Download Magnetic nanoparticles for MRI applications in medicine

Magnetic nanoparticles for MRI applications in medicine
Marc-André Fortin*
Laboratory for Biomaterials in Imaging, Axe médecine régénératrice, Centre de
recherche du Centre hospitalier universitaire de Québec (CR-CHUQ), 10 rue de
l’Espinay, Québec, G1L 3L5, Canada, Département de génie des mines, de la métallurgie
et des matériaux and Centre de recherche sur les matériaux avancés (CERMA)
Université Laval, Québec, G1V 0A6, Canada
E-mail: [email protected]
Abstract
Magnetic resonance imaging (MRI) has developed at an exponential rate since the last
decades, and is now widely used as an anatomical and functional medical imaging
modality. The development of magnetic nanoparticles as contrast agents for vascular,
molecular, and cellular MRI applications has followed this trend. Nanoparticles that
generate either “positive” or “negative” contrast in MRI figure among one of the most
important direct applications of nanotechnology. The present chapter is an introduction to
the principles underlying the performance of nanoparticulate-based MRI contrast agents.
It addresses the main considerations guiding the design, the synthesis, and the physicochemical characterization of magnetic nanoparticles based on the elements iron,
manganese and gadolinium (Fe, Mn, Gd). The fundamental aspects of nanoparticle
magnetism and relaxometric characterization are introduced, as well as examples of
applications in biological models.
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Table of contents
1. 2. 3. 4. INTRODUCTION TO NANOPARTICLES FOR MRI APPLICATIONS ........................................................ 3 THE BASICS OF MRI IN MEDICINE .................................................................................................... 7 RELAXIVITY: THE PERFORMANCE OF MRI CONTRAST AGENTS ......................................................... 8 SYNTHESIS AND CHARACTERIZATION OF MAGNETIC NANOPARTICLES ............................................ 9 A. SYNTHESIS OF MAGNETIC NANOCRYSTALS ..................................................................................................... 9 B. NANOPARTICLE COATINGS FOR MRI APPLICATIONS ...................................................................................... 14 C. PHYSICO-­‐CHEMICAL CHARACTERIZATION ..................................................................................................... 16 5. PHYSICAL PROPERTIES OF MAGNETIC NANOPARTICLES ................................................................ 16 6. MR RELAXATION PROPERTIES OF MAGNETIC NANOPARTICLES ..................................................... 22 A. RELAXIVITY OF PARAMAGNETIC CAS .......................................................................................................... 23 B. RELAXIVITY OF SUPERPARAMAGNETIC CAS .................................................................................................. 26 C. RELAXOMETRIC PERFORMANCE OF MRI CAS AT CLINICAL MAGNETIC FIELD STRENGTHS ....................................... 28 7. BIOLOGICAL PERFORMANCE OF MAGNETIC NANOPARTICLES FOR MRI ......................................... 29 A. IN VIVO BARRIERS .................................................................................................................................. 29 B. IMPACT OF NANOPARTICLE SIZE AND SURFACE ON COLLOIDAL STABILITY AND BLOOD RETENTION ............................ 30 C. DIRECTING NANOPARTICLES IN VIVO .......................................................................................................... 32 D. TOXICITY .............................................................................................................................................. 33 8. CONCLUSION ................................................................................................................................ 34 9. REFERENCES ................................................................................................................................. 35 2
1. Introduction to nanoparticles for MRI applications
Nanoparticles (NPs) made of the elements iron (Fe), gadolinium (Gd) or
manganese (Mn) are currently used in many diagnostic applications performed under
magnetic resonance imaging (MRI). In fact, MRI scanners do not directly detect
magnetic nanoparticles (MNPs): they reconstruct images from the electromagnetic
signals generated by the stimulation, and relaxation, of the large pool of hydrogen (1H)
protons found in biological tissues. By interactions taking place at the atomic and
molecular level, MNPs influence the relaxation time of hydrogen protons contained in
small, mobile molecules such as water. Thereby, MNPs induce contrast enhancement
effects on the reconstructed MR images, either as a signal increase (“positive” contrast
agents, CAs) or a signal decrease (“negative” CAs).
In fact, MRI is one of the most reliable, high resolution (75 – 300 µs), and multifunction imaging modalities of modern medicine. Compared to optical imaging and
echography, it allows the acquisition of in-depth, whole-body images, from the mouse
model to the humans. Contrarily to compute tomography (CT), another anatomical
imaging modality, it does not rely on ionizing radiation. MRI also enables the tracking of
cells, molecules, and drug delivery vehicles in the human body. These applications
require the development of appropriate biomedical imaging probes (CAs or tracers). Such
probes must be sensitively detected at the nanomolar concentration level, to allow the
efficient detection of a reasonable amount of molecules, or cells, in the body.[1-3]
Compared to MRI, positron emission tomography (PET) and luminescence/fluorescence
imaging do not provide anatomical images, and neither are they considered as high
resolution imaging modalities in general. However, both are truly efficient “molecular”
imaging modalities (nanomolar detection). Unfortunately, MRI has a relatively poor
sensitivity compared with such techniques. To fully exploit the wide range of advantages
that MRI has to offer in molecular and cellular imaging applications, it has been
necessary to design CAs which have the capacity to very efficiently interact with
hydrogen protons.
Contrast agents (CAs) have been developed since the inception of MRI, to
selectively change the longitudinal and transverse relaxation times (T1 and T2) of 1H in
biological tissues. Such energy transfers mainly occur though interactions taking place
between the magnetic elements and the spins of hydrogen protons in their vicinity. Most
clinically approved CAs are based on small molecules that sequestrate the paramagnetic
ion Gd3+.[4, 5] They are mainly used as non-specific agents, to enhance the general
contrast of organs, enabling thereby a better identification of anatomical changes
occurring in the body. They are also applied to blood-pool and blood perfusion
procedures. Gadolinium has 7 unpaired electrons in its 4f orbitals giving it a very large
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magnetic moment; a relatively slow electronic relaxation rate compared with other
paramagnetic elements, also enhances its proton relaxation properties.[6] Manganese can
also be exploited in MRI applications, although its magnetic moment is weaker than that
of Gd3+ (Mn2+ has 5 unpaired electrons on its 3d orbital). After the first marketing
authorization of Gd-DTPA (Magnevist™) in the US, Europe and Japan in 1988, other
Gd-based chelates were introduced to the market: Dotarem™, based on the DOTA
chelator, as well as Omniscan™, Prohance™, Optimark™, and Gadovist™.[7] Today,
contrast media are applied in approximately 30% to 40% of all MRI procedures.
In parallel with the development of paramagnetic CAs, progress in the field of MRI
CAs also occurred through the unique properties of iron oxide NPs. By opposition to
paramagnetic molecules, which produce “positive” contrast in MRI, iron oxide NPs are
well known for their potential to increase the signal. Iron oxide NPs are, in general,
divided in two classes: small particles of iron oxide (SPIO), and ultra-small particles of
iron oxide (USPIO). SPIOs are made of iron oxide cores of mean diameters in the range 3
– 20 nm; however they form agglomerates of hydrodynamic diameter typically superior
to 50 nm. The concept of hydrodynamic diameter (Figure 1) refers to the total effective
diameter of a particle suspended in a fluid and forming a colloid. The hydrodynamic
diameter is generally measured by means of laser analysis (dynamic light scattering,
DLS), by evaluating the Brownian motion of the suspended NPs. USPIOs on the other
hand, refer to the class of iron oxide NPs that are made of iron oxide cores (3 – 20 nm
diam.) surrounded by a coating of molecules that preserve their individuality (mean
diameter < 50 nm).
Figure 1: Schematic representations of a) the structure of functional magnetic nanoparticles (MNPs) for
MRI applications and b) a colloidal nanoparticle suspended in biological media.
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SPIOs and USPIOs have very different mean hydrodynamic diameters, as well as
different “relaxometric” properties. Such properties can be advantageously exploited
either in cell labeling (preferentially with SPIOs), liver cancer diagnostic (preferentially
with SPIOs), molecular targeted imaging (preferentially with USPIOs), or blood
pool/angiography procedures (preferentially with USPIOs). In fact, a very large fraction
of molecular and cellular MRI applications are based on the design and fabrication of
MNPs consisting of a) an inorganic nanocrystal core (e.g. Fe2O3/Fe3O4, Gd2O3, MnO,
NaGdF4), b) a coating made of small ligands, or biocompatible organic molecules, and c)
a functional anchor that is used to graft specific molecules, complementary imaging
functionalities (e.g. radioactive atoms or fluorescent molecules), or medicinal compounds
for drug delivery. Injected intravenously, SPIOs are captured by macrophages. They end
up in the Kupffer cells of the liver (part of the reticuloendothelial system, RES), where
they are expected to degrade. As a result, an important shortening of the transverse
relaxation time (T2/T2*) is observed in the liver tissue, which is reflected by a strong
signal loss.[8-10] An example of this is provided in Figure 2.
Figure 2 : Intravenous injections of polymer-coated iron oxide NPs in the mouse model (unpublished data,
Fortin et al.). The presence of USPIOs in the blood enhance the vascular signal during at least 20 minutes
(a : arrowhead, b : blood signal-enhancement ratio) : USPIOs can be used as blood-pool agents. After
injection, NPs are gradually sequestrated by the macrophages, and follow the retiduliendothelial (RES)
route, mainly in the liver (a : t = 24 h., full arrow; and c). After several hours, a significant signal
enhancement appears in the gall bladder, an indication of NPs excretion through the hepatobiliairy way.
By this mechanism, iron oxide NPs “passively”, but rather homogenously
accumulate in the liver, where they are found to lower the MR signal. However, tumors
and metastases are void of Kupffer cells, and they do not internalize iron oxide NPs. This
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strategy has been used in the diagnostic of liver cancer and liver metastasis, as well as for
imaging the spleen and of the lymph nodes.[11, 12] Because of their elimination by the
liver, SPIOs are not efficient in applications requiring longer blood half-lives (MR
angiography, tissue perfusion imaging, functional imaging of the brain). For vascular
applications, USPIOs are better candidates since they stay in the blood for longer times.
Their relaxation properties also allow them for use as positive CAs in T1-weighted
imaging (see next section) and angiography.[13-17] In the recent years, the research on
targeted MRI contrast agents has focused on the development of targeted USPIOs that
could enable the efficient, sensitive and selective detection of atherosclerosis, apoptosis,
and amyloid deposition in Alzheimer’s disease.[18-25] This concept is frequently
referred to as molecular targeted imaging.
In parallel with iron oxide NPs, a new generation of paramagnetic NPs has been
developed, based on Gd and Mn-containing inorganic nanocrystals.[26, 27] These are
based on the synthesis of Gd2O3[28-32], NaGdF4[33-36] and MnO [37-39] nanocrystals.
When adequately covered with biocompatible ligands and suspended in aqueous
solutions, these are “positive” CAs in MRI. Because the core of inorganic nanoparticles
is very small (< 20 nm), and because they contain hundreds or thousands of paramagnetic
atoms, they can be used to label molecules or cells that are then much more sensitively
detected than if they were labeled by traditional MRI contrast agents containing only one
paramagnetic atom per unit of contrast agent (e.g. commercially available Gd-DTPA, or
Gd-DOTA).[4, 5] Although they are not as sensitively detected as their “negative” CA
counterparts based on iron oxide NPs, they provide contrast enhancement effects that are
largely devoid of magnetic susceptibility image artifacts. In addition to this, the detection
of signal-enhancement effects generated by “positive” CAs could enable more
quantitative molecular and cellular imaging studies. Until now, CAs based on Gd and Mn
nanocrystals, have been largely constrained to pre-clinical studies because of the inherent
risks associated with the leaching of potentially toxic Mn2+ and Gd3+ ions. They
nonetheless represent very useful contrast agents for research purposes in the field of cell
labeling and tracking, as well as for molecular imaging with MRI.
This chapter first describes the general considerations that must be taken into
account in the design and preparation of MNPs for MRI in vivo applications. Synthesis
and coating procedures to prepare suspensions of MNPs of narrow particle size,
appropriate magnetic and relaxometric properties, optimal colloidal stability, and good
biocompatibility, are presented. The contrast-enhancement characteristics of
superparamagnetic and paramagnetic NPs are discussed, in particular their respective use
either as “negative” or “positive” CAs in T2/T2*-weighted or T1-weighted MRI sequences.
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2. The basics of MRI in medicine
A brief introduction to the basic principles of MRI is provided at this step, to clarify
important concepts guiding the design and use of MNP for MRI applications in molecular
and cellular imaging. The reader is invited to expand his knowledge by referring to a
selection of readings on the general topic of MRI. [40, 41] In brief, the strength of the
magnetic field in common MRI scanners, ranges from 1 to 3 Tesla, which is 20 000 to 60
000 times stronger than the Earth’s magnetic field. When a patient is introduced in the
gantry, the hydrogen protons (1H) align their spins along the direction of the magnetic
field (B0, Figure 3.a). The sum of each one of the magnetic moments of these spins
represents the “macroscopic magnetization vector” 𝑀 of the biological tissue. This
vector is globally oriented along the main magnetic field of the scanner. In their initial
state, the spins “precess” at a certain frequency (ω0: the Larmor frequency). They are not
coherent in phase. Then, using a transmitter coil, a radiofrequency (RF) wave tuned at
Larmor conditions, is applied to the biological tissue. This causes the excitation of 1H
spins. At this step, spins lose their preferential orientation along the main magnetic field
(Figure 3.b). After application of the RF excitation, 𝑀 oscillates around the main
magnetic field (the “z” axis of the MRI scanner), and the spins are phase-coherent in the
x-y plane (Figure 3.b).
Figure 3: Schematic representation of the macroscopic magnetization vector generated by MR
excitation.
From this moment, the oscillation motion along the x-y plane is detected by a
receiver coil, and recorded. This represents the “MRI signal”. Then, the “x-y” phase
coherence (Mxy) is gradually lost (within milliseconds) as the magnetic moments of
neighboring 1H protons exert a mutual influence on each other (Figure 3.c). The time
constant used to quantify this loss of phase coherence is called the transversal relaxation
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time (T2, Figure 3.d). Independently from this mechanism, the excited spins progressively
release their energy and recover their initial orientation along the main magnetic field of
the scanner (Mz recovery, Figure 3.c). This return to the initial macroscopic
magnetization state occurs within a time constant that is referred to as the longitudinal
relaxation time (T1, Figure 3.d). Both T1 and T2 are intrinsic characteristics of any
biological tissue and, together with the density of 1H spins in the tissue (ρ), are the most
important parameters influencing the signal. For instance, the signal recorded for a given
tissue (S), using a basic spin-echo sequence, is given by the following equation:
S = ρ (1− e
(−TR/T1 )
)(e
−TE/T2
)
Equation 1
where TR and TE are the repetition and the echo times (parameters in the spin-echo
sequence, [41]). Interactions between the paramagnetic elements Fe3+, Gd3+, Mn2+ take
place at the molecular level, causing T1 and T2 to decrease and, in turn, to modify the MR
signal according to Equation 1. Such interactions accelerate the release of energy
communicated to 1H protons during the RF excitation.
3. Relaxivity: the performance of MRI contrast agents
MNPs made of Fe, Gd, or Mn, influence the macroscopic relaxation times (T1 and T2)
of H contained in neighboring mobile molecules. In turn, this effect modulates the MRI
signal (Equation 1), which translates into contrast effects in anatomical MR images. The
efficiency of MRI CAs to decrease both T1 and T2 of hydrogen protons contained in the
solution, is refereed to as the “relaxivity”. Relaxivity depends on the concentration, as
well as on the physico-chemical characteristics of the CAs: hydrodynamic diameters,
number of water binding sites at the magnetic ions, etc. The relaxivity of water protons in
CAs also depends on temperature, pH, and on the magnetic field strength that is used to
perform the MRI scans. Here is a list of suggested readings on the topic of MRI CA
relaxivity.[4, 5, 42-44]
1
In brief, the effect of CAs on the relaxation time of protons (1H protons), are usually
measured by relaxometric analysis, which is the technical term that refers to the
measurement of T1 and T2 of mobile 1H species in aqueous suspensions or in biological
tissues. In MRI, the impact of CAs on the relaxation rate of protons, measured in fixed
conditions of magnetic strength and temperature, is described by the following equation:
!
𝑅! = ! =
!
!
!!!
+ 𝑟! 𝐶 Equation 2
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where Ri=1,2 is the relaxation rate of the aqueous solution, Ti0 is the relaxation time of the
aqueous media in the absence of the CA, ri=1,2 is the relaxivity (usually at T = 20 or 37oC,
pH = 7, and B0 = 1.5 or 3.0 Tesla), and C is the CA concentration (in mM of Gd, Fe, or
Mn). The concentration of magnetic elements is measured by spectroscopic and
spectrometric elemental analysis techniques (e.g. atomic absorption spectroscopy – AAS;
inductively coupled plasma optical emission spectroscopy (ICP-OES); mass
spectrometry, ICP-MS). Therefore, the relaxometric performance of MRI CAs is assessed
first by measuring their relaxation rates (1/T1 and 1/T2), followed by normalizing the data
to the paramagnetic elemental concentration. Relaxivity values (r1 and r2) are extracted
from the slope of the graph given by Equation 2. These are often referred to as relaxivity
curves, and they figure among the most fundamental aspects of MRI CA quantification.
A more detailed explanation of CA relaxivity mechanisms is found in section 6. Finally,
CAs can be divided into two categories, based on the r2/r1 ratio. First, CAs having a
similar impact on both T1 and T2, result in low r2/r1 ratios (close to 1), and a capacity to
enhance the MR signal. This is in agreement with Equations 1 and 2. Such CAs are
referred to as “positive”. On the other hand, CAs that more preferentially decrease T2,
with r2/r1 ratios superior to 5 and often as high as 100, are called “negative”. Examples of
the performance of CAs are found in section 6.
4. Synthesis and characterization of magnetic nanoparticles
Comprehensive reviews have been written, describing the different ways to
synthesize MNPs. These mainly report on iron oxide NPs, and most of these colloidal
synthesis routes can also be adapted to other metal ions such as Mn2+ and Gd3+. The most
important criteria guiding the selection and optimization of a particular colloidal NP
synthesis route, with MRI applications as an objective, is a good control over nanocrystal
size, shape, and NP size distribution (as narrow as possible). A reasonable synthesis yield
is also critical: CA products must be concentrated enough to produce efficient contrast
enhancement effects in MRI procedures; the colloidal synthesis technique must also
minimize the loss of paramagnetic materials and surfactants used during the production,
in order to enable industrial upscale of the process.
a. Synthesis of magnetic nanocrystals
Until now, NP cores have been made from different materials and with varying sizes,
shapes, uniformities, and magnetic properties[45-49]. Apart from MRI applications,
MNPs have been formed from iron and cobalt [50]. Procedures enabling the synthesis of
CoPt3[51] and FePt [52], as well as oxides [44] such as magnetite (Fe3O4) and maghemite
(γ-Fe2O3), have been successfully developed and mastered [44, 53, 54]. Iron oxide NPs
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have also been doped to enhance their magnetic properties to form MFe2O4 structures
where M is a +2 cation such as Mn, Fe, Co or Ni[55, 56]. These MNPs make excellent
MR contrast agents; because their magnetic susceptibility is relatively high, they can also
be manipulated by external magnetic fields. However, NPs containing Co and Ni are
toxic, making them poor candidates for clinical use. To a lesser extent, this also the case
for Gd and Mn-based nanocrystals. However, these have their own advantages and, as
preclinical CAs, can advantageously applied to molecular and cellular imaging in animal
models. For clinical use, the biocompatibility profile of iron oxide NPs is well
established, and adequately documented with a large range of clinical studies. Iron oxide
NPs injected in vivo eventually degrade to their non-toxic iron and oxygen components,
making them particularly attractive as clinical MRI CAs[57].
The present section reports the major colloidal nanoparticle synthesis procedures that
must be selected to enable efficient, rapid, and high-yield production of small particles of
well-controlled and narrow size distributions. For both paramagnetic and
superparamagnetic nanocrystals, core size dictates both the magnetic and the
relaxometric performance of CAs. A good control over this parameter is therefore critical
in all aspects, and bottom-up approaches using chemical colloidal synthesis techniques
are almost invariably privileged. For a matter of concision, and because Fe, Gd and Mnbased materials account for the huge majority of MRI NPs in pre-clinical and clinical
applications, only synthesis routes enabling the production of nanocrystals containing
these three elements, are described here. A selection of references is provided to the
reader who wishes to read on the different variants to these colloidal synthesis
techniques.
Iron oxide nanoparticles: Among the most important colloidal synthesis methods
that are used to fabricate iron oxide NPs, figure co-precipitation, co-precipitation in
constrained environments, thermal decomposition and/or reduction, hydrothermal
synthesis, and polyol synthesis. [44, 49, 58-61]. Each one has its own specific
advantages. The most common, the simplest and possibly the most efficient is that of coprecipitation. It is based on the use of an ageing stoichiometric mixture of ferrous and
ferric ions (Fe2+/Fe3+) in aqueous solutions. The chemical reaction for the formation of
magnetite (Fe3O4) is:
Fe2+ + 2Fe3+ + 8OH- ⇒ Fe3O4 + 4H2O
According to the thermodynamics of this reaction, complete precipitation of Fe3O4 occurs
at basic pH, with a stoichiometric ratio of 2:1 (Fe3+:Fe2+).[62] Therefore, the reaction
takes place with the addition of base under an inert atmosphere.
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In the presence of oxygen, magnetite oxidizes into maghemite, as follows:
Fe3O4 + 2H+ ⇒ γ−Fe2O3 + Fe2+ + H2O
In the co-precipitation technique, NPs are formed by a nucleation and growth mechanism.
NPs of relatively narrow size distributions can be synthesized, provided a short
nucleation event takes place, followed by a slower growth phase. The type of salts
employed (e.g. chlorides, sulfates, nitrates), the ratio between of ferrous and ferric ions,
the temperature, pH and ionic strength figure are all parameters that must be finely tuned
to yield NPs of desired size and narrow distribution.[49] Although they are totally
appropriate to synthesize large amounts of MNPs, standard co-precipitation methods do
not consistently deliver the same products (size, shape, and polydispersity). The presence
of impurities and surface defects also affect the magnetic properties of such NPs.[46]
Adaptations to the co-precipitation approach have been investigated to improve the
uniformity and stability of USPIOs and SPIOs. This was made through addition of
polymers or polyelectrolytes to the ferric/ferrous ion solutions, with various
improvements on size, shape, and crystallinity.[46, 63-66] The addition of chelating
organic ions (carboxylate or α hydroxyl carboxylate ions, such as citric acid, gluconic, or
oleic acid) or polymer surface complexing agents (dextran carboxydextran, starch, or
polyvinyl alcool) during the formation of the magnetite crystals, help controlling NP size.
The chelation of these charged organic molecules on the iron oxide surface can either
prevent the nucleation and lead to larger particles, or inhibit the growth of crystal nuclei,
leading to small nanoparticles. Adding polymers such as poly (acrylic acid) directly into
the synthesis solution and in various concentrations, allows to tune the particle diameter
between 7−14 nm[67]. Alternatively, polyethylene glycol (PEG) compounds such as
PEG-g-poly(glycerol monoacrylate) are also used to modulate the size of USPIOs[68].
Importantly, these polymers may act as surface coatings when the nucleation and growth
processes are complete. These are called in situ coating processes.
To achieve smaller sizes, narrower particle distributions and higher magnetic
properties, new synthesis techniques have been developed, based on high-temperature
decomposition methods using organic iron precursors[69]. For instance, a hightemperature reaction of iron (III) acetylacetonate, Fe(acac)3, in phenyl ether in the
presence of alcohol, oleic acid, and oleylamine, yielded monodisperse, hydrophobic
magnetite NPs with tunable sizes of 4–20 nm[70]. Because one of the major prerequisites of MNPs for MRI applications, is to achieve a good dispersion in aqueous
solvents, an additional step must be introduced in the synthesis procedure, to replace the
hydrophobic coating with an amphiphilic and biocompatible surfactant.
11
The polyol process is an alternative to thermal decomposition methods, for the
synthesis of NPs with well-defined shapes and controlled sizes.[71, 72] Owing to their
high dielectric constants, solvents as polyethylene glycol (PEG) are able to dissolve
inorganic compounds on a large range of temperatures. Polyols also serve as stabilizers to
control particle growth, and o prevent particle aggregation. In polyol synthesis, the
precursor compound is suspended in a liquid polyol, then heated to a given temperature.
During this reaction, the solubilized metal precursor forms an intermediate, and is then
reduced into metal nuclei that lead to NP growth. Good examples of iron oxide NP
synthesis in different polyols (di-,tri-,tetra-) ethylene glycol, can be found in a selection
of recent articles.[73-75] Overall, the nanoparticles synthesized by polyol routes have the
smallest and narrowest size distributions, high water dispersion rates, and higher
magnetization compared with particles produced by more conventional methods.
Paramagnetic nanocrystals (Gd2O3 and NaGdF4): the main advantage of
paramagnetic CAs compared to superparamagnetic NPs, is the possibility to generate
“positive” signal at the accumulation of labeled molecules or labeled cells. Signal
enhancement is in general easier to quantify than signal voids resulting from T2/T2* (e.g.
magnetic susceptibility) effects (generated by iron oxide NPs). A good example of this is
illustrated in Figure 4.[76] PEG-coated Gd2O3 NPs are used to label different types of
cells[77-83], and, in this specific example, F98 brain cancer cells. A total of 3x105
labeled F98 cells were injected in mice brains (caudoputamen), to produce a brain tumor.
This is a frequently used animal model in the field of brain cancer and oncology research.
The same amount of SPIO-labeled cells was injected, to compare the capacity of both
contrast media to allow cell tracking in MRI. Figure 4.a reveals the possibility to
efficiently visualize the area of brain cancer cell implantation, at least 48 hours following
the injection. Because this is a “positive” CA scanned with a T1 -weighted MR sequence,
the anatomical information surrounding the area of cell implantation, is preserved. After
one week, the tumor contours can be efficiently delineated without the use of a CAs. On
the other hand, SPIO-labeled cells are much more sensitively detected than Gd2O3labeled ones (Figure 4.b). However, the susceptibility artifact characteristic of iron oxide
NPs largely exceeds the exact location of the implanted cells. Even after one week, the
image artifact still obliterates important anatomical information directly in the area of the
growing tumor. This is a good example of the potential of paramagnetic NPs, to replace
SPIOs and UPSIOs in niche applications where the preservation of anatomical details is
an issue (such as in cell implantation and tracking studies).
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Figure 4 : Mice brains implanted with 300 000 brain cancer cells (F98) labeled with a) PEG-Gd2O3 (T1weighted MR imaging) and b) SPIOs, (T2-weighted MR imaging).[76]
The very high density of Gd atoms per unit of contrast agent, is an advantage over
macromolecules containing Gd chelates: Gd2O3 and NaGdF4 ultra-small nanoparticles
(2−5 nm diameter) have narrow particle size distributions, and contain hundreds of Gd
atoms per CA unit.[31, 76, 84] The production of Gd2O3 NPs, a pre-clinical CA, has
required the development of advanced colloidal synthesis techniques in high boiling point
alcohols (e.g. di-, tri-, poly-ethylene glycol) [31, 85-88]. The surface of Gd2O3
nanocrystals forms hydroxide in contact with water, and this form is susceptible to leach
potentially toxic Gd3+ ions, as in the case prolonged dialysis procedures.[30] Concerns
related to the potential toxicity of Gd2O3 nanoparticles, even for pre-clinical applications
in animal models, have led to the development of other forms of potentially more stable
paramagnetic nanocrystals. NaGdF4, in particular, is an attractive material due to its high
concentration of Gd atoms in a crystalline form that is less susceptible to degrade in
water. Sodium fluoride NPs are synthesized by thermal decomposition in oleic acid and
octadecene, leading to hydrophobic surfaces that must be subsequently transferred to
aqueous suspensions by using appropriate ligand exchange procedures.[89-91] Rare-earth
fluorides doped with series of lanthanides, have very promising luminescent upconversion properties. In particular, it has been demonstrated that ultra-small NaGdF4
nanocrystals (3 nm diam.) doped with Tm and Tb, can be used for dual MRI and nearinfra-red optical imaging, with a wide array of applications in biomedical research.[34]
Antiferromagnetic MnO nanocrystals: Although Mn2+ ions are paramagnetic, MnO
NPs express an antiferromagnetic behavior.[92] They overall behave as “positive” CAs.
Thermal decomposition has been one of the most widely used and reliable routes to
produce relatively large NP synthesis batches of small and narrow particle size
13
distributions.[38, 93-95] One-pot synthesis techniques have also been developed in high
boiling point solvents, enabling the production of 1–3 nm diameter MnO particles.[87]
b. Nanoparticle coatings for MRI applications
NPs must form stable colloids in physiological media (blood, plasma, lymph, urine)
for being considered in MRI applications. They must also show good biocompatibility,
and prolonged vascular retention. This is particularly critical in the case of to targeted
CAs, that must bring sufficient contrast at the molecular site where they are expected to
bind to a molecular biomarkers (e.g. of atherosclerosis or Alzheimer’s disease). The
ligands and polymers, which are used as particle coatings, must efficiently cover the
particles and promote their individualization. They must also limit the adsorption of
plasma proteins, which is the phenomenon that is precursor to their retrieval from the
blood. Indeed, “opsonized” NPs are quickly recognized and ingested by macrophages.
NPs must be coated with a surface ligand or a polymer that either induces an electrostatic
repulsion between the particles, or exert a strong steric repulsion between them.
Otherwise, particles are susceptible to agglomerate when they are submitted to strongly
ionic conditions. Hence, any injection of magnetic nanoparticles in vivo, implies their
preparation in a milieu that is close to the osmolality conditions of biological media. In
general, the smaller the particle size, the more neutral and hydrophilic its surface, and the
longer its plasma half-life.[96]
The selected biocompatible coatings should not influence the cell viability. They
must be grafted with a functionality (e.g. carboxylates, phosphates and sulfates) that
strongly binds to the NP surface. It should not be too pH-sensitive. For instance, acidic
functions (-COOH) can bind to metal oxide surfaces. However, if the bonding is
monodendate only, their attachment at the particle surface is relatively weak and not
strong enough for many applications. Citric acid is a small and very effective ligand that
binds to the surface of iron oxide NPs through carboxylic binding.[97] In particular, this
strategy was used to synthesize the commercial product VSOP C184 (4 nm core
size).[98] However, citric acid cause a strong degradation of the surface, which affects
the magnetic properties of iron oxide NPs.[99] It is also very rapidly and strongly
deflected to the liver, due to its high surface charge. Other ligand molecules can be used
for the stabilization of iron oxide NPs in aqueous medium (e.g. gluconic acid,
dimercaptosuccinic acid, phosphorylcholine, as well as phosphate and
phosphonates).[100, 101]
14
Among the polymer coatings that most efficiently enhance the blood retention of
SPIOs and USPIOs, figure dextran and (carboxy, carboxymethyl)dextran.[102, 103] In
particular, commercial USPIO products such as ferumoxtran-10 (AMI-227), ferumoxytol,
and Supravist (SHU-555C), are all based on iron oxide of core diameters in the range 4 –
8 nm, and of hydrodynamic sizes not larger than 30 nm.[104] Dextran and starch-coated
iron oxide nanoparticles have shown very good relaxometric properties; however, such
coatings are relatively unstable when submitted to biological media. In order to improve
colloidal stability, biocompatibility and blood retention, PEG is used at the surface of
iron oxide nanoparticles.[105-107] Feruglose (Clariscan) is a commercial product that
was developed using a PEGylated starch coating.[108] Such coatings are hardly
recognized by the macrophage-monocytic system. Paramagnetic Gd2O3 and NaGdF4, as
well as MnO nanoparticles, have also been coated using similar strategies (e.g. citric acid,
dimercaptosuccinic acid, glucoronic acid, PEG).[28, 30, 34, 109-112] PEG is widely
applied as coatings for paramagnetic nanoparticles, through –OH[86, 113], –COOH[31,
114], -silane[30, 109, 115] and –phosphate grafting.[116] Silane coatings, for instance,
could delay the degradation of Gd2O3 nanoparticles submitted to acidic environments.
Unfortunately, silane coatings restrict the optimal water exchange with paramagnetic ions
at the surface of NPs, which is the most important relaxation mechanism of “positive”
CAs (see section 6). PEG-phosphate molecules can be used instead, with the significant
advantage that they are not susceptible to homocondensation such as for silane-based
products. [116] Recently, PEGylated phosphonate dendrons were successfully used to
cover MnO and iron oxide nanoparticles, and this coating strategy provided enhanced NP
excretion profiles through the urinary and gastrointestinal pathways.[117, 118]
Figure 5: a) Electron microscopy study of MnO nanoparticles produced by thermal decomposition, with b)
particle-size distribution and c) hydrodynamic size measurements (from [39]).
15
c. Physico-chemical characterization
After synthesis and ligand exchange, NPs must be carefully cleaned from the residual
magnetic ions (Fe2+, Fe3+, Gd3+, Mn2+ and other) that could contaminate the physicochemical, magnetometric, and relaxometric measurements. This procedure can be
achieved either by dialysis in saline water (e.g. 10 - 154 mM NaCl), by centrifugationfiltration cycles, or by size-exclusion chromatography. Dynamic light scattering (DLS)
measurements are performed directly after the purification process, in order to assess the
overall colloidal stability and to demonstrate the absence of large-size aggregates. No
drastic divergence should be noted between “intensity”- and “number”-weighted results,
as diverging results point to the presence of large-size agglomerates that must be
eliminated prior to further use of the particles. Then, the NPs must ideally demonstrate
the presence of only one major hydrodynamic diameter peak, for their magnetic and
relaxometric properties to be adequately predicted and controlled. To demonstrate the
colloidal stability of the particles, a weeklong DLS assay experiment is necessary, in
parallel with zeta potential measurements (electrostatic charge of NPs). Then, the
particles are submitted to a comprehensive set of physico-chemical characterization
measurements using high-resolution electron microscopy (HRTEM: particle size,
morphology, crystallographic parameters), X-ray photoelectron spectroscopy (XPS:
surface elemental analysis), Fourier transform infrared spectroscopy (FTIR: molecular
groups at the surface of particles; molecular grafting assessment), thermogravimetric
analysis (TGA: mass ratio between inorganic cores and organic coating), 1H-NMR
relaxometry (measurement of T1 and T2), and magnetometric measurements, only to
name the most important techniques. Finally, the elemental concentration (Fe, Gd, Mn)
of colloidal NPs is measured by spectrometric or spectroscopic measurements by AAS,
GF-AAS, ICP-OES, ICP-MS (mentioned in section 3), after careful digestion in
appropriate acidic conditions.
5. Physical properties of magnetic nanoparticles
NPs based on SPIOs, USPIOs, Gd2O3, NaGdF4, and MnO, have core diameters
typically in the range 2 – 20 nm. However, superparamagnetic and paramagnetic
nanoparticles have very different magnetic behaviors, whose influence on their
relaxometric performance is major. USPIOs and SPIOs are said to be superparamagnetic,
because they have no remanence after being introduced into, and retracted from, the
strong magnetic field. Upon introduction in the scanner, the global magnetic moment of
superparamagnetic NPs aligns in the direction of this magnetic field. As soon as the
magnetic field is set back to zero (e.g. when the patient is removed from the scanner), the
magnetic moment of the NP also goes down to zero. This is not the same behavior as for
“bulk” ferromagnetic magnetite/maghemite, which are materials that clearly show a
16
strong remanence. The absence of residual magnetization is a very critical and useful
aspect of superparamagnetism applied to biomedicine. Macroscopically, the magnetic
behavior of superparamagnetic particles is similar to paramagnetism (e.g. Gd2O3,
NaGdF4), except that they feature an exceptionally high magnetic moment per unit of CA
(the “core” of iron oxide particles). This strong magnetization is largely responsible for
the remarkable “negative” CA properties of iron oxide NPs. In fact, the absence of
magnetic remanence for USPIOs and SPIOs, is due to the return to equilibrium of the
magnetic moments through Néel relaxation. On the other hand, paramagnetic NPs, and to
a lesser extent antiferromagnetic MnO nanocrystals, do not develop strong magnetization
at clinical magnetic field strengths. Instead, they generate signal enhancement (“positive”
contrast) mainly through direct interactions taking place between Gd and Mn ions, and
mobile 1H protons in their surroundings, as described in section 6. Here are introduced
the basic differences in magnetism between superparamagnetic and paramagnetic NPs.
a) Superperparamagnetic NPs, are very efficient “negative” CAs in MRI.[119] Because
of their strong impact on the transverse relaxation times (T2/T2*) of aqueous solutions,
they have for been considered very useful products for molecular and cellular MRI.[103]
As mentioned above, each superparamagnetic iron oxide NP features a high magnetic
moment, and Curie constants much higher than for paramagnetic NPs. As a result, they
respond quickly to the application of an external magnetic field, and their magnetization
quickly becomes saturated at relatively low magnetic field strengths (Figure 6).
Figure 6 : Magnetometric measurements of USPIOs and Gd2O3 nanoparticles (adapted from [44, 120]).
17
Superparamagnetism only occurs when NPs are small enough to belong to single
magnetic domains. It is worth mentioning that suspensions of iron NPs (and not iron
oxide nanoparticles) would have a much higher magnetization than magnetite/maghemite
(about 5 times higher); however, iron NPs are every quickly oxidized into iron oxide NPs
in aqueous media, and this potential technology, until now, could not be applied in
biomedicine. Magnetite (Fe3O4) and maghemite (γ-Fe2O3) are two relatively similar
forms of iron oxide (crystal structure and magnetic properties).[121, 122] Both are
present in superparamagnetic iron oxide NPs, and it is often difficult to distinguish
between them using common X-ray diffraction analysis techniques. Magnetite is
typically preferred due to its superior magnetic properties [44]. Maghemite (Fe3+ [Fe2+
Fe3+ ]O4) often results from the oxidation of magnetite (Fe3+[Fe3+5/3 V1/3]O4, where V
represents a cation vacancy). Bulk magnetite is ferromagnetic. The occurrence of an
oxygen-mediated coupling mechanism aligns all the magnetic moments of the iron ions
located in the tetrahedral sites of the crystal (8 crystallographic sites per unit structure),
whereas all the magnetic moments of the octahedral ions (16 crystallographic sites per
unit structure) are aligned in the opposite direction. It is assumed that the magnetic
properties of magnetite are provided by uncompensated Fe2+ ions, whereas for
maghemite, they are provided by that of Fe3+ ions.[123]
The magnetic energy of iron oxide NPs, depends upon the direction of its
magnetization vector, and this vector in turn depends on the crystallographic directions
(the magneto-crystalline anisotropy field).[121] The directions that minimize the
magnetic energy are called anisotropy directions, or easy axes (Figure 7.a, from [121]).
The resulting magnetic moment of a magnetite/maghemite crystal is preferentially
aligned along such specific directions. The magnetic energy increases with the tilt angle
between the magnetic vector of the easy directions.[124] For the sake of simplicity, the
anisotropy of magnetite particles is often assumed to be uniaxial, with a single anisotropy
axis. In fact, there are several anisotropy axes dictated by the oxide’s crystallographic
structure. The anisotropy energy (the amplitude of the curve), is given by the product of
the crystal volume (V) times a constant (Ka : the anisotropy constant).
𝐸! = 𝐾! 𝑉
Equation 3
The anisotropy energy, proportional to V, determines the Néel relaxation time.
Large samples of bulk ferrimagnetic magnetite/maghemite, are divided into Weiss
domains (represented in Figure 7.b) Inside each one of these volumes, the magnetic
moments are aligned into one preferential direction; however, between each one of these
domains, the magnetic moment is not oriented along the same direction. As iron oxide
18
nanocores (such as in USPIOs) are smaller than one of these domains, each NP is
composed of a single domain whose magnetic moment is oriented in a specific direction.
Figure 7 : Magnetic behavior of iron oxide NPs (radius = 5 nm): a) uniaxial anisotropy for
magnetite/maghemite nanoparticles (i.e. the probability of alignment of the magnetic moment in one
direction with respect to the angle between this direction and the anisotropy axis); b) representation of
Weiss domains in a large magnetite/maghemite crystal, compared with the dimensions of a typical NP (the
small circle), much smaller than a Weiss domain; c) magnetization versus magnetic field for
magnetite/maghemite NPs. With permission from [100, 121, 125].
In these single domains, the direction of the magnetic moment can flip from one
orientation to the other. When the thermal energy, given by kT (k : Boltzman constant; T:
absolute temperature) is sufficient to overcome this anisotropy energy barrier, the
magnetization fluctuates between the different anisotropy directions, according to a
characteristic time: the Néel relaxation time (τN).[126] Although τN relaxation influences
the hydrogen relaxation times by inducing changes to the magnetic moment of MNPs, it
is a phenomenon entirely distinct from the nuclear relaxation mechanisms of hydrogen
protons (1H) described in section 6. Néel relaxation refers to the relaxation of the global
electronic moment of a superparamagnetic crystal constituted by a ferri, ferro, or
19
antiferromagnetic compound. For dry powders of monodomain iron oxide NPs, τN
indicates the time it takes to the magnetization to come back to a state of equilibrium,
after it was submitted to a strong magnetic field. For highly anisotropic crystals, the
crystal magnetization is “locked” in the easy axes. The Néel relaxation defines the rate of
fluctuations that arise from the jumps of the magnetic moment between the different easy
axes (Figure 7.c). In order to flip from one easy direction to the other, the magnetization
of a NP must jump over an anisotropy energy hump. For a superparamagnetic NP of
specific V and Ka, the Néel relaxation time (τN) is given by an Arrhenius law that is
similar to that describing the activation energy for a chemical reaction [127]:
!!
𝜏! = 𝜏! 𝐸! 𝑒 !"
Equation 4
where τ0(Ea) is the pre-exponential factor of the Néel relaxation time expression, that
depends on factors such as the volume (V), the specific magnetization of the nanocrystal,
and the gyromagnetic ratio of the electron.[44, 128, 129] Whereas the pre-exponential
factor decreases as the value of anisotropy energy increases, τN increases as an
exponential function of V because of the second factor of Equation 4. For small values of
the anisotropy energy and at high temperatures, Ea << kT (the exponential term tends to
1), and τN is mainly determined by the preexponential term. These conditions are
fulfilled, for instance, with individual ultra-small NPs of iron oxide of r < 4 nm. On the
other hand, for high anisotropy energies, when Ea >> kT the evolution of τN is mainly
dictated by the exponential factor (fast increase with Ea).
According to equation 4, the flipping of the magnetic moment of
magnetite/maghemite crystals is observed only for NPs of size r < 12 nm. Indeed, for
magnetite (τ0 ≈ 10−9 s, K ≈ 13500 J m−3), τN goes from ~500 years for particles of r =15
nm, down to the ms for particles of about r = 10 nm. Practically, this means that for
particles having a Néel relaxation time (τN) longer than the measurement time, the
magnetization curve of the NP system is irreversible and shows an hysteresis loop. These
are referred to as “frozen single domain” NPs. NPs that demonstrate Néel relaxation
times of several years, or centuries, can be used in computer hard disks. Indeed, such
long-term data storage applications require as less as possible magnetic material, with a
maximal anisotropy constant (Ea).
Such NPs cannot be used in MRI applications. First, the Néel relaxation of larger
nanoparticles showing high Ea, has no effect on the nuclear relaxation of neighboring
water protons. Indeed, the diffusive rotation time of the particles in water, and the
diffusion time of water around the particles are much shorter than 1 ms. This is
incompatible with the long Néel relaxation times of large particles.
20
For magnetic iron oxide NPs to provide efficient properties as MRI CAs, the
condition τN < 1 s should be respected. Then, for NPs dispersed in a liquid media (a
colloid), the return of the magnetization to equilibrium after application of a strong
magnetic field is determined by both τN and the Brownian relaxation τB of the particles.
The latter characterizes the rotation of the particle in the fluid, and takes into account the
viscosity of the solvent (Figure 7.c). The global magnetic relaxation rate is a sum of two
processes:
!
!
!
!
!
!
=! +!
Equation 5
where τ is the global magnetization time and τB is the Brownian relaxation time, given
by:
𝜏! =
!!"
Equation 6
!"
where η is the viscosity of the solvent. For large particles, τB < τN because the Brownian
component of the magnetic relaxation is proportional to the crystal volume (Equation 6)
and the Néel relaxation is an exponential function of the volume (Equation 4). Therefore,
in suspensions of large NPs, the viscous rotation of the particles in the liquid becomes the
dominant process determining the global magnetic relaxation properties (Figure 7.d) As a
result, the magnetic relaxation of suspensions of magnetic nanoparticles is much faster
than for dry powders of iron oxide NPs.
In resume, superparamagnetism refers to a specific magnetic condition for which
ultra-small particles of a size well inferior to typical Weiss domains of
magnetite/maghemite materials, can be submitted to high magnetic field strengths,
without evidences of magnetic remanence once they are retrieved (Figure 6).
Macroscopically, the magnetization of iron oxide NPs suspensions is described by a
Langevin function, whose shape depends on the saturation magnetization (Msat) and the
size of the magnetite crystals:
𝑀 𝐵! = 𝑀!"# 𝐿(𝑥)
Equation 7
where Msat is the magnetization at saturation, and L(x) is the Langevin function as:
!
𝐿 𝑥 = coth 𝑥 − !
Equation 8
with :
21
𝑥=
!! ! !!!
!"
Equation 9
Magnetization curves of ultra-small iron oxide NPs are perfectly reversible
because the fast magnetic relaxation allows the system to be always at thermodynamic
equilibrium.[130] Finally, because iron oxide superparamagnetic NPs obey this law, it is
possible to estimate the NP core size from the fit of the magnetization curves (Figure 6).
b) Paramagnetic nanoparticles also respond to the application of an external magnetic
field, by developing a magnetization vector oriented along the direction of this field, that
slightly increased the local magnetic field strength.[131] Paramagnetic nanocrystals
follows Curie’s law:
!
𝑀 = 𝜒𝐻 = 𝐻
!
Equation 10
where χ is the magnetic susceptibility, H the applied magnetic field (e.g. that of the MRI
scanner), T the absolute temperature, and C a material-specific Curie constant. Unlike
ferromagnets, paramagnetic materials do not retain magnetization in the absence of an
external magnetic field, and the thermal energy is sufficient to randomize the induced
magnetization. At similar concentrations of metal elements (Fe, Gd, or Mn), the “positive
contrast effect” of paramagnetic NPs is less efficiently detected that the “negative
contrast” effect generated by superparamagnetic agents. The more limited magnetization
response of the rare-earth ions, compared with USPIOs and SPIOs, is due to their
magnetic moment that is not saturated at magnetic fields strengths typical used in
MRI.[6, 132] The difference of magnetization response between iron oxide (USPIOs) and
paramagnetic Gd2O3 nanoparticles is shown in Figure 6.
6. MR relaxation properties of magnetic nanoparticles
For a given concentration of magnetic atoms, the magnetization of paramagnetic
CAs is much lower than that of superparamagnetic nanocrystals. However, Gd chelates
are still, by far, the most widely used CAs in routine MRI. This is mainly due to their
exceptional relaxometric properties. Both paramagnetic and superparamagnetic CAs
interact with the small, hydrogen-containing mobile molecules contained in biological
media. In fact, one of the most important factors influencing signal enhancement in MRI,
is the binding of water molecules to the paramagnetic ions. Then, the motion of contrast
agents in the fluid also has an impact, as well as the magnetic field strength-dependent
electronic relaxation of the paramagnetic elements Finally, Gd3+ chelates are, in general,
very efficiently excreted by the kidneys, and this also accounts for their widespread use
in clinical MRI. This section does not aim at bringing a comprehensive explanation of the
22
complex relaxometric mechanisms taking place between paramagnetic – or
superparamagnetic – CAs, and the water molecules. Several books and reviews have
described this subject.[5, 6, 44, 133-135] Here is a resume of the most important aspects
that must be taken into account when designing optimal nanoparticulate contrast agents
for MRI applications.
a. Relaxivity of paramagnetic CAs
Theories to explain the relaxivity of paramagnetic CAs, have been developed since the
inception of MRI. In fact, the efficiency of CAs is linked to molecular motions of the CA
unit, as well to the motions of small molecules containing the “spins” (1H protons). They
are also linked to intrinsic properties of the nuclei (magnetic moment, gyromagnetic ratio,
spin). The relaxation of paramagnetic solutions is mainly explained by two mechanisms:
the “inner sphere” (IS) and “outer sphere” (OS) contributions.[6] From Equation 2, it is
possible to dissociate two contributions for Ti, one from the water protons in contact with
the paramagnetic elements (p), and one for the rest of the water protons in the matrix
(diamagnetic contribution, d):
!
!!,!"#
=
!
!!,!
+
!
Equation 11
!!,!
Where, i = 1, 2. From the paramagnetic contribution, it is possible to dissociate the IS and
the OS contributions, as follows:
!
!!,!"#
=
!
!!,!
!"
+
!
!!,!
!"
Equation 12
The IS relaxation relies on the exchange of energy between the spins and the electrons of
the paramagnetic elements, which is facilitated when water molecules bind to the
paramagnetic ions (Figure 8). The water molecules that bind to the paramagnetic centers
(Gd3+, Mn2+) ions, and denoted “p” water molecules (red circles in Figure 8), rapidly
leave the first coordination sphere, and are immediately replaced by “fresh” molecules
from the matrix (“d” water molecules: orange circles in Figure 8). In contact with the
paramagnetic ions, hydrogen relax faster. The water residence time in the inner sphere
(τM) is in the order of ∼1 ns and this means the relaxation effect propagates very fast to
the rest of the solution (to “d” protons). Each one of the water protons that relax energy,
participate in the decrease of the overall longitudinal relaxation time of the water solvent.
The IS model has been described by the Solomon-Bloembergen-Morgan theory
(SBM).[136, 137]
23
Figure 8: Schematic representation of the inner sphere (IS), outer sphere (OS), chemical exchange, and
rotational correlation times guiding the paramagnetic relaxation.
From the inner sphere contribution, T1 in the first coordination sphere is:
! !"
!!
= 𝑓𝑞
!
Equation 13
!!! !!!
where f is the relative concentration of paramagnetic complex over water molecules, and
q is the number of water molecules in the first coordination sphere. The calculation of
T1M is based on a model that includes the amplitude of the magnetic interaction, its
temporal modulation and the effect of the strength of the external magnetic field, as
follows:
!
!
!!!
= !"
!! !
!!
!
𝛾!! 𝛾!! ℏ! 𝑆(𝑆 + 1) ! !
!!!!
!! !! !!! !
!!
+ !! ! !!!
! !!
!
Equation 14
where:
!
!!"
!
!!!
!
!!!
!
!
!
!
!
!"
=! +! +!
!
= !!
!
!"
!
= !"!
!"
!
!!!!! !!
Equation 15
!
+ !!!!! !!
Equation 16
! !
!
!
3 + !!!! !! + !!!!! !!
! !
! !
Equation 17
24
and where γS and γH are the gyromagnetic ratio of the electron (S) and of the proton (H),
respectively, ωS,H is the angular frequencies of the electron and of the proton, r is the
distance between coordinated water protons and unpaired electros spins, τc1,2 is the
correlation tie modulating the interaction (defined by Equation 15), τR is the rotational
correlation time of the hydrated complex, τs1,2 is the longitudinal and transverse
relaxation of the electron, τS0 is the value of τs1,2 at zero field, and τv is the correlation
time characteristic of the electronic relaxation times.
The second term of Equation 12, the outer sphere relaxation (OS), is explained by the
dipolar interaction at long-distance between the magnetic moment of the paramagnetic
substance, and the nuclear spin of hydrogen protons. In fact, paramagnetic centers
influence the local magnetic field around the 1H protons flowing in their vicinity. The
complete equations explaining the outer sphere contribution of the longitudinal relaxation
! !"
rate
!!
, are too complex to be detailed here, and the reader is invited to refer to the
description of the OS model by Freed.[44, 138] The most important aspect to retain from
the OS contribution in the case of paramagnetic substances, is the fact that the dipolar
intermolecular mechanism is modulated by the translational correlation time (τD) that
takes into account the relative diffusion (D) of the paramagnetic center and the solvent
molecule, as well as their distance of closest approach (d), as follows:[138]
𝜏! =
!!
!
Equation 13
This expression indicates that viscous solvents and large particles lead to high
translational correlation times. In resume, the equations describing IS and OS
contributions to the relaxivity of paramagnetic CAs are rather complex, and a large
number of parameters influence the overall relaxometric performance of CAs: τM, q, τR,
D, r, d, τV, τS0). Because of this high number of parameters, it is often difficult to perform
an accurate theoretical estimation of the performance of MRI CAs at different magnetic
field strengths. Relaxometric measurements must be performed experimentally, by using
a technique called proton nuclear magnetic relaxation dispersion (NMRD). NMRD
curves characterize the efficiency of CA at different magnetic fields. For more
information about the interpretation of NMRD curves, the reader is invited to read a
selection of references on the subject.[6, 44, 134]
Here are a few points that are important to know prior to the interpretation of NMRD
profiles for paramagnetic solutions:
•
The rotational correlation time (τR): it characterizes the reorientation of the
vector between the paramagnetic ions and the protons of the water molecule. For
25
•
•
•
•
a low molecular weight complex, τR limits the relaxivity of paramagnetic CAs at
magnetic field strengths used in clinical MRI (0.5 – 3 Tesla). The value of τR
cannot be measured by proton NMRD, but instead by 17O NMR measurements
(longitudinal relaxation of the nucleus 17O), and other methods.[6]
Electronic relaxation times (τS1 and τS2): Longitudinal and transverse electronic
relaxation times describe the process of return to equilibrium of the magnetization
associated to electrons that transit between electronic levels of the paramagnetic
center. These transitions produce fluctuations that allow the relaxation of protons;
τS1 and τS2 are magnetic field-dependent.
Number of coordinated water molecules (q) strongly influences the IS
contribution. For small complexes such as Gd-DTPA, the number of coordinated
water molecules is equal to 1. It means that, in general, only one water molecule
can bind to the paramagnetic Gd3+ ion sequestrated in the DTPA molecule. The
value of q can be estimated either in solid phase (X-rays or neutron diffraction) or
in solution [fluorescence of Eu or Yb complexes, LIS (lanthanide-induced shift)
method in 17O-NMR].
Proton-metal distance (r): The efficiency of the IS dipolar mechanism is
proportional to 1/r6, where r is the metal-proton distance. Small changes to this
distance have a considerable impact on the relaxivity.
Coordinated water residence time (τM): The mechanism of IS relaxation is
based on an exchange between water molecules surrounding the complex and the
water molecules coordinated to the lanthanide ion. Consequently, the exchange
rate (kex = 1/τM) is an essential parameter for transmitting the relaxation effect to
protons in the water matrix.[6]
b. Relaxivity of superparamagnetic CAs
For superparamagnetic particles, the IS contribution to the relaxation is minor and
often completely negligible as compared to the dominant OS contribution. As mentioned
in the previous section, this contribution is largely dependent on the movement of water
molecules near the local magnetic field gradients generated by the superparamagnetic
nanoparticles. The relaxation of superparamagnetic NP suspensions, is generally
governed by Freed’s equations when τS1 is the Néel relaxation time.[139] When τD is
much shorter than the Néel relaxation time, Freed’s equation are simplified. In fact, the
ability of a fluctuation to relax the 1H proton spins depends upon whether its correlation
time is longer or shorter then the precession period of the spins within the external
magnetic field B0. If the global correlation time τC (τC-1 = τD-1 + τN-1) is longer than this
period, the fluctuation is averaged by the precession and it is inefficient. It is efficient in
the opposite situation. Other parameters such as electron polarization and crystal
anisotropy have also a considerable influence on the relaxation times of water protons
26
submitted to different magnetic fields. As in the case of paramagnetic CAs, the
theoretical models explaining the relaxometry of aqueous suspensions of iron oxide NPs,
must be validated by NMRD profiles (Figure 9). First, it is possible to precisely measure,
with NMRD, the relaxometric potential of NP as contrast agents for MRI.[140] Then,
NMRD profile analysis also represents a powerful tool to control the reproducibility of
synthetic MNP suspensions, as well as for optimizing the parameters of nanomagnet
synthesis.[141] Finally, the fitting of NMRD profiles by adequate theories provides
information about the average radius (r) of the nanocrystals, their specific magnetization
(Ms), their anisotropy energy (Ea), as well as their Néel relaxation time (τN).[119]
Figure 9 : NMRD profile for magnetite particles in colloidal solution (from [134]).
The main information extracted from NMRD profiles, for superparamagnetic NP
suspensions, is:
•
•
The average radius (r): at high magnetic fields, the relaxation rate depends only
on τD and the inflection point corresponds to the condition ωI·τD ∼ 1 (Figure 14).
According to Equation 13, the determination of τD gives the crystal size r (good
complement to HRTEM measurements).
The specific magnetization (Ms): at high magnetic fields, Ms can be obtained
from the equation Ms ≈ [(Rmax/ (C · τD)]1/2, where C is a constant and Rmax the
maximal relaxation rate.
27
•
•
The crystal anisotropy energy (Ea): the absence or the presence of an inflection
point at low magnetic field strengths (10-2 – 1) is an indication of the anisotropy
energy. For crystals characterized by a high Ea compared to the thermal agitation,
the low field dispersion disappears. This was confirmed in a previous work with
cobalt ferrites, a high anisotropy energy material.[142]
The Néel relaxation time (τN): the relaxation rate at very low fields (R0) is
governed by a “zero magnetic field” correlation time τc0, which is equal to τN if τN
<< τD. However, this situation is often not met, and in this case τN is olny reported
as qualitative information.
c. Relaxometric performance of MRI CAs at clinical magnetic field strengths
Among polymer coatings that have been successfully used to enhance the blood
retention of USPIOs, figure dextran and (carboxy, carboxymethyl)dextran. In particular,
commercial products such as Ferumoxtran-10 (AMI-227), Ferumoxytol, and Supravist,
are all based on iron oxide of core diameters in the range 4 – 8 nm, and of hydrodynamic
sizes not larger than 30 nm (Table 1). At 1.5 Tesla, the longitudinal relaxivities (r1) of
those particles range between 9 and 15 mM-1s-1.[143] As mentioned in the previous
sections, the relaxometric ratio at clinical magnetic strengths (typically at 1.5 T) is used
to classify the behavior of CAs between “positive” (i.e. r2/r1 < 5), and “negative” ones
(i.e. r2/r1 >> 10). The relaxometric ratio of dextran and carboxy/carboxylmethyldextrancoated USPIOs, is found between 2 and 5 (1.5 T). Finally, the blood half-lives of these
products, from 6 to 36 hours, figure among the longest of all iron oxide NP systems. For
such reasons, and in spite of the potential instability of dextran grafting at their surface,
ferumoxtran, feromoxytol and supravist particles have been widely applied to MR
molecular imaging. The following table summarizes the performance of a selection of
commercial and pre-clinical products, based on both superparamagnetic and
paramagnetic NPs.
Table 1 : Relaxometric properties of MNPs measured at clinical MRI field strengths
Product, commercial
name, coating
SPIO (Ferumoxides,
Endorem, Feridex), (AMI25), dextran T10
SPIO, Resovist,
Ferucarbotran (SHU-555A),
Resovist
NP core
diam.,
TEM
(nm)
∼5
Hydrodyn.
size, DLS
(nm)
(1.5T,
37oC)
r1
r2/r
120 - 180
10.1
12
∼10
60
9.7
19.5
1
Blood
half-life,
T1/2
(species)
2h
(humans)
Use
Refs.
LI, CL
[144,
145]
2.4 – 3.6 h
(humans)
LI, CL
[146]
28
USPIO, Ferumoxtran-10
(AMI-227), dextran T10
4.5
15 - 30
9.9
6.57
24-36 h
(humans)
USPIO, Ferumoxytol-7228,
carboxlymethyldextran
6.7
30 - 35
15
5.93
10-14 h
(humans)
USPIO, Supravist (SHU555C), carboxydextran
USPIO, Feruglose
NC100150, Clariscan,
PEGylated starch111
VSOP (iron oxide), C184
citrate
ESION (iron oxide),
PO-PEG
USPIO
bis-phosphonate-PEG
Gd2O3
PEG-diacid
NaGdF4
citrate
3-5
21
10.7
3.55
6.43
11.9
∼18
n.a.
6h
(humans)
6h
(humans)
14
2.4
4.77
6.12
(* at 0.5 T)
7
14
3
(* at 3 T)
9.5
2.97
5.5 ± 0.6
24 ± 3
2
9 – 11
14.2
1.2
<5
6-9
3.37
1.18
(*at 3T, RT)
MLNI,
MI,
BPA,
CL
MI,
BPA,
CL
BPA,
CL
[102]
BPA
[17,
149]
0.6 – 1.3 h
(humans)
> 10 min.
(rat)
< 40 min.
(mouse)
_
BPA,
CL
[150]
BPA
[151]
BPA
[152]
CL
[76]
> 90 min.
(mouse)
BPA,
MRIOI
BPA
[34]
[144,
147]
[148]
[118]
MnO
6-8
13.4 –
4.4
8.6
< 20 min.
bis-phosphonate-PEG
16.2
(mouse)
dendrons
*Use: LI: liver imaging; CL: cellular labeling; MLNI: metastatic lymph node imaging; MI: macrophage
imaging; BPA: blood pool agent; MRI-OI: MRI-optical imaging
7. Biological performance of magnetic nanoparticles for MRI
MNPs for MRI applications are complex pharmaceutical constructs that must
navigate the body either to provide a general contrast enhancement effect, or in search of
a target. They are made of at least two, if not four or five different components (Figure
1): a central magnetically active core, a stabilizing shell, or coating, made of one or many
types of biocompatible molecules, to which targeting ligands and additional imaging
modalities are anchored. Therapeutic agents can also be embedded in the structure. MNPs
must be biocompatible, and should not harm the patient. The behaviour of the
nanoconstruct in vivo, as well as that of each one of the different components (blood
retention, clearance kinetics, possible degradation resulting in metal ion leaching,
polymer or drug elution, etc.), must be comprehensively investigated prior to approval by
the health authorities. Finally, MNPs coated and functionalized, should preserve their
relaxometric properties.
a. In vivo barriers
MNPs injected in vivo, must overcome several biological barrier either to reach their
target, or to remain in the blood for prolonged times. Most MRI CAs are administrated
through intravascular (i.v.) injections. MNPs immediately encounter blood, a high ionic
29
strength, heterogeneous solution that contains high concentrations of organic molecules.
Chemical binding and electrostatic interactions occur between these molecules and the
MNPs, that can lead to dramatic changes in their hydrodynamic diameter, relaxometric
properties, and colloidal stability. Agglomeration and surface charge effects can also
accelerate the opsonization of MNPs by the immune system, which results in stronger
and more rapid uptake by the macrophages. Depending on their molecular coating, MNP
can interact more or less strongly with the extracellular matrix of cells and, in the case of
binding, this can cause the NPs to be taken up by cells prematurely before they reach
their target tissue.[153].
The NPs must also overcome different anatomical size restrictions which limit
their access to target tissues (e.g. extravasation of lymph node-targeting NPs from the
blood).[154] These size limitations are very stringent in the case of certain organs such as
the brain and the kidneys.[155] For instance, only NPs of sufficient small sizes and
appropriate physicochemical properties may pass the blood brain barrier, a structural and
metabolic barrier consisting of endothelial cells and reinforcing astrocyte cells that
protect the brain.[156] In addition to biological barriers present in the extracellular space
(blood vessels, lymphatic conducts), intracellular barriers may also restrict the function of
several biomedical NP systems. NPs that bind at the surface of targeted cells, are
typically taken-up by such cells through a receptor-mediated endocytosis mechanisms.
Upon ingestion in the cells, NPs are entrapped and “trafficked” through endosome
compartments where they progressively degrade by acidification. The endosomes are
progressively translocated into lysosomes, compartments in which hydrolytic and
enzymatic reactions metabolise or evacuate macromolecules and NP debris.[157]
Different strategies, more or less complex or potentially cytotoxic, have been developed
to facilitate the escape of NPs from the endosomes.[158] Overall, each one of these
biological obstacles illustrates the different levels of complexity that must be addressed
when designing optimal MNPs as CAs for blood pool, cell labeling and tracking, targeted
imaging, and drug delivery procedures.
b. Impact of nanoparticle size and surface on colloidal stability and
blood retention
The ability of MNPs to remain in the blood for prolonged times and to pass
biological barriers is largely related to their main physico-chemical characteristics:
hydrodynamic size, surface charge, and type of coating.[155, 159, 160] The retention of
NP in the blood vessels, the mechanisms of NP clearance, and the permeability of NPs
from the vasculature, all strongly depend on these three parameters.[161-163] In
particular, NP charge and hydrophobicity can affect NP biodistribution by limiting or
enhancing interactions of NPs with the adaptive immune system, with plasma proteins,
30
with extracellular matrices and with non-targeted cells.[153] Specifically, hydrophobic
and strongly charged NPs have short circulation times due to the adsorption of plasma
proteins such as the opsonins, which are responsible for the recognition of MNPs by the
reticuloendothelial system (RES). As a next step, the macrophages recognize opsonins
and remove MNPs from the blood circulation.[160] Most of particles end up in the liver
and in the spleen, where the particles eventually degrade. The sequestration of MNPs by
the RES, can been advantageously used in the diagnosis of hepatic lesions (e.g. liver
cancer). However, particles that are too rapidly sequestered and removed from the blood
stream, cannot optimally fulfill targeted applications such as targeted imaging (e.g. for
cancer, atherosclerotic plaque, and Alzheimer’s disease), and drug delivery to specific
organs.
Hydrodynamic size strongly affects the clearance of NPs from the vascular
circulation [37,47–51]. In general, small NPs (< 20 nm) can be more efficiently excreted
by the kidneys[164-166], whereas large NPs (> 50 nm) are mostly found in the liver and
spleen.. However, strongly charged particles such as citrate-capped iron oxide and
paramagnetic nanocrystals, follow the RES route and are cleared in the liver.[98] As a
consequence, and a minor fraction of them find their way through the urinary tracts.[150]
Most endothelial barriers allow nanoparticles < 150 nm hydrodynamic diameter to pass;
however, other barriers, such as the BBB, are far more restrictive. Because the liver can
metabolize very large amounts of iron, USPIOs and SPIOs are in general well-tolerated
products. As a matter of fact, Ferumoxides, Feromoxtran, Ferucarbotran, Supravist,
Clariscan and a few other iron oxide-based NP systems have either been commercialized,
or at least investigated in the clinics (see Table 1). In the case of paramagnetic
nanoparticles that are not based on Gd or Mn chelates (MnO, Gd2O3, NaGdF4), the
toxicity risk that represents the massive injection of Mn and Gd-based crystals potentially
leaching toxic Mn2+ and Gd3+ ions, greatly restricts their eventual transfer to clinical
applications. However, emerging coating strategies, in particular those based on silanePEG or phosphate/phosphonate-PEG (Figure 10), have proved to promote the rapid
excretion of paramagnetic nanocrystals through the gastrointestinal and urinairy
routes.[118, 166] It is too soon for the moment to predict if these strategies could
potentially allow the clinical use of paramagnetic nanocrystals in the future. However,
they facilitate the development of more complex and specific “positive” CAs for
molecular and cellulair pre-clinical research.
31
Figure 10 : Injections of MnO NPs PEGylated with phosphonate dendrons, and scanned in 1 T MRI; a)
after 1h, strong evidences of CA elimination is found in the gall bladder (hepatobiliairy way); b) blood
signl-enhancement persists at lest 20 minutes after injection, while evidences of CA clearance are found in
the kidneys (b, 1h and c). Adapted from [118].
In order to be truly efficient for targeted molecular imaging applications, it is
necessary for the MNP CAs to reach blood half-lives of many hours (such as for
Feromuxtran, Ferumoxytol and Feruglose, Table 1). This way, the CAs have more
chances to effectively bind to molecular epitopes and receptors that are expressed at the
surface of vascular cells, and to which the targeted CAs are designed to reach. Surface
modification with hydrophilic molecules dextran and polyethylene glycol (PEG), have
been shown to reduce opsonization, leading to prolonged NP circulation times (Table 1,
and [167]).
c. Directing nanoparticles in vivo
The specificity of NPs for select tissues is critical in MRI-based diagnostic
imaging. [16,74,75]. NPs can be engineered to have an affinity for target tissues through
passive, active, and magnetic targeting approaches. Passive targeting uses the
predetermined physicochemical properties (size and surface charge) of a given NP to
specifically migrate to a given tissue region. In particular, the enhanced permeation and
retention effect (EPR) can be used to target solid tumour tissue.[76] Tumor cells, in an
effort to grow rapidly, stimulate the production of new blood vessels (termed
« neovasculature »). Such vessels are poorly organized and have leaky fenestrations. This
enables the extravasation of NPs out of the vasculature, into the tumor tissue [77,78].
Then, because the lymphatic drainage is relatively inefficient in solid tumors, NPs tend to
accumulate at this site[79,80]. However, the EPR effect is limited to specific metastatic
solid tumors, and the successful implementation of CA systes relying on the EPR effect,
32
is dependent upon a number of factors including the degree of capillary disorder, blood
flow, and the lymphatic drainage rate. As a result, it is not possible to base a therapeutic
treatment, or a MR diagnostic, only by relying on the EPR effect.
However, passive targeting does work only for very specific biomedical
application, such as for solid tumor of leaky vasculature. Neither does it guarantee their
internalization by targeted cells. To achieve this, it is necessary to modify the NPs with
molecular targeting ligands, to provide an “active” approach. NPs functionalized with
molecules that specifically bind to molecular epitopes expressed at the surface of cancer
cells, or other diseases tissues, allow a more efficient way to concentrate NPs selectively
to the tissue to be diagnosed, or treated (81, 82). In particular, a number of iron oxide
nanoparticle systems have been developed and tested in vivo, with varying success.
Among targeting molecules that have been used so far to achieve active targeting with
iron oxide NPs, figure small organic molecules [81,83,84], peptides [71,85–88], proteins
[89], antibodies [90–92], and aptamers [93–95].
In addition to engineering NPs for tissue targeting, external magnetic fields can be
used to assist the diffusion of MNPs to given organs. This strategy is referred to as
“magnetic targeting” [18,104]. It consists in focusing high fields, high gradients, or
strong rare earth magnets on the target organ, or biological site. This is a good strategy to
accumulate high-susceptibility MNPs at specific sites, to conduct hyperthermia
treatments.[100] This technique was successfully implemented in a clinical trial to deliver
the chemotherapeutic doxorubicin to hepatocarcinoma cells [105]. The effectiveness of
magnetic targeting is unfortunately limited to target tissues that are located close to the
surface of the body (rapid loss of magnetic field strength away from the magnets).
d. Toxicity
In order to generate significant signal-enhancement effects, NPs used as MRI CAs
must be injected at relatively high dosage compared with common tracers used in nuclear
medicine (e.g. PET). Therefore, NPs must be demonstrated safe for cells and for different
tissues, in particular when high quantities find their way to critical organs (liver, kidneys,
spleen). The impact of MNPs on the proliferation and on the viability of different cell
types is always demonstrated in vitro prior to injection in vivo. Depending on the
concentration, type of particle, surface charge, and class of coating ligand, the presence
of high concentration of nanoparticles in the vicinity of cells, can have a transitory effect
on their cell division cycle, and sometimes influence their viability. Apoptose
measurements can also be performed at high dosages of MNPs, to evaluate the risk
induced by high concentrations of MNPs to cells. The cells selected for in vitro tests
should ideally represent either the tissues that are expected to receive the highest
33
concentrations of MNPs, or the cells that are the most susceptible to be affected by
MNPs. As an example, for assessing the biocompatibility of ultra-small MNPs for
targeted vascular imaging, epithelial vascular, kidney, and hepatic cell lines could be
used.
Finally, MNPs are not intact when they are excreted by body. They are confronted
to different biological mechanisms that impact their integrity. Nanosystems for medical
applications face as very important challenge: it is necessary to evaluate the toxicity of
both the intact products, and of their different components. For instance, the potential
leaching on metal ions from MNPs in different organs must be carefully and
comprehensively quantified. Also, the impact of the different NP coatings, as well as
their degradation products, on specific cells and organ in which they could potentially
accumulate, is also an important aspect of the toxcicity evaluation of MNPs. The nature
of the degradation by-products, must also be addressed.[168] Nanotoxicology is an
emerging and expanding research area, and a selection of works have been written to
specifically address this topic.[155, 169, 170]
8. Conclusion
The recent advances in synthesis and characterization of MNPs as MRI CAs, has allowed
the emergence of a variety of new biomedical applications: stem cell labeling and
tracking in vivo, imaging-assisted drug and gene delivery, molecular targeting of chronic
diseases such as atherosclerosis and cancer. Because of their very strong impact on the
transverse relaxivity, superparamagnetic NPs in particular, have been used in a variety of
clinical applications (liver, spleen, lymph nodes imaging). In this chapter were reviewed
the basic principles of MNPs for MRI applications. The main parameters and conditions
guiding their optimal design, use, and performance in biological applications, were
presented. Because the relaxometric potential of MNPs is very dependent on their size,
fine particles of narrow size distributions are developed, then coated with biocompatible
molecules that provide enhanced colloidal stability in physiological environments
through electrostatic and steric repulsion mechanisms. The biocompatibility of
nanoparticles must be assessed in vitro, prior to measuring their magnetic and
relaxometric properties. Finally, the biological kinetics (blood retention, organ uptake,
clearance) and contrast-enhancement effects of each new nanoparticulate system, must be
carefully studied in vivo. NPs are complex systems, and the medical health authorities
enforce very strict requirements over the design, production reproducibility, potential
toxicity, and pharmacokinetics performance of such injectable products. Finally, the
expansion of hybrid imaging modalities (MRI/PET, MRI/luminescence, MRI/SPECT,
MRI/echography), call for the development of multifunctional and increasingly complex
imaging tracers. For instance, stem cell therapies could advantageously been conducted
34
using MRI/PET, which would enable a more sensitive detection and quantitation of areas
of implanted stem cells. The delivery of targeted drugs through nanovectors could also be
performed under MRI/PET guidance, to provide quantitative measurements of the
residual concentration of drug delivery vehicles still present in the blood, then
accumulating into specific organs. The use of MNP-labeled nanosized targeted drug
carriers would enable to record real-time, personalized pharmacokinetic profiles of drugs
that more efficiently finding their ways to target organs.
9.
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