Funtionalization of magnetic nanoparticles for bio applications

FUNCTIONALIZATION OF MAGNETIC NANOPARTICLES FOR
BIO-APPLICATIONS






WUANG SHY CHYI
(B. Eng (Hons), NUS)





A THESIS SUBMITTED
FOR THE DEGREE OF DOCTOR OF PHILOSOPHY
DEPARTMENT OF CHEMICAL AND BIOMOLECULAR
ENGINEERING
NATIONAL UNIVERSITY OF SINGAPORE
2008
ii
ACKNOWLEDGEMENTS

I would like to express my heartfelt gratitude to my supervisors, Professors K. G.
Neoh, Daniel Pack, E T. Kang and Deborah Leckband for their continued guidance,
invaluable suggestions and profound discussion throughout this work. Without their
enthusiasm and help, this project would not be possible. The knowledge gained under
their supervision and the research experiences pave the way for my lifelong study.

I would also like to thank the other members of my committee for their help and time,
as well as the research staff and laboratory officers, both in the National University of
Singapore and the University of Illinois at Urbana-Champaign.

Finally I thank my family, colleagues and numerous friends for their love, support and
encouragement.
iii
TABLE OF CONTENTS
ACKNOWLEDGEMENT
ii
TABLE OF CONTENTS
iii
SUMMARY
v
LIST OF TABLES
vi
LIST OF FIGURES
vii
NOMENCLATURE
x
CHAPTER 1: INTRODUCTION
1
CHAPTER 2: LITERATURE REVIEW
2.1 Magnetic nanoparticles
2.2 Magnetic drug delivery
2.3 Cancer targeting
2.4 Magnetic Fluid Hyperthermia
2.5 Magnetic nanoparticles in imaging
2.6 Biocompatible polymers

4
8
12
14
15
18
CHAPTER 3: HEPARINIZED MAGNETITE NANOPARTICLES
3.1 Introduction
3.2 Methods and materials
3.3 Results and discussion
3.4 Chapter 3 Conclusion

20
21
28
43
CHAPTER 4: DOXORUBICIN-ATTACHED MAGNETITE
NANOPARTICLES
4.1 Introduction
4.2 Methods and materials
4.3 Results and discussion


44
45
51
iv
4.4 Chapter 4 Conclusion
63
CHAPTER 5: ANTIBODY-ATTACHED MAGNETITE

NANOPARTICLES
5.1 Introduction
5.2 Methods and materials
5.3 Results and discussion
5.4 Chapter 5 Conclusion


65
66
70
80
CHAPTER 6: POLYPYRROLE-MAGNETITE NANOSPHERES
6.1 Introduction
6.2 Methods and materials
6.3 Results and discussion
6.4 Chapter 6 Conclusion

81
82
91
124
CHAPTER 7: SUMMARY CONCLUSION
126
CHAPTER 8: RECOMMENDATIONS FOR FUTURE WORK
129
REFERENCES
132
LIST OF PUBLICATIONS
149
LIST OF CONFERENCES

150





v
SUMMARY
Magnetite nanoparticles were modified to render them suitable for bio-applications,
namely drug delivery and hyperthermia, using two different approaches. The first
approach is to graft polymers onto the nanoparticles using surface-initiated atom
transfer radical polymerization, followed by chemical linking of biomolecules onto
the grafted polymers. The monomers used include N-isopropylacrylamide, N-
vinylformamide and methacrylic acid while the immobilized biomolecules include
heparin, folic acid, doxorubicin and anti-HER2/neu antibodies. It was found that the
heparinized nanoparticles could reduce macrophage uptake, and at the same time
inhibit plasma clotting. The doxorubicin-bearing nanoparticles were able to release a
greater amount of the drug under acidic conditions as opposed to physiological pH,
and could potentially serve as drug depots. Particles that were functionalized with
anti-HER2/neu antibodies showed a preferential binding to cancer cells and may be
useful for imaging purposes.

The second approach was to encapsulate these magnetite nanoparticles into
polypyrrole nanospheres via emulsion polymerization for potential use as
hyperthermia causing agents. The nanospheres were then functionalized with folic
acid or herceptin to impart onto them cancer cell-targeting properties. These
functionalized nanospheres target cancer cells in vitro and possess good
magnetization which is useful for magnetic fluid hyperthermia.

vi
LIST OF TABLES
Table 3.1 Dispersion of as-synthesized and functionalized magnetite in various
solvents at 25ºC.

Table 3.2 Comparison of PRT obtained in the absence and presence of magnetic
nanoparticles.

Table 6.1 Characteristics of the PPY nanospheres (Scale bar = 200nm).

Table 6.2 Properties of NS(PVA).

Table 6.3 Properties of NS(HA) and NS(HA)-HER2.

Table 6.4 Amount of iron associated with SK-Br-3 and MDA-MB-231 cells after
incubation with NS(HA) and NS(HA)-HER2.
vii
LIST OF FIGURES
Figure 3.1 Schematic representation of the process for preparing MNP-NP-He.
Figure 3.2 XPS C 1s and S 2p core-level spectra of as-synthesized MNP (a, c),
magnetite-Cl (b, d) and N 1s core-level spectra of magnetite-Cl (e) and
MNP-NP (f).

Figure 3.3 XPS C 1s and S 2p core-level spectra of MNP-NP (a, c) and MNP-NP-
He (b, d).

Figure 3.4 FTIR spectra of MNP (a), magnetite-Cl (b), MNP-NP(c) and MNP-
NP-He (d)

Figure 3.5 Room temperature magnetization curves of as-synthesized and
functionalized magnetite nanoparticles as a function of applied
magnetic field: MNP (a), MNP-NP (b) and MNP-NP-He (c).

Figure 3.6 Optical microscopy images of macrophages cultured with control cells
(a, b), MNP (c, d), MNP-NP (e, f) and MNP-NP-He (g, h) after 2 and
24 h respectively. Figure scale bar = 50μm, inset scale bar = 25μm.

Figure 3.7 Total uptake of as-synthesized and functionalized magnetite
nanoparticles by macrophages after 2, 8 and 24 h.

Figure 3.8 Cytotoxicity of as-synthesized and functionalized magnetite
nanoparticles, as measured by the viability of macrophages grown in
media containing 0.2 mg/ml of these nanoparticles relative to the non-
toxic control. T represents the results obtained with the toxic control.
Results are represented as mean ± standard deviation.

Figure 4.1 Schematic for synthesis of doxorubicin-conjugated particles.

Figure 4.2 XPS C 1s core-level spectra of MNP-P(MAA)-NHNH
2
(a), MNP-
P(MAA)-NH-N=Dox (b) and doxorubicin hydrochloride (c).

Figure 4.3 Magnetization profiles of MNP-P(MAA)-NH-N=Dox in the solid state
(a), and as dispersed in 1% agarose (b).

Figure 4.4 In vitro doxorubicin release from MNP-P(MAA)-NH-N=Dox at 37°C
or 42°C in various pHs as indicated.
.

Figure 4.5 In vitro doxorubicin release from MNP-P(MAA)-NH-N=Dox at 37°C.
Arrow indicates point of pH change from 7.4 to 5.5 or 6.6.

Figure 4.6 In vitro doxorubicin release from MNP-P(MAA)-NH-N=Dox. Arrow
indicates point of temperature change from 37°C to 42°C and pH
change from 7.4 to 5.5 or 6.6.

viii
Figure 4.7 Figure 4.7: Viabilities of MDA-MB-231 cells (relative to non-toxic
controls) incubated with medium containing free doxorubicin (Dox) or
MNP-P(MAA)-NH-N=Dox. “*”, “#”, “**” and “##” denote statistical
differences (P < 0.05) between the similarly marked samples.

Figure 4.8 Optical microscopy images of prussian blue staining of MDA-MB-231
cells cultured with MNP-P(MAA)-NH-N=Dox in pH 5.5 (a) and pH
7.4 (b). Scale bar = 25μm.

Figure 5.1 Schematic for synthesis of MNP-NVAM-PEG-Ab.

Figure 5.2 FTIR spectra of MNP-NVAM (a), MNP-NVAM-PEG (b) and MNP-
NVAM-PEG-Ab (c)

Figure 5.3 Optical microscopy images of SK-Br-3 cells after a 4 h incubation with
MNP-NVAM-PEG-Ab (a) and co-treatment with MNP-NVAM-PEG-
Ab and free antibody (b). Scale bar = 25 µm.

Figure 5.4 Prussian blue staining of the liver (a), bladder (b), heart (c), kidney (d),
spleen (e), lung (f), tumors (g-i) and the site of injection, tail (j). Scale
bar = 50 µm.

Figure 5.5 Prussian blue staining of the liver (a), bladder (b), heart (c), kidney (d),
spleen (e), lung (f), tumors (g-i) and the site of injection, tail (j). Scale
bar = 50 µm.

Figure 5.6 MMOCT spectra of negative control (a), and phantom with an
equivalent particle concentration of 23 µg Fe/ml (b). Scale bar = 250
μm.

Figure 6.1 Schematic representation of the preparation route NS(PVA)-FA.
Figure 6.2 (a) FTIR spectra of Fe
3
O
4
, PPY nanospheres and NS(PVA) (b) XRD
spectrum of NS(PVA).

Figure 6.3 Room temperature magnetization curves of NS(PVA) as a function of
applied magnetic field.

Figure 6.4 FESEM and TEM images of NS(PVA) with (a, b) 0 %, (c, d) 23.5 %,
(e, f) 28.0 % and (g, h) 38.8 % Fe
3
O
4
content respectively.

Figure 6.5 XPS C 1s core-level and wide scan spectra of (a, b) NS(PVA) and (c, d)
NS(PVA)-FA.
Figure 6.6 Viabilities of 3T3 fibroblasts incubated with medium containing 0.2
mg/ml of NS(PVA) with the indicated Fe

3
O
4
content. FA-(28.0%)
denoted NS(PVA)-FA with 28.0% of Fe
3
O
4
. “*” denotes statistical
differences (P < 0.05) compared to the control experiment.

ix
Figure 6.7 Viabilities of MCF 7 cells incubated with medium containing 0.2
mg/ml of nanospheres containing 28.0 % of Fe
3
O
4
. “*” denotes
statistical differences (P < 0.05) compared to the control experiment.
Figure 6.8 Optical microscopy images of MCF-7 cells cultured with: (a) no
nanospheres (control) (b) NS(PVA) and (c) NS(PVA)-FA after 24 h.
Figure scale bar = 50μm.

Figure 6.9 Schematic for preparation of NS(HA) and subsequent functionalization
with herceptin.

Figure 6.10 Uptake of the nanospheres by HCC1954 cells: (a) Optical microscopy
images of cells after 24 h incubation with (i) NS(HA), and (ii)
NS(HA)-HER1. Figure scale bar = 50 μm. (b) Amount of iron in cells
after 2 h and 24 h incubation with NS(HA) and NS(HA)-HER1

determined using ICP.

Figure 6.11 Schematic representation of the preparation of NS(HA)-HER2.
Figure 6.12 XPS C 1s core-level spectra of (a) NS(NH
2
), (b) NS(HA)-HER2 and (c)
herceptin.

Figure 6.13 Amount of iron associated with SK-Br-3 cells after 2, 4 and 24 h
incubation with NS(HA), NS(HA)-HER2 and NS(HA)-HER2 with free
herceptin. Three sets of duplicates were done for each data point. The
iron association of NS(HA)-HER2 is significantly higher (P < 0.01)
than those for NS(HA) and NS(HA)-HER2 with free herceptin at all
time points.

Figure 6.14 Transmission electron micrographs of cells cultured with (a) NS(HA)-
HER2 (b) NS(HA)-HER2 with pre- and co-treatment of 200 µg/ml free
herceptin, for 4h.

Figure 6.15 Cytotoxicities of NS(HA)-HER2 and NS(HA) with various
concentrations of herceptin (HER), as measured by the viabilities of
SK-Br-3 cells grown in media containing 0.2 mg/ml of these
nanospheres relative to the non-toxic control. Results are represented
as mean ± standard deviation. “*” denotes statistical differences (P <
0.05) compared to the control experiment.

Figure 6.16 Plot of viability of cells versus iron uptake by breast cancer cells.
Tested cell lines include SK-Br-3 (■), MDA-MB-231 (♦) and MCF-7
(▲).

Figure 6.17 Magnetization curves of NS(HA)-HER2 in different environments (a)
NS(HA)-HER2 solid, (b) NS(HA)-HER2 dispersed in culture medium
with 1% agarose and (c) endocytosed NS(HA)-HER2 dispersed in
culture medium with 1% agarose.

x
NOMENCLATURE
ATRP Atom-transfer radical polymerization
Bpy 2-2’-bipyridyl
CPA 3-chloropropionic acid
CT Computed tomography
CTCS 2-(4-chlorosulfonylphenyl) ethyltrichlorosilane
DMF Dimethyl formamide
DMSO Dimethyl sulphoxide
EDC 1-ethyl-3-(3-dimethylamino)-propyl
carbodiimide
EGFR Human epidermal growth factor receptor
FMOC-NH-PEG-SCM Fluorenylmethoxycarbonyl-poly(ethylene
glycol)-succinimidyl carboxymethyl
FTIR Fourier-transform infrared
HA Hyaluronic acid
HER2 Human epidermal growth factor receptor 2
ICP Inductively coupled plasma spectroscopy
LCST Lower critical solution temperature
mAb Whole monoclonal antibodies
MFH Magnetic fluid hyperthermia
MMOCT Magnetomotive OCT
MNP Magnetite nanoparticles
MNP-NP poly(NIPAAM)-grafted MNP
MNP-NP-He Magnetite-poly(NIPAAM)-Heparin

MNP-P(MAA) P(MAA)-grafted MNP
xi
MNP-P(MAA)-NH-N=Dox Doxorubicin conjugated P(MAA)-grafted MNP
MNP-NVAM poly(N-vinylamine)-grafted MNP
MNP-NVAM-PEG-Ab Antibody-linked PEGylated MNP-NVAM
MPS Mononuclear phagocytic system
MRI Magnetic resonance imaging
M
s
Specific saturation magnetization
MTT (3-[4,5-dimethyl-thiazol-2-yl]-2,5-
diphenyltetrazolium bromide)
NaDS Sodium dodecylbenzene sulfonate
NHS N-hydroxysuccinimde
NS(HA) HA-stabilized PPY-Fe
3
O
4
nanospheres
NS(HA)-HER1, NS(HA)-HER2 Herceptin-functionalized NS(HA)
NS(PVA) PVA-stabilized PPY-Fe
3
O
4
nanospheres
NS(PVA)-FA Folic acid –functionalized NS(PVA)
OCT Optical coherence tomography
PBS Phosphate-buffered saline
PEG Polyethylene glycol
PEO Polyethylene oxide

P(MAA) Poly(methacrylic acid)
poly(NIPAAM)) poly(N-isopropylacrylamide)
PPP Platelet-poor plasma
PPY Polypyrrole
PRT Plasma Recalcification Time
PVA Poly(vinyl alcohol)
RF Radiofrequency
xii
SAR Specific power absorption rate
scFv Single chain Fv
SDS Sodium dodecyl sulfate
SPIO Superparamagnetic iron oxide
SQUID Superconducting Quantum Interference Device
T
1
Longitudinal relaxation time constant
T
2
Tranverse relaxation time constant
T
2
*
T
2
in cases of magnetic inhomogeneity
TEM Transmission electron microscopy
TG Thermogravimetric
THF Tetrahydrofuran
SEM Scanning electron microscopy
XPS X-ray photoelectron spectroscopy

XRD X-ray powder diffraction
VSM Vibrating sample magnetometer



Chapter 1
- 1 -

CHAPTER 1: INTRODUCTION

In recent years, magnetic nanoparticles have been proposed for use in a number of
biomedical applications such as drug delivery, hyperthermia and chemotherapy and as
radiotherapy enhancement agents because of their special physical properties [1-3].
Magnetic nanoparticles have controllable sizes, smaller than cells and comparable to
proteins and other biological entities, and hence they can be modified for interaction
with these biological entities. Through manipulation by an external magnetic field,
magnetic nanoparticles are potentially very useful in the transport and immobilization
of magnetically tagged biological cargoes, and also in transferring energy from the
exciting field. The focus of this research project is to functionalize or modify
magnetic nanoparticles to render them suitable for biomedical applications such as
drug delivery, tumor-targeting and hyperthermia. Two different approaches were used.

Iron oxides, namely Fe
3
O
4
and γ-Fe
2
O
3

, have high saturation magnetization values,
and their potential uses in biomedical applications are widely investigated. Magnetite
(Fe
3
O
4
) nanoparticles offer distinct advantages in localized drug delivery. They are
superparamagnetic (i.e. they retain no magnetic properties when the magnetic field is
removed) and can be guided to the targeted area with external magnetic fields [4].
They exhibit low toxicity [5-7] and can be made biocompatible. Therefore, magnetite
nanoparticles were chosen for use in this project.

The first approach was to coat individual magnetite nanoparticles of about 6-8 nm
with polymers to improve their biocompatibility, followed by chemical linking of
biomolecules onto the grafted polymers. In Chapter 3, the surface modification of
Chapter 1
- 2 -

magnetite nanoparticles with heparin to increase their circulation time and for
potential delivery of heparin locally to prevent the formation of blood clots was
investigated. Preliminary work has shown that the direct immobilization of heparin on
the magnetite nanoparticles could not be easily achieved. As such, a poly(N-
isopropylacrylamide) (poly(NIPAAM)) layer was first formed on the surface of the
magnetite nanoparticles via atom-transfer radical polymerization (ATRP) followed by
the immobilization of heparin onto the poly(NIPAAM) shell. The bioactivity of
heparinized magnetite nanoparticles was assessed via their uptake by macrophages
and plasma recalcification time. Uptake by the mononuclear phagocytic system (MPS)
or reticuloendothelial system and complement activation leading to clot formation are
two major challenges confronting the in vivo use of nanoparticles as drug delivery
vehicles. Heparin is hydrophilic and can inhibit coagulation by binding and thereby

inhibiting thrombin. Thus, the tested hypothesis was that heparin immobilized on the
surface of nanoparticles would reduce or even eliminate the processes of macrophage
uptake and clot formation.

In Chapter 4, a cancer drug, doxorubicin, was functionalized onto poly(methacrylic
acid)-grafted magnetite nanoparticles through the use of acid-sensitive hydrazone
linkages. With doxorubicin conjugated to the magnetic carriers, an external magnet
can then direct the drug to the target site (usually cancerous tumors). With such a site-
specific drug delivery system, the local concentration of the cytotoxic drug at the
target site can be maintained at appropriate levels while reducing the overall dosage
or systemic concentration. Moreover, the use of acid-sensitive linkages allows a
greater amount of doxorubicin to be released at the acidic conditions of the tumor
environment. Next, using a similar approach, antibodies (namely anti-HER2/neu
Chapter 1
- 3 -

antibodies) were attached onto poly(N-vinylamine)-grafted magnetite nanoparticles
using a poly(ethylene glycol) bifunctional linker (Chapter 5). This imparts targeting
capability onto the nanoparticles which can then be potentially used in the imaging of
tumors.

Another approach used in the modification of magnetite nanoparticles is to
encapsulate these particles inside polypyrrole (PPY) nanospheres. The PPY-Fe
3
O
4

nanospheres retained high levels of magnetization and electrical conductivity and
hence are potentially useful for hyperthermia by deploying both the magnetic and
conductive heating capacities. To impart tumor-targeting capability onto these

magnetic nanospheres, folic acid (a vitamin) and herceptin (a cancer antibody) were
surface-immobilized onto the nanospheres. As discussed in Chapter 6, these targeting
moieties lead to an increased uptake of the nanospheres by cancer cells. The
encapsulation of magnetic nanoparticles inside electrically conducting PPY
nanospheres imparts magnetic property onto the nanospheres. In addition, the heating
effect deployed in hyperthermia may be enhanced with conductive heating in addition
to magnetic heating. When the magnetic nanospheres used in hyperthermia can
specifically target cancer cells, the effectiveness of hyperthermia would be further
enhanced.
Chapter 2
- 4 -

CHAPTER 2: LITERATURE REVIEW

2.1 MAGNETIC NANOPARTICLES
Magnetic nanoparticles are of great interest for researchers from a wide range of
disciplines, including catalysis [8, 9], data storage [10], environmental remediation
[11, 12] and more recently in biotechnology/biomedicine [13]. Some of the more
specific biomedical applications of magnetic nanoparticles include their use as
magnetic contrast agents in magnetic resonance imaging (MRI), hyperthermia agents,
where the magnetic particles are heated selectively by application of a high frequency
oscillating magnetic field, and magnetic drug delivery.

In most biomedical applications, magnetic nanoparticles perform best when the size
of the nanoparticles is around 10–20 nm [14]. Each nanoparticle then becomes a
single magnetic domain and shows superparamagnetic behavior when the temperature
is above the Curie temperature. Superparamagnetism is a phenomenon by which
magnetic materials may exhibit a behavior similar to paramagnetism even when at
temperatures below the Curie or the Néel temperature. The energy required to change
the direction of the magnetic moment of a superparamagnetic particle is comparable

to the ambient thermal energy. At this point, the rate at which the particles will
randomly reverse direction becomes significant. Such individual nanoparticles have a
large constant magnetic moment and behave like a giant paramagnetic atom with a
fast response to applied magnetic fields with negligible remanence (residual
magnetism) and coercivity (the field required to bring the magnetization to zero).
These features make superparamagnetic nanoparticles very attractive for a broad
range of biomedical applications because the risk of forming agglomerates is
Chapter 2
- 5 -

negligible at room temperature. Superparamagnetic iron oxide (SPIO) nanoparticles
are small synthetic γ-Fe
2
O
3
or Fe
3
O
4
particles with a core size of ~10 nm and an
organic or inorganic coating. Superparamagnetic magnetization is, compared to
normal paramagnetic materials, much higher and can reach nearly the magnetization
saturation of ferromagnetic iron oxide [15]. This behavior allows the tracking of such
particles in a magnetic field gradient without losing the advantage of a stable colloidal
suspension.

Nanotechnology has allowed for the production, characterization and
functionalization of magnetic nanoparticles for specialized clinical applications.
Extensive research has been done on use of magnetic nanoparticles for bio-
applications in recent years. For instance, dendrimer modified magnetic nanoparticles

have been synthesized to improve protein binding [16]. Iron oxide nanoparticles
coated with insulin have been prepared for exact drug delivery [17], while modified
metal oxide-based nanoparticles were developed for conjugation with cell targeting
agents [18]. Others also demonstrated the use of magnetic nanoparticles in imaging
[19] and gene delivery [20, 21]

applications.

The main advantages of magnetic nanoparticles in biomedicine are that they can be (i)
visualized by magnetic resonance imaging due to their ability to change the T
1
or T
2

relaxation times of the surrounding tissues; (ii) manipulated by means of a magnetic
field (i.e. in magnetic drug delivery); and (iii) heated in a magnetic field to trigger
drug release or to produce hyperthermia/ablation of tissue.

Chapter 2
- 6 -

The common methods of synthesis of magnetic nanoparticles include co-precipitation,
thermal decomposition, microemulsion and hydrothermal synthesis [14]. Co-
precipitation is a facile and reproducible way to synthesize iron oxides (γ-Fe
2
O
3
or
Fe
3

O
4
) from aqueous Fe
2+
/Fe
3+
salt solutions by the addition of a base under inert
atmosphere at room temperature or at elevated temperature. The size, shape, and
composition of the magnetic nanoparticles depends greatly on the type of salts used,
the Fe
2+
/Fe
3+
ratio, as well as the reaction conditions (temperature, pH and ionic
strength). It is, however, difficult to control the size and shape of the particles with the
co-precipitation method.

Monodisperse magnetic nanocrystals of smaller size can essentially be synthesized
through the thermal decomposition of organometallic compounds in high-boiling
organic solvents containing stabilizing surfactants. Sun et al. have demonstrated the
synthesis of mondisperse MnFe
2
O
4
and Fe
3
O
4
nanoparticles via high-temperature
organic phase decomposition [22, 23] whereas synthesis of CoFe

2
O
4
, MnO, CuO and
other monodisperse nanocrystals of transition metal oxides was reported by Park and
co-workers [24, 25]. In principle, the ratios of the starting reagents including
organometallic compounds, surfactant, and solvent are the decisive parameters for the
control of the size and morphology of magnetic nanoparticles. The reaction
temperature, reaction time, as well as aging period may also be crucial for the precise
control of size and morphology.

Reverse micelles, which are water-in-oil droplets stabilized by a monolayer of
surfactant, have been applied as nanoscale reactors for the synthesis of various
nanoparticles [24, 26]. In water-in-oil microemulsions, the aqueous phase is dispersed
Chapter 2
- 7 -

as microdroplets surrounded by a monolayer of surfactant molecules in the continuous
hydrocarbon phase. The size of the reverse micelle is determined by the molar ratio of
water to surfactant [27]. On mixing two identical water-in-oil microemulsions
containing the desired reactants, the microdroplets will continuously collide, coalesce,
and break again, resulting in mixing of the reactants; finally a precipitate forms in the
micelles. By the addition of solvent, such as acetone or ethanol, to break the
microemulsions, the precipitate can be extracted by filtering or centrifuging the
mixture. In this sense, a microemulsion can be used as a nanoreactor for the formation
of nanoparticles. Using the microemulsion technique, metallic cobalt, cobalt/platinum
alloys, and gold-coated cobalt/platinum nanoparticles have been synthesized in
reverse micelles of cetyltrimethylammonium bromide, using 1-butanol as the
cosurfactant and octane as the oil phase [28]. MFe
2

O
4
(M: Mn, Co, Ni, Cu, Zn, Mg, or
Cd, etc.) are among the most important magnetic materials and have been widely used
for electronic applications.

Under hydrothermal conditions a broad range of nanostructured materials can be
formed. Li et al. reported a generalized hydrothermal method for synthesizing a
variety of different nanocrystals by a liquid–solid–solution reaction [29]. The system
consists of metal linoleate (solid), an ethanol–linoleic acid liquid phase, and a water–
ethanol solution at different reaction temperatures under hydrothermal conditions. The
strategy is based on a general phase transfer and separation mechanism occurring at
the interfaces of the liquid, solid, and solution phases present during the synthesis.
Fe
3
O
4
and CoFe
2
O
4
nanoparticles can be prepared in very uniform sizes of about 9
and 12 nm, respectively. Hydrothermal synthesis is a relatively little explored method
for the synthesis of magnetic nanoparticles, although it allows the synthesis of high-
Chapter 2
- 8 -

quality nanoparticles. To date, magnetic nanoparticles prepared from co-precipitation
and thermal decomposition are the best studied, and they can be prepared on a large
scale.

2.2 MAGNETIC DRUG DELIVERY
Various organic materials (polymeric nanoparticless, liposomes, micelles) have been
investigated as drug delivery nanovectors using passive targeting, active targeting
with a recognition moiety or by a physical stimulus (e.g. magnetism in
magnetoliposomes) [30]. However, these organic systems still present limited
chemical and mechanical stability, swelling, susceptibility to microbiological attack,
inadequate control over the drug release rate [31], and high cost. Polymer
nanoparticles also suffer from the problem of high polydispersity. As-synthesized
particles with a broad size distribution and irregular branching could lead to
heterogeneous pharmacological properties. Due to the disadvantages of organic
nanoparticles for drug delivery, inorganic vectors are gaining much attention in
research.

In a general case of magnetic drug delivery, a drug or therapeutic radionuclide is
bound to a magnetic compound, introduced in the body, and then concentrated in the
target area by means of a magnetic field (using an internally implanted permanent
magnet or an externally applied field). Depending on the application, the particles
then release the drug or give rise to a local effect [30]. Drug release can proceed by
simple diffusion or take place through mechanisms requiring enzymatic activity or
changes in physiological conditions such as pH, osmolality, or temperature [32]; drug
Chapter 2
- 9 -

release can also be magnetically triggered from the drug-conjugated magnetic
nanoparticles.

In biomedicine, one major hurdle that underlies the use of nanoparticle therapy is the
problem of getting the particles to a particular site in the body [1]. A potential benefit
of using magnetic nanoparticles is the use of localized magnetic field gradients to

attract the particles to a chosen site, to hold them there until the therapy is complete
and then to remove them. The particles may be injected intravenously, and blood
circulation used to transport the particles to the region of interest for treatment.
Intravenous administration of drugs is the most versatile method to reach target
organs and tissues since the blood circulation supplies all vital cells. Alternatively, in
many cases the particle suspension may be injected directly into the general area
where treatment is desired. Either of these routes has the requirement that the particles
do not aggregate and block their own spread.

The size, charge, and surface chemistry of the magnetic particles are particularly
important and strongly affect both the blood circulation time as well as bioavailability
of the particles within the body [33]. In addition, magnetic properties and
internalization of particles depend strongly on the size of the magnetic particles. For
example, following systemic administration, larger particles with diameters greater
than 200 nm are usually sequestered by the spleen as a result of mechanical filtration
and are eventually removed by phagocytic cells, resulting in decreased blood
circulation times. On the other hand, smaller particles with diameters of less than 10
nm are rapidly removed through extravasation and renal clearance. Particles ranging
from circa 10 to 100 nm are optimal for in vivo injection and demonstrate the most
Chapter 2
- 10 -

prolonged blood circulation times [34]. The particles in this size range are small
enough both to evade the MPS of the body as well as penetrate the very small
capillaries within the body tissues and, therefore, may offer the most effective
distribution in certain tissues [35].

The MPS is a cell family consisting of bone marrow progenitors, blood monocytes
and tissue macrophages. These macrophages are widely distributed and strategically
placed in many body tissues to recognize and clear senescent cells, invading micro-

organisms or particles [36]. After particles are injected into the bloodstream they are
rapidly coated by components of the circulation, such as plasma proteins. This process
is known as opsonization, and is critical in dictating the fate of the injected particles
[37]. Normally opsonization renders the particles recognizable by the body’s MPS
and results in their subsequent clearance by the macrophages. As a result, the
application of nanoparticles in vivo or ex vivo would require surface modification that
ensures particles are non-toxic, biocompatible and stable to the MPS.

As conventional colloidal drug delivery systems are rapidly removed from the blood
stream after their intravenous administration, the fate of nanoparticles in the blood
compartment have therefore been widely discussed, and the development of surface-
modified nanoparticles appears crucial to increase circulation time [38]. The first
requirement for active targeting is minimizing or delaying the phagocytosis of the
nanoparticles by the MPS. Macrophage-evading particles increase the probability of
attaining the desired target. Modified nanoparticles should present surfaces which
inhibit complement activation and opsonization by plasma proteins as these are the
key factors involved in uptake of particles by the MPS. Particles that have a largely
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hydrophobic surface are efficiently coated with plasma components and thus rapidly
removed from the circulation, whereas particles that are more hydrophilic can resist
this coating process and are cleared more slowly [39]. This has been used to an
advantage when attempting to synthesize MPS-evading particles by sterically
stabilizing the particles with a layer of hydrophilic polymer chains. The most common
coatings are derivatives of dextran, polyethylene glycol (PEG), polyethylene oxide
(PEO), poloxamers and polyoxamines [40]. The role of the dense brushes of polymers
is to inhibit opsonization, thereby permitting longer circulation times.

Besides uptake by the MPS, there are other fundamental problems associated with the

use of magnetic directed drug targeting. Targeting to a specific cell type, for example,
may be possible with directed coatings. However, retaining the particles at the cell
membrane for exact drug localization or magnetic cell separation and recovery [17]
for any length of time is difficult as the cell often instigates receptor-mediated
endocytosis. In addition, the ability of magnetic particles to concentrate will depend
on both the blood flow rate and the intensity of the magnetic field. The success
therefore depends to a large extent on the construction of strong magnets able to
produce high magnetic field gradients at the target sites. It has been shown that most
of the available fields are only strong enough for the manipulation of particles against
the diffusion and bloodstream velocities found in living systems over a distance of a
few centimeters from the sharp end of a magnet pole [41]. This means that it is
difficult to build up and sustain field strength sufficient to counteract the linear blood
flow rates in tissues so as to effectively retain the drug carrier at a specific location.
Improvements are needed to make magnetic directed drug targeting effective.


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2.3 CANCER TARGETING
Many bio-applications require magnetic particles to possess a cell targeting property
especially in the case of cancer diagnosis and treatment. Active targeting is based on
the over or exclusive expression of particular epitopes or receptors in tumoral cells,
and on sensitivity or response to stimuli such as temperature, pH, electric charge, light,
sound or magnetism. Active targeting may also be based on species which bind to
over-expressed receptors. These species include low molecular weight ligands (folic
acid, thiamine, sugars), peptides (RGD, LHRD), proteins (transferrin, antibodies,
lectins), polysaccharides (hyaluronic acid), polyunsaturated fatty acids, DNA, etc.
Efforts in conferring cell targeting property onto the magnetic nanoparticles include
the modification of the particles with chemotherapeutic agents [32, 42], ligands [42,

43] and antibodies [44]. These agents usually enter the target cells via endocytosis.

Endocytosis is an important pathway by which drugs and chemotherapeutic agents
enter cells. It is also the most studied mechanism for cell targeting and uptake. Cells
will endocytose solutes from their extracellular environment through one of the
following processes [45, 46]: (i) fluid-phase pinocytosis, where the solute to be
endocytosed is present within the extracellular fluid bathing the cell surface and some
of the extracellular fluid is captured within the lumen of the budding endocytic vesicle;
(ii) adsorptive endocytosis, where the solute that is to be endocytosed binds to the cell
surface through non-specific mechanisms; or (iii) receptor-mediated endocytosis,
where a solute will bind to its cognate cell membrane receptor to elicit either a
constitutive or ligand-stimulated internalization.

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Human CD38 antigen, a 42–45 kDa type II transmembrane glycoprotein which is
upregulated on cell surfaces in many lymphoid tumors, is a promising candidate in
antibody therapy. Orciani and co-workers have demonstrated that coupling liposome
and an anti-CD38 antibody does not interfere with CD38 functionality [47]. Their
results showed a specific mechanism owing to CD38-mediated endocytosis of the
immunoliposome.

The human epidermal growth factor receptor (EGFR) is a group of transmembrane
receptors consisting of four related members: HER1 (EGFR), HER2 (also known as
c-erbB-2 and neu), HER3, and HER4. EGFR and HER2 are two important receptors
frequently employed in the treatment of metastatic breast cancers. The recombinant
humanized monoclonal antibody trastuzumab (trade name: herceptin) was developed
as an immunotherapeutic agent against the HER2 extracellular domain. It binds to the
HER2 receptor and reduces tumor cell proliferation by interrupting the cellular

pathway [48]. When conjugated to poly(lactic acid) nanoparticles, an efficient
internalization of the particles was observed in human ovarian carcinoma cells
(SKOV-3) overexpressing HER2 [49]. In a separate study by Germershaus et al.,
transfection experiments using human breast cancer cells (SK-Br-3) showed up to
seven-fold higher luciferase expression with trastuzumab-conjugated complexes as
compared to the complexes without trastuzumab at N/P=3.5 [50]. Reporter gene
expression was significantly inhibited by increasing trastuzumab concentrations. This
efficient inhibition with free trastuzumab verified the HER2-receptor dependency of
the reporter gene expression.