanoparticle aqueous suspensions (typically ~45°C)56. SAR
was calculated from the initial slope of the temperature-time curve (Figure 3.4b), and was
found to be 134.3 W/g. The SAR value could be optimised by varying parameters such as
frequency, applied field and physical properties of the nanoparticles57, 58. The field strength had
a considerable effect on the heating capability. Figure 3.4b shows that at an applied field of
309Oe (400 A), the temperature of the aqueous suspension with high concentrations of iron
oxide nanoparticles (5 mg/ml) reaches 45 °C within 2.3 min. When the applied field was
reduced to 270 Oe (350A) and 232 Oe (300 A), the time taken to reach 45 °C increased to 3.3
and 4.5 min, respectively.
The colloidal stability and dispersion of the nanoparticles are associated with the
electric charge of the particle surface. The zeta potential of the different Glu–Fe3O4 and Fe3O4–
DGX (x = 3, 5 and 6) nanoparticle systems were measured and found to be −26.4 5.6, −19.0
2.0, −20.4 4.9 and −23.9 6.4, respectively. Apparently, the attachment of glutamic acid
on the surface of Fe3O4 nanoparticles resulted in a negatively charged surface (−26.4 mV).
Glutamic acid is a bi-carboxylic amino acid, that is, it has two –COOH groups per molecule.
Some of these carboxylate groups bind with the Fe cations in situ, while those that remain
unbound extend freely into the surrounding aqueous medium. These free carboxylate groups
not only impart an anionic surface to the nanoparticles but also play an important role in the
enhanced aqueous stability of the suspension of the nanoparticles. Zeta potential analysis also
showed that all the Fe3O4–DGX nanoparticles exhibited positive potential until pH 4; however,
their surface charge became negative with increasing pH. Around the physiologically relevant
pH of 7.4, the surface charges of the nanoparticles were measured and found to be −19.0, −20.4
or −23.9 mV for G3, G5 and G6, respectively. This change in the surface charge of the three
generations was due to the abundance of PEG-moiety on the dendrimers. The large negative
surface charge on the surface is believed to have enhanced the colloidal stability of the
nanoparticles.
3.3.2 DOX Loading and Release
Fluorescence spectroscopy was used to investigate the interactions of DOX with
Fe3O4–DGX nanoparticles (Figure 3.5). The interactions of DOX with the nanoparticles were
54
evident by the dropping of fluorescence in the presence of the nanoparticles. The loading
interactions were evaluated at λex = 490 nm and λem = 590 nm of the DOX. The drug loading
efficiencies of the Fe3O4–DGX nanoparticle systems were calculated to be 86.9, 93.2 and 96.1
%, respectively, for generation 3, 5 and 6, using eqn. 3.1 (Figure 3.5d). It was found that the
average loading efficiency of DOX onto Fe3O4–DGX increased with the increase in the
generation of PAMAM, and was up to ~96 % for Fe3O4-DG6 carriers, which differs from
previous work18, 59. However, we also noticed that the increments were only statistically
significant between DG3 and DG6.
Figure 3.5 (a,b,c) Fluorescence intensity of DOX-loaded Fe3O4–DGX nanoparticles. (d) Drug
loading efficiency versus dendrimer generation. The difference between DG3 and DG6 was
significant (p < 0.05). There were no significant differences between DG3 and DG5, and DG5
and DG6 (p > 0.05)
It has been reported that dendrimers of higher generations are more capable of
encapsulating drugs than lower generation ones, while lower generations facilitate the
conjugation of drugs18, 59. In the previous works, the conjugation capacities of higher
generations of dendrimers were impaired by their more compact packing and steric hindrances
that were posed by the surface amine groups18. The amine groups present on the surface of the
PAMAM dendrimers play an important role in interacting with the drug molecules. It was
55
observed that the electrostatic interactions between the −NH2 (δ+) groups of DOX with the –
OH (δ−) groups of the PEG-PAMAM are responsible for the high drug loading capacities. The
number of available PEG chains increase with an increase in the generation of the dendrimer.
Thus, the enhanced conjugation capacity of higher generations of dendrimers could be
attributed to the PEGylation on the surface of the PAMAM dendrimer The benefit from the
loosely packed long chains of PEG is two-fold. First they offer more space to host drug
molecules; second, they reduce steric hindrance when compared to the bare PAMAM
dendrimers.
Absorbance spectroscopy was utilised to evaluate the interactions of EGCG molecules
with Fe3O4–DGX nanoparticles and to calculate their loading efficiencies. Figure 3.6 shows the
drug loading efficiency obtained from the spectral changes in the EGCG with increasing
amounts of Fe3O4–DGX nanoparticles, calculated against absorbance of the free EGCG
solution. The drug molecules are expected to interact with the dendrimers electrostatically,
resulting in a drug-nanoparticle conjugate60. The loading interactions were evaluated at the
absorption maxima of EGCG, λmax = 274 nm. As the amount of Fe3O4–DGX was increased in
the solution, a slow decrease in the absorbance peak of the pure EGCG was observed, thereby
suggesting drug-nanoparticle interaction. A saturation point was reached when no further
decrease in absorbance was observed even after the addition of more Fe3O4–DGX
nanoparticles. The absorbance value of the supernatant of the particles was used to calculate
the drug loading capacities.
Figure 3.6 (a) Loading efficiencies of EGCG on to Fe3O4–DGX nanoparticles (b) Loading
efficiencies of three groups of Fe3O4–DGX (x = 3, 5 and 6) nanoparticles against dendrimer
generation. The difference between DG3 and DG6 was significant (p < 0.05).
56
The loading of EGCG on the Fe3O4–DGX nanoparticles could be explained by the
formation of hydrogen bonds between the –OH groups of EGCG and PEG-PAMAM. At the
same time, a repulsive force exists between the –OH groups of both the molecules, which limits
the drug loading capacity to a certain extent. Statistical analysis revealed that the difference
between the loading capacities of Fe3O4–DG3 and Fe3O4–DG6 was significant (p<0.05), while
no significant difference was seen between the loading capacities of Fe3O4-DG3 and Fe3O4-
DG5, and Fe3O4-DG5 and Fe3O4-DG6 (p>0.05).
Three main types of interactions are expected between the dendrimer-nanoparticle
drug, namely, electrostatic interactions between the dendrimer and the charged functional
groups present on the drug; hydrophobic interactions between relative non-polar cavities of
dendrimers and hydrophobic end groups of the drug; and hydrogen-bond interactions between
dendritic cavities and drug molecules. Among these interaction mechanisms, the electrostatic
interactions contribute more to the solubility enhancement of the drugs than do the hydrophobic
and hydrogen-bond interactions. The loading of DOX on Fe3O4–DGX is higher than EGCG,
which could be explained on the basis of attractive interactions between the functional groups
of the drug molecules and PEG-PAMAM. Electrostatic interactions develop between the −NH2
(δ+) groups of DOX and the –OH (δ−) groups of the PEG-PAMAM, while the –OH groups of
EGCG interact through the formation of hydrogen bonds with PEG-PAMAM.
The drug release experiments were carried out in reservoir-sink conditions at 37 °C,
which mimics the cellular environment (reservoir: pH 5 and sink: pH 7.3). Figure 3.7a shows
that the release of DOX from Fe3O4–DGX was collectively faster at pH 5 than at pH 7.3. The
electrostatic conjugation is through the interaction of the cationic ends of DOX with the anionic
ends on the surface of PEG-PAMAM. The lowering in the pH value leads to protonation of the
PEG-modified dendrimers, which results in the cleavage of the bond of PEG with the DOX
molecules. It also shows that at pH 5, DOX was released rapidly for the initial 10 h before a
plateau was reached. The plateau percentages of DOX release observed over a period of 24h
were 608, 689 and 809 % by the Fe3O4−DG3, Fe3O4−DG5 and Fe3O4−DG6, respectively.
The release efficacy of DOX from Fe3O4–DGX increased with the generation of PAMAM,
which again contradicted previous work18, 59.
57
Figure 3.7 Drug-release profile of (a) DOX-loaded Fe3O4–DGX nanoparticles at pH 5.0 and
7.4 and (b) EGCG-loaded Fe3O4–DGX nanoparticles at pH 7.4
In general, EGCG molecules were released slowly in the acidic environment, although,
in appreciable amounts (Figure 3.7b). It can be seen that up to 40–60% of EGCG molecules
were released over the first 24 h. The release of EGCG was slow and the release profile was
seen to follow a linear trend, which did not attain saturation/plateau even after 24 h. The release
mechanism of EGCG in an acidic environment involves the protonation of the dendrimers,
which results in weakening of the interactions between the EGCG molecules and the PEG-
PAMAM dendrimers. Hence, the slow releasing profile indicates that the cleavage of the
binding interactions is a slow process. The release of both of these drugs is stimulated by the
high [H+] concentration in the surrounding medium. The ease of protonation of DOX molecules
enhances their release in comparison to the EGCG molecules, which are not protonated but
released only when the dendrimeric chains are protonated. The PEGylation of PAMAM chains
also plays a role in the slow release of EGCG molecules as against the higher DOX release.
3.3.3 In vitro Evaluation of Cellular Toxicity and the Therapeutic Effect
The Fe3O4–DGX nanoparticles with or without the drugs were incubated with HeLa
cells in order to evaluate their compatibility or toxicity to the cell proliferation activity.
Quantitative evaluation showed that at physiological pH (7.2-7.4), the drug-free Fe3O4–DGX
nanoparticles had no significant effect on the proliferation of HeLa cells (Figure 3.8). Also, no
change in cellular morphology was observed, suggesting that the nanoparticles alone were
biocompatible. Although the lower generations were biocompatible, the higher generations of
PEG-PAMAM revealed moderate levels of cytotoxicity. The degradation of PAMAM in a
cellular environment could lead to the generation of toxic products, namely, methyl
58
methacrylate and ethylene diamine, leading to elevated toxicity levels, depending on the
generation of the dendrimer.
Figure 3.8 Percentage relative cell viability versus the concentration of Fe3O4–DGX
nanoparticles. A decreasing trend of biocompatibility of these nanoparticles is seen with
increase in dendrimer generation and concentration.
The internalisation of the nanoparticles was examined using DOX-loaded Fe3O4–DG5
nanoparticles under confocal laser scanning microscopy. HeLa cells (2×103) were incubated
with DOX-loaded Fe3O4–DG5 for 5 h. Due to the inherent fluorescence of DOX, no tagging
moiety was required. The confocal images (Figure 3.9a) demonstrated that the control cells
showed no discernible changes in cytoplasm, nucleus, nuclear membrane and nucleoli. Figure
3.9b demonstrates that the particles were internalised by the cells. Though the incubation time
was not long enough for the drug to kill the cell (Figure 3.9b), the DOX-loaded nanoparticles
had caused significant changes in cellular morphology and related features, namely,
granulation of cytoplasm and degradation of the nuclear membrane.
59
Figure 3.9 Confocal laser scanning images of (a) control HeLa cells and (b) treated HeLa
cells after 5 h of incubation (scale−20 μm) (c) Viability of HeLa cells in a medium that
contained either Fe3O4–DG5 or DOX–Fe3O4–DG5 nanoparticles in various amounts. The
difference between any pair of data at each concentration was insignificant (p > 0.05).
Under physiological conditions, PAMAM–Fe3O4–drug nanoparticles would cause a very mild
side effect to normal tissues due to low levels of drug release. Hence, the drug molecules can
be carried in a reasonably benign manner and not released until the triads are taken up into the
cell. The inhibition of cell proliferation activity is observed with both the drug molecules.
EGCG is seen to have less toxicity when compared to DOX molecules. At high doses of
PAMAM–Fe3O4–drug nanoparticles, EGCG reduced the cell population by 40%, while DOX
reduces the population by 80%, reiterating the effectiveness of DOX over EGCG. The IC50 was
calculated to be 11.8±5.3 µg/ml (R2 = 0.966) and 56.3±2.7 µg/ml (R2 = 0.982) of DOX-loaded
and EGCG-loaded Fe3O4-DGX nanoparticles, respectively.
3.4 Summary
The last chapter concluded that Fe3O4 nanoparticles were good drug delivery vectors,
but their properties and performances can be improved further by using macromolecules. In
this chapter, we have presented the dendrimer–Fe3O4 system as an improved platform for the
delivery of both cationic and anionic molecules. The chapter established the fabrication
procedures of the PAMAM-Fe3O4-drug nanosystem and the therapeutic efficiency of pH-
responsive triads by using different generations (G3, G5 and G6) of PEG-PAMAM. The
nanocarriers were synthesised by using the PEG-modified PAMAM dendrimers to encapsulate
60
glutamic acid-modified Fe3O4 nanoparticles, and the drug molecules were electrostatically
conjugated to the surface of the nanocarriers. It was observed that the PAMAM of higher
generations possesses more favourable qualities in terms of both higher conjugation capacity
and greater efficiency in the release of the drugs. These nanocarriers are evidently capable of
carrying both cationic and anionic therapeutic molecules. At pH 7.4, the drug release was as
low as 15 % with sustained (DOX) and linear (EGCG) drug release profiles that were obtained
in acidic environment. The nanocarriers alone (that is, drug-free) are biocompatible and have
little or no adverse effect on cellular proliferation and morphology when incubated with HeLa
cells. The PAMAM-Fe3O4-drug triad showed controlled anticancer activity with a sigmoidal
(DOX) and approximately linear (EGCG) dose-response profile. Moreover, the
superparamagnetic behaviour and calorimetric measurements of nanocarriers suggest their
probable use in the hyperthermic treatment of cancer. These dendrimer-Fe3O4 nanosystems
showed excellent physico-chemical properties, and drug carrying and delivery performances,
but raised demands of improved biocompatibility, when higher generations of dendrimers were
in question. Towards this end, we aimed to synthesise a peptide dendrimer with an internal
chemical environment and physico-chemical properties similar to PAMAM. This peptide
dendrimer is expected to match PAMAM in its drug delivery performances with enhanced
biocompatibility.
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65
Chapter 4
Section A
4.1 Synthesis of Peptide Dendrimer and its in vitro Characteristics
4.1.1 Introduction
The physico-chemical properties of dendrimers are primarily governed by their
repetitive units of monomers (internal chemical environment) and the functional groups that
are present on their surfaces. To achieve a high degree of precision and structural order,
dendrimers are synthesised in a stepwise fashion. Generally, two different methods, namely,
divergent and convergent, are adopted for the synthesis of dendrimers. These two synthetic
methods have inherent advantages and disadvantages. Using the divergent synthesis method,
the dendritic molecule is formed from a central core, which then extends radially outwards
through addition of branching molecules. The main advantage of the divergent method is that
highly molecular nanoscaffold architecture is attained with desired repetitive branching
monomers. Thus, the dendrimer can be tailored to achieve maximum functionalities and
properties. However, two major challenges are encountered in divergent synthesis. First, the
number of reaction points increases in geometric progression with every generation, followed
by an increase in molecular weight. This compromises the reaction kinetics, making it slower;
the synthesis of high generation dendrimers then becomes difficult, further lowering the yield
of the desired product. The addition of each branching unit requires care and precision in order
to prevent structural defects and asymmetry in the structure of the dendrimer. Secondly, the
separation of the desired dendrimer from the by-products is hindered due to the molecular
66
similarity exhibited by the by-product as well as the desired dendrimer. On the other hand, the
convergent method employs synthesis of small dendrites from the exterior, and the reaction
then proceeds inwards to the central core. The convergent procedure results in lesser structural
defects and easy purification of dendrimers, resulting in a high degree of monodispersity.
Despite the possibility of purer and flawless dendrimers, the convergent method falls short in
the synthesis of higher generation dendrimers. This choice is limited due to the steric forces
that crowd the dendrites around the central core molecule.
Despite the difficulties, these macromolecules have gained interest over classical
polymers due to the varied options presented by dendritic macromolecules. The vast pool of
molecules offers freedom of choice for the central core, the branching monomeric units and the
surface functional groups and gives rise to a multivalent system. This, in turn, has a wide
variety of possible chemical compositions, internal chemical environments, tailorable surface
groups, structural properties and architecture. Multi-functionalised dendrimers have also been
successfully synthesised, mainly, by modifying the terminal functional groups, which offers
adjustments in physical and chemical properties as required1, 2. Methyl acrylate alternating with
ethylene diamine forms the most widely synthesised, studied and used class of polyamidoamine
(PAMAM) dendrimers3; the internal amide groups provide an abundance of lone pairs of
electrons. Another popular class of amine-terminated dendrimers is the poly-(propylene imine)
(PPI) synthesised by Michael's addition of primary amines to acrylonitrile, followed by
subsequent hydrogenation by Raney cobalt or Raney nickel catalyst4. The interiors of PPI
dendrimers are the tertiary nitrogen atoms with lone pairs of electrons that contribute to their
reactive cavities. Both the classes of dendrimers have primary amine groups on the surfaces,
which govern the surface properties, reactivity and surface charge. These polyamidoamine
(PAMAM), polyethylene imine (PEI), polypropylene imine (PPI) dendrimers have found their
way into biomedical applications such as drug delivery 5-7, gene delivery8-10, antimicrobials11,
bioseparation12, biosensing13-15 and magnetic resonance imaging16-19. The world of biomedical
applications of dendrimers is currently dominated by the PAMAM dendrimers.
PAMAM surpasses other dendritic systems due to the ease of its preparation, desirable
chemical and physical properties, surface functional groups, and comparatively lower toxicity
to other dendrimers. Despite the lower cytotoxicity in comparison with other dendrimers and
dendritic systems, the degradation products of PAMAM show significant toxicity, which limits
their use in biological systems 20, 21. However, when various factors (such as biocompatibility,
haemocompatibility, immunogenicity, and biodistribution) are taken into consideration,
67
cytotoxicity remains the most critical factor in the successful use of any dendritic system in
any living system. The cytotoxicity of these dendrimers is not only attributed to the charges of
the surface functional groups but also to the number of generations, and has been well reported
20-23. Thus, the primary aim of this chapter was to fabricate dendrimers with enhanced
biocompatibility without compromising their chemical composition, internal and external
environment, physical and chemical properties, and efficacy in biomedical applications.
4.1.2 Experimental and Characterisation Techniques
4.1.2.1 Synthesis of Peptide Dendrimer
The structure of the dendrimer was tailored by the use of ethylene diamine as the core
molecule along with L-lysine and L-arginine as its branching monomers. The synthesis of the
peptide dendrimer was undertaken as described elsewhere, with minor modifications24. For a
typical synthesis, 1 mmol of ethylene diamine and 4 mmol of ester-activated di-Boc-L-lysine
were dissolved in dimethyl formamide. The reaction was maintained at 0–5 °C for 24 h under
constant stirring and a nitrogen atmosphere. The synthesis scheme of peptide dendrimer is
elaborately depicted in Figure 4.1. Compound 1 precipitates at the end of the coupling reaction
and is washed with 15% sodium chloride, 5% aqueous citric acid, 5% sodium bicarbonate
solution and ultrapure water successively, and subsequently dried under high vacuum. It is then
purified by column chromatography by using 3% ethyl acetate in PET ether as the eluent, and
subjected to deprotection to remove the boc- groups in order to yield compound 2 (generation
1: 72.1% yield). For the synthesis of the second generation dendrimer, 8 mmol of di-Fmoc-L-
Arginine was dissolved in anhydrous dichloromethane and activated in the presence of 1-ethyl-
3-(3-dimethylaminopropyl) carbodiimide and N-hydroxysuccinimide for 45 minutes.
Simultaneously, 1 mmol of compound 2 was dissolved in anhydrous dichloromethane and
stirred under a nitrogen atmosphere in an ice bath. The activated arginine solution was added
to the solution under continuous stirring under a nitrogen atmosphere, and maintained between
0–5 °C for 24 hours. After completion of the reaction, a white compound 3 was precipitated,
dried under reduced pressure and washed thoroughly in a manner similar to compound 1.
Compound 3 was then purified by column chromatography with 7% ethyl acetate in PET ether
as the eluent, and subjected to deprotection to remove the Fmoc- groups, in order to yield
compound 4 (EDA-KR2 dendrimer) (generation 2: 61.7% yield).
68
Figure 4.1 Schematic representation of the synthesis of the EDA-KR2 dendrimer
4.1.2.2 Evaluation of Biocompatibility of the Dendrimer
Cell Culture: The toxicity of functional nanomaterials in a living system is an important factor
that limits their use in biomedical applications. To meet this end, an as-prepared peptide
dendrimer was assessed for its biocompatibility with murine fibroblast (L929), human cervical
cancer (HeLa), human oral carcinoma (KB), human breast adenocarcinoma (MCF-7), and
human prostate cancer (PC-3) cell lines. The cell lines were cultured in appropriate growth
medium supplemented with 10% FBS and antibiotics at 37 °C in a 5% CO2 environment in a
humidified incubator. When confluent, the cells were treated with trypsin-EDTA solution for
detachment and counted by the trypan blue exclusion method25.
For a typical experiment, 2×104 cells/well were seeded in a 96-well plate and were
incubated in a suitable tissue culture medium for 24 h at 37 °C in a 5% CO2 environment. The
formulations were suspended in the appropriate growth medium and serially diluted (20.0,
10.0, 5.0, 2.0, 1.0, 0.5, 0.25, 0.125, 0.625 and 0.3125 mg/ml). The exhausted media in the wells
was replaced with 200 µl of this mixture and the cells were allowed to grow for an additional
24 h in the presence of the nanoparticles in the growth media under similar growth conditions.
69
The cells were then washed with 1X PBS (pH 7.3) carefully and the viable cell population was
determined by the sulforhodamine-B colorimetric assay26 described in section 2.2.4.
4.1.2.3 Characterisation Techniques
The 1H and 13C NMR spectra were recorded by the Brüker Avance 500WB, 500 MHz
spectrometer. The chemical shifts, δ, are denoted in ppm (parts per million). The X-ray
Photoelectron Spectroscopy (XPS) data was recorded by Multilab 2000, Thermo VG
Scientific, USA, using 250 W micro-focused monochromatic Al Kα as an X-ray source
(hν=1486.6 eV). The binding energies obtained from the XPS analyses were standardised using
C1s core levels at 284.6 eV. Fourier-transform infrared (FTIR) spectroscopy measurements,
thermal degradation profiles, measurements of hydrodynamic diameter and the zeta potential
and the absorbance measurements of the SRB assay were recorded as described in section 2.2.5.
4.1.3 Results and Discussions
4.1.3.1 Characterisation of the peptide dendrimer
Compound 2 was a white solid (4.6 g, yield 72.1%) after purification and deprotection
of compound 1 (Figure 4.1). 1H NMR (δ, CDCl3): 1.277 – 1.809 (m, -CH2; Lys), 2.531 (m,
NCH2NHCO; amide), 2.853 – 3.279 (s, CH2N; EDA and m, CH2NH; Lys &NCH2CH2NHCO),
3.813 – 3.967 (m, COCH(R)NH), 5.11 (s, -NH2; Lys). 13C NMR (δ, CDCl3): 21.77 (4C5, -CH2),
29.204 (4C6, -CH2), 31.115 (4C4, -CH2), 40.840 (4C7, -CH2-CH2-NH2), 47.006 (2C1, -CH2
EDA), 52.230 (4C3, -CO-CH(NH2)CH2), 176.199 (4C1, CH2(NH2)CON).
Compound 4 was also obtained as a white solid (2.8 g, yield 61.7%) after deprotection
of compound 3 (Figure 4.1). 1H NMR (δ, d6-DMSO): 1.034 – 1.702 (m, -CH2; Lys & Arg),
1.938 (d, CH2NHC(R); Arg), 2.563 (t, CH2CH2NH; Arg), 3.363 (t, CH2CH(NH2)CO; Lys &
Arg), 3.689 (s, CH2N; EDA), 4.120 (t, CH2CH2N(R), Lys), 5.264 (d, COCH(NH)CH2, Lys
&Arg), 8.653 (s, NHC(NH)NH2, Arg). 13C NMR (δ, CDCl3): 22.868 (4C4, CH2CH2CH2, Lys),
24.342 (8C4, CH2CH2CH2, Arg), 25.310 (8C3, CH(NH2)CH2, Arg), 29.183 (4C5, CH2CH2CH2,
Lys), 34.264 (4C3, CH(NH2)CH2, Lys), 40.840 (8C5, CH2CH2NH, Arg), 47.006 (4C6,
CH2CH2N(R), Lys), 48.034 (2C, CH2CH2, EDA), 53.838 (8C2, COCH2(NH2)CH2, Arg),
60.446 (4C2, COCH2(NH2)CH2, Lys), 156.909 (8C6, NHC(NH)NH2, Arg), 175.864 (12C1,
NOCH2(NH2), Lys & Arg).
70
XPS was used to study and confirm the formation of the amide bonds between the
monomers to form the EDA-KR2 (Figure 4.2 (a, b and c)). The C1s core level of EDA-KR2
shows three deconvoluted peaks at 284.6, 286.5, and 289.6 eV, respectively. The peak at 284.6
eV corresponds to the C-H and C-C skeleton of the peptide dendrimer, while the intermediate
one at 286.5 eV is associated with the C-N bond. The peak towards higher binding energy
could be assigned to N-C=O and –O-C=O bonds of EDA-KR2. The N1s core levels of EDA-
KR2 show a sharp peak that reveals two deconvoluted peaks at 400.1 and 401.7 eV, which
correspond to the C-N-H and N-C=O bonds, respectively. The O1s core level of EDA-KR2
shows a single broad peak that is further deconvoluted into three components that appear at
532.4, 533.7, and 535.2 eV and are attributed to the N-C=O, –O-C=O and –O-C=O bonds,
respectively. The XPS spectrum of the EDA-KR2 dendrimer shows asymmetric features, which
indicates the presence of multiple oxygen bonds. All of these chemical bonds show a slight
shift in their binding energies, which confirms the successful synthesis of the peptide
dendrimer.
Figure 4.2 High-resolution X-ray Photoelectron spectra showing C1s, N1s and O1s core
levels of the EDA-KR2 dendrimer
The FTIR spectrum of the EDA-KR2dendrimer is shown in Figure 4.3a. The spectral
bands of the monomeric units of lysine and arginine separately are appropriately resolved,
whereas, the bands seen in the peptide dendrimer are less resolved. The FTIR of generation
one (G1) dendrimer primarily shows vibrational bands due to the presence of lysine, which
masks the bands of EDA. The spectrum of EDA-KR2 depicts the vibrational bands of both
lysine and arginine, which occur at 3333 cm-1 (3151 cm-1) due to NH- stretching, 2969 cm-1
(2928 cm-1) due to asymmetric stretching of CH3 and 1710 cm-1 (1680 cm-1) due to out-of-
plane bending of -NH2, while the band that appears at 1581 cm-1 (1574 cm-1) represents the
stretching of the C=O present in the dendrimer. The bands observed at 1450 cm-1 (1464 cm-1)
arise due to asymmetric bending of CH3, and at 1410 and 1376 cm-1, due to symmetric, in-
71
plane bending of CH3 groups, respectively. The characteristic peaks of the secondary amide
(1544 cm-1) and the primary amide (1645 cm-1) confirmed the successful binding of arginine
to the lysine end groups through amide/peptide bonds. The values in parentheses represent the
values reported in earlier published works27, 28.
Figure 4.3 (a) FTIR spectra (b) thermal degradation profile of the EDA-KR2 dendrimer
The thermal analysis profile of the EDA-KR2 dendrimer (Figure 4.3b) showed a weight
loss of ~0.6% with a corresponding endothermic peak at ~50 °C, which could be ascribed to
the removal of adsorbed moisture. The weight loss of ~7.6% with a sharp endothermic peak
seen at ~180 °C is associated with the disintegration and removal of arginine molecules from
the surface of the peptide dendrimer. A further weight loss of 50.2% by 250 °C, with a broader
endothermic peak is attributed to the complete degradation of the constituent molecules of the
dendrimer. Therefore, the thermal degradation profile of the EDA-KR2 molecule suggested
that the synthesised dendrimer is stable up to the temperatures of ~150 °C. In addition, the zeta
potential measurements showed that the EDA-KR2 dendrimer carries a positive surface charge
of +45 mV at a near-physiological pH (pH 7.3), which arises due to the resonating secondary
amine groups of arginine, free on the surface of the EDA-KR2 dendrimer and extending into
the surrounding aqueous medium.
4.1.3.2 In vitro assessment of the peptide dendrimer
The cells were exposed to very high concentrations of the EDA-KR2 dendrimer in their
growth medium. Concentrations of as high as 20 mg/ml of the dendrimer showed minimal
effect on L929, PC3 and MCF-7 cells, with an approximate of 10% of decrease in the viable
cell population of HeLa and KB cells. As the concentration of the dendrimer was reduced, it
was observed that there was no hindrance to the cell proliferation activity, irrespective of the
72
cell line. The viable cell population and the cellular morphology also remained unchanged and,
therefore, it was safely concluded that these dendrimers can be safely used with a wide variety
of cancer cells in in vivo settings.
Figure 4.4 Cell viabilities incubated with peptide dendrimer. The dendrimer is seen to be
biocompatible for a variety of cell lines, even at concentrations of as high as 20 mg/ml.
4.1.4 Summary
In the current chapter, an attempt was made to synthesise a biocompatible peptide
dendrimer that had amide interiors and amine exteriors. NMR, FTIR, XPS, Zeta potential
measurements, and thermogravimetry were utilised to confirm the synthesis and to characterise
these peptide dendrimers. Since amino acids were used as branching units, the dendrimer was
found to be biocompatible even at higher concentrations. Amino acids were chosen as the
branching units because, they could be utilised by the cells for their metabolism after
degradation of the dendrimer, thereby reducing the toxicity. The amide bond established
between amino acids is a peptide bonds and thus, these dendrimers can also be categorised as
peptide dendrimers. The synthesised peptide dendrimers have an edge over the widely used
PAMAM dendrimers due to better biocompatibility.
Based on these results, we expect this peptide dendrimer to behave in a fashion similar
to that of PAMAM dendrimers. This necessitates a thorough comparison of their performances
as drug delivery vectors and other biomedical applications. To this end, the forthcoming section
demonstrates a comparative study of PAMAM (commercial) and EDA-KR2 (as-prepared) in
73
their ability to conjugate with Fe3O4 nanoparticles and successfully deliver DOX. These
platforms were also compared for combinatorial thermo-chemo therapy in in vitro settings.
4.1.5 References
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Vessières and G. Jaouen, Macromolecules, 2007, 40, 8568-8575.
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Wiley & Sons, Ltd, 2002, DOI: 10.1002/0470845821.ch25, pp. 587-604.
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5. K. Rouhollah, M. Pelin, Y. Serap, U. Gozde and G. Ufuk, Journal of Pharmaceutical
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Lee, Pharm Res, 2012, 29, 1627-1636.
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Release, 2007, 117, 137-146.
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Vasu, N. Madhusudhan, P. K. Maiti, A. K. Sood, S. Das and N. Jayaraman,
Bioconjugate Chemistry, 2013, 24, 1612-1623.
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B. Voit, S. Rozalska and K. Lisowska, New Journal of Chemistry, 2012, 36, 2215-2222.
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Leimkühler, U. Wollenberger and R. Haag, ACS Applied Materials & Interfaces, 2014,
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2, 302-311.
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Tan, A. W. Griffioen, J. M. A. van Engelshoven and W. H. Backes, NMR in
Biomedicine, 2006, 19, 133-141.
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ONE, 2015, 10, e0137104.
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Chemistry, 2015, 87, 3949-3956.
20. N. Malik, R. Wiwattanapatapee, R. Klopsch, K. Lorenz, H. Frey, J. W. Weener, E. W.
Meijer, W. Paulus and R. Duncan, Journal of Controlled Release, 2000, 65, 133-148.
21. R. Duncan and L. Izzo, Advanced Drug Delivery Reviews, 2005, 57, 2215-2237.
22. R. B. Kolhatkar, K. M. Kitchens, P. W. Swaan and H. Ghandehari, Bioconjugate
Chemistry, 2007, 18, 2054-2060.
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75
Chapter 4
Section B
4.2 Dendritic Fe3O4 Nanoparticles for Combinatorial Therapy:
Peptide Dendrimers with Enhanced Efficiency as Alternative
Platforms for PAMAM Dendrimers
4.2.1 Introduction
In the earlier chapters, we demonstrated the need for surface engineering of Fe3O4
nanoparticles and their successful use in drug delivery applications. Our work in chapter three
of this thesis demonstrated that the encapsulation of these nanoparticles by PAMAM dendrimer
presents itself as a promising alternative platform in drug delivery. We evaluated the drug
delivery performance of these nanoparticles thoroughly, and it was observed that they were
capable of carrying both cationic and anionic therapeutic payloads without any loss in the
anticancer activity of the drug molecules. In the last section, we synthesised a peptide
dendrimer in order to contest its biomedical performance against PAMAM dendrimer. Also,
we intend to look into alternative therapeutic strategies with this peptide dendrimer. The
classical therapies that combat cancer in clinics are radiotherapy and chemotherapy.
Radiotherapy utilises high-energy radiations that induce cell death by sensitising the cancer
cells1. Chemotherapy, on the other hand, involves the use of a variety of chemical molecules
(singularly or in combination) to kill cancer cells2, 3. Also, surgical removal of the tumour mass
is undertaken but is limited to solid tumours. These therapies do not destroy cancer cells alone,
76
but are detrimental to normal cells. This lack of differentiation between cancerous and normal
cells in classical therapies has forced researchers to look for alternative therapies that
selectively act on tumour cells only. This search has led to the development of multifunctional
nanomaterials that could be exploited for targeted cancer therapeutics that have improved
clinical significance.
The magnetic nanoparticles in colloidal suspensions generate heat when exposed to an
alternating magnetic field4. Both ferromagnetic as well as superparamagnetic nanoparticles are
capable of generating heat and could be used for magnetic hyperthermia. But, so far, only
superparamagnetic nanoparticles have been clinically used. Two major factors dictate this
choice of particles. First, on application of ACMF, due to the high hysteresis losses the
ferromagnetic particles would heat uncontrollably, which in clinical settings will prove lethal
to normal cells in addition to the cancer cells. Second, on removal of the external magnetic
field, the particles will be still magnetised (remnant magnetisation) and will agglomerate due
to interparticle magnetic interactions. This agglomeration is highly undesirable inside a
biological system. On the other hand, superparamagnetic particles heat only due to rotational
losses, that tend to saturate and, therefore, cannot generate heat beyond a certain temperature.
Also, due to zero hysteresis, these particles show no residual magnetisation after removal of
ACMF and do not agglomerate.
When the difficulties mentioned above are taken into account, multifunctional
superparamagnetic nanomaterials enter the picture of cancer therapeutics. These ‘smart’ or
‘multifunctional’ nanomaterials are widely used in various biomedical applications5-10. These
nanomaterials could be targeted to the cancer site by using an external magnet (MDT), which
makes them selective to some extent. Due to the external heating of these nanoparticles, the
rise in temperature can be successfully limited to the area in which these particles are
accumulated. Studies about the optimising of the working parameters of magnetic
hyperthermia hypothesise that high magnetisation and higher SAR values are the prerequisites
for these nanoparticles, towards their clinical use. However, after accumulation in target cells,
these superparamagnetic nanoparticles behave very differently, and the observed SAR and the
resultant heating then are lower than required. Thus, the working parameters need to be
redefined for successful magnetic hyperthermia11-13. Since there are no existing strict protocols
for the clinical use of hyperthermia, this limitation is dealt with by using a high concentration
of nanoparticles and amplified strengths of applied magnetic fields. In addition to being
successful drug carriers, superparamagnetic magnetic nanoparticles have also been widely used
as hyperthermic agents in cancer therapeutics14. Shah et al.15 developed core-shell
77
hyperthermic nanoplatforms, which were successfully targeted to the mitochondria inside the
cell. The presence of pro-apoptotic amphipathic tail-anchoring peptide (ATAP)-targeting
ligand improved tumour targeting and cellular uptake in malignant brain cancer cells
(glioblastomamultiforme, U87vIII) and metastatic breast cancer cells (MDA-MB-231).
Hyperthermia was seen to sensitise the cells towards ATAP and to act synergistically with the
ATAP moieties, inducing apoptosis by mitochondrial membrane permeation.
The enhanced sensitivity towards chemotherapeutic drugs by magnetic hyperthermia
opened up avenues of combination therapy, which involved a variety of therapeutic molecules
in combination with the magnetic hyperthermia for superior anticancer treatment. The use of
magnetic hyperthermia by Barick et al.16 to sensitise human cervical cancer (HeLa) cells
towards doxorubicin effected results that were quite interesting. The peptide-mimicking cross-
linked macromolecules on the surface of nanocarriers were seen to behave like amphiphilic
cell-penetrating peptides; these macromolecules facilitated cellular internalisation, enhanced
accumulation of doxorubicin molecules, and in combination with ACMF, showed significantly
high cell toxicity. Kim et al.17 fabricated copolymer of N-isopropylacrylamide (NIPAAm) and
N-hydroxy-methylacrylamide (HMAAm) (poly(NIPAAm-co-HMAAm)) with magnetic
nanoparticles and used them for ‘switchable’ delivery of DOX. The population of viable human
melanoma cells (COLO 679) decreased by 30% just after the exposure to ACMF for 5 min.
This decrease was attributed to the induction of apoptosis by synergistic action of DOX and
hyperthermia. Apart from doxorubicin, hyperthermia was seen to sensitise and enhance the
anticancer capability of other therapeutic drugs as well. Rao et al.18 used thermally-responsive
polymeric nanoparticle-encapsulated curcumin against human prostate adenocarcinoma. The
application of ACMF and the subsequent hyperthermia was seen to significantly enhance the
anticancer capability of curcumin, decreasing its inhibitory concentration 7-fold. Recently,
researchers have also combined magnetic hyperthermia with the delivery of other therapeutic
molecules such as oligonucleotides19, 20; this technique is in its nascent stage.
A variety of multifunctional nanoplatforms have been developed and studied for
magnetic hyperthermia and combinatorial therapy but dendrimer functionalised magnetic
nanoparticles have not yet been explored much. The current chapter thus demonstrates an
improved performance in anticancer combinatorial therapy through the use of dendritic
nanoparticles as carrier and hyperthermic platforms. The work undertaken in this chapter
demonstrates the fabrication of the dendrimer-modified Fe3O4 nanoparticles through the use of
as-prepared peptide dendrimers, which were then examined for their efficacy in chemotherapy
78
and combinatorial thermo-chemotherapy in comparison with the widely used PAMAM-coated
Fe3O4 nanoparticles.
4.2.2 Experimental Techniques
4.2.2.1 Synthesis of Dendrimer-coated Fe3O4 Nanoparticles
Fe3O4 nanoparticles were synthesised by the conventional co-precipitation method and
stabilised by glutamic acid, as described in chapter three of this thesis. The surface of the
glutamic acid-modified Fe3O4 nanoparticles (Glu–Fe3O4) was further modified with
commercially available PAMAM dendrimers and as-prepared peptide dendrimers to yield
PAMAM-modified Fe3O4 (PAMAM–IO) and EDA-KR2-modified Fe3O4 (KR2–IO)
nanoparticles, respectively. For dendrimer modification, 500 µg of Glu–Fe3O4 nanoparticles (2
mg/ml) were incubated with varying amounts of both the dendrimers, PAMAM and EDA-KR2
(w/w ratios of 1, 2, 4, 6, 8 and 10); the volume of the mixture was made up to 1 ml. This
reaction mixture was incubated under shaking/rocking overnight to yield PAMAM–IO and
KR2–IO nanoparticles. These dendritic nanoparticles were then collected over a permanent
magnet and washed with ultrapure water 3–4 times. The absorption spectra of the supernatants
(and washings) were analysed at 280 (PAMAM) and 230 nm (EDA-KR2) in order to calculate
the bound dendrimers on the surface of the nanoparticles against the standard plot that was
prepared under similar conditions (R2=0.999 and 0.997, respectively). Each experiment was
performed in triplicates.
4.2.2.2 Drug Loading and Assessment of Binding Interactions
The carrier efficiency of the PAMAM–IO and KR2–IO nanoparticles was evaluated
using DOX as the therapeutic molecule. The drug loading efficiency of both the dendritic
nanoparticles was investigated by recording the fluorescence spectrum (ex=490 nm and
em=560 nm) of DOX. Each experiment was performed in triplicates and the standard deviation
was calculated. For a typical drug loading experiment, varying amounts of dendritic
nanoparticles (PAMAM–IO & KR2–IO; Stock: 2 mg/ml) were added to 1 ml of an aqueous
DOX solution (10 g/ml), mixed gently and then incubated for 10 min; the fluorescence
spectrum of the supernatant was then recorded from 520 to 800 nm after magnetic
sedimentation. The spectrum of the pure DOX solution (10 g/ml) was recorded for
79
comparative evaluation and was used as a standard initial intensity. The addition of dendritic
nanoparticles showed a steady decrease in the fluorescence intensity of the DOX solution. The
process was continued until no further decrease in the intensity was observed. The DOX-loaded
dendritic nanoparticles were washed with ultrapure water 3–4 times over a permanent magnet,
and the washings were collected and analysed for loss of any superficially bound drug. The
loading efficiencies were calculated using eqn. 4.1.
The binding interactions of DOX molecules with both the dendritic nanoparticles were
further studied and understood by using a modified Stern–Volmer plot. For calculation of the
binding constant, 10 g of DOX was dissolved in 1 ml of ultrapure water and the fluorescence
spectrum was recorded. After 15 min, 20 g of PAMAM–IO and KR2–IO from the stock (2
mg/ml) were added and incubated for 15 min. The fluorescence spectrum of the supernatant
was recorded by magnetic sedimentation. The fluorescence intensities of the DOX molecules
were plotted against the corresponding total concentration of the nanoparticles. The data were
plotted and analysed according to the following relation21:
log [ΔF 𝐹⁄ ] = log 𝐾 + 𝑛 log [𝑄] (Eqn. 4.1)
where ΔF is the difference between the initial and final fluorescence intensity of DOX
(fluorophore) in the absence and presence of nanoparticles (quencher), K is the binding
constant, n is the binding affinity and [Q] is the total concentration of nanoparticles.
4.2.2.3 Drug Release Studies
The DOX-loaded dendritic nanoparticles were quantified according to their loading
efficiencies in order to perform the drug release experiments. The release behaviour of DOX
was assessed under reservoir-sink conditions using low pH as a stimulus. For evaluation of
drug release profiles, each of the DOX-loaded dendritic nanoparticles was suspended in 5 ml
of a 0.1X sodium acetate buffer (pH 5.0) in a dialysis bag (separately) that acted as reservoir.
The dialysis bag was then suspended in 200 ml 0.1X PBS (pH 7.3) in order to represent the
reservoir-sink conditions. Aliquots of 1 ml from the sink (PBS) were withdrawn and replaced
with an equal amount of fresh PBS simultaneously at fixed time intervals in order to maintain
the concentration gradient across the semi-permeable membrane. The aliquots were then
subjected to the analyses, and the amount of DOX released (cumulative release) was
determined by the measurement of fluorescence against the standard plot. Each experiment was
performed in triplicates and the standard deviation was calculated. For evaluation of drug
80
release under neutral pH conditions, the DOX-loaded dendritic nanoparticles were suspended
in PBS (pH 7.3), and an experiment similar to the one described above was carried out.
4.2.2.4 Evaluation of Biocompatibility and Therapeutic Efficacy
The biocompatibility of PAMAM, EDA-KR2, PAMAM–IO and KR2–IO were assessed
with murine fibroblasts (L929), human cervical cancer (HeLa), human oral carcinoma (KB),
human breast adenocarcinoma (MCF-7), and human prostate cancer (PC-3) cell lines. The
formulations were suspended in an appropriate growth medium and serially diluted (2.0, 1.0,
0.5, 0.25, 0.125, 0.625 and 0.3125 mg/ml). 200 µl of this mixture was used to replace the spent
media in the wells, and the cells were allowed to grow for an additional 24 h in the presence of
the nanoparticles in the growth media under similar growth conditions. To evaluate the
potential of DOX-loaded dendritic nanoparticles to release DOX in the cancer cells, a dose-
dependent study was undertaken over 24 h. The therapeutic efficacy was evaluated with
different cancer cell lines (HeLa, KB, MCF-7 and PC-3). The concentration of the DOX-loaded
dendritic nanoparticles that reduced the cell population by 50% was referred to as the inhibitory
concentration (IC50) values of the said formulation, and was calculated by a dose-responsive
sigmoidal curve fitting by using Origin 8.0 software. The live cell population was determined
by the SRB assay22 and represented as relative percent viability against untreated cells as the
control (Eqn. 3.3).
4.2.2.5 Evaluation of Calorimetric Potential of Dendritic Nanoparticles
In order to evaluate time-dependent heat generation by the nanoparticles, 1 ml of an
aqueous suspension of different concentrations (1, 2 and 5 mg/ml) of the systems of both the
dendritic nanoparticles was placed along the radial centre of a solenoid coil of a radio frequency
generator; the suspension was then exposed to an ACMF of 232, 271 and 309 Oe (operating at
a current of 300, 350 and 400 A, respectively). The set-up was appropriately insulated in order
to minimise the heat loss. The time-dependent increase in the temperature of the suspension
was recorded, and the SAR was calculated by the initial slope using eqn. 3.2. The
concentrations of both, the dendritic nanoparticles and the applied ACMF, were optimised in
order to heat the suspension to 45±0.5 °C and to maintain it at that temperature for subsequent
magnetic hyperthermia and combinatorial therapeutic studies.
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4.2.2.6 In vitro Magnetic Hyperthermia (MHT) Studies
The effect of both the dendritic nanoparticles, with and without the exposure to ACMF,
was assessed on the live cell population using human cervical cancer (HeLa) cells as a model
system. To a cell suspension that contained 1×106 cells, 500µl (2 mg/ml) of each of the sterile
dendritic nanoparticles was added. The volume was made up to 1 ml by a supplemented growth
medium and mixed very gently to lessen the cell agglomeration. The cell suspensions were
then exposed to an ACMF of 271 Oe (350 A) initially for 8 min. (PAMAM–IO) and 12 min.
(KR2–IO) in order to reach 45 °C. The temperature was maintained at 45±0.5 °C for different
time intervals (5, 10, 15, 20, 25 and 30 min) by reducing the applied field strength to 193 Oe
(250 A). After this, the time for which the cell suspension was exposed to the ACMF (after
reaching 45 °C) is referred to as the treatment time. The treated cell suspension was then
incubated for 6 h at 37 °C in a 5% CO2 environment and then counted for viable cell population:
% 𝑉𝑖𝑎𝑏𝑙𝑒 𝐶𝑒𝑙𝑙𝑠 =𝑁𝑜. 𝑜𝑓 𝑙𝑖𝑣𝑒 𝐶𝑒𝑙𝑙𝑠
𝑇𝑜𝑡𝑎𝑙 𝑐𝑒𝑙𝑙𝑠 × 100 (Eqn. 4.2)
4.2.2.7 ACMF-triggered Combinatorial Therapy
The combined therapeutic effect of DOX-loaded dendritic nanoparticles and magnetic
hyperthermia was evaluated using HeLa cells. The experimentation performed was similar to
the one followed for magnetic hyperthermia studies. An aqueous suspension of DOX-loaded
dendritic nanoparticles (at IC50 concentration) was added to a cell suspension of 1×106 cells
and then exposed to ACMF for a treatment time of 10 min. (LD50). The cell suspension was
then incubated for 6 h and the viable cell population was calculated using eqn. 4.2. The results
were represented as the mean ± standard deviation of the mean (SD) from three independent
experiments that were performed at different times. The analyses were done using the statistics’
module of Origin 9.1 software in order to evaluate significant differences between a pair of
results. The levels of significance that were evaluated were p< 0.05 (*), p< 0.01 (**) and p<
0.001 (***).
4.2.3 Results and Discussions
4.2.3.1 Characterisation of Dendrimer-coated Fe3O4 Nanoparticles
The characterisation of the prepared nanomaterials was undertaken as described in
section 2.2.5. The XRD patterns of PAMAM–IO and KR2–IO nanoparticles were analysed to
82
identify the crystal structure (Figure 4.5a). The diffraction peaks correspond to the inverse
spinel cubic crystal structure of Fe3O4. It was also observed that the characteristic XRD pattern
of Fe3O4 was not affected by the conjugation of dendrimers, maintaining the crystallinity of
Fe3O4 nanoparticles. Figure 4.5b shows the FTIR spectra of the glutamic acid, Glu–Fe3O4,
PAMAM, PAMAM–IO, EDA-KR2 and the KR2–IO nanoparticles. The FTIR of PAMAM
dendrimers show characteristic vibrations, as reported in earlier works23. The FTIR spectrum
of EDA-KR2 dendrimers is in accordance with the spectrum described earlier in chapter five.
The Glu–Fe3O4 nanoparticles showed few broad bands upon being coated with the dendrimers.
Thus, the results confirm the successful conjugation of the dendrimers onto the surface of Glu–
Fe3O4 nanoparticles. UV-Vis absorption spectroscopy was used to quantify the amount of
dendrimers that were utilised in the surface modification of Glu–Fe3O4 nanoparticles. The
percent conjugation of PAMAM and EDA-KR2 dendrimers with the surface of Glu–Fe3O4 was
found to be 34.0±5.3% and 32.5±5.6%, respectively, as calculated against the standard curves.
Figure 4.5 (a) XRD pattern; (b) FTIR spectra; (c) thermal degradation profiles and (d, e, f)
electron micrographs of Glu–Fe3O4, PAMAM–IO, and KR2–IO nanoparticles, respectively (σ
≤ 15%)
The thermal analyses profiles of both the dendritic magnetic nanoparticles show weight
loss, which is attributed to the loss of adsorbed moisture, the dendrimers (PAMAM and EDA-
KR2 each), and the phase transition of magnetite to maghemite (Figure 4.5c). PAMAM–IO
nanoparticles show an initial weight loss of 1.3%, which corresponds to the water molecules
(d) (e) (f)
83
that are physically adsorbed onto the nanoparticles; a further loss of 7.7% in weight is observed
up to a temperature of 400 °C and is attributed to the degradation of PAMAM molecules that
are conjugated with the surface of Glu–Fe3O4 nanoparticles. The degradation profile of KR2–
IO nanoparticles reveals a higher weight loss of approx. 13.7%, when compared to the weight
loss of PAMAM–IO. This is due to the bulky architecture and molecular weight of the EDA-
KR2 dendrimer as opposed to the PAMAM dendrimer. The transmission electron micrographs
of Glu–Fe3O4, PAMAM–IO, and KR2–IO nanoparticles show spherical morphology with slight
agglomeration. The particle sizes of Glu–Fe3O4, PAMAM–IO and KR2–IO nanoparticles range
from 8–12 nm, 10–15 nm and 10–18 nm, respectively (Figure 4.5d, e and f).
XPS was used to analyse the binding energies of Glu–Fe3O4, PAMAM–IO and KR2–
IO nanoparticles (Figure 4.6). The spectrum of Glu–Fe3O4 nanoparticles reveals that the C1s
levels deconvoluted into three peaks that correspond to the C–H and C–C, C–N and N–C=O, -
O–C=O, respectively. The XPS spectra of PAMAM–IO and KR2–IO nanoparticles show that
the C1s core levels deconvoluted in two (284.6 and 287 eV) and three (283.8, 284.7 and 286.8
eV) peaks, respectively. The peaks that arise at lower energies correspond to the C–H and C–
C bonds of the dendrimers, while the intermediate peak in KR2–IO (284.7 eV) is associated
with the C–N bond (not visible in the PAMAM–IO). The higher binding energy peaks arise
due to the N–C=O and -O–C=O bonds present in the interiors of the dendrimers. The Fe2p
levels of Glu–Fe3O4 reveal the Fe3+ and Fe2+ oxidation states of magnetite, which is consistent
even after the conjugation of the dendrimers. This confirms that surface engineering by the
dendrimers does not alter the crystal structure of Glu–Fe3O4. The O1s core levels of both these
dendritic nanoparticles show asymmetric features, indicating the presence of multiple oxygen
bonds. The O1s levels show a single sharp peak with a shoulder, which is further deconvoluted
into two components that mark the presence of iron oxide-bound oxygen and the oxygen in the
dendrimer skeleton, respectively. The N1s levels of both the dendritic magnetic nanoparticles
show sharp peaks, which further reveal two deconvoluted peaks that represent the C–N–H and
N–C=O bonds, respectively. All of the binding energies show a slight shift, confirming the
successful conjugation of dendrimers to Glu–Fe3O4.
84
Figure 4.6 High resolution XPS spectra for C1s, Fe2p, O1s and N1s of the Glu–Fe3O4,
PAMAM–IO and KR2–IO nanoparticles
4.2.3.2 Drug Loading and Assessment of Binding Interactions
Figure 4.7a and b shows the spectrum of the pure DOX solution against the spectra
obtained by incubating with different amounts of PAMAM–IO and KR2–IO, respectively. The
conjugation of DOX to these nanoparticles leads to the reduction in the number of free electrons
Glu-Fe3O4 PAMAM-IO KR2-IO
279 282 285 288 291 294
C1s
In
ten
sit
y (
a.u
.)
Binding Energy (eV)
(284.6 eV)
C-H, C-C
C-N
(285.9 eV)
N-C=O, -O-C=O
(288.3 eV)
278 280 282 284 286 288 290 292 294
C1s
Inte
ns
ity
(a
.u.)
Binding Energy (eV)
(284.6 eV)
C-C, C-H
(287 eV)
N-C=O, -O-C=O
282 283 284 285 286 287 288
C1s
Inte
nsit
y (
a.u
.)
Binding Energy (eV)
(283.8 eV)
C-C, C-H
(284.7 eV)
(286.8 eV)
N-C=O, -O-C=O
C-N
700 707 714 721 728 735 742
Fe2p
Inte
ns
ity
(a
.u.)
Binding Energy (eV)
(711.4 eV)
Fe in Fe3O4
700 705 710 715 720 725 730 735 740
Fe2p
Inte
ns
ity
(a
.u.)
Binding Energy (eV)
710.2 eV
Fe in Fe3O4
700 710 720 730 740
Fe2p
Inte
ns
ity
(a
.u.)
Binding Energy (eV)
527 528 529 530 531 532 533 534 535
O1s
Inte
ns
ity
(a.u
.)
Binding Energy (eV)
Fe3O
4 - Oxide
Hydroxides
(530 eV)
(531.6 eV)
524 526 528 530 532 534 536 538
O1s
Inte
nsit
y (
a.u
.)
Binding Energy (eV)
(529 eV)
Fe3O4 - Oxide
N-C=O
(530.5 eV)
524 526 528 530 532 534 536 538
O1s
Inte
nsit
y (
a.u
.)
Binding Energy (eV)
(530 eV)
N-C=O
(528.7 eV)
Fe3O4
396 397 398 399 400 401 402 403 404
N1s
Inte
nsit
y (
a.u
.)
Binding Energy (eV)
(399.9 eV)
C-N-H
(401.2 eV)
N-C=O
392 394 396 398 400 402 404
N1s
Inte
nsit
y (
a.u
.)
Binding Energy (eV)
C-N-H(398.8 eV)
N-C=O
(400.4 eV)
392 394 396 398 400 402 404
N1s
Inte
ns
ity
(a
.u.)
Binding Energy (eV)
(398.4 eV)
C-N-H
(399.3 eV)
N-C=O
85
available for transition, which in turn is responsible for the lowering of the intensity of its
fluorescence. No further decrease even after the addition of nanoparticles indicated the
saturation of the drug loading capacities of the nanoparticles. The drug loading capacity of
PAMAM–IO was comparable to that of KR2–IO and found to be 90.1±2.5 % and 88.7±2.7%,
respectively. The slight difference in the loading efficiency of KR2–IO may be due to the steric
hindrance offered by resonating secondary amine groups of the arginine present on the surface
of EDA-KR2. The results show that the peptide dendrimer (EDA-KR2) is as effective as the
PAMAM dendrimers in conjugating with DOX.
Figure 4.7 Drug loading profiles of (a) PAMAM–IO and (b) KR2–IO nanoparticles;
Gaussian profiles of (c) DOX-PAMAM–IO and (d) DOX-KR2–IO; and Stern–Volmer plots
of DOX with (e) PAMAM–IO (R2=0.973) and (f) KR2–IO (R2=0.995)
86
In order to confirm the conjugation of DOX with the dendrimers, the fluorescence
spectra were further deconvoluted using the Gaussian curve fitting (Figure 4.7c and d). The
emission spectrum of DOX exhibits three maxima peaks at 560 nm (P1), 590 nm (P2) and 630
nm (P3). The nature of the interactions of DOX with the dendritic nanoparticles is signified by
a change in the peak maxima position, the shape of the spectrum, and the ratios of the area
under each deconvoluted peak. From the spectrum of pure DOX, a gradual decrease is observed
in the values of A1/A2 and A3/A2 for the DOX–nanoparticles, which indicates their
conjugation (Table 4.1).
Table 4.1 The area under the peaks of DOX-loaded PAMAM–IO and KR2–IO nanoparticles
(σ≤10%)
PAMAM–IO KR2–IO
A1/A2 A3/A2 A1/A2 A3/A2
DOX 0.658 0.869 0.728 1.004
DOX-Dend. nanoparticles_25% 0.614 0.691 0.614 0.670
DOX-Dend. nanoparticles_50% 0.532 0.649 0.538 0.646
DOX-Dend. nanoparticles_75% 0.527 0.633 0.529 0.629
DOX-Dend. nanoparticles_90% 0.517 0.606 0.515 0.615
In order to understand the binding interactions between DOX and the dendritic
nanoparticles better, the Stern–Volmer plot was used to calculate the binding constant of the
drug loading process (Figure 4.7e and f). The y-intercept of the plot represents the logarithmic
value of the binding constant, which was 1.018 and 1.0009 (μg/ml)-1 for PAMAM–IO
(R2=0.973) and KR2–IO (R2=0.995), respectively. The slope of the plot is a representation of
the fraction of DOX that takes part in the binding or the binding affinity, which was 0.0442
and 0.0512 for PAMAM–IO and KR2–IO, respectively. The binding affinity of KR2–IO
towards DOX is seen to be more than that of the PAMAM–IO, confirming that KR2–IO
nanoparticles have better drug-carrying properties. The appropriate linear fitting of the
modified Stern–Volmer equation leads to the conclusion that the quenching observed is a
bimolecular process that involves DOX and dendritic nanoparticles.
87
4.2.3.3 Drug Release Studies
Figure 4.8 shows the DOX release profiles from the DOX-loaded dendritic
nanoparticles under low pH conditions over a period of 48 h. This reservoir-sink system is
expected to behave and interact in a fashion similar to the lysosome-cytoplasm interface in a
living cell, because it mimics their [H+] environments that are separated by a semi-permeable
membrane. Thus, a low pH acts as an external trigger for on-demand delivery of the DOX
molecules. Large amounts of DOX are released (> 50%) within the initial 6–7 h, followed by
sustained release that attains a plateau after ~12 h. The increase of [H+] in the
microenvironments results in the protonation of DOX, which in turn initiates the weakening of
the non-covalent binding between the DOX and the dendritic nanoparticles. This weakening
results in the release of DOX molecules into the immediate environment and a subsequent slow
diffusion into the sink through the dialysis membrane.
Figure 4.8 Drug release profile of (a) PAMAM–IO and (b) KR2–IO nanoparticles under the
stimulus of pH 5.0 and 7.3
The release efficiency of DOX from KR2–IO is found to be 81.7±1.6%, which is substantially
more than the 64.4±3.2 % of PAMAM–IO, recorded over an experimental period of 48 h. This
indicates that, for similar amounts of DOX carried by these dendritic nanoparticles, the release
efficiency of the KR2–IO is significantly higher than PAMAM–IO nanoparticles. This elevated
drug release observed from KR2–IO could also be due to the acidic hydrolysis of peptide bonds
in the EDA-KR2 dendrimer and the disintegration of the whole dendrimeric architecture, which
may not occur in PAMAM–IO. On the other hand, at near physiological pH, the DOX released
from both the carriers is less than 10%, which confirms the pH responsive behaviour of these
88
dendritic nanoparticles. When the DOX-loaded dendritic nanoparticles are delivered to the
cancer site, which has lower pH than normal body fluids, the release of DOX is expected to
initiate. This DOX release would be facilitated further upon the internalisation of the
nanoparticles in the lysosomes of the cancer cells.
4.2.3.4 Evaluation of Biocompatibility and Therapeutic Efficacy
The Glu–Fe3O4, EDA-KR2, PAMAM–IO, and KR2–IO nanoparticles are seen to be
non-toxic. No hindrance in cell proliferation activity and viable cell population was observed.
The cellular morphology was also unchanged and, therefore, it was concluded that these
nanoparticles can be safely used with a wide variety of cancer cells in in vivo applications.
Therapeutic efficacy studies reveal that the DOX-loaded nanoparticles are competent drug
carriers and can efficiently reduce the population of cancer cells. The IC50 values of DOX-
PAMAM–IO and DOX-KR2–IO were calculated by a dose-responsive sigmoidal curve fitting
from the Origin 8.0 software, and are listed in Table 4.2.
Table 4.2 IC50 values of (a) DOX-PAMAM–IO and (b) DOX-KR2–IO nanoparticles with
various cancer cells. The values are represented in µg/ml of formulations with ≤ 15% (IC50
of pure DOX was 1.1-1.5μM with different cell lines)
Cell Line/s DOX-PAMAM–IO DOX-KR2–IO
HeLa 90.0 125.7
MCF-7 59.9 67.0
PC-3 74.4 89.6
KB 63.0 77.3
4.2.3.5 Evaluation of Calorimetric Potential of Dendritic Nanoparticles
Superparamagnetic nanoparticles are known to generate heat when exposed to an
ACMF due to Brownian and Néel relaxations, which largely depends upon the anisotropy and
size of the nanoparticles4.The zero remanence in superparamagnetic nanoparticles ensures
minimal aggregation of the nanoparticles in the absence of the magnetic field, when present
within a biological system. This heat-generating property was investigated as a time-dependent
study for PAMAM–IO and KR2–IO. The SAR was calculated to be 110.36 W/g (PAMAM–
IO) and 106.12 W/g (KR2–IO), as against the ~134 W/g of Glu–Fe3O4. The observed decrease
89
in the heating capability of the dendritic nanoparticles is attributed to the non-magnetic
dendrimer entities on their surfaces.
However, the SAR is dependent on various physical parameters such as the strength of
the ACMF (H), the frequency of the applied ACMF (ω), and the physical characteristics of the
magnetic nanoparticles, such as magnetisation, particle size and size distribution4, 24-26.
Nevertheless, practically, the heating proficiency of nanoparticles cannot be raised solely by
increasing the parameters mentioned above. Therefore, the SAR value needs to be optimised
by varying these parameters (frequency, field and physical properties of the nanoparticles)
(Figure 4.9A and B). The concentration of 5 mg/ml of dendritic nanoparticles in an aqueous
suspension was seen to be very efficient in raising the surrounding temperature to the desired
levels. However, this concentration is potentially toxic to normal cells. Hence, the subsequent
evaluation of the heating performance was focused on the aqueous suspensions of dendritic
nanoparticles at concentrations of 1 or 2 mg/ml. Under the conditions of applied fields, the
suspension of 2 mg/ml effectively reached 45°C in 7.5 min (309.33 Oe) and 10.5 min (271 Oe),
but failed to achieve this temperature at 232 Oe even after 25 min. The concentration of 1
mg/ml was found to be insufficient under any applied field strength, even after 30 min of
exposure. In summary, the 2 mg/ml concentration of the dendritic nanoparticle suspension and
the ACMF of 271 Oe were considered to be optimal working parameters that would achieve
the maximal hyperthermia treatment effect without being detrimental to the neighbouring
normal cells and were optimised for subsequent magnetic hyperthermia studies.
Figure 4.9 Time-dependent calorimetric measurements of (A) PAMAM–IO and (B) KR2–IO
nanoparticles at different ACMF and concentrations
90
The time required to attain 45 °C (hyperthermia temperature) was observed to be
broadly dependent on the strength of the applied field (H) and the concentration of
nanoparticles in the suspension. Also, at a given concentration, the time required to reach 45
°C shows an inverse proportionality in relation to applied field and increases with a decrease
in the field. This behaviour is also in agreement with the fact that the quantity of heat produced
is directly proportional to the square of the applied ACMF (H) (Figure 4.10).
Figure 4.10 Dependence of SAR on the strength of applied ACMF for (a) PAMAM–IO and
(b) KR2–IO
4.2.3.6 In vitro Magnetic Hyperthermia (MHT) Studies
Hyperthermia induces three main cellular responses: cytotoxicity, radio-sensitisation,
and thermal tolerance27, 28. Cancer cells undertake the apoptotic route in direct response to
applied heat, while the normal cells adapt to the temperature changes much easily. The cells
that do not undergo immediate apoptosis become sensitised towards other forms of therapy,
such as certain chemotherapeutic drugs or radiations. The heat-generating property of magnetic
nanoparticles has been widely exploited in order to find an effective treatment for cancer11, 29-
31. The exposure of radio waves to a biological system is limited in terms of the SAR of the
nanomaterials, and is given below32:
𝑆𝐴𝑅 = 𝜎𝐸𝑖
2
𝜌 (Eqn. 4.3)
where Ei represents the rms value of the applied electric field strength in the tissue (V/m), σ is
the conductivity of body tissue (S/m), and ρ is the density of the body tissue (kg/m3). Its effect
on the cancer cells is assessed as the rate at which electromagnetic energy is absorbed by a unit
mass of the tissue.
91
There are two possible means to induce heat in the cancer cells that are intended for
therapeutics; by heating its (a) extracellular and its (b) intracellular environment. Both the
methods have their own advantages and disadvantages, which are mainly based on their ability
to affect the viable cancer cell population. Extracellular heating was followed in the current
work and the results were encouraging. Two conditions served as the control for this study: (a)
cell suspension with the dendritic nanoparticles, but without the ACMF and (b) cell suspension
without dendritic nanoparticles, but with application of ACMF. The control cells do not show
any reduction in cell proliferation activity and cell population. This confirms that the
application of the ACMF and the administration of nanoparticles separately is not enough to
cause the desired cell death.
On the other hand, the nanoparticles in combination with ACMF led to much more
interesting and promising results (Figure 4.11a). The effect of MHT was established from the
change in the cellular morphology and integrity. Though the MHT (under the above mentioned
parameters) was not enough to kill the cells, the treatment caused discernible changes in the
cellular morphology, the plasma membrane and the cytoplasm, which led to cell death. Figure
4.11b depicts the effect of elevated temperatures on the HeLa cell population under the
influence of the ACMF for various treatment times. The reduction of ~20% of viable cells
instantaneously in the viable cell population is observed after a treatment time of 5 min. Even
a small treatment time of 10 min is sufficient to reduce the number of live cells by ~50%. The
treatment time at which the cell population is reduced to 50% is considered as its median lethal
dose (LD50) and is further used to evaluate the effects of combinatorial therapy in vitro. This
lethal effect of heat on the proliferation and the viability of cells attains saturation at 20 min
and shows minimal effects with further increase in the treatment time.
92
Figure 4.11 (a) Synergistic effects of DOX-loaded nanoparticles and hyperthermia on HeLa
cells (b) Viable HeLa cell population in the presence of ACMF for varying treatment times
for both PAMAM–IO and KR2–IO nanoparticles
4.2.3.7 ACMF-triggered Combinatorial Therapy
The combinatorial effect of DOX-loaded dendritic nanoparticles and hyperthermia was
evaluated in HeLa cells. The treatment time at LD50 values and the formulation concentration
at IC50 values were used in the combining of chemotherapy with magnetic hyperthermia. Figure
4.12 summarises the effect of DOX-loaded nanoparticles in combination with the ACMF-
induced elevated temperatures. The enhanced cell death that is observed is a result of the fatal
synergistic contribution of DOX and high temperatures. The combinatorial treatment is seen to
reduce the viable cell population to 2.5 and 3.6 % for DOX-PAMAM–IO and DOX-KR2–IO,
respectively.
Figure 4.12 Synergistic effects of DOX and magnetic hyperthermia (MHT) on viable HeLa
cell population after exposure to both, DOX-loaded dendritic nanoparticles and ACMF.
93
The laser scanning confocal microscopy (LSCM) images (Figure 4.13) of the control HeLa
cells against the treated cells confirm the detrimental effects of combinatorial therapy. LSCM
images were recorded using an inverted confocal microscope by Olympus, model IX 81. Due
to the fluorescent property of DOX, no tagging molecule was required; the nucleus was stained
using a 4′, 6-Diamidino-2-phenylindole (DAPI) solution. The images show the confocal images
as well as the graphical representation of the fluorescent intensity. These images confirm that
the DOX-loaded dendritic nanoparticles were successfully internalised by the HeLa cells and
caused detrimental changes. The images clearly depict that the plasma membrane is completely
disrupted after the combinatorial treatment, which is due to the toxic effect of DOX in addition
to the MHT. The nucleus is seen to be intact, which gradually degrades as the cells undergo
apoptosis. The graphical analyses of the fluorescent intensity of the control cells have a smooth
and collective intensity profile. In the treated cells, however, the intensity is scattered and
higher. This increase in the intensity of green fluorescent peaks suggested the degradation and
fragmentation of cells due to the released DOX molecules, which lead to the desired cell death.
FITC-labelled nanoparticles were used to tag the control cells. The DOX-loaded dendritic
nanoparticles in synergism with hyperthermia are seen to be successful in reducing the live cell
population to a minimum.
94
Figure 4.13 Laser Scanning Confocal microscopy images of treated HeLa cells with FITC-
nanoparticles (control), DOX-PAMAM–IO and MHT, and DOX-KR2–IO and MHT. The
scale bar of the images is 50 μm.
95
4.2.4 Summary
This section elaborates on our attempt to combine dendrimers with superparamagnetic
Fe3O4 nanoparticles, which resulted in the ‘best of both worlds’ and opened up new arenas in
the field of improved cancer therapeutics. This work demonstrated the fabrication of aqueous-
stable, magnetic and biocompatible formulation of dendritic nanoparticles through the use of
PAMAM and peptide dendrimers. These dendritic nanoparticles showed good SAR values
hence, better efficiency in magnetic hyperthermia. The peptide dendrimer is substantially
efficient than the PAMAM dendrimer in both loading/carrying and delivering DOX to the
desired cancer cells. These formulations were also successful in combining chemotherapy with
magnetic hyperthermia. The comparative evaluation of these dendritic nanoparticles for drug
delivery, magnetic hyperthermia and combination therapy established the superiority of the
peptide dendrimer over PAMAM dendrimers. Also the physico-chemical properties of peptide
dendrimer were minimally compromised as compared to its PAMAM counterpart. These
promising in vitro results established the dendritic-Fe3O4 nanoparticles attractive platforms for
further in vivo evaluations.
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Chapter 5
Assessment of Doxorubicin-loaded Dendritic Fe3O4
Nanoparticles for Magnetic Drug Targeting in Murine
Melanoma Model
5.1 Introduction
The previous chapters of this thesis have led to the conclusion that peptide dendrimer
based Fe3O4 nanoparticles (a) have superior drug delivery efficacy and (b) can synergistically
combine thermo- and chemotherapy for improved therapeutic performances. While the Fe3O4
nanoparticles actively contribute in magnetic hyperthermia, they are also capable of playing a
dual role as a targeting platform. A major challenge in stepping up the successful in vitro
therapies to in vivo scenarios is the uncontrolled biodistribution of the delivery vectors1, 2. This
random systemic dispersal of nanoparticles leads to the release of drug molecules in non-
specified locations causing adverse effects to otherwise healthy tissues. Towards this end,
specific targeting of these delivery vectors have been attempted using a variety of molecules
like antibodies3-6, ligands7, 8, peptides9, 10 and the like. These molecules are known to attach
specifically to the receptors present on the cell membranes and are thus, selectively internalised
by the cancer cells11. This reduces their uptake by the neighbouring normal cells thereby
reducing the side effects. But these targeting moieties also have critical disadvantages as they
are specific to only a certain kind of receptors expressed in cancer cells, that is, a receptor
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present on a prostate cancer cell may not be expressed by a breast cancer cell. Thus, one type
of targeting ligand would be specific to deliver drugs to the tumour depending on the type of
cancer dealt with.
To overcome this limitation, magnetic drug targeting presents itself as a promising
attempt. As the name suggests, this method involves passive targeting of the delivery vectors,
bound to the drug molecules, to the tumour region by the use of a strong external magnetic
field12. The magnetic nanoparticles are of immense importance because the movement of these
nanoparticles can be controlled and guided by an external magnetic field. In concept,
chemotherapeutic agents are bound to these nanoparticles and introduced into the arterial
circulation, supplying to the tumour, while guided externally by a strong magnetic field. It not
only leads to higher accumulation of drug in the tumour but also reduces the overall dose
administered in the body thus minimising the side effects. In order to be used for MDT, one of
the mandatory requirement for these nanoparticles is to exhibit magnetism only in the presence
of an external magnetic field. This requirement should be fulfilled in order to avoid any
undesirable magnetic agglomeration in the absence of magnetic fields. Thus,
superparamagnetic nanoparticles are chosen over the ferromagnetic nanoparticles. Many
attempts have been made recently to address the achievability of magnetic drug targeting13.
Nowicka and co-workers fabricated a doxorubicin-iron oxide conjugate and applied a
magnetic field to guarantee its efficient delivery to the desired human urinary bladder
carcinoma cell line14. Pourmehran et al.15, 16 used discrete phase modelling and one-way
coupling of particle–fluid phases to simulate the air flow and magnetic particle deposition and
tracking in the presence of an external non-uniform magnetic field in human lung model. Their
computational results showed that MDT technique is promising and has a direct relation with
increasing the particle diameter. Lunnoo et al.17 simulated and investigated the parameters,
which govern the capture efficiency of the drug carriers by MDT in mimicked arterial flow.
Their study concluded that the capture efficiency of small particles decreased with decreasing
size and was less than 5% for magnetic particles in the superparamagnetic regime. Though the
thickness of non-magnetic coating materials did not significantly influence the capture
efficiency, it was difficult to capture smaller nanoparticles (D<200 nm) in the arterial flow.
These results provided new insight into the mechanism and governing parameters of the MDT
for clinical settings. Arias et al.18 fabricated Fe3O4/chitosan nanocomposites, which not only
enhanced the intravenous delivery of gemcitabine to the cancer tissue but was also seen to be
hyperthermia responsive in an alternating magnetic field. The in vivo proof of concept, using
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Prussian blue staining, further confirmed the magnetic targeting capabilities of this
magnetite/chitosan core/shell nanocomposite. On the other hand, Hornung and co-workers
employed Fe3O4 nanoparticles to carry and deliver the cytotoxic drug mitoxantrone to human
colon carcinoma cells19. They developed three-dimensional multicellular tumour spheroids
from the colorectal cancer cells and studied the accumulation and delivery efficiency of these
nanoparticles. In an entirely different approach, Kempe et al.20 demonstrated the successful use
of tissue plasminogen activator (tPA) modified Fe3O4 nanoparticles in thrombolytic therapy by
employing magnetic drug targeting. Alexiou and co-workers found quite interesting and
encouraging results when they employed the MDT in the local accumulation of mitoxantrone
in the squamous cell carcinoma implanted in New Zealand White rabbits21. They administered
their nanoparticles in the femoral artery while an external magnetic field was applied on the
tumour. They found a significant (p < 0.05), complete, and permanent remission of the
squamous cell carcinoma as compared to the control group of animals. They successfully
combined intratumoural nanoparticle accumulation and locoregional cancer treatment without
systemic toxicity.
In the studies elaborating the application of MDT, the researchers have directly loaded
the therapeutic molecule on to the nanoparticle surface and administered in the artery supplying
the tumour for better accumulation. Though a variety of magnetic nanosystems are being
explored, dendrimer functionalised magnetic nanoparticles have not yet been evaluated as a
platform for MDT. The following sections thus, evaluates the performance of delivery of DOX-
loaded dendritic nanoparticles by magnetic drug targeting to the syngeneic melanoma mouse
model in C57BL/6 black mice. This chapter explores the bio-distribution and biocompatibility
of the dendritic nanoparticles, while assessing the therapeutic potential of DOX-loaded
nanoparticles in the presence and absence of external magnetic field.
5.2 Experimental and Characterisation Techniques
5.2.1 In vitro Evaluation
The biocompatibility of PAMAM–IO and KR2–IO was assessed with murine
melanoma (B16F10; skin cancer). The melanoma cells were grown in a 96-well microplate for
24 h prior to addition of the nanoparticles. The formulations were suspended in an appropriate
growth medium and serially diluted. 200 µl of this suspension was used to replace the spent
media in the wells, and the cells were allowed to grow for an additional 24 h and 48 h in the
101
presence of the nanoparticles in humidified CO2 environment at 37 °C. To calculate the
therapeutic concentrations of DOX-PAMAM–IO and DOX-KR2–IO, a dose-dependent
evaluation was undertaken over 24 and 48 h. After the set time period (24 and 48 h), the MTS
colorimetric assay (CellTiter 96® Cell Proliferation Assay, Promega) was performed according
to the manufacturer’s protocol. The percentage viable cell population was estimated and
represented against control cells according to equation 2.3. The inhibitory concentration (IC50)
values of each of the nanoparticles was calculated by dose-responsive sigmoidal curve fitting
using Origin 9.1 software.
The effect of the magnetic field on the internalisation of these nanoparticles was
assessed. Towards this end, 1 × 106 cells were grown in 35 mm petridish for 24 h under
physiological conditions. After 24 h, the spent media was replaced by a suspension of
PAMAM–IO and KR2–IO nanoparticles (in the growth medium) of concentration 1 mg/ml.
These petridishes were then placed on a universal magnetic plate and the cells were incubated
in presence of magnetic field for 3, 6, 12 and 24 h. Another set of treated cells were grown for
similar time periods but without the magnetic plate. The untreated cells, unexposed to magnetic
field, were used as control. After the specific time periods, the cells were washed with PBS
(multiple times), trypsinised, centrifuged and the cell pellet was collected. The pellet was then
digested using 1 ml of conc. hydrochloric acid for 30 min. The volume of the cell lysate was
then made up to 10 ml with ultrapure water and the amount of internalised nanoparticles was
determined by the inductively coupled plasma-atomic emission spectroscopy (ICP-AES)
analysis. It should be noted that while rinsing the cells with PBS, some amount of nanoparticles
remains adhered to the cell membrane and cannot be washed out completely. These
nanoparticles are erroneous but still contribute to the final iron content calculated by ICP-AES.
5.2.2 In vivo Therapeutic Efficacy Studies
The animal studies were conducted in compliance with the Institutional Animal Ethical
Committee guidelines. 4-6 weeks old C57BL/6 mice were purchased, maintained and
experiments were carried out at National Toxicology Centre (NTC, Pune, India). Animals were
fed with food & water ad libitum and housed in a constant temperature and humidity conditions
with 12 h light and dark cycles. C57BL/6 black male mice (4-6 weeks’ age) were used for the
biocompatibility and bio-distribution study of the PAMAM–IO and KR2–IO (samples
described in earlier chapters). In vivo tumour inhibition studies were carried out in syngeneic
allograft of melanoma tumour model grown in the same species of black mice.
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5.2.2.1 Biocompatibility and Bio-distribution Studies
The animals were divided into three major experimental groups; Control, Sample I
(PAMAM–IO) and Sample II (KR2–IO). The detailed information about the grouping of
animals, dosing and sacrifice is listed in table 5.1.
Table 5.1 Animal denomination, dosage of dendritic nanoparticles and time points of sample
collection for biocompatibility and biodistribution studies (h=hour; d=day)
Grouping of animals Number of
animals
Time of
sample
collection
Group Sub-group
PAMAM–IO I 3 24 h
II 3 48 h
III 3 7 d
IV 3 14 d
KR2–IO V 3 24 h
VI 3 48 h
VII 3 7 d
VIII 3 14 d
Control IX 3 24 h
X 3 14 d
The dose for the biocompatibility and biodistribution studies were determined based on the
existing literature1 (15 mg/kg body weight). 200 µl of this concentration was administered
through the tail vein using a 27-gauge needle. The animals of the control group were injected
with 200 µl of sterile PBS. The animals were euthanised by cervical dislocation after the set
time points post dose-administration. Their tissues (heart, kidneys, lungs, liver, spleen, brain,
stomach and thigh muscles from right hind leg) were collected, weighed and stored at -20ºC
for further analysis and their blood was collected for immediate analysis for the biochemical
and haematological parameters. A complete analysis of the blood was performed (Mindray
BC2800 Vet Auto hematology Analyser) which included examination of white blood cells
(WBC), lymphocytes (LYM), monocytes (MON), granulocytes (GRAN), red blood cells
(RBC), haemoglobin (Hb), haematocrit (HCT), mean corpuscular volume (MCV), mean
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corpuscular haemoglobin (MCH), mean corpuscular haemoglobin concentration (MCHC) and
platelets (PLT). The activity of enzymes serum glutamic pyruvic transaminase (SGPT), serum
glutamic oxaloacetic transaminase (SGOT) and alkaline phosphatase (ALP) were analysed
from the plasma using aSmart-7 Semi auto Chemistry analyser, Pathozyme. The levels of
creatinine and blood urea nitrogen (BUN) were also evaluated. The activity of the above
mentioned enzymes, creatinine and BUN reflect on the activity and health of the liver, which
is a direct indicator of the biocompatibility of these nanoparticles. For sample preparation for
ICP-AES, all the tissues were dried at 55-60 °C, their dry organ weight was recorded, and
crushed to obtain powder. To a known amount of tissue powder, conc. hydrochloric acid was
added and kept overnight for complete dissolution and used for the estimation of their iron
content through ICP-AES. The iron accumulation in the vital organs were represented per mg
corresponding dry organ weight.
5.2.2.2 Tumour Regression and Therapeutic Efficacy Studies
Development of animal tumour model: The mice bearing subcutaneous melanoma were used
as the model. To male mice (4-6 weeks), 1×106 B16F10 melanoma cells (suspended in sterile
saline solution) were subcutaneously injected on the right flank and observed for tumour
growth. On the 12th – 15th day of inoculation, a palpable tumour appeared and the animals with
uniform tumour volume (~ 100 mm3) were selected for further studies.
Anti-tumour Efficacy and Magnetic Drug Targeting: Our earlier studies established that
KR2–IO nanoparticles showed superior performances over PAMAM–IO nanoparticles.
Henceforth, DOX-loaded KR2–IO nanoparticles were assessed for in vivo therapeutic efficacy
studies. The effect of magnetically guided passive targeting was evaluated in terms of
nanoparticle accumulation and corresponding tumour regression. The tumour bearing mice (on
15th day of tumour inoculation) were randomly divided into four groups with three animals
each. The details about the grouping of animals, dosage and magnetic exposure are listed in
table 5.2.
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Table 5.2 Animal denomination, dosage of dendritic nanoparticles and magnetic exposure for
therapeutic efficacy studies (h=hour; MF=magnetic field)
Group Sample Number of
animals Dose
Magnetic
exposure
I Control 3 Sterile PBS -
II Pure DOX 3
10 mg of
DOX per kg
body weight
-
III DOX-loaded
KR2–IO 3 -
IV DOX-loaded
KR2–IO + MF 3
6 h after every
dosage
Group I mice served as control and were injected with sterile PBS, group II with pure
DOX, group III with DOX-loaded KR2–IO and group IV with DOX-loaded KR2–IO in
presence of a neodymium-iron-boron based permanent magnet (250 mT for 6 h). The samples
were administered in the tail vein of the animals at a DOX concentration of 10 mg/kg body
weight22, 23 and a half dose was administered every 6th day. The animals of group IV were
exposed to the magnetic field along with every dosage. After every third day, from the start of
the study, the tumour dimensions were measured using a digital Vernier caliper and the tumour
volume was calculated according to equation 5.124.
𝑇𝑢𝑚𝑜𝑢𝑟 𝑠𝑖𝑧𝑒 (𝑚𝑚3) = 𝜋
6 × 𝑚𝑎𝑗𝑜𝑟 𝑑𝑖𝑚𝑒𝑛𝑠𝑖𝑜𝑛 × 𝑚𝑖𝑛𝑜𝑟 𝑑𝑖𝑚𝑒𝑛𝑠𝑖𝑜𝑛2 (Eqn. 5.1)
It is noteworthy that caliper measurements of the subcutaneous allograft are superficial and are
affected by the contribution from the epidermis and underlying adipose tissue, as well as fur,
each of which adds some error and inconsistency in the volume calculations.
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Figure 5.1 Melanoma tumour bearing C57BL/6 mouse being administered intravenously
with DOX-loaded KR2–IO nanoparticles. The skin around the tumour region was shaved and
neodymium magnet was stuck using a medical tape for 6 h.
The treated animals were then kept under observation till all the animals of control
(group I) died. The tumour size of the animals of group II, III and IV was measured on regular
intervals and tumour regression/inhibition, if any, was duly recorded. The animals were
subjected to euthanasia by cervical dislocation and tissues (heart, kidneys, lungs, liver, spleen,
brain and tumour) were collected and stored for analyses.
Histopathological examination was carried out of the collected tissues in order to
confirm the iron accumulation visually. The tissue pieces from the vital organs, of suitable size,
from all the experimental animals were collected group-wise in 10% formalin in labelled plastic
container for fixation immediately after necropsy. The collected tissues were then processed
for histological procedure using alcohol-xylene protocol for 24 h in order to dehydrate and
clear the tissues. The tissues were initially placed in 50% absolute alcohol for 4 to 6 h for
removal of fixative, then in ascending grades of alcohol for 2 h each, and finally kept in
absolute alcohol for 2 h. After clearing with xylene, the tissues were embedded in paraffin
blocks, sectioned at 4–5 µm thickness using automated microtome (Leica, Germany). The
tissues’ sections were stained by standard Prussian Blue staining protocol and fixed for
observation. All the fixed slides were examined under a Binocular Light Microscope
106
(Olympus, Japan) and the images were captured using a digital camera (Sony, Japan).
Histopathological evaluation was carried out by an expert qualified veterinary pathologist, who
was blinded to the groups. The iron estimation in the dried tissue powder was also undertaken
as described earlier.
Statistical Analyses: All the results were expressed as mean ± standard deviation. Statistical
significance of the results was computed by two sample t-test using Origin 9.1 software and
p<0.05 confidence intervals were considered significant.
5.3 Results and Discussion
5.3.1 In vitro Evaluation
The PAMAM–IO, and KR2–IO nanoparticles were seen to be non-toxic as the cell
proliferation activity and viable cell population remained unaffected (Figure 5.2 a, b). After 48
h of incubation, the higher concentrations of the PAMAM–IO nanoparticles showed meagre
loss in percent viable cells (~10%) while KR2–IO nanoparticles essentially maintained the
same viable cell population as control. The melanoma cells showed neither change in the
cellular morphology nor any other loss of structural integrity. Therefore, it was concluded that
these nanoparticles were safe to be used for in vivo studies. The DOX-loaded dendritic
nanoparticles are competent to release DOX which efficiently reduced the population of the
melanoma cells (Figure 5.2 c). The IC50 values of DOX-PAMAM–IO and DOX-KR2–IO are
summarised in table 5.3.
Figure 5.2 (a, b) Biocompatibility profile of PAMAM–IO and KR2–IO with B16F10
melanoma cells after 24 and 48 h respectively (c) Dose-dependent cell viability profile of
DOX-loaded dendritic Fe3O4 nanoparticles after 24 and 48 h. Their IC50 values showed
significant difference for 24 h and 48 h (p < 0.05) while there was no significance between
the DOX-PAMAM–IO and DOX-KR2–IO at same time point (p > 0.05).
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Table 5.3 IC50 values of DOX-PAMAM–IO and DOX-KR2–IO nanoparticles with B16F10
melanoma cells. The values are represented in mg/ml of formulations with ≤ 15% (IC50 of
pure DOX was 0.11 – 0.17 μM25)
24 h 48 h
DOX-PAMAM–IO 2.4±1.9 (R2 = 0.999) 0.7±0.1 (R2 = 0.992)
DOX-KR2–IO 3.2±1.5 (R2 = 0.999) 0.6±0.2 (R2 = 0.955)
These dendritic Fe3O4 are seen to be successfully internalised by the melanoma cells (Figure
5.3). The positively charged surface of the dendrimers present on these nanoparticles play a
very important role in facilitating their uptake. These nanoparticles interact electrostatically
with the negatively charged cell membrane and compromise its structural integrity which forms
the passage for their internalisation. Figure 5.3a schematically illustrates the concept of
magnetically-guided cellular internalization of nanoparticles. Figure 5.3b shows confocal
micrographs (A, C) of the melanoma cells and their corresponding fluorescence profile (B, D).
The green fluorescence of FITC-tagged dendritic Fe3O4 nanoparticles is seen to be evenly
distributed throughout the cytoplasm while blue fluorescence marked the nucleus. Though both
of the dendritic Fe3O4 nanoparticles are positively charged, KR2–IO show better uptake than
the PAMAM–IO. This behaviour could be attributed to the resonating amine groups of arginine
on the surface of KR2–IO which makes it more positively charged than PAMAM–IO (discussed
in chapter 4B).
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Figure 5.3 (a) Schematic representation of magnetically-guided cellular internalization of
dendritic Fe3O4 nanoparticles, (b) (A, C) Confocal micrographs and (B, D) fluorescence
profiles of PAMAM–IO and KR2–IO nanoparticles internalized by B16F10 melanoma cells
after 24 h (c) Amounts of iron internalized by the melanoma cells estimated using atomic
emission spectroscopy
This enhanced cellular uptake is further improved by the magnetic field, which aids in better
adherence of these nanoparticles on the cell membrane. The effect of presence of magnetic
field was evaluated on the amount of nanoparticles taken up by the melanoma cells as a function
of the magnetic field exposure time (Figure 5.3 c). The AES analyses show that control cells
have an iron content of 0.0625±2.24 µg and the internalisation of both the dendritic
nanoparticles increased the iron content of the cells substantially. The amount of iron was
represented as µg per mg of dendritic Fe3O4 nanoparticles. After 3h of incubation, there is no
significant effect of the magnetic field in the iron content of the cells. Only after 6 h of
incubation, a substantial increase in iron content is observed, i.e. ~33 µg/mg for PAMAM-IO
and ~38 µg/mg for KR2-IO, in the cells exposed to the magnetic field. When compared between
the formulations with and without MF, 6 h of incubation on the magnetic plate shows ~7-fold
increase in the amount of iron internalised by the cells which reduced to ~3-fold increase at 12
h and further decreased to ~1.6-fold after 24 h in comparison to the control cells. This is
indicative of the fact that magnetically-guided uptake of the nanoparticles could only be
109
enhanced to a certain limit which is dependent on the time for which the magnetic field is
applied. This result proved to be the pivotal basis of our in vivo magnetic drug targeting study
in the mouse tumour model.
5.3.2 In vivo Therapeutic Efficacy Studies
5.3.2.1 Biocompatibility and Bio-distribution Studies
Haematological Parameters: A complete blood evaluation (CBC) were performed to study
the variations, if any, after the intravenous sample administration. Figure 5.4 shows changes
observed in counts of red blood cells (RBC), haemoglobin (Hb), platelets (PLT), white blood
cells (WBC), lymphocytes (LYM), monocytes (MON) while figure 5.5 depicts the summarised
values for mean corpuscular volume (MCV), mean corpuscular haemoglobin (MCH) and mean
corpuscular haemoglobin concentration (MCHC). The levels of RBC and haemoglobin were
minimally affected by the nanoparticles suggesting their haemocompatibility. The platelets’
count in the blood is a very critical parameter as these cells are responsible for haemostasis.
PAMAM–IO caused a sudden drop of platelets’ level in the first 24 h (p<0.01) which regained
its normal level within a week and is maintained thereafter. On the other hand, KR2–IO did not
show any significant effect on the count of platelets through the period of 14 days post-
administration. The introduction of these nanoparticles in the systemic circulation generates an
immune response, which is indicated by the elevated levels of white blood cells. This immune
response triggers specific cells (lymphocytes and monocytes) in blood which are responsible
for the clearing of the nanoparticles from the blood stream and take them to organs like liver,
lungs and spleen to be degraded and cleared out of the body. The levels of WBCs take
approximately 7 days to return to their normal levels, which is also in congruence with the iron
estimation analyses discussed later in this chapter. The administration of PAMAM–IO and
KR2–IO did not cause any significant variation in the MCV, MCH and MCHC.
110
Figure 5.4 Histograms depicting variations in blood parameters for (a) red blood cells, (b)
haemoglobin, (c) Platelets, (d) white blood cells, (e) lymphocytes and (f) monocytes. The
study showed variations across 14 days of study after intravenous administration of dendritic
Fe3O4 nanoparticles. The results are expressed as mean ± s.d. (n=3) with statistical
significance *p < 0.05 and **p < 0.01 with respect to control animals.
Figure 5.5 Histograms showing (a) mean corpuscular volume, (b) mean corpuscular
haemoglobin and (c) mean corpuscular haemoglobin. The results are expressed as mean ± s.d.
(n=3) with statistical significance *p < 0.05 and **p < 0.01 with respect to control animals.
Serum Biochemistry: After intravenous administration of dendritic Fe3O4 nanoparticles to the
mice, the serum was analysed for the biochemical parameters. The activity of enzymes SGPT,
SGOT and ALP were evaluated over the course of the study (Figure 5.6 a,b,c). These enzymes
are suggestive of healthy liver functions. Any fluctuations in their levels primarily indicate the
inflammatory response of the body towards a foreign moiety, such as nanoparticles. Both
PAMAM–IO and KR2–IO cause an instant elevation in the levels of SGPT and SGOT (p <
111
0.05) for the initial 24 h and resume their normal levels by the 7th day. This behaviour is also
suggestive of little or no toxicity of these nanoparticles to organs and regular bodily functions.
Another important parameter is the level of creatinine and urea in the blood, which is an
indicator of healthy renal function. No significant changes were observed in the creatinine
levels of the treated animals compared to the control (Figure 5.6 d). On the other hand, slightly
elevated levels of blood urea were observed for a period of 48 h (Figure 5.6 e) which declined
during the course of next 7 days and returned to their physiological levels (p<0.005). These
results clearly indicated healthy renal function in the mice after treatment with nanoparticles.
The results directly suggest the biocompatibility of both PAMAM-IO as well KR2-IO
nanoparticles.
Figure 5.6 Histograms depicting variations in serum biochemical parameters for (a) SGPT,
(b) SGOT, (c) ALP, (d) Creatinine and (e) blood urea nitrogen (over different time points
during the 14-day study period). The results are expressed as mean ± s.d. (n=3) with
statistical significance *p < 0.05 and **p < 0.01 with respect to control animals.
Bio-distribution assessment by iron estimation: The bio-distribution of dendritic Fe3O4
nanoparticles, administered at 10 mg per kg body weight, was evaluated after intravenous
injection in various vital organs (heart, kidneys, lungs, liver, spleen, brain, stomach and thigh
muscles). Post intravenous administration these dendritic Fe3O4 nanoparticles are freely
available to every organ in the body. Analysis and estimation of iron in the vital organs could
be a direct indication of the bio-distribution of these nanoparticles. A time-dependent
estimation of iron was carried out and is summarised in figure 5.7. Except lungs, liver and
112
spleen, the level of iron in all the tested organs was very low (<10 µg/mg of dry organ weight),
suggesting non-specific uptake in these organs. Lungs, liver and spleen show substantially high
level of iron (> 20 µg/mg of dry organ weight) for the first 24 h which gradually declined over
the next 13 days attaining their least value (~5 µg/mg of dry organ weight). This trend of high
iron level in these vital organs provided a better understanding and suggested that these organs
are responsible for biodegradation of the nanoparticles with the passage of time. This suggests
that these organs play a very important role in degradation and clearance of the nanoparticles.
The macrophages and Kupffer cells are responsible for uptake of the nanoparticles by active
phagocytosis in the lungs and liver, which is responsible for their rapid clearance from the
systemic circulation and accumulation in the organs. The size of the nanoparticles plays a very
important role in their phagocytosis and clearance. Ultra-small nanoparticles (<20 nm) have
been known to evade these cells resulting in a longer circulation time as compared to their
bigger counterparts1. The brain exhibited an extremely low accumulation of iron. This could
be due to the blood brain barrier (BBB), which prevents the accumulation of unwanted foreign
matter in the brain. It should also be noted that a meagre amount of iron extravasated to the
thigh muscles extracted from the right flank of the mice. This trend suggests that the non-
specific uptake of dendritic Fe3O4 by the thigh muscles is negligible and displays a major
difference when compared to the magnetically-guided accumulation (discussed later). By the
end of 30th day, the residual iron content of all the analyzed organs was insignificant (<2 µg/mg
of organ) and similar to control animals which confirmed the biodegradability and clearance
of both our dendritic Fe3O4 nanoparticles. Figure 5.7c depicts the variations in the body weight
of the mice, and shows negligible changes in the body weight post-administration. From the
above results, it could be concluded that both of these dendritic Fe3O4 nanoparticles are
completely compatible and safe to be used with mice and do not have any detrimental effects
on the body.
113
Figure 5.7 Biodistribution of (a) PAMAM–IO and (b) KR2–IO nanoparticles through
quantification of iron accumulated in various vital organs represented at different time points
post intravenous administration. The quantified iron is represented as mean ± s.d. after
control values were deducted from the plot. (c) Body weight profiles of mice across 14 days
of study. The changes in weight of animals are non-significant (p > 0.05) and does not show
any critical changes.
5.3.2.2 Tumour Regression and Therapeutic Efficacy Studies
Tumour regression and Magnetic Drug Targeting: The in vivo bio-distribution of the
PAMAM–IO and KR2–IO nanoparticles indicated that these nanoparticles are biocompatible
and could be successfully used as drug delivery vectors. Towards this end, the DOX-loaded
KR2–IO nanoparticles were assessed for in vivo anti-tumour therapeutic efficacy using a
syngeneic melanoma model in C57BL/6 mice. The administration of free DOX in group II
mice was seen to reduce the tumour size in comparison to the control animals (p<0.05) but also
affected the body weight after 14 days’ post-administration (Figure 5.8a). This result was in
accordance to the fact that when the free drug is administered, it is non-specifically taken up
not only by the tumour cells but also by all the other vital organs and thus, have unhindered
detrimental side effects on the bodily functions. DOX-loaded KR2–IO nanoparticles also
successfully delivered DOX to the tumour region which resulted in significant tumour growth
inhibition when compared to the untreated control animals (Figure 5.8b). In the animals of
group III, the tumor size remained static till 14th day of the treatment (tumor growth inhibition)
and only showed insignificant growth by the end of 35th day. Even after significant tumor
regression is observed and 35 days of treatment, the tumor is still tangible in the animals of
group III. The rate of survival of these animals was much higher than control and free DOX
114
treated animals as 60% of animals (group III) were alive after 60 days with a small tangible
tumor.
Figure 5.8 (a) Body weight profiles of mice across 35 days of study. The changes in weight
of control animals are due to the uncontrolled tumour size. The weight of animals of group II,
III and IV are significantly decreased (*p < 0.05) in comparison to the control animals, (b)
tumour volume profiles of untreated control animals against the treated groups II, III and IV.
Tumour growth was significantly inhibited with DOX-loaded KR2–IO nanoparticles as
against animals treated with free DOX. (c) Quantification of iron accumulated in various vital
organs. In absence of magnetic field, the thigh muscles (tumour) tend to accumulate small
amount of iron which is significantly elevated with application of magnetic field.
The accumulation of iron in the tumour is observed in the animals of group III where
the amount of iron was found to be below 25 µg/mg dry organ weight (Figure 5.8 c). This
meagre increase could be attributed to the nanoparticle accumulation due to leaky tumour
vasculature. On the other hand, the application of magnetic field around the tumour region in
animals of group IV, caused a substantial increase in the amount of nanoparticles (161 µg of
Fe/mg dry organ weight) in the tumour. The application of magnetic field attracts the
nanoparticles from the blood stream to the tumour region and the accumulation is further
facilitated by the leaky vasculature at the tumour site. This elevated level of nanoparticles in
the tumour leads to high concentration of DOX released which in turn improves the tumour
inhibition efficacy of our DOX-loaded KR2-IO nanoparticles (Figure 5.8 c). Therefore, the rate
of tumour size reduction and growth regression was seen highest in the mice of group IV. By
the end of second dosing (14th day), the average tumour volume was 55 ± 8.3 mm3 as compared
to the control animals in which the tumour volume was seen to be 4794 ± 844 mm3 (~ 88-fold
decrease). The tumour was seen to disappear by the end of 20th day post-treatment but the mice
were kept under observation for evaluation of their survival and tumour re-lapse, if any. The
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mice of all the groups were kept under continuous observation until the control mice died which
marked the termination point of the study. We observed death of only one animal from the
treated group IV in entire study period on 37th day while,~100% survival rate was observed
with no tumor re-lapse even after 60 days.
Figure 5.9 shows the histopathological images of the tissue sections with Prussian blue
staining. The stained slides did not exhibit any signs of toxicity in terms of tissue damage and
structural integrity in the organs viz. kidney, lungs, heart, brain, spleen and liver. Further, the
tumor sections of the treated group without the magnetic field exhibited very less nanoparticle
accumulation against the group exposed to the magnetic field along with the nanoparticles.
These results further support the magnetically-guided nanoparticle accumulation data and
tumor regression, confirming the improved anti-tumor potential of peptide dendrimer-based
Fe3O4 nanoparticles.
Figure 5.9 Histological analyses of the excised organs after 24 h. Optical micrographs of the
treated tumor exhibited zones with compromised cellular cohesion and structural integrity
with significantly different iron content in absence and presence of magnetic field (MF).
Other vital organs showed minimal accumulation of iron causing little or no toxicity to
normal tissues and bodily functions.
5.4 Summary
After validating the in vitro efficacy of PAMAM–IO and KR2–IO nanoparticles, this
chapter focussed on in vivo investigation of these particles in a subcutaneous syngeneic murine
melanoma model. The systemic exposure of these nanoparticles caused changes in various
blood and serum parameters. This assisted with the information about their toxicity and non-
specific uptake by various organs. The variations in the complete blood counts and the enzymes
(SGPT and SGOT) showed only an initial change which resumed to their physiological levels
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within 7 days of treatment. Both the dendritic nanoparticles showed a tendency to localise in
lungs, liver and spleen which are responsible for their rapid clearance from the systemic
circulation. However, in tumour bearing mice, localisation of nanoparticles is manipulated by
the application of an external magnetic field. We observed ~6-fold higher nanoparticle
localisation in magnetically targeted tumour (right flank) as compared to passive localisation
in thigh muscles (right flank) of non-tumour bearing mice. This elevated levels of iron
accumulation in the tumour is due to the erratic angiogenesis causing leaky vasculature in the
tumour, which is further facilitated by the application of an external magnetic field. This high
localisation pattern of DOX-loaded KR2–IO led to high concentrations of DOX in the tumour
and thus was effective in arresting the tumour growth significantly. It was seen that lower
number of doses are sufficient to suppress the tumour growth in combination with magnetic
field than that required without the magnetic field. This elaborate study confirmed that KR2–
IO nanoparticles not only surpass PAMAM-IO as drug delivery vectors, but can also act
successfully as platforms for magnetic targeting of cancer.
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2015, 44, 8576-8607.
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Rodriguez, R. Lujan Rangel, D. Romanovicz and M. Jose-Yacaman, Metallomics,
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Chemistry, 2004, 15, 1174-1181.
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118
Chapter 6
MR Contrast Properties of Dendritic Fe3O4 Nanoparticles
6.1 Introduction
Magnetic Resonance Imaging (MRI) is a powerful technique used widely in clinical
settings to image interiors of human body. It is a non-invasive, tomographic technique, without
the dangers of ionising radiation that offers good spatial resolution1, 2. MRI is based on the
principles of nuclear magnetic resonance (NMR) which can also provide us with the chemical
and physical information about molecules. The human body comprises of mainly fat and water
molecules, contributing approximately 63% hydrogen atoms, magnetic relaxation of which
generates the NMR signal. When the proton is placed in an external magnetic field, the nuclear
spin vector of the particle aligns itself with the external field (either parallel or antiparallel to
the field), just like any other magnet would. This alignment initiates a precession at a constant
frequency. In a simplistic sense, MRI is just the application of proton NMR to biological
systems, which yields intensity maps of water proton spin relaxation in tissues. Contrast in the
images generated by MRI, is the consequence of proton density and relaxation time constants
that vary throughout the tissues.
However, MRI suffers from relatively low sensitivity that limits its utility as it relies
exclusively upon the above mentioned inherent contrast mechanisms. This flaw can be
addressed using exogenous agents that influence local proton spin relaxation dynamics thereby
119
enhancing the contrast and ability to distinguish the structures within the images. There are
several different types of contrast agents currently in use, and the two main classes are based
on chelated paramagnetic ions, such as gadolinium (Gd)3, 4, and superparamagnetic
nanoparticles like Fe3O45, 6. Particulate superparamagnetic contrast agents offer advantages
over their paramagnetic counterparts in both efficiency and mechanism of action. Henceforth,
Fe3O4 nanoparticles are discussed as an exemplary contrast agent for the superparamagnetic
nanomaterials. Paramagnetic chelates produce local effects that are mediated by exchange of
water protons, but Fe3O4 nanoparticles produce magnetic field gradients which results in the
changes in overall tissue susceptibility thereby affecting the contrast. These magnetic field
gradients impact a larger region without the need for direct contact with water flux. Many
factors make Fe3O4 excellent materials for developing targeted MRI contrast agents. They have
been known to naturally occur in many organisms and are biodegradable, exhibiting neither
acute nor chronic toxicity. These properties make them the only inorganic particulate contrast
agent currently approved for clinical settings. In clinical scenario, the contrast arises due to the
disproportionate accumulation of Fe3O4 nanoparticles resulting from the reduced or eliminated
reticulo-endothelial system (RES) function of pathologic versus healthy tissue. Once these
nanoparticles reach the tumour site, they extravasate through the porous endothelium (leaky
vasculature) and are retained by the tumour due to its compromised clearance mechanisms
(poor drainage). Relaxivity is an extrinsic property of these nanomaterials and is dependent on
a variety of factors7-9. These parameters can be modulated accordingly to improve their MR
performance and quality of the images in the clinical scenarios.
To meet this end, a variety of agents have been evaluated and used to encapsulate the
Fe3O4 nanoparticles to the region of interest for MRI. Jaiswal et al.10 have developed thermo-
responsive polymeric nanohydrogels to encapsulate Fe3O4 nanoparticles to enhance the
contrast properties exhibiting relaxivity rate of 173 mM-1s-1. An et al.11 synthesised Fe-Co
nanoalloy system modified with dextran molecules as sensitive contrast agents for MRI. These
nanoalloy systems had a relaxivity of 17.4 mM-1s-1 and gave a negative contrast in in vitro and
in vivo systems. Mandal and co-workers attempted to develop multimodal nanocarriers capable
of carrying therapeutic load in addition to their MR contrast properties12. They combined FITC
labelled Fe3O4 nanoparticles with antibody and gemcitabine drug to Hep2 cells. Yang et al.13
attempted to develop a dual-contrast agent by conjugating superparamagnetic silica-coated
Fe3O4 core-shell nanoparticles with gadolinium complex. This hybrid material was seen to be
water-dispersible, stable, and biocompatible while displaying both T1- and T2- relaxivities.
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Relaxivity measurements show that they have a T1 relaxivity (r1) of 4.2 mM−1s−1 and T2
relaxivity (r2) of 17.4 mM−1s−1, suggesting a possibility to use them as both T1 positive and T2
negative contrast agents. Fe3O4 nanoparticles have been commercialised as various
formulations, such as Ferrixan® (Bayer Schering Pharma AG) which shows a negative
relaxivity of 150 mM-1s-1 and Feridex® (AMAG Pharmaceuticals, Inc.) having a negative
relaxivity of 160 mM-1s-1.
This chapter deals with the use of the formulations discussed earlier (in second section
of chapter four of this thesis) as potential MR contrast agents. These nanoformulations have
been shown to successfully deliver doxorubicin to cancer cells and useful in magnetic
hyperthermia. To establish the multimodality of these nanoparticles in diagnostic imaging,
their performance as contrast agents in MR imaging was assessed. Also, as the working
environment of these nanoparticles would be hyperthermic (therapeutics), the dependence of
their MR relaxivity was studied under the conditions of varying buffer ions and temperatures.
6.2 Experimental Techniques
The samples prepared in section two of chapter five of this report (PAMAM–IO and
KR2–IO) were used to study and evaluate their comparative MR relaxivity properties. The
samples were suspended in 0.05% agarose gel (in various buffer solutions) and fixed. The
liquid microenvironments used were ultrapure water, phosphate buffer and simulated body
fluid (SBF). The temperatures inside the magnetic bore were maintained at 25, 37 and 45 ±0.5
°C by air blower placed near the bore and monitored digitally. The samples were scanned using
a multi-echo T2-weighted fast spin echo imaging sequence (TR/TE=3500/30, 45, 60, 75, 90
and 105 ms, slice thickness = 2 mm) by a 9.4T Agilent small animal scanner. For in vitro T2-
weighted MR imaging of HeLa cells, the cells were incubated with PAMAM–IO and KR2–IO
in the presence and absence of magnetic field. The magnetic field was applied and maintained
using a universal magnet plate. After incubation, the cells were gently washed by phosphate
buffered saline and fixed in 0.05% agarose gel for MR imaging.
6.3 Results and Discussions
6.3.1 Relaxivity studies
The T2-weighted MR images of PAMAM–IO and KR2–IO at different iron
concentrations were recorded at 9.4T. A significant signal attenuation was seen with increase
121
in the iron concentration of the recorded phantoms. The relaxation rates showed a linear
increase with an increase in iron concentrations. At 25 °C, in ultrapure water, the r2 values of
PAMAM–IO and KR2–IO were comparable and were calculated to be 302 and 321 mM-1s-1
respectively (Figure 7.1). On increasing the temperature to physiological (37 °C) and then to
hyperthermia (45 °C) temperature, a gradual decrease in relaxivity is seen. The values
decreased to 177.4 and 178.5 mM-1s-1 for PAMAM–IO and KR2–IO respectively.
Figure 6.1 Plots of transverse relaxivity (r2) values of (a) PAMAM–IO and (b) KR2–IO in
ultrapure water
A similar trend was observed when these samples were evaluated in phosphate buffered
saline of pH 7.3 (Figure 7.2). At 25 °C, the r2 values of PAMAM-IO and KR2-IO were
comparable and were calculated to be 306 and 326 mM-1s-1 respectively. On increasing the
temperature to 45 °C, the values decreased to 149 and 183 mM-1s-1 for PAMAM-IO and KR2-
IO respectively. The difference between these values suggest that KR2-IO nanoparticles would
show improved performance in buffered environment than in pure water environment. The
relaxivity of these nanoparticles is sufficiently high to provide good MR contrast than the
PAMAM-IO nanoparticles.
122
Figure 6.2 Plots of transverse relaxivity (r2) values of (a) PAMAM-IO and (b) KR2-IO in
phosphate buffered saline (pH 7.3) (c) phantom images showing substantial reduction in
transverse relaxation times of both the nanoparticles
The obtained phantom images showed a significant signal reduction with increasing
concentration of Fe from 0.001 to 0.25 mM (Figure 7.3). At 25 °C, in simulated body fluid
(SBF), the r2 values of PAMAM-IO and KR2-IO were comparable and were calculated to be
284 and 287 mM-1s-1 respectively. On increasing the temperature to physiological 37 °C, the
values decreased to 221 and 213 mM-1s-1 respectively. At 45 °C, a further reduction in
relaxivity is seen. The values now decreased to 134 and 165 mM-1s-1 for PAMAM-IO and KR2-
IO respectively.
123
Figure 6.3 Plots of transverse relaxivity (r2) values of (a) PAMAM-IO and (b) KR2-IO in
simulated body fluid (pH 7.4) (c) phantom images showing substantial reduction in
transverse relaxation times of both the nanoparticles
This decrease in relaxivity with increase in temperature is due to the thermal activation and
movement of the water molecules14. At elevated temperatures the magnetic spins are also
disordered leading to reduction in magnetic strength of the Fe3O4 nanoparticles. This in
accordance with the results observed with PAMAM-IO and KR2-IO. Under the working
conditions of hyperthermia, i.e. 45 °C and in actual physiological conditions, KR2-IO assures
better MR performance than PAMAM-IO nanoparticles.
6.3.2 In vitro T2 weighted imaging
The T2 weighted images reveal that both the dendritic Fe3O4 nanoparticles are efficient
in providing improved contrast for imaging the HeLa cells (Figure 6.4). It was seen that the
presence of magnetic field and the duration of exposure to the magnetic field, during the
incubation, influenced the cellular uptake of these nanoparticles. The exposure of MF for
longer duration caused substantial increase in the amounts of nanoparticles internalised by the
cells (discussed in detail in chapter 5). This led to the observation that larger number of
124
nanoparticles inside the cell provided better MR contrast as compared to the cells with lower
nanoparticle content.
Figure 6.4 T2-weighted MR images of HeLa cells with PAMAM-IO (a,c) and KR2-IO (b,d)
in the absence (a,b) and presence (c,d) of magnetic field
6.4 Summary
This chapter focussed on the evaluation of transverse relaxation times of PAMAM–IO
and KR2–IO nanoparticles and their MR contrast properties. It was seen that the transverse
relaxation time of these nanoparticles was greatly dependent on the ambient temperature. As
these dendritic nanoparticles are intended to be used in hyperthermic environment, their MR
contrast at 45 °C plays a critical role in their MRI applications. A decreasing trend in the
relaxivity values of the nanoparticles was observed when the temperature was increased from
ambient (25 °C) to physiological (37 °C) and then to hyperthermia (45 °C) temperatures. This
observation was attributed to the disordering in the water molecules and the magnetic moments
caused by the elevated temperatures. Despite the reduction, these nanoparticles were expected
to be efficient under physiological environment and high temperatures due to their high
transverse relaxivity. This was also confirmed by the T2 imaging of the cervical cancer cells.
6.5 References
1. M. S. Shiroishi, G. Castellazzi, J. L. Boxerman, F. D'Amore, M. Essig, T. B. Nguyen,
J. M. Provenzale, D. S. Enterline, N. Anzalone, A. Dörfler, À. Rovira, M. Wintermark
and M. Law, Journal of Magnetic Resonance Imaging, 2015, 41, 296-313.
125
2. D. T. Ginat, B. Swearingen, W. Curry, D. Cahill, J. Madsen and P. W. Schaefer, Journal
of Magnetic Resonance Imaging, 2014, 39, 1357-1365.
3. W.-L. Zhang, N. Li, J. Huang, S.-F. Luo, M.-X. Fan, S.-Y. Liu, B. Muir and J.-H. Yu,
Journal of Applied Polymer Science, 2011, 121, 3175-3184.
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2015, 53, 349-357.
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Chemistry C, 2008, 112, 8127-8131.
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– A European Journal, 2002, 8, 1040-1048.
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Physical Chemistry B, 2007, 111, 832-840.
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and V. P. Dravid, ACS Applied Materials & Interfaces, 2014, 6, 6237-6247.
11. L. An, Y. Yu, X. Li, W. Liu, H. Yang, D. Wu and S. Yang, Materials Research Bulletin,
2014, 49, 285-290.
12. A. Mandal, S. Sekar, M. Kanagavel, N. Chandrasekaran, A. Mukherjee and T. P. Sastry,
Biochimica et Biophysica Acta (BBA) - General Subjects, 2013, 1830, 4628-4633.
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Biomaterials, 2011, 32, 4584-4593.
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Journal of materials science. Materials in medicine, 2002, 13, 113-117.
126
Chapter 7
Conclusions and future scope
7.1 Conclusions
The nanoparticle platforms have emerged as promising drug delivery vehicles for
carrying a variety of cargo in cancer therapeutics. Among a vast variety of nanoparticulate
systems, very few could be actively used in biomedical applications because of their poor
aqueous stability, off-site toxicity, poor biocompatibility and toxic degradation profile. Fe3O4
nanoparticles have been thoroughly and critically evaluated for biomedical applications.
Recent works have also established dendrimers as promising nanocarriers as they show
improved properties and efficiencies when compared to the classical polymers. To improve the
aqueous stability of Fe3O4 nanoparticles, a variety of organic molecules are used to engineer
their surface which also plays an important role in linking to the cargo molecules. Towards this
end, we evaluated surface functionalised Fe3O4 nanoparticles for their ability to carry and
deliver doxorubicin to cancer cells. These DOX-loaded nanoparticles (loading efficiency, 90%,
w/w) were stable, biocompatible with good specific absorption rate. They exhibited a pH-
dependent release pattern and released DOX in appreciable amounts in mild acidic
environment. As most of the DOX molecules were bound to the surface of the nanoparticles,
their release was uncontrolled as they were vulnerable to the change in pH in their immediate
environment. This could lead to undesirable, off-site release of DOX. To circumvent this
potential drawback, the Fe3O4 nanoparticles were surface modified with PAMAM dendrimers.
127
The modification of Fe3O4 nanoparticles by different generations of PAMAM (G 3, 5
and 6) improved their material properties which were thoroughly characterised. Their
capability to carry both cationic (DOX) as well as anionic (EGCG) drugs was evaluated. It was
seen that the drug loading capacity of PAMAM-Fe3O4 increased with increase in the generation
of the dendrimer. It could be explained by the fact that as the generation of dendrimer is
increased, the surface functional groups also increase along with the size of their internal
cavities. The cargo molecules are known to be bound either by the surface groups or
encapsulated in the cavities. Thus, the results indicated that higher generations of dendrimer
were more efficient in carrying as well as releasing the drug molecules as compared to the
lower generations. Overall, these nanosystems showed no loss of cell proliferation activity with
the mouse fibroblast (L929), human cervical carcinoma (HeLa), oral carcinoma (KB), prostate
cancer (PC-3) and human breast cancer (MCF-7) cells. But it was noted that the higher
generations of the PAMAM dendrimer showed noticeable amount of toxicity towards cells and
was not entirely biocompatible. PAMAM is the most studied and widely used dendrimer in
various biomedical applications due to its excellent physico-chemical properties and tailorable
architecture. It application in clinical settings is limited by its compromised biocompatibility,
which is a limiting factor and thus raises a demand of improved biocompatibility. Thus, we
synthesised a biocompatible peptide dendrimer that had amide interiors and amine exteriors
(chemically similar to PAMAM). The structural characterization of peptide dendrimer was
carried out by various sophisticated techniques as nuclear magnetic resonance (NMR), FTIR
and X-ray photoelectron spectroscopy (XPS). Since amino acids were used as branching units,
the peptide dendrimer was found to be biocompatible even at higher concentrations (20 mg/ml).
When the performances of as-prepared peptide dendrimer were compared to PAMAM, it was
seen that PAMAM was still more efficient in conjugating with Fe3O4 nanoparticles. The KR2–
IO nanoparticles exhibited similar drug loading efficiency as the PAMAM–IO, however, their
drug release capacity was significantly improved. The comparative evaluation of these
dendritic nanoparticles for drug delivery, magnetic hyperthermia and combinatorial therapy
established that peptide dendrimer have an edge over the PAMAM dendrimers. Also the
physico-chemical properties of peptide dendrimer were minimally compromised as compared
to its PAMAM counterpart.
After validating the in vitro efficacy of PAMAM–IO and KR2–IO nanoparticles, their
in vivo investigation in a subcutaneous syngeneic murine melanoma model was performed. It
was seen that both the dendritic nanoparticles were haemocompatible and had no adverse
128
effects on the normal functioning of the body. Both the dendritic nanoparticles showed a
tendency to localise in lungs, liver and spleen which was responsible for their rapid clearance
from the systemic circulation. These Fe3O4-based nanosystems also have the advantage of
being responsive to external magnetic field while circulating in the bloodstream. This was
exploited as magnetically targeting these nanoparticles to the tumour site. Approximately 60-
fold increase in the accumulation of iron in the tumour region was observed when compared to
the tumour of mice which were not exposed to the magnetic field. This high localisation of
DOX-loaded nanoparticles led to high concentrations of DOX in the tumour and thus the rate
of tumour regression was faster than other mice groups. It was also seen that lower number of
doses were sufficient to suppress the tumour growth in combination with magnetic field than
that required without the magnetic field. Though the tumour was completely supressed by 20th
day, it reappeared after 35 days of treatment. This necessitated the revision of parameters for
improved therapeutic efficacy of these nanoparticles.
Furthermore, the dendrimer functionalized Fe3O4 nanoparticles were also seen to be
MR active and showed high transverse relaxivities. Various physico-chemical parameters
potentiate and affect their MR contrast properties. A decreasing trend in the relaxivity values
of the nanoparticles was observed when the temperature was increased from ambient (25 °C)
to physiological (37 °C) and then to hyperthermia (45 °C) temperatures. Despite this decreasing
behaviour, the relaxivity of KR2–IO in simulated body fluid at 45 °C show optimally high value
of 165 mM-1s-1. In vitro T2 weighted MR imaging of dendrimer functionalized Fe3O4 treated
(for various treatment times) HeLa cells showed increased contrast when compared to the
untreated cells. Owing to the promising drug delivery performances, high magnetisation, high
specific absorption rate and shorter transverse relaxation times, the dendrimer functionalized
Fe3O4 nanoparticles not only prove to be suitable delivery vectors, but can also act as anchors
for passive targeting and contrast agents for MR imaging and multimodal cancer theranostics.
7.2 Future scope
The studies presented in this report demonstrate the potential of peptide dendrimer-
coated Fe3O4 nanoparticles in drug delivery, hyperthermia, combinatorial therapy and in vitro
MR imaging studies against PAMAM coated nanoparticles. This also opens up avenues of
using peptide dendrimer-coated Fe3O4 nanoparticles in a variety of other biomedical
applications.
129
(a) Higher generations of this peptide dendrimers could be evaluated for their biomedical
applications. It is expected that higher generations will prove better vehicles as compared
to the lower generations.
(b) These dendritic nanoparticles could be used in combining chemotherapy and MDT with
magnetic hyperthermia for tumour model in small animals and if possible taken up further
to clinical settings.
(c) Attachment of a targeting ligand or antibody to these formulations might add on to the
specificity and improved efficacy of these formulations in addition to their magnetic
targeting.
(d) Initial MR imaging studies reveal their potential for in vivo tumour imaging in animal
models.
(e) These dendrimer-coated nanoparticles could also be used as platforms for a variety of
other biomedical applications such as cell sorting and separation, immunodetection,
electrochemical sensing towards early detection of cancer etc.
This chapter has been published in J. Nanopharma. Drug Del., 2013, 1, 365-375.
Appendix I
Thermally-activated delivery of curcumin using magnetic
liposomes
AI.1 Introduction
Chemotherapy involves the use of chemicals to damage DNA, RNA and proteins in order
to trigger cell cycle arrest or apoptosis in cancer cells. However, such agents generally induce
apoptosis in both cancer and normal cells1. Hence, it is highly desirable to develop a drug delivery
system that has maximum chemotherapeutic efficiency and minimal side effects. An approach
toward the development of such a drug delivery system is the combining of a chemical that
selectively induce apoptosis in targeted cancer cells, with a delivery vehicle that can release the
drug to the targeted tumour in a controlled manner. Curcumin (diferuloylmethane, a natural,
hydrophobic polyphenol, and the primary constituent of the rhizome of turmeric) is one of the few
agents that selectively induce apoptosis in highly proliferative cells. Cellular apoptosis induced by
curcumin is significantly higher in cancer cells than in non-cancerous cells. Curcumin first
attracted attention for its antioxidant, anti-inflammatory2, antimicrobial3, 4 and antineoplastic
activities5, 6. The efficacy of curcumin as a therapeutic agent in cancer was reported later5, 7, 8.
Curcumin has since been reported to increase apoptotic death, control cell proliferation and down-
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regulate the oncogenic phenotype by controlling the signalling cascades involved in the cell cycle9,
10. However, the low aqueous solubility, poor chemical stability and lipophilic nature of curcumin
have limited its oral bioavailability and delivery efficiency to targeted cancer sites and have, thus,
impaired its anticancer therapeutic potential.
Certain types of polymer nanoparticles have been used in the encapsulation of curcumin in
order to enhance its chemical stability11-15; poly(lactic-co-glycolic) (PLGA) is the most studied
among these polymer nanoparticles. PLGA is a biodegradable polymer which generates lactic and
glycolic acid as degradation products, which are metabolised by cells through the Krebs cycle.
Although PLGA nanospheres as vehicles can enhance the delivery of curcumin when compared to
curcumin alone, the amphiphilic nature of the polymer results in low levels of encapsulation, which
results in a need for the administration of higher dosages in order to achieve pharmaceutical
activity at the targeted site14, 16-18. Saccharide-based nanoparticle vehicles have also been used to
deliver curcumin for oral administration in order to enhance the release efficiency. However, the
fast water solubility of saccharide compromises the chemical stability of the drug system19. In
short, the delivery vehicles mentioned above have failed to address the current issues that are
associated with the curcumin delivery.
Liposomes, which are bilayer vesicles of lipid molecules, have been regarded as promising
alternatives to polymeric systems20-22. The amphiphilic nature of lipid molecules causes them to
form a closed vesicular structure in aqueous solutions, with the apolar regions being oriented away
from the aqueous phase and the polar regions being in contact with the water. The hydrophilic
surface of liposomes renders them highly water soluble, while the amphiphilic nature of the lipid
bilayer facilitates the anchoring of both hydrophilic and lipophilic molecules. A liposomal carrier
that is made up of soy phosphatidylcholine and 1,1-diphenyl-2-picrylhydrazyl has been reported
to enhance the antioxidant (therapeutic) effects of the encapsulated curcumin23. However, it
remains to be elucidated whether the liposome carrier affects the anticancer properties of
curcumin. Therefore, an objective of the present work was to investigate the curcumin-
encapsulation efficiency, curcumin-releasing profile and the anticancer therapeutic effect of
curcumin when delivered by liposomes. Another advantage of liposomes is that they can release
encapsulated drugs under external stimuli such as a changing pH or temperature, or ultrasound
waves, allowing control over the release profile and the pharmacokinetics of the drug molecules.
Among these stimuli, a changing temperature is clinically the most feasible because it is virtually
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impossible to control the pH values in the body, and ultrasound has the potential to damage the
phospholipid membranes of healthy cells while destroying the liposomes. In fact, hyperthermia
treatment is already used in some forms of cancer treatment, especially in solid tumours, in order
to kill cancer cells directly or to make them more sensitive to radiation and certain anticancer
drugs24, 25. There are a number of techniques to deliver heat remotely, which include infrared
sources, focused microwaves, magnetic fields and infusions of warmed liquids26-28. In the case of
magnetic hyperthermia, nanoparticles can be subjected to an alternating magnetic field, which
induces electron flow in the nanomaterial, thereby producing localised heat. To use magnetic
hyperthermia to control the release of curcumin from liposomes was another objective of this work.
Superparamagnetic Fe3O4 nanoparticles have been employed in hyperthermia treatment24,
29, 30, and as drug delivery vectors31, 32 such as curcumin delivery17, 18. When compared with other
ferrite systems (iron, cobalt, manganese and nickel), superparamagnetic Fe3O4 nanoparticles have
shown greater potential due to their better aqueous stability and biocompatibility33. These
nanomaterials also exhibit a relatively large surface-area-to-volume ratio due to their small size.
The downside of superparamagnetic Fe3O4 nanoparticles is that they have a high surface energy,
which leads to a strong tendency to agglomerate in colloidal suspensions via van der Waals forces
and magnetic dipole–dipole interactions. This agglomeration necessitates colloidal stability of
these nanoparticles in aqueous suspensions for the biomedical applications mentioned above.
Surface engineering of Fe3O4 nanoparticles for biomedical applications has also been performed
using a variety of organic molecules34, 35, in addition to polymers36, 37 and polysaccharides38, 39.
This results in bioactive surfaces that can be used to anchor different molecules of interest, which
will enhance their colloidal and surface chemistry35. Small organic molecules of biological origin
have been considered to be promising candidates for nanoparticle surface engineering because
their biocompatible degradation products minimise the risk of toxicity to biological systems.
Saccharides and oligosaccharides, for example, readily degrade into sugar monomers, which are
actively metabolised by cells. Due to their colloidal solubility, amino acids40, monosaccharides38,
and oligosaccharides39 are three of the most frequently used small organic molecules.
Oligosaccharides such as dextran and β-cyclodextrin, have been reported to significantly improve
the colloidal stability of nanoparticles41, 42. Monosaccharides are also well known for their ease of
conjugation with various biomolecules such as proteins, which forms peptidoglycans and
proteoglycans, or with lipids, which forms glycolipids and aminoglycans. Given the above
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properties, dextrin was selected as the surface engineering moiety for the Fe3O4 nanoparticles in
the present study.
After being surface engineered with small organic molecules, magnetic nanoparticles can
be further modified by various macromolecules43, polymers44 and biomolecules45 in order to
improve their recognition in biological systems and to lower their immunogenicity46. In this
respect, lipids47, 48 have an advantage over other macromolecules because they can aid in cellular
uptake of nanoparticles by facilitating transport across the phospholipid membranes of cells.
Fabrication of magnetic liposomes (MLs) by coating magnetic nanoparticles with a lipid bilayer
has been explored for the possible delivery of a number of drugs, except for curcumin45. In
summary, the objectives of this chapter are to fabricate and characterise an ML formulation, to
evaluate its curcumin encapsulation and delivery performance, and finally, to assess the anticancer
effects of the ML/curcumin formulation within cancer cells.
AI.2 Experimental Techniques
AI.2.1 Synthesis of Fe3O4 Nanoparticles and Magnetic Liposomes
To synthesise the Fe3O4 nanoparticles, 4.44 g of ferric chloride and 1.732 g of ferrous
chloride were dissolved in 80 ml of ultrapure water, and the temperature was increased to 60 °C
under a nitrogen atmosphere with mechanical stirring at 1000 rpm. The temperature was then held
at 70 °C for 30 min, which was followed by the addition of 30 ml of ammonia solution and the
maintenance of the temperature at 70 °C for another 30 min. Finally, 10 ml of an aqueous solution
of dextrin (0.07 g/ml) was added to the reaction mixture and the temperature was raised to 90–95
°C under reflux and maintained for 90 min with continuous stirring. A black precipitate of dextrin-
coated Fe3O4 nanoparticles (Dx–Fe3O4) was obtained and was washed thoroughly with ultrapure
water. During each wash step, the precipitate was separated from the supernatant through the use
of a permanent magnet.
The MLs were prepared by a thin-film hydration technique that was reported previously45.
In a typical synthesis, 200 mg of soy-PC was dissolved in a solvent mixture of
chloroform/methanol (2:1 v/v). The solvent was evaporated under vacuum (120 mbar) at 35 °C in
a rotary evaporator in order to form a thin and uniform lipid film on the walls of the round-
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bottomed flask. In order to optimise the lipid-to-nanoparticle ratio, the lipid film was hydrated
with varying amounts of Dx–Fe3O4 nanoparticles (20, 40, 60, 80, 100, 150 and 200 µg) using a
water bath-type sonicator for approximately 30 min. The temperature was maintained below 40
°C until the lipid film was transferred to the aqueous suspension, yielding the MLs. Control
liposomes were synthesised by hydrating the lipid film in PBS (pH 7.34). A lipid-to-nanoparticle
ratio of 10:3 was used for the drug delivery vesicle because it showed the maximum encapsulation
efficiency of the Dx–Fe3O4 nanoparticles and the most stable aqueous suspension.
AI.2.2 Analysis of Drug Loading and Release
The absorbance spectra of curcumin were studied in order to investigate the loading
efficiency of curcumin. In a typical loading experiment, 200 mg of soy-PC was dissolved in a
solvent of chloroform/methanol (2:1 v/v), which contained varying amounts (20, 40, 60, 80, and
100 µl) of curcumin (stock solution - 1 mg/ml in chloroform) and had ML-to-curcumin ratios of
10, 5, 3.33, 2.5 and 2. The solvent was evaporated under vacuum (120 mbar) at 35 °C in a rotary
evaporator in order to form a thin and uniform lipid film on the walls of the round-bottomed flask.
The lipid film was then hydrated with a previously optimised amount (lipid-to-nanoparticle
ratio=10:3) of Dx–Fe3O4 nanoparticles under ultrasonication for approximately 30 min. The
temperature was kept below 40 °C until the lipid film was transferred to the aqueous suspension
to yield the drug-loaded MLs. The absorbance spectrum of the supernatant (after magnetic
sedimentation of the curcumin-loaded MLs) was recorded with a UV-Vis spectrophotometer in
order to determine the amount of the drug loaded into the MLs. The curcumin-loaded MLs were
re-suspended in water and the loading efficiency (w/w%) was calculated as follows:
% 𝐿𝑜𝑎𝑑𝑖𝑛𝑔 𝑒𝑓𝑓𝑖𝑐𝑖𝑒𝑛𝑐𝑦 = 𝐴𝐶𝑢𝑟− 𝐴𝑆𝑢𝑝
𝐴𝐶𝑢𝑟 × 100 (Eqn. AI.1)
where ACur is the absorbance of pure curcumin and ASup is the absorbance of the supernatant.
The drug release study was carried out at temperatures of 37 °C and 45 °C. The amount of
curcumin released from loaded MLs was quantified on the basis of the loading efficiency. The
curcumin-loaded MLs (10 mg) were suspended in PBS at pH 7.3 and placed in a dialysis bag.
Dialysis was performed with 200 ml of PBS (pH 7.3) under continuous stirring at a physiological
temperature (37 °C) and at a hyperthermic temperature (45 °C). At different time intervals, aliquots
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were withdrawn and replaced with fresh PBS. The fluorescence intensity of these aliquots was
analysed at λex= 425 nm and λem=540 nm. Drug release was calculated from the fluorescence
intensity of the aliquots according to a standard curve prepared under similar conditions. The
thermos-sensitivity of these curcumin-loaded MLs was evaluated by the calcein release assay,
which utilizes the fluorescence property of calcein49. At a high concentration, calcein shows self-
quenching, resulting in a decrease in fluorescence, and an increase, which is observed when calcein
is released from the liposomal carrier into the surrounding medium. To establish the effect of pH
sensitivity on the release of the drug by the curcumin-loaded MLs, similar experiments were
performed with a sodium acetate buffer (pH 4.8) as a stimulus at 37 °C. The results showed that
negligible amounts of the drug were released from the ML system at the decreased pH, thereby
establishing the independence of pH sensitivity on the curcumin-loaded MLs.
Time-dependent calorimetric measurements to evaluate the heating ability of the Dx-Fe3O4
suspensions were performed using a radio frequency generator. A total of 1 ml (10 mg/ml of iron)
of the Dx-Fe3O4 colloidal suspension was placed in an AC magnetic field (7.64, 8.82, 9.41 and
10.0 kA/m) with a fixed frequency of 425 kHz and with arrangements to minimise heat loss. The
specific absorption ratio (SAR) was calculated using eqn. 3.2.
AI.2.4 Evaluation of Biocompatibility and Therapeutics
The biocompatibility of the MLs was established with a mouse fibroblast cell line (L929)
and cervical cancer cell lines (HeLa). The toxicity of the curcumin-loaded MLs was evaluated with
cervical cancer cell lines (HeLa) by an SRB colorimetric assay50. In order to establish the potential
of these carriers in delivering curcumin, a dose-dependent study was undertaken to evaluate the
50% inhibitory concentration(IC50) values of free curcumin and the curcumin-loaded MLs over 48
h. The cells were seeded in 96-well plates at a cell density of 2×104 cells per well and incubated
in a tissue culture medium for 24 h at 37 °C in a 5% CO2 environment. After 24 h, different
concentrations of the MLs (2, 1, 0.5, 0.25, 0.125, 0.0625 and 0.03125 mg/ml) and the curcumin-
loaded MLs (10, 8, 4, 2, 1, 0.5, 0.25, 0.125, 0.0625 and 0.03125 mg/ml) were mixed with the
growth media, and the cells were incubated for an additional 24 h(MLs) or 48 h (curcumin-loaded
MLs). After 24 h (MLs) or 48 h (curcumin-loaded MLs), the cells were carefully washed with PBS
(pH 7.3), and an SRB assay was performed in order to determine cell viability. For the assay, cells
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were fixed with cold 10% trichloroacetic acid (at 4 °C) and stained with 0.4% SRB (in 1% acetic
acid). After one hour of incubation in the dark, the unbound dye was thoroughly washed with 1%
acetic acid, and the cell-bound dye was later extracted with 10 mM Tris buffer (pH 10.5).
Absorbance was recorded at 560 nm using a Thermo Scientific Multiskan EX multiplate reader.
The relative cell viability was calculated according to eqn. 3.3.
AI.3 Results and Discussion
AI.3.1 Characterisation of the Synthesised Nanoparticles
The characterisation of the prepared nanomaterials was undertaken as described in section
2.2.5. The crystalline structure and crystallite size of the Dx–Fe3O4 nanoparticles were investigated
by powder XRD (Figure AI.1a). The corresponding diffraction planes of the indices showed good
agreement with the reported values for magnetite (JCPDS Card No. 19-0629, a = 8.3967 Å). The
XRD pattern revealed the formation of single-phase magnetite, which has an inverse spinel
structure with a crystallite size of ∼8.48 nm, as calculated by the Scherrer formula. A high degree
of crystallinity of the nanoparticles was indicated by the presence of the sharp and intense peaks
of the Dx–Fe3O4 nanoparticles.
Figure AI.1 (a) XRD pattern of Dx–Fe3O4, TEM micrographs of (b) Dx–Fe3O4 (inset shows the
selected area diffraction pattern of Dx–Fe3O4 and (c) magnetic liposomes (inset shows magnified
image of MLs)
Figure AI.1b shows the TEM micrograph of the Dx–Fe3O4 nanoparticles. The Dx–Fe3O4
nanoparticles were mostly spherical and smaller than 15 nm in diameter. The inset in figure AI.1b
shows the electron diffraction pattern of a selected area. The pattern, which has been indexed with
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the inverse spinel magnetite crystal structure, is consistent with the XRD results. Figure AI.1c
shows the morphology of the MLs. The inner and outer diameters of the MLs averaged ∼100 and
∼200 nm, respectively, with a bilayer thickness of 50 nm. The MLs were stable and well dispersed
in the aqueous solution (the inset shows an image of an ML at a scale of 50 nm). The bilayers were
formed uniformly. The intact liposome is spherical and has varying amounts of Fe3O4 encapsulated
in its hydrophilic core.
Figure AI.2a shows the FTIR spectra of pure dextrin and the Dx–Fe3O4 nanoparticles. The
spectrum of dextrin is well resolved and contains a few broad bands in addition to narrow bands.
The very broad peak at 3297 cm−1 is characteristic of the aromatic sp2 C–H stretch and is attributed
to the pyranose ring vibrations of dextrin. Characteristic peaks of the α-D-glucose units of the
polysaccharide were visible at 1203 cm−1, which occur due to the in-plane C–H and O–H
vibrations51. The multiple bands that appear in the region between 1150 and 930 cm−1 coincide
with the in-plane C–H bending vibrations of the pyranose ring and the C–O stretching vibrations.
The vibrational bands present at 1150 and 1077 cm−1 are attributed to valent vibrations of the C–
O–C bond of the glycosidic bridge, whereas, the peak at 1025 cm−1 is due to the substantial chain
flexibility of dextrin around the glycosidic bonds. The vibrational bands of Dx–Fe3O4 are relatively
broad and are fewer compared with those of pure dextrin. The peaks that are present in the
spectrum of dextrin at 1150 cm−1, 1077 cm−1 and 1025 cm−1 were also present in the spectrum of
Dx–Fe3O4, indicating the successful coating of dextrin on the surface of the nanoparticles. Other
peaks are interpreted as follows: The small and narrow bands around 3000 cm−1 are due to the
presence of H-bonded –OH groups from water molecules, which are physically adsorbed on the
surface of the Dx–Fe3O4 that is suspended in an aqueous medium. A new peak at 3009 cm−1 is
visible in the O–H stretch region. This may be the result of the formation of hydrogen bonds
between the free –OH groups on dextrin and the oxygen from the Fe–O core of the magnetite
nanoparticles, which results in the formation of an iron-oxy-hydroxide monodentate coordination
complex. The peaks at 2934 and2986 cm−1 present in the dextrin spectrum are retained in the Dx–
Fe3O4 spectrum, indicating that the C–H bonds were unaltered during the conjugation and did not
play a role in the coating process. The appearance of the peak at1660 cm−1 indicates the retention
of the cyclic ring structure of dextrin after its conjugation with the Fe3O4 nanoparticles. As seen in
figure AI.2a, the peaks at 1365 and 1417 cm−1 in dextrin shift to 1400 and 1535 cm−1 in the Dx–
Fe3O4 spectrum. This might be due to the formation of O–H bonds in the iron-oxy-hydroxide
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complex and the overlap with C–H bending vibrational bands of the dextrin molecules. The
appearance of the peak at 864 cm−1 (boat conformation) and the disappearance of the peak at 930
cm−1 (chair conformation) could be attributed to the loss of the chair conformation or to a
conformational change to the boat conformation of the glucopyranose units after interaction with
the surface of the Fe3O4. The peak at 575 cm−1 can be attributed to the stretching vibrational modes
of Fe–O.
Figure AI.2 FTIR spectra of (a) dextrin and Dx–Fe3O4 nanoparticles (b) MLs and Cur-MLs
Figure AI.2b shows the FTIR spectra of the soy-PC, MLs, curcumin and curcumin-loaded
MLs. The spectrum of soy-PC reveals characteristic vibrational bands consistent with previously
reported studies52. The peak at 1739 cm−1 is due to C=O vibrations of the fatty acid chains, and the
bands at 1643, 1460 and 1232 cm−1 are due to C–NH3 symmetric scissoring of the choline entity
of lipid, C–H scissoring and P=O vibrations of the phosphatidyl group, respectively. The shift in
the vibrational bands at 1739 cm−1 to 1737 cm−1 in the spectrum of the MLs is indicative of the
interaction of the C=O groups of soy-PC with the Dx–Fe3O4 nanoparticles by either electrostatic
or van der Waals interactions. The bands at 1643 and 1460 cm−1 do not exhibit any shifts, pointing
to the non-participation of the –NH3 groups of the fatty acid in any type of bond formation with
the Dx–Fe3O4 nanoparticles. The band representing the P=O of the choline group of soy-PC (1232
cm−1) exhibits a minor shift (1234 cm−1), indicating the interaction of the group with the free –OH
groups present on the Dx–Fe3O4 nanoparticles. The overlap of soy-PC vibrations on dextrin masks
the dextrin frequencies in the spectrum of the MLs.
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The FTIR spectrum of curcumin-loaded MLs shows sharp, intense characteristic peaks of
soy-PC due to the high ratio of ML to curcumin. Characteristic peaks of curcumin are also present
in the FTIR spectrum of curcumin-loaded MLs, which are in good agreement with the vibration
spectrum of curcumin, which was reported by Mohan et al53. According to their interpretations,
the band at 3502 cm−1 is due to phenolic –OH stretching vibrations, the peak at 1427 cm−1 is
attributed to olefinic bending vibration of the C–H bound to the benzene ring of curcumin, the
peak at 1272 cm−1 due to an enol vibration (C=O), the band at 958 cm−1 due to benzoate trans –
CH vibrations, and the vibrations at 709 cm−1 due to the cis-CH of an aromatic ring. These results
further indicate the successful loading of curcumin into the MLs. The successful conjugation of
curcumin to the MLs was indicated by the shift of some peaks in the FTIR spectrum of curcumin-
loaded MLs, as discussed as follows. A shift observed in the band at 1737 cm−1 in MLs to 1741
cm−1 in the curcumin-loaded MLs is attributed to the attachment of the drug to the –COO groups
of the fatty acid. Other bands shifted from 1643 to 1651 cm−1, 1460 to 1457 cm−1, 1234 and 1232
to 1238 cm−1 and 1056 to 1067 cm−1 in the spectrum of the curcumin-loaded MLs. The shift in
these bands indicates that the –CH3 side chains and the free –OH groups may be involved in the
conjugation of curcumin to the MLs, confirming the drug loading into the MLs. These results
confirm the conjugation of dextrin onto the Dx–Fe3O4 nanoparticles, encapsulation by the lipid
bilayer and the subsequent conjugation of curcumin.
TGA-DTA analysis was conducted to determine the decomposition profile of the
formulations and the relative mass of dextrin (Figure AI.3a) and liposome (Figure AI.3b) coatings.
The thermal profile of the Dx–Fe3O4 nanoparticles revealed that the degradation of the
nanoparticles occurred in three primary steps. Initially, a weight loss of 3.5% occurred, which
corresponded to the endothermic DTA peak at 80 °C. This loss could be attributed to the water
molecules which were superficially adsorbed by the dextrin molecules. A gradual 5.1% loss in the
weight of the nanoparticles was then observed in the temperature range of 100–430 °C, which
corresponded to two DTA peaks at 285 and 400 °C, respectively. Dextrin molecules underwent a
decomposition step at 320 °C. Thus, this observed decrease in weight could be due to the surface-
conjugated dextrin molecules54. The third weight loss step of 0.4% was observed at 570 °C, which
had a corresponding sharp exothermic DTA peak. This decrease occurs due to the loss of FeO
molecules during degradation of magnetite to maghemite.
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Figure AI.3 Thermal degradation profiles of (a) Dx–Fe3O4 and (b) magnetic liposomes
The thermogram of MLs indicated that weight loss occurred in three distinct steps. In the
first step, an initial weight loss of approximately 3.7% occurred, which corresponded to a small
endothermic DTA peak at 60 °C due to the removal of physically absorbed water. In the second
step, a steep decline in the curve was observed. This drop in the TGA plot represents a weight loss
of approximately 9.0% and has a corresponding broad exothermic DTA peak at 280 °C, again due
to the removal of the organic dextrin and lipid molecules by thermal decomposition from the
surface of the Fe3O4 nanoparticles. This large and well-defined weight loss may be due to the
larger surface concentration and high molecular weight of the molecules, which were made up of
carbon, oxygen, hydrogen and phosphorus. In the third step, a negligible weight loss of
approximately 0.3% occurs. There is a corresponding exothermic DTA peak at 600 °C due to the
removal of iron (II) oxide (FeO) during the phase transformation of magnetite to maghemite.
Hence, the total weight loss percentage was approximately 13.7%. This mass loss represents the
weight percentage of organic coatings on the surface of the Fe3O4 nanoparticles and the
encapsulating liposome bilayer.
The colloidal stability and dispersion of the nanoparticles is associated with the electric
charge of the particle surface. Figure AI.4 shows the zeta potential of the Dx–Fe3O4 and MLs at
different pH values, indicating that the conjugation of the dextrin molecules onto the surface of
the Fe3O4 nanoparticles creates a highly negative surface charge with an isoelectric point at pH 5.
These zeta potential values confirm the presence of negatively charged dextrin groups on the
surface of the Fe3O4 nanoparticles. The addition of the lipid bilayer to the Dx–Fe3O4 lowered the
surface charge of the MLs and shifted the isoelectric point to pH 4.5. At pH 7.0, the MLs showed
a negative surface charge of approximately −30 mV, which is highly desirable, as it ensures high
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aqueous stability of the MLs in a physiological environment before reaching the targeted cancer
cells. The dynamic light scattering (DLS) measurements on the MLs (Inset of Figure AI.4) showed
that the MLs had a mean hydrodynamic diameter of 1.2 µm, which can be attributed to the presence
of the water that is associated with the hydrophilic heads of the lipid layer.
Figure AI.4 Zeta potential of Dx–Fe3O4 and MLs as a function of pH. Inset shows the
hydrodynamic diameter of the MLs obtained from DLS measurements (1.2 µm)
Figure AI.5a shows the field-dependent magnetisation (M versus H) plot of the Dx–Fe3O4
nanoparticles and the MLs at room temperature. Both exhibited superparamagnetic behaviour, that
is, zero magnetic hysteresis and remanence. The maximum magnetisation of the Dx–Fe3O4 and
the MLs was 54.2 emu/g and 24.3 emu/g, respectively, at a field of 20 kOe. The observed
magnetisation of Dx–Fe3O4 is comparable to that (approximately 60 emu/g) of bare Fe3O4
nanoparticles that have been reported in previous work55. Dextrin chains with more than 10 glucose
units tend to adopt a helical structure in an aqueous solution that contains cavities that easily form
an inclusion complex with low molecular weight compounds such as Fe3O4 nanoparticles and
ions56. These helical cavities may interfere with the domain alignment of the Fe3O4 nanoparticles
with the applied magnetic field, thus lowering the value of its magnetisation in comparison to bare
nanoparticles. The strong magnetic response of these aqueous-stable nanoparticles could be
exploited for various applications such as magnetic targeting, hyperthermia treatment, bio-sensing
and magnetic resonance imaging. The substantial drop in the magnetisation values of the MLs is
attributed to the high molecular weight of the non-magnetic soy-PC bi-layer. This bilayer masks
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the domain alignment and restricts the response of the Dx–Fe3O4 nanoparticles that are trapped
within the bilayers to the applied magnetic field.
Figure AI.5 (a) Field-dependent magnetisation (M vs. H) plot of Dx–Fe3O4 and MLs at room
temperature (b) Time-dependent specific absorption measurements of Dx–Fe3O4 nanoparticles
(Inset depicts the dependence of heat generation on the applied ACMF)
Figure AI.5b shows the time-dependent SAR of the Dx–Fe3O4 nanoparticles in response to
the application of varying ACMFs. The calorimetric measurement was used to determine the
heating rate of the Dx–Fe3O4 suspension. The SAR values of the nanoparticles were 9.51, 19.50,
22.03 and 32.90 W/g of iron with an applied field (H) of 7.64, 8.82, 9.41 and 10 kA/m,
respectively. Figure AI.5b illustrates that with the increment of the applied field, the time required
to reach a temperature of 45 °C from an initial ambient temperature was reduced, which is
consistent with the relationship between heat generation and the applied ACMF (the inset of Figure
AI.5b).
AI.3.2 Encapsulation (Loading) of Curcumin
The absorbance spectra of the pure curcumin and the supernatant (after magnetic
sedimentation) that was obtained from the MLs are given in Figure AI.6a, which shows that the
absorbance of the supernatant decreased with an increase in the ML-to-curcumin ratio. This
decrease that is observed in absorbance is attributed to the conjugation of the curcumin molecules
with the phosphatidyl choline groups, due to the increase in the concentration of the latter relative
to that of the former. In other words, the decrease of absorbance indicates an increased
encapsulation of curcumin. The hydrophobic curcumin and the lipophilic ends of soy-PC in an
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aqueous environment interact with each other in a way that facilitates the encapsulation of the drug
within the lipid bilayer. This been reported to be strongly dependent on the ratio of phosphatidyl
choline to curcumin during the thin-film formation step, because a single molecule of curcumin
binds with six molecules of phosphatidylcholine57, 58. In the present work, a maximum
encapsulation efficiency of 97.6% was achieved when the ratio of phosphatidyl choline to
curcumin was 1:0.6, with the curcumin-loaded MLs showing a high degree of homogeneity and
stability in aqueous suspensions. The encapsulation efficacy observed was higher than values
reported for similar nanocarrier systems14, 59.
Figure AI.6 (a) Absorbance spectra of curcumin-loaded MLs against pure curcumin (b) Drug
release profiles of curcumin-MLs at physiological and hyperthermic temperatures
AI.3.3 Release of Curcumin at Elevated Temperatures
Figure AI.6b shows the release profile of the curcumin-loaded MLs under physiological
(37 °C) and elevated (45 °C) temperature in reservoir-sink conditions. The sink (PBS) was spiked
with 0.1% (v/v) chloroform to facilitate the diffusion of hydrophobic curcumin in an aqueous
environment. Under normal physiological conditions, curcumin release was lower than 6%, which
is negligible. In contrast, it was released rapidly under hyperthermia temperatures following an
initial short period (25 min) of slow release. At 45 °C, the curcumin release reached a plateau at
150 min and had a total release percentage of ∼56%. The release of the curcumin molecules could
be attributed to the disruption of the heat sensitive lipid bilayer at elevated temperatures, which
may have weakened and, thus, disrupted the hydrophobic interactions between the curcumin and
the phosphatidyl choline groups in the bilayer, thereby exposing curcumin to the surrounding
144
environment. The phase transition temperature is an important parameter in determining the
fluidity of the bilayer of liposomes. The fluidity, in turn, affects the release of curcumin from the
MLs. Since the physiological temperature is below the transition temperature of soy-PC, the
release of curcumin in physiological environments is inhibited. In contrast, a hyperthermic
temperature of 45 °C is above the transition temperature and fluidises the bilayer, enhancing the
release of the curcumin60.
Another factor that must be considered in the evaluation of anticancer drug release is the
acidic microenvironment around a tumour. Tumours, due to their hypoxic (low oxygen) conditions
and high lactic acid secretion, tend to cause acidity in their surrounding microenvironment with a
pH value below physiological (7.2–7.4). Hence, a pH value that is lower than 7.0 is typically used
to evaluate the pH sensitivity of cancer drug release. In this work, the effect of pH sensitivity on
the release of the curcumin-loaded MLs was investigated with a sodium acetate buffer (pH 4.8).
The result showed that little curcumin was released from the ML system at the decreased pH value,
thereby establishing the independence of pH sensitivity on the curcumin-loaded MLs. Considering
this result, this work focused on the temperature-stimulated drug release of curcumin-loaded MLs
at pH 7.0.
AI.3.4 Biocompatibility of Magnetic Liposomes and Anticancer Therapeutic
Effect of Curcumin-Loaded Magnetic Liposomes
Figure AI.7a shows the cell viabilities of mouse fibroblast (L929) and cervical cancer
(HeLa) cell lines that were incubated in a growth medium that contained MLs. An SRB assay was
performed in order to quantify the viable cell population, thereby indicating the effect of the MLs
on the growth and phenotype of the L929 and HeLa cells. The results suggested that the MLs alone
are biocompatible and, thus, are safe for in vivo studies. The relative cell viability, which was
calculated from eqn. 3.3, exceeded 100% when the concentration of the nanoparticles was low;
this was attributed to the facilitation of cell growth by the iron released by the cellular degradation
of Fe3O4 61.
145
Figure AI.7 (a) Percentage of the cell viability of MLs incubated with mouse fibroblasts and
cervical cancer cells for 24 h (b) Dose-dependent evaluation of Curcumin-MLs for determination
of IC50 with HeLa cells
Figure AI.7b illustrates the results of the dose-dependent study that was performed with
HeLa cells in order to evaluate the half maximal inhibitory concentration (IC50) of the curcumin-
loaded MLs. The results indicate that the formulations were able to inhibit cell proliferation by
approximately 65% in HeLa cells. The amount of pure curcumin that reduced the cell population
by half (IC50) was 125 µg/ml over a period of 48 h. The IC50 value of curcumin has been reported
as ranging from 15 to 30 µM with various cancer cell lines14, 62, 63. The dose-response curve fitting
with Origin 8 software showed that the curcumin-loaded MLs hindered cell growth and reduced
the cell population by half (IC50) at a concentration of 2.09 mg/ml (R2=0.982). However, cell
viability was decreased to 70% at a concentration ∼62.5 µg/ml, as indicated by Figure AI.7b. The
IC50 value of 2.09 mg/ml for curcumin-loaded MLs was higher than that (125 µg/ml) of pure
curcumin, which is due to the loading of the Dx–Fe3O4 and the encapsulative lipid bilayers, along
with the ML release profile over 48 h. In a physiological environment, the anticancer therapeutic
effect of pure curcumin is hampered by its low aqueous solubility and poor chemical stability,
which limits its bioavailability at the targeted cancer site. Hence, the MLs would enhance the
therapeutic effect by increasing the delivery efficiency of curcumin when compared with pure
curcumin.
146
AI.4 Summary
The controlled release of curcumin using the ML system and the therapeutic effect of this
drug formula have been explored for the treatment of cervical cancer. The surface-engineered
magnetic nanoparticles and the dextrin-coated iron oxide (Dx–Fe3O4) were fabricated using a
single-step facile co-precipitation approach. The Dx–Fe3O4 nanoparticles were successfully
encapsulated within the hydrophilic core of liposomes, producing a magnetic liposome (ML)
system. The conjugation of dextrin, lipid and curcumin molecules onto the Fe3O4 nanoparticles
was also achieved. The MLs exerted little or no toxic effects on normal fibroblast cells (L929),
showing potential for being used as a drug carrier to deliver hydrophobic drugs in aqueous
environments without harming normal cells. The release of curcumin molecules can be controlled
by varying the temperature, which can be achieved by applying an ACMF. The fabricated
curcumin-loaded MLs can release curcumin in appreciable amounts (up to 56.3±3.9%) at 45 °C
under the control of magnetic stimulation. More significantly, the curcumin-loaded MLs can
effectively inhibit cancer cell growth and viability and have an IC50 value of 2.09 mg/ml. In short,
the work has demonstrated the ability of MLs to efficiently deliver curcumin and to maintain its
anticancer activity, in an aqueous environment for the treatment of cervical cancer.
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List of Publications
Research Articles
1. Saumya Nigam, K. C. Barick, D. Bahadur, Development of citrate-stabilized Fe3O4
nanoparticles: Conjugation and release of doxorubicin for therapeutic applications, J.
Magn. Magn. Mater., 2011, 323, 237-243. (Citations - 149)
2. S. Chandra, Saumya Nigam, D. Bahadur, Combining unique properties of dendrimers and
magnetic nanoparticles towards cancer theranostics, J. Biomed. Nanotech., 2014, 10, 32-
49. (Citations - 13)
3. Saumya Nigam, S. Chandra, D. F. Newgreen, D. Bahadur, Q. Chen, Poly (ethylene
glycol)-modified PAMAM-Fe3O4-Doxorubicin triads with the potential for improved
therapeutic efficacy: Generation-dependent increased drug loading and retention at neutral
ph and increased release at acidic pH, Langmuir, 2014, 30, 1004-1011. (Citations - 18)
4. Saumya Nigam, A. Kumar, G. A. Thouas, D. Bahadur, Q Chen, Curcumin delivery using
magnetic liposomes, J. Nanopharma. Drug Del., 2013, 1, 365-375. (Citations - 1)
5. Saumya Nigam, S. Chandra, D. Bahadur, Dendrimers based electrochemical biosensors,
Biomed. Res. J., 2015, 2, 21-36. (Citations - 1)
6. Saumya Nigam, D. Bahadur, Dendrimerized magnetic nanoparticles as carriers for the
anti-cancer compound, epigallocatechin gallate, IEEE Trans. Magn., 2016, 52, 1-5.
7. Saumya Nigam, D. Bahadur, Dendritic Fe3O4 nanoparticles for combinatorial therapy:
Peptide dendrimers with enhanced efficiency as alternative platforms for pamam
dendrimers (Under Review)
8. Saumya Nigam, D. Bahadur, Temperature-induced MR contrast properties of dendritic
Fe3O4 nanoparticles (To be communicated)
9. Saumya Nigam, D. Bahadur, Assessment of doxorubicin-loaded dendritic Fe3O4
nanoparticles for magnetic drug targeting in murine melanoma model (Under Review)
Jeotikanta Mohapatra, Saumya Nigam, J. Gupta, A. Mitra, D. Bahadur, M. Aslam,
Enhancement of magnetic heating efficiency in size controlled MFe2O4 (M = Mn, Fe, Co
and Ni) Nanoassemblies, RSC Adv., 2015, 5, 14311-14321. (Citations - 8)
152
Eswara Vara Prasadarao K, Saumya Nigam, M. Aslam and D. Bahadur, Novel Mg-Al
layered double hydroxide-Fe3O4 magnetic nanohybrids for efficient thermo-chemo therapy
of cervical cancer, New J. Chem., 2016, 40, 423-433. (Citations - 7)
Swati, Saumya Nigam, A. S. Khanna, R.K. Singh Raman, Silane coated magnesium
implants with improved in vitro corrosion resistance and biocompatibility, J. Mater. Sci.
Surf. Engg., 2016, 13, 29.
Presentations at National/International Conferences
1. Saumya Nigam, D. Bahadur, Assessment of Doxorubicin-loaded Dendritic Fe3O4
Nanoparticles for Magnetic Drug Targeting in Murine Melanoma Model, Presented at
IITB-NTU Joint Symposium on Healthcare Technologies 2016, Sept 26 – 27, 2016,
Mumbai, India (Won the Best Poster award)
2. Saumya Nigam, D. Bahadur, Peptide Dendrimer-Fe3O4 Nanoparticles with Enhanced
Efficiency as Platforms for Combinatorial Thermo-Chemo Therapy, Presented at 43rd
Annual Meeting & Exposition of the Controlled Release Society, July 17 – 20, 2016,
Seattle, Washington, U.S.A.
3. Saumya Nigam, D. Bahadur, Peptide dendrimers in non-invasive magnetic resonance
imaging, Presented at International Symposium on Nanotechnology and Cancer
Theranostics, Feb 19 – 21, 2015, Mumbai, India
4. Saumya Nigam, D. Bahadur, Peptide dendrimers in non-invasive magnetic resonance
imaging, Presented at In-house symposium of IITB-Monash Research Academy, Aug 7,
2015, Mumbai, India (Nominated for Best poster award)
5. Saumya Nigam, Q. Z. Chen, D. Bahadur, Comparative therapeutic performances of drug
loaded magnetic dendritic nanoparticles towards various cancer cell lines, Presented at In-
house symposium of IITB-Monash Research Academy, Aug 8, 2014, Mumbai, India
6. Saumya Nigam, Asmita Kumar, Qizhi Chen, D. Bahadur, Therapeutic effect of curcumin
loaded magnetic liposomes on cervical cancer, Presented at 7th International Conference
on Materials for Advanced Technologies, June 30 – July 5, 2013, Singapore, Singapore
7. Saumya Nigam, Qizhi Chen, D. Bahadur, Therapeutic performance of PEG-PAMAM
dendrimers functionalized magnetic nanoparticles towards cervical cancer, 2013 Spring
153
Meeting of European Materials Research Society, May 27-31, 2013, Strasbourg, France –
Presented by Prof. D. Bahadur
8. Sudeshna Chandra, Saumya Nigam, Akshaya K Swain, Shanta S Naorem, Anand Prakash,
Dipa Dutta, D. Bahadur, Advanced Functional Nanostructures: From Synthesis to
Applications, Presented at Nano India, Feb 19-20, 2013, NIIST, Trivandrum, India
9. Attended one-day workshop on “Fluorescence Steady State and Lifetime Analysis”, on 7th
Dec, 2010 at Mumbai University
154
Acknowledgments
I am under no illusion that my PhD work would have been possible without the support of
others, so I am sincerely grateful to a long list of people and the various roles they played during
my PhD. To begin with, I wish to express my utmost gratitude to my parents and friends for always
supporting me and being the steadfast pillars of my life.
I would like to acknowledge the tireless work of my supervisor Prof. D. Bahadur (IIT
Bombay) for his valuable and expert guidance and patience for bearing with me for the last few
years. All these years have been a constant learning process under his supervision and he shall
always be a constant source of inspiration for me. I would also like to thank Dr. Q. Z. Chen (ex-
supervisor; Monash University) for her pivotal suggestions towards my research work during my
tenure at Monash University and later Dr. X. Chen. I would also like to thank my entire research
progress committee for their individual guidance, suggestions and support during my doctoral
studies. Thank you Prof. Rohit Srivastava and Dr. Lizhong He for your scientific guidance and
insightful suggestions.
I am also thankful to the staff members of my department (both at IIT Bombay as well as
Monash University) and the technical staff at SAIF (Sophisticated Analytical Instrument Facility)
for their help in various characterisation facilities and other helps during my work. I would like to
thank Ms. Aradhana Pant, Ms. Pradnya Nikam and Mr. Sachin Tawde at IIT Bombay with a special
mention of Dr. Qizhu Wu of Monash Biomedical Imaging facility for all his efforts and help
towards my MRI studies.
Last but not the least, I would like to extend my gratitude to all of my group members (both at IIT
Bombay and Monash University), past and present, especially, Dr. Sudeshna Chandra, Dr. K. C.
Barick, Dr. Pallab Pradhan and all the co-workers for their insightful and elaborate research
discussions and providing an amiable environment in the lab throughout my PhD.
(Saumya Nigam)