Formulation of superparamagnetic iron oxides by nanoparticles of biodegradable polymer for magnetic resonance imaging (MRI)

FORMULATION OF SUPERPARAMAGNETIC IRON OXIDES BY NANOPARTICLES OF BIODEGRADABLE POLYMER FOR MAGNETIC RESONANCE IMAGING (MRI) NG YEE WOON (B.Eng.(Hons.), NUS) A THESIS SUBMITTED FOR THE DEGREE OF MASTER OF ENGINEERING NUS NANOSCIENCE AND NANOTECHNOLOGY INITIATIVE 2006 Acknowledgements I would like to thank NUS and EDB for awarding me the scholarship, thus making it possible for me to pursue a postgraduate course in nanoengineering. During the past two years, I have learnt a lot and I would take to take this chance to show my gratitude to those who have helped me with this project in one way or another. First and foremost, I would like to thank my supervisor A/P Feng Si-Shen for his guidance and support. I would also like to show my appreciation to A/P Wang ShihChang and A/P Ding Jun who has helped me on the study of MRI and magnetic properties respectively. Next, I would like to express my gratitude to Dr. Borys Shuter for taking time to assist me in the use of MRI machine for my experiments. I would also like to show my appreciation to Dr. Chen Yan for her advice and encouragement. Another important person I want to thank is Ms. Wang Yan, my fellow schoolmate, who has extended a helping hand to me whenever I met with difficulties in my experiments. Last but least, I would like to thank everyone in the Chemotherapy Laboratory who has given me help when I need them and make my life in NUS a memorable one. i Table of Contents Acknowledgements i Table of Contents ii Summary v List of Tables viii List of Figures ix List of Symbols xi Chapter 1 Introduction 1 1.1 Background 1 1.2 Objectives 5 1.3 Organization of thesis 7 Chapter 2 Literature Review 8 2.1 Nanoparticles of Biodegradable Polymers 8 2.2 Introduction to MRI 13 2.3 Introduction to MRI contrast agent 16 2.4 Research done on IO encapsulated polymeric nanoparticles 20 Chapter 3 Materials and Methods 23 3.1 Materials 23 3.2 Preparation of the nanoparticles 23 3.3 Physicochemical characterization of the nanoparticles 25 3.4 MR Characterization of the nanoparticles 29 3.5 Biodistribution 30 ii Chapter 4 Physicochemical Characterization 32 4.1 Crystalline structure and surface chemistry 32 4.2 Size Distribution and Iron loading 34 4.3 Surface Morphology 36 4.4 Surface Charge 37 4.5 Stability 38 4.6 In vitro release profile 39 4.7 Summary 40 Chapter 5 Magnetization properties 42 5.1 Characteristics of superparamagnetic materials 42 5.2 Magnetization – temperature dependence 43 5.3 Blocking temperature TB 45 5.4 Summary 47 Chapter 6 In vitro MR studies 48 6.1 Relaxivity plots 48 6.2 Qualitative analysis 51 6.3 Investigations on encapsulation effects 52 6.4 Theories behind relaxivity enhancement 54 6.5 Summary 55 Chapter 7 Animal studies 56 7.1 Biodistribution studies 56 7.2 Ex vivo MRI 57 7.3 Summary 58 iii Chapter 8 Conclusions and Recommendations 60 8.1 Conclusions 60 8.2 Recommendations 63 References 64 iv Summary Magnetic resonance imaging (MRI) is an imaging technique used primarily in medical settings to produce high quality images of the inside of the human body. Iron oxides (IOs) which increase the R2 relaxation rate of the surrounding medium to create signal voids on MR images, have been used as an MRI contrast agent. Their major applications include imaging of the liver, spleen, and breast. For future applications such as imaging of specific molecular targets to allow for earlier recognition and characterization of disease, earlier and direct evaluation of treatment outcomes, and a deeper understanding of disease development, there is a need to develop special contrast agents with greater ability to amplify the MRI signals [1]. This can only be achieved if contrast agents are accumulated in the target cells by passive endocytosis, or by active transporter systems such as transferring receptors that shuttle contrast agents into targeted cells [2]. A feasible way of enabling active targeting is to employ a nanoparticulate structure, which can serve as a scaffold for targeting ligands and magnetic labels [3]. Therefore, much attention has been paid to the research and development of nanoparticles to further enhance the contrast efficiency of IOs. The main objective of this project is to develop a novel formulation of MR contrast agent by encapsulating IOs with biodegradable polymer, methoxy poly(ethylene glycol)-poly(lactide-co-glycolide) (PLGA-mPEG). The IOs used are commercial MR contrast agent Resovist®. The IO loaded PLGA-mPEG nanoparticles, prepared by v water in oil in water (w/o/w) double emulsion technique, were characterized by several techniques including laser light scattering (LLS) for the particle size, field emission scanning electron microscopy (FESEM) for the surface morphology, transmission electron microscopy (TEM) for qualitative determination of IOs loaded, inductively coupled plasma-optical emission spectroscopy (ICP-OES) and/or inductively coupled plasma-mass spectrometer (ICP-MS) for quantitative determination of IOs loaded, superconducting quantum interference device (SQUID) for magnetization measurement, and MRI for contrast effect determination. In addition, in vitro release study to determine the release kinetics profile and stability tests to evaluate the resistance of the IO loaded PLGA-mPEG nanoparticles towards aggregation and iron leakage upon exposure to osmotic agent NaCl (sodium chloride) were carried out. These nanoparticles were spherical with an average diameter of 233.0 nm and a relatively narrow size distribution of ±12.5 nm. The iron loading was 1.37%. They showed enhanced saturation magnetization, improved r2 and r2* relaxivities, and increased contrast effect of both in vitro and ex vivo MR images. The feasibility of the enhancement effect achieved can be substantiated by MR theories such as motional averaging regime (MAR) and static dephasing regime (SDR). The signal amplification achieved may be due to agglomeration of IOs inside the polymer matrix. vi In summary, the remarkable increase in the MR contrast efficiency of the developed IO loaded PLGA-mPEG nanoparticles over the commercial IO contrast agent Resovist®, suggests that these nanoparticles could be potential MRI contrast agent. vii List of Tables Table 3.1 The TE and TR parameters for measuring relaxivities of the IOs and IO loaded PLGA-mPEG nanoparticles. 29 Table 4.1 Properties of the IOs and IO loaded PLGA-mPEG nanoparticles. 35 Table 4.2 Zeta potential of the IO loaded NPs 38 Table 6.1 r1, r2 and r2* relaxivities of the IOs and the IO loaded PLGA-mPEG nanoparticles. 50 viii List of Figures Figure 3.1 Schematic of the preparation of IO loaded PLGA-mPEG nanoparticles by w/o/w double emulsion. 24 Figure 4.1 Peaks in XRD patterns of the IOs correspond to spinel Fe3O4 phase peaks. 33 Figure 4.2 Fe 2p XPS of the IOs showing Fe3+ and Fe2+ peaks. 33 Figure 4.3 TEM images of (a) the IOs (bar = 20 nm) and (b) the IO loaded PLGA-mPEG nanoparticles (bar = 50 nm). 34 Figure 4.4 Particle size distribution of IO loaded PLGA-mPEG nanoparticles. 36 Figure 4.5 FESEM images of the IO loaded PLGA-mPEG nanoparticles (bar = 1 µm). 37 Figure 4.6 Stability of the 233 nm IO loaded PLGA-mPEG nanoparticles in 39 NaCl solution at 37◦C. Figure 4.7 In vitro release profile of the IO loaded PLGA-mPEG 40 nanoparticles in PBS at 37◦C. Figure 5.1 Magnetization curve for IOs and IO loaded PLGA-mPEG nanoparticles at 300 K. 43 Figure 5.2 Magnetization as a function of temperature for the IOs and the IO loaded PLGA-mPEG nanoparticles (applied field 20 kOe). 44 Figure 5.3 Blocking temperature of (a) IOs and (b) IO loaded PLGA-mPEG nanoparticles. 46 ix Figure 6.1 (a) r1, (b) r2 and (c) r2* relaxativities of the IOs and the IO loaded 49 PLGA-mPEG nanoparticles. Figure 6.2 Comparison of IOs and IO loaded PLGA-mPEG nanoparticles at TE=7 ms 51 Figure 6.3 Relaxation rate (a) R2 and (b) R2* of blank PLGA-mPEG nanoparticles, IOs, and mixtures of them with different concentrations of blank PLGA-mPEG nanoparticles. 53 Figure 7.1 Biodistribution of iron in various organs (1 hr after injection) 57 Figure 7.2 MR imaging of the livers of the rats (upper is the control; bottom is that of the rat injected with IO loaded PLGA-mPEG nanoparticles). 58 x List of Symbols B0 Longitudinal magnetic field BBB blood brain barrier DCM Dichloromethane EPR enhanced permeability and retention FC Field cooling Fe Iron FESEM field emission scanning electron microscopy FID Free induction decay FOV field of view ICP-MS inductively coupled plasma-mass spectrometer ICP-OES inductively coupled plasma-optical emission spectroscopy IOs iron oxides LLS laser light scattering M0 equilibrium magnetization MAR Motional averaging regime PLGA-mPEG methoxy poly(ethylene glycol)-poly(lactide-co-glycolide) MRI Magnetic resonance imaging Ms Saturation magnetization MW molecular weight MXY transverse magnetization xi MZ longitudinal magnetization NaCl sodium chloride NaOH sodium hydroxide NEX number of excitations NMR nuclear magnetic resonance PBS Phosphorus buffer solution PEG Polyethene glycol PGA poly(glycolide) PLA poly(D,L lactide) PLGA poly(D,L latide-co-glycolide ) PS-AAEM poly(styrene/acetoacetoxyethyl methacrylate) PVA Polyvinyl alcohol RF Radio frequency rms root-mean-square SDR static dephasing regime SPIOs Superparamagnetic iron oxides SQUID superconducting quantum interference device TB blocking temperature TE time to echo TEM transmission electron microscopy TR time of repetition USPIOs ultrasmall superparamagnetic iron oxides w/o/w water in oil in water xii XPS X-ray photoelectron spectroscopy XRD X-ray diffraction ZFC Zero field cooling xiii Chapter 1 Introduction 1.1 Background Magnetic resonance imaging (MRI) is a popular non-invasive method for clinical diagnosis of soft tissue or cartilage pathologies with new ideas of considerable potential surfacing on a regular basis [4]. It produces image contrast based on the different relaxation times of hydrogen nuclei, provides great technical flexibility, and is free of the hazards related to ionizing radiation. It is well known that the presence of magnetic particles within tissue allows a very large MRI signal to be obtained. The MRI signal is affected by the interaction of the total water signal (proton density) and the magnetic properties (R1 [the longitudinal relaxation rate (1/s)] and R2 [the transverse relaxation rate ([(1/s)]) of the tissues being imaged. The most frequently used nonspecific contrast agents are gadoliniumbased. Their paramagnetism manipulates R1 of the surrounding molecules to increase the total signal. In recent years, superparamagnetic iron oxides (SPIOs) that enhance R2 of the surrounding medium to produce signal voids on magnetic resonance images have been developed [4]. Iron oxides (IOs) are the most-studied materials for magnetic targeting because of their favorable magnetic properties and high biocompatibility. Superparamagnetic 1 magnetite and maghemite have the highest saturation magnetizations (Ms) among the IOs [5]. SPIO contrast agents are small synthetic γ-Fe2O3 or Fe3O4 particles with a core size of less than 10 nm and an organic or inorganic coating. They have no remnant magnetic moment once the external field is withdrawn. The suitability of the IOs as a contrast agent for MRI depends upon: a) Their magnetic susceptibility to achieve magnetic enhancement [6]; b) Their sizes should ideally be in the range of 6-15 nm [7]; c) The exhibition of their superparamagnetic characteristics [8]; d) Customized surface chemistry for precise biomedical applications [9]. The efficacy of IOs as MR contrast agent can be assessed through their abilities to alter the relaxation rates. The MR properties of the IOs were characterized and quantified by relaxivity, which is defined by R = R0 + r ⋅ C (1.1) where R is the proton relaxation rate (1/T, s-1) in the presence of the contrast agent, R0 is the relaxation rate in the absence of the contrast agent and C is the contrast agent concentration (mM). The constant of proportionality, r is the T-relaxivity ( mM −1 ⋅ s −1 ) [10]. Two main factors that influence the relaxation rates are the magnetization of the IOs and the diffusion of the water molecules in the surrounding medium. The 2 magnetization of the IOs is directly correlated to its size. In other words, the larger the particle size of the IOs, the stronger the magnetization. The diffusion time τ D is the time during which the protons of the water molecules experience the magnetic field of the IOs and is given by τ D = rp / D where rp is the radius of the IOs and D is 2 the diffusion coefficient. Depending on the rate of diffusion of the water molecules and size of IOs, they can be operating in the motional averaging regime (MAR) or static dephasing regime (SDR). In both regimes, the R2 relaxation rate (measured using single (Hahn) spin-echo sequence) is considered to be equal to R2* relaxation rate (measured using gradient echo sequence) because the time to echo (TE) is too long for the 180° refocusing pulse in the spin-echo sequence to be effective. Briefly speaking, when the radius of the IOs is small and the diffusion time taken for the water molecules to diffuse a distance of 2rp in any specified direction is short, the IOs are said to be in the MAR. In this regime, relaxation rates increase linearly with particle size. When the IOs are large enough, it can be assumed that the diffusion time is so long that the water molecules are effectively motionless and the IOs are in the SDR. In this regime, the maximum relaxation rates are achieved. However, we should note that in situations where R2 ≠ R2* and IOs are very large, R2 relaxation rate actually decreases as particle size increases. 3 Presently, a range of SPIO contrast agents have been developed, with variations in hydrodynamic particle sizes (from 10 to 500 nm) and coating materials used (such as dextran, starch, albumin, silicones, poly(ethyleneglycol)). Some of them have been approved for clinical use and are marketed under the trade names such as Lumirem®, Endorem®, Sinerem® and Resovist®. Their major applications include imaging of the liver, spleen, and breast. For future applications such as imaging of specific molecular targets to allow for earlier recognition and characterization of disease, earlier and direct evaluation of treatment outcomes, and a deeper understanding of disease development, there is a need to develop special contrast agents with greater ability to amplify the MRI signals [1]. Significant signal amplification can be achieved if the contrast agent is allowed to accumulate in the target cells by passive endocytosis, or by an active transporter system such as a transferring receptor that shuttles targeted contrast agent into the cell [2]. In order to do so, the current IOs have been improved to enable active targeting. A feasible way of doing so is to employ a nanoparticulate or complex macromolecular structure such as liposomes and dendrimers. In general, nanoparticulates offer large surface area, which can serve as a scaffold for targeting ligands and magnetic labels [3]. Therefore, much attention has been paid to the research and development of IO encapsulated nanoparticles. IO loaded nanoparticles made from biocompatible and biodegradable polymers such as poly D,L lactide (PLA), poly(D,L latide-co-glycolide) (PLGA), poly(styrene/acetoacetoxyethyl methacrylate) (PS-AAEM) and polystyrene were reported in the literature [11, 12, 13, 14, 15, 16, 17]. These works had already 4 addressed issues such as cytotoxicity, the influence of physicochemical properties (e.g. size and surface morphology), chemical composition of polymer matrix and iron entrapment efficiency, and conduct magnetization measurements. The magnetization values of the nanoparticles are important but not a direct indicator of efficacy of these nanoparticles as MRI contrast agents. So far, none of the research groups have carried out MRI measurements to determine the relaxivities of the IO loaded biocompatible and biodegradable polymeric nanoparticles developed. Though Pouliquen et al [18] carried out a very comprehensive study which included in vitro and in vivo MRI measurements, the magnetization measurements had not been conducted yet. In addition, their developed composite particles were in the micron range and produced decreased MR relaxivities. 1.2 Objectives As part of a programme to develop multi-functional nanoparticles that enable controlled and targeted MRI for diagnostic and therapeutic purposes, we would like to produce composite particles in the nano range that can increase the MR relaxivities. The main objective of this project is thus to develop a novel formulation of MR contrast agent by encapsulating IOs with biodegradable polymer, methoxy poly(ethylene glycol)-poly(lactide-co-glycolide) (PLGA-mPEG). Our studies were conducted with comparison to commercially available IOs (Resovist®). 5 Complete characterizations of the IO encapsulated polymeric nanoparticles are required to determine whether they are suitable for MRI applications. Their physicochemical and magnetization properties were first characterized. The IO loaded PLGA-mPEG nanoparticles, prepared by water in oil in water (w/o/w) double emulsion technique, were characterized using several techniques including laser light scattering (LLS) for evaluating the particle size, field emission scanning electron microscopy (FESEM) for measuring the surface morphology, transmission electron microscopy (TEM) for qualitative determination of IOs loaded, inductively coupled plasma-optical emission spectroscopy (ICP-OES) and/or inductively coupled plasmamass spectrometer (ICP-MS) for quantitative determination of IOs loaded, and superconducting quantum interference device (SQUID) for magnetization measurements. In addition, in vitro release study to determine the release kinetics profile and stability tests to evaluate the resistance of the IO loaded PLGA-mPEG nanoparticles towards aggregation and iron leakage upon exposure to osmotic agent sodium chloride (NaCl) were also carried out. To assess the efficacy of IO loaded PLGA-mPEG nanoparticles as MRI contrast agents, in vitro MRI was first conducted to measure relaxation properties of both the IOs and IO loaded PLGA-mPEG nanoparticles. After which, ex vivo MRI studies were carried out by imaging the organs of rats injected with IO loaded PLGA-mPEG nanoparticles. Biodistribution of IO loaded PLGA-mPEG nanoparticles in rats were studied as well. 6 1.3 Organization of thesis The thesis consists of (i) thorough literature review; (ii) description of materials and methods used in the novel formulation of biodegradable IO loaded PLGA-mPEG nanoparticles; (iii) results and discussions of their physicochemical characterization; (iv) magnetization properties and MRI studies; and (v) conclusion and recommendations. The literature review covers the basics of biodegradable polymers, their manufacture techniques, the working principle behind MRI, its contrast agents, and previous work done on IO encapsulated polymeric nanoparticles. Under the materials and methods section, detailed descriptions of materials and methods used in the preparation of biodegradable IO loaded PLGA-mPEG nanoparticles are given. The results of the characterization experiments, magnetization measurements, in vitro MRI and animal studies are presented and discussed in four separate chapters. In the concluding section, the results are summarized, and some suggestions for future directions of this research are given. 7 Chapter 2 Literature Review 2.1 Nanoparticles of Biodegradable Polymers 2.1.1 Basic information of Biodegradable Polymers Recently, there has been increased interest in developing long-circulating nanoparticles as a drug carrier. The studies using polymeric biodegradable nanoparticles to encapsulate anti-tumor drugs such as paclitaxel, doxorubicin and 5fluoruracil have demonstrated promising results for the treatment of cancer in animal models. Besides being a potential drug delivery system, nanoparticles can be used for fluorescent biological labels, gene delivery, separation and purification of biological molecules and cells, MRI contrast enhancement, and detection of proteins [19]. Furthermore multi-functional nanoparticles can also be developed to encapsulate both drug and MRI contrast agent to achieve simultaneous diagnostic and therapeutic effects. One of the factors determining the particle size and the size distribution of nanoparticles is the preparation methods used such as solvent extraction/evaporation and spontaneous emulsification/solvent diffusion. Nanoparticles manufactured using solvent evaporation tend to be larger (300 nm and above) while those prepared using solvent diffusion can be made to be smaller than 100 nm. Nanoparticles can also be 8 prepared by polymerization of monomers. Hydrophilic nanoparticles with diameters less than 100 nm and narrow size distribution have been prepared by using the aqueous core of the reverse micellar droplets as nanoreactors [20]. An advantage of nanoparticles is that due to their small sizes, they can pass through smaller capillaries and be taken up by cells, thereby allowing efficient drug and/or IOs accumulation at the target sites. Also, being made of biodegradable materials, they can achieve sustained drug release at the target site. Nanoparticles may offer protection to the drug molecules during transportation in the circulation and nanoparticle formulation can be developed into a platform technology applicable to a wide range of drugs, either hydrophilic or lipophilic. Drugs and/or IOs may be bound to nanoparticles in various forms, such as a solid solution, dispersed or adsorbed on the surface or chemically attached. The surface of nanoparticles can be modified to prolong their blood circulation and coated or attached with targeting ligands to achieve site-specific drug delivery. However, nanoparticles tend to be removed rapidly from the blood circulation following intravenous administration. The rate of nanoparticle removal is related to both particle size and surface characteristics. Ideally, the size of the long-circulating rigid particles should not exceed 200 nm, preferably in the range of 120-200 nm in diameter, in order to decrease clearance by the reticuloendothelial system (RES). Nanoparticles used for drug delivery to the brain are generally the diameters of 60 – 400 nm. Efforts have been made to modify the surface of nanoparticles to increase their systemic circulation time, by either physical adsorption of a hydrophilic polymer on the particle surface or chemical 9 grafting of polymer chains onto particles. To date, the most successful longcirculating biologically stable nanoparticles have been coated with PEG [21]. 2.1.2 Manufacture techniques of nanoparticles There are many ways to manufacture the nanoparticles, for instance, dispersion of the preformed polymers or by polymerization of monomers [20]. Some other more commonly used methods are briefly described in this section. Solvent extraction/evaporation In the solvent extraction/evaporation technique, the polymer is dissolved in an organic solvent such as dichloromethane, chloroform or ethyl acetate. The hydrophobic anticancer drug is dissolved or dispersed into the preformed polymer solution, and the resulting mixture, after emulsification by high-speed homogenization or sonication, is added into an aqueous solution to make an oil-inwater emulsion with the aid of an amphiphilic surfactant emulsifier/stabilizer/additive (single emulsification). If the anticancer drug is hydrophilic, the technique is slightly modified to form a water-in-oil-in-water (w/o/w) emulsion (double emulsification) [22]. After the formation of a stable emulsion, the organic solvent is evaporated by continuous stirring in an increased temperature or a decreased pressure (vacuum) environment, with or without the aid of an inertial gas flow. Centrifugation or filtration is applied to collect the formed particles, which can then be freeze-dried to 10 form dry powders for storage. However, this method is only suitable for small-scale production. Spray-drying Technologies such as spray-dry and spray-freeze-dry have been developed for mass production of drug-loaded nanoparticles. In brief, the drugs are suspended or dissolved in organic solution where the polymer is also dissolved, and then the mixture is spray dried to form particles. The challenges for spray-drying include how to produce particles with sufficiently small size and how to increase the drug encapsulation efficiency [23]. Spontaneous emulsification/solvent diffusion This technique, in which a water-soluble solvent (e.g., acetone or methanol) and a water-insoluble organic solvent (e.g., dichloromethane or chloroform) are used, employs low-energy emulsification [24]. Due to the spontaneous diffusion of the water-soluble solvent, an interfacial turbulent flow is created between the two phases, leading to the formation of nanoparticles. As the concentration of water-soluble solvent increases, a considerable decrease in particle size can be achieved [25]. Supercritical fluid spraying Production of polymeric nanoparticles by supercritical fluid spraying does not required the use of any toxic organic solvent and surfactant. The drug and the polymer of interest are solubilized in a supercritical fluid, and the solution is 11 expanded through a nozzle. The supercritical fluid is evaporated in the spraying process and the solute particles eventually precipitate. This technique is clean because the precipitated solute is completely solvent-free [26]. Polymerization of monomers Polymerization includes emulsion polymerization and interfacial polymerization. Emulsion polymerization builds up a chain of polymers from single monomers. When the monomer-contained organic phase and aqueous phase are brought together by mechanical force, interfacial polymerization will take place. Couvreur et al [27] reported the production of nanoparticles of about 200 nm diameter by polymerizing mechanically the dispersed methyl or ethyl cyanoacrylate in aqueous acidic medium in the presence of polysorbate-20 as a surfactant. The cyanoacrylic monomer is added to an aqueous solution of the surface-active agent under vigorous mechanical stirring to polymerize alkylcyanoacrylate at ambient temperature. The drug is dissolved in the polymerization medium either before the addition of the monomer or at the end of the polymerization reaction. The nanoparticle suspension is then purified by ultracentrifugation or by resuspending the particles in an isotonic medium. During polymerization, various stabilizers such as dextran and poloxamer are added. In addition, surfactants such as polysorbate are also used. 12 2.2 Introduction to MRI MRI is an imaging technique that generates images of the body using nuclear magnetic resonance (NMR). When a patient is placed into the cylindrical magnet, a magnetic steady state is first created within the body by using a strong magnetic field. Then the body is stimulated with radio waves to change the steady-state orientation of protons and the electromagnetic signals emitted from the body is used to construct detailed internal images of the body using a computer program. This technique is non-invasive, and free of the hazards associated with ionizing radiation. 2.2.1 Basic principles of MRI Nuclear spin is the basis of NMR. When a nucleus contains an even number of protons and neutrons, the individual spins of these particles pair off and cancel out, leaving the nucleus with zero spin. However, if a nucleus has an odd number of protons or neutrons, there is incomplete pairing and the net spin is ½. All such nuclei experience NMR, but in clinical MRI the hydrogen nucleus, comprising of a single proton, is used because of its high NMR sensitivity and its natural abundance in the human body. For clinical applications, a powerful magnet is used to provide a strong uniform constant ‘longitudinal’ magnetic field (B0) in the z-direction. Its magnetic field strength is typically 4000 to 60 000 times that of the Earth. It generates a macroscopic 13 magnetisation due to alignment of hydrogen nuclei with the field. However, to obtain MR images, an external magnetic field has to be applied to excite the hydrogen nuclei. The radio frequency (RF) coils are used to transmit RF pulses required for excitation, and also to detect the emitted MR signal which is known as free induction decay (FID). Following excitation, the nuclei return to their equilibrium state either through the loss of energy from the spin system or simply exchange of energy between spins. These two types of relaxation processes are known as spin–lattice and spin–spin relaxation, and are characterized by the relaxation times T1 and T2, respectively. The MRI signal is thus the product of interaction between the total water signal (proton density) and the magnetic properties (1/T1 [the longitudinal relaxation rate (1/s)] and 1/T2 [the transverse relaxation rate [(1/s)]) of the tissues being imaged. 2.2.2 T1 process At equilibrium, the net magnetization vector lies along the direction of the applied magnetic field Bo and is called the equilibrium magnetization M0. In this case, the longitudinal magnetization MZ equals M0 and there is no transverse (MXY) magnetization. The time constant which describes how MZ returns to its equilibrium value is called the spin-lattice relaxation time (T1). The equation governing this behavior as a function of the time t after its displacement is: M Z (t ) = M 0 (1 − e − t / T1 ) (2.1) 14 2.2.3 T2 process The time constant which describes the return to equilibrium of the transverse magnetization, MXY, is called the spin-spin relaxation time, T2. It is given by: M XY = M XY 0 e − t / T2 (2.2) T2 is always less than or equal to T1. The net magnetization in the XY plane goes to zero and then the longitudinal magnetization grows until we have M0 along Z. The two factors that contribute to the decay of transverse magnetization are molecular interactions (pure T2 molecular effect) and spatial variations in B0 (inhomogeneous T2 effect) within the body. The combination of these two factors is what actually results in the decay of transverse magnetization. The combined time constant is called T2* and is given as follows. 1 / T2 * = 1 / T2 + 1 / T2 in hom o 2.2.4 (2.3) Imaging Techniques In this project, the single (Hahn) spin-echo sequence and the gradient echo sequence are used to obtain the R2 (=1/T2) and R2* (=1/T2*) relaxation rates, respectively. The spins are refocused to compensate for local magnetic field inhomogeneities in T2 imaging, but not in T2* imaging. This sacrifices some image resolution but provides 15 additional sensitivity to the relaxation processes that cause incoherence of transverse magnetization. Spin-echo Sequence The time between repetitions, is called the repetition time (TR), of the sequence. The TE defined as the time between the 90o pulse and the maximum amplitude in the echo. In brief, the spin-echo sequence begins with a 90o pulse and produces a FID that decays according to the T2* relaxation time. After a delay time of TE/2, a 180o refocusing pulse is applied to invert the spins, it reestablishes phase coherence and generates an echo at TE. The inhomogeneities of external magnetic field are cancelled and the peak amplitude of the echo is determined by T2 decay. Gradient echo sequence Unlike the spin-echo sequence, it does not have a 180o refocusing pulse. The spins are refocused by reversing the direction of the spins rather than flipping them over to the other side of the XY plane. Gradient refocusing of the spins takes considerably less time than 180 o RF pulse refocusing. The disadvantage of gradient echo sequences is the loss of signal due to magnetic field inhomogeneity. 2.3 Introduction to MRI contrast agent 2.3.1 Types of contrast agents 16 The most commonly used contrast agents are gadolinium-based. Their paramagnetism changes the R1 relaxation rate of the surrounding molecules to give an increase in total signal. In recent times, iron oxides were developed as MR contrast agents. They work by enhancing the R2 relaxation rate of the surrounding medium to reduce signal intensity on MR images. 2.3.2 Classification of IOs To date a wide variety of IOs have been produced, differing in particle sizes (hydrodynamic particle size varying from 10 to 500 nm) and types of coating materials used (such as dextran, starch, albumin, silicones, poly(ethyleneglycol)). They tend to be classified into two main groups according to their size, as this affects plasma half-life and biodistribution. The first group are termed superparamagnetic iron oxides (SPIOs) where nanoparticles have a size greater than 50 nm (coating included) and the second type termed ultrasmall superparamagnetic iron oxides (USPIOs) where nanoparticles are smaller than 50 nm. Both types of particles are commercially available. Some examples of SPIOs are Lumirem®, silicon-coated particles with 300 nm diameter, and Endorem®, magnetite particles with a 150 nm diameter. They are used for gastro-intestinal tract imaging and for liver and spleen disease detection, respectively. The USPIOs can act as blood pool agents for perfusion imaging of brain or myocardial ischemic diseases. For example, Sinerem®, 17 which is currently being used for tumour detection, consists of magnetite particles with a 30 nm diameter [28]. The particle size also affects the relaxation rates of IOs. USPIO can be considered as a single ferrite crystal, so a uniform distribution of the magnetic crystals within the solvent can be assumed for the calculation of its nuclear magnetic relaxation rate [29, 30, 31]. However, for SPIO which contain several ferrite crystals per particle, this assumption is no longer valid. The transverse relaxation is affected by the agglomeration and determined by two components. The first is the SPIO crystal itself and the second is the assumption of the entire particle as one large sphere [32]. 2.3.3 Relaxation rates of IOs The two main factors that influence the relaxation rates are the magnetization of the IOs and the diffusion of the water molecules in the surrounding medium. Depending on the rate of diffusion of the water molecules and the size of IOs, they can be operating in the MAR or SDR. MAR In the MAR, the relaxation rate can be obtained from the quantum mechanical outer sphere theory: R2 = (4 / 9)vτ D (∆ωr ) 2 (2.4) 18 where v is the volume fraction occupied by the magnetized spheres, τ D is the diffusion time, and ∆ωr is the rms angular frequency shift at the particle surface. It is 3 given by ∆ω r = 4 γBeq = 4 γµ / rp = (8π / 3)γM / 5 where γ is the proton 5 5 gyromagnetic ratio, Beq is the equatorial magnetic field of the particle, µ is its magnetic moment and M is its magnetization. Equation (2.4) is valid if the particles are small enough to satisfy the motional averaging condition ( ∆ω rτ D < 1 ), and relaxation is not affected by the refocusing echo pulse [33]. In this regime, the relaxation rate increases linearly with particle size and R2 = R2*. SDR In the SDR, there is dephasing of motionless magnetic moments of the protons by the randomly distributed IOs in a non-uniform field. There exists an upper limit on the R2 relaxation rate that can be reached in the absence of a refocusing pulse and R2 = R2*. This limit is given by: R2 = π 15v∆ωr / 9 * (2.5) Though equation (2.5) is formulated based on the assumption of motionless spins, it remains valid for slow motion as long as the particles are large enough to satisfy the condition ∆ω rτ D > 1 [34, 35]. However, we should note that these two regimes are applicable only for cases where the 180° refocusing pulse used in the spin echo sequence is not effective to recover signal loss due to field inhomogeneities, thus R2 = R2*. In situations where R 2 ≠ R2* 19 and the particles are very large, the R 2 relaxation rate actually decreases as particle size increases [36]. PRESS 2.4 Research done on IO encapsulated polymeric nanoparticles Ideally, these polymeric magnetic carriers should be small enough (less than 1µm) to pass through capillaries to reach the targeted site, have adequate magnetic sensitivity to magnetic fields in physiological environments, evoke minimum toxicity and immunological response, and be also biodegradable with no or little toxicity of degradation products [37]. Some of the popular biocompatible and biodegradable polymers researched on are poly(D,L latide-co-glycolide ) (PLGA), poly(D,L lactide) (PLA), and poly(glycolide) (PGA) [38, 39, 40]. A considerable amount of work has also been done to demonstrate that biodegradable polymers are ideal as carriers because of their minimum toxicity and immunological response [41, 42, 43, 44]. The combination of biocompatible and biodegradable polymer with SPIOs enables the minimization of systemic side effects while sustaining local higher concentrations of the contrast agent [45]. Several researchers have described the methods on how to prepare these IO loaded nanoparticles of biodegradable polymers. Muller et al [11] produced magnetite loaded PLA and PLGA nanoparticles, sizes of which were between 456 and 890 nm with a theoretical magnetite content up to 50% (w/w). These magnetite loaded polymeric nanoparticles have relatively low cytotoxicity, qualifying them as potential 20 formulation for intravenous injection. Okassa et al [12] achieved the incorporation of modified magnetite/maghemite nanoparticles into PLGA nanoparticulate matrix, but did not report any magnetization properties of these composite nanoparticles. GomezLopera et al [13] also synthesized composite particles by coating a magnetic nucleus (magnetite) with a biodegradable PLA polymer, but they found these composite particles had decreased saturation magnetism. Lee et al. [14] prepared ferrofluidic PLGA nanoparticles and suggested that a decrease in particle size may increase the magnetic susceptibility of nanoparticles as a result of the increase in packing density or volume fraction of the nanoparticles. They also reported MRI image enhancement in the kidney of rabbit after injection of their composite nanoparticles. Other polymers were also used to encapsulate IOs. Dresco et al. [15] synthesized magnetite and polymer magnetite nanoparticles using methacrylic acid and hydroxyethyl methacrylate, but they assumed that the magnetic susceptibility of magnetite did not change after the encapsulation into the polymer matrix. Pich et al. [16] prepared composite poly(styrene/acetoacetoxyethyl methacrylate) (PS-AAEM) particles with encapsulated magnetic IO, and Zheng et al [17] incorporated up to 40 % (w/w) of 8 nm superparamagnetic magnetite particles into polystyrene nanospheres with an average diameter of 80 nm. These works had addressed issues of cytotoxicity, investigated the influence of physicochemical properties such as size and surface morphology, chemical composition of polymer matrix and iron entrapment efficiency, and conducted magnetization measurements. The magnetization values of the nanoparticles are important but not a direct indication of efficacy of these nanoparticles as MRI contrast agents. So far, none of the research groups have carried 21 out MRI measurements to determine the relaxivities of the IO loaded biocompatible and biodegradable polymeric nanoparticles they developed. Though Pouliquen et al [18] had carried out a very comprehensive study which included in vitro and in vivo MRI measurements; they did not carried out magnetization measurements. In addition, their developed composite particles were in the micron-range and produced decreased MR relaxivities. Encapsulation of SPIOs with biodegradable polymers allows surface modification of the nanoparticles to prolong their blood circulation, and coating or attachment of targeting ligands leads to achieving site-specific drug delivery. Long circulating nanoparticles can be obtained by coating with polyethene glycol (PEG). Drugs encapsulated in these nanoparticles have been shown to passively target the tumour tissue through enhanced permeability and retention (EPR) effect [46, 47]. Cellspecific targeting of contrast agents allows early MRI detection of tumour cells. For potential active targeting through surface modification, much research had been conducted on targeted drug delivery through the attachment of ligands such as folic acid [48] and lectins [49] which are over expressed in certain tumour cells. The coating of the particle surface may also help nanoparticles to cross physiological barriers. One such example is the use of polysorbates to coat poly(butylcyanoacrylate) nanoparticles to enhance their drug delivery cross the blood brain barrier (BBB) [50, 51, 52, 53]. 22 Chapter 3 Materials and Methods 3.1 Materials Resovist®, a commercial MRI contrast agent, was purchased from Schering AG for used as IOs in this project. It is a stable, aqueous solution of SPIOs coated with carboxydextran in an approximate ratio of 1:1.1 (w/w). The PLGA-mPEG polymer with 4.75 % (w/w) PEG and lactide:glycolide molar ratio of 80:20 was a kind gift from Curtin University of Technology, Australia. The PEG polymer has molecular weight (MW) of 2,000 Da while the PLGA polymer has MW of 30,000 - 50,000 Da, Polyvinyl alcohol (PVA) with MW of 30,000~70,000 was purchased from SigmaAldrich Co., USA. Milli-Q water with resistivity of 18.2 MΩ•cm was obtained from a Milli-Q Plus System (Millipore Corporation, Breford, USA). Dichloromethane (DCM) was purchased from Merck & Co., Inc.,USA, concentrated (>69.5%) nitric acid was from Sigma-Aldrich Co., USA, and 31.0% hydrogen peroxide was from Kanto Corporation, USA. 3.2 Preparation of the nanoparticles The IO loaded PLGA-mPEG nanoparticles were prepared by w/o/w double emulsion technique as shown in Figure 3.1. Briefly, 0.17 ml of IO aqueous suspension was added to 2.5 ml of 2% PLGA-mPEG DCM solution and sonicated using a 23 MICROSONICTM ultrasonicator equipped with a microtip probe (XL2000, Misonix Incorporated, NY) for 60s at 25W, to obtain an water-in-oil emulsion. Then, this water-in-oil emulsion was poured into an aqueous PVA (as an emulsifier) solution (1% (w/v)) and sonicated for 90s at the same energy output. The organic solvent was rapidly removed by evaporation under mechanical stirring at room temperature overnight (for 12h). The formed nanoparticles were collected by centrifugation (Eppendorf 5810R) at 12,000 rpm for 15min at 20◦C and washed with Milli-Q water for three times to remove excessive emulsifier and free IOs. To obtain fine powder of nanoparticles, nanoparticle suspension was freeze dried using a freeze dryer (Christ, Alpha-2, Martin Christ, Germany). Nanoparticle suspension was used for all characterization work. Blank PLGA-mPEG nanoparticles were prepared in the same way by replacing the IO aqueous suspension with water. Figure 3.1 Schematic of the preparation of IO loaded PLGA-mPEG nanoparticles by w/o/w double emulsion. 24 3.3 Physicochemical characterization of the nanoparticles 3.3.1 XRD Analysis Crystallographic analysis of the IOs was performed by XRD machine (Bruker, Advance D8, USA) with a Cu kα radiation (λ=1.54056 Å) to identify the dominant phase of the IOs in order to estimate the maximum theoretical relaxation rate that the IOs can achieve. The phase was determined using standard powder diffraction files of Joint Committee for Powder Diffraction Studies (JCPDS). 3.3.2 Surface chemistry X-ray photoelectron spectroscope (XPS, AXIS His-165 Ultra, Kratos Analytical, Shimadzu, Japan) was used to determine the surface chemistry of the IOs. Curve fitting of the experimental data was performed using the software supplied by the manufacturer. 3.3.3 Particle Size analysis The particle size and size distribution of the prepared IO loaded PLGA-mPEG nanoparticles were determined by LLS with a particle size analyzer (90 Plus, Brookhaven Inst, Huntsville, US) at a fixed angle of 90◦ at 25◦C. In brief, the 25 nanoparticles were suspended in Milli-Q water and sonicated to produce homogenous suspension of nanoparticles. 3.3.4 Surface morphology The surface morphology of the IO loaded PLGA-mPEG nanoparticles was observed by FESEM (JSM-6700F, JEOL, Japan) at an accelerating voltage of 10 kV after platinum coating of the nanoparticles by a sputter coater (JFC-1300, JEOL, Tokyo, Japan) for 30 s in a vacuum at a current intensity of 30 mA. The nanoparticles were immobilized on metallic studs with double-sided conductive tape. 3.3.5 TEM Measurement TEM (JEM 2010F, JEOL, Japan) examination of the IOs and IO loaded PLGAmPEG nanoparticles was carried out with an electron kinetic energy of 200kV. A drop of well dispersed nanoparticle aqueous suspension was placed on a Formvar/carbon 200 mesh copper grid and then dried at ambient condition before it was attached to the sample holder on the microscope. 3.3.6 ICP-MS and ICP-OES measurements The iron contents of both IOs and IO loaded PLGA-mPEG nanoparticles were determined by either ICP-MS (Elan 6100, Perkin-Elmer, USA) or ICP-OES (Optima 26 3000DV, Perkin-Elmer). For iron concentrations in dilute solutions (less than 10 parts per million), the ICP-MS was used. In solutions with higher iron concentrations (more than 10 parts per million), the ICP-OES was employed. To completely digest the samples to release iron before ICP-MS or ICP-OES analysis, the particles were pre-treated using microwave digestion system (1200 MEGA, Milestone, Leutkirch, Germany). In brief, 10mg particles, 3ml of Milli-Q water, 2ml of concentrated nitric acid and 1.5ml of 31.0% hydrogen peroxide were added to each digestion vessel and digestion was performed with the program developed by Krachler et al [54]. The amount of iron loading (% w/w) was calculated as the ratio of the mass of iron (mg) that can be detected using ICP analysis to the sum of the mass of iron (mg) and the mass of polymer (mg). 3.3.7 Magnetic properties The saturation magnetization of the IOs and the IO loaded PLGA-mPEG nanoparticles was determined by vibrating sample magnetometer (VSM, Lakeshore 7300 Series, US) and SQUID (MPMS XL5, Quantum Design, US). The temperaturedependent magnetization of the samples was obtained by measuring the magnetization in the temperature range of 2-400K with maximum applied field of 20 kOe. Blocking temperatures (TB) could be read from ZFC (zero field cooling) and FC (field cooling) curves taken under the applied magnetic field of 100Oe between 2 and 400K. To obtain the ZFC graph, the samples were cooled from 400 K to 2 K without applying an external field. After reaching 2 K, a 100 Oe field was applied and the 27 magnetization was recorded as the temperature increased. For measuring FC, the samples were first cooled from 400 K under an applied field of 100 Oe, and then the magnetization was recorded as the temperature increased. 3.3.8 Stability study The IO loaded PLGA-mPEG nanoparticles were evaluated for their resistance to osmotic agent NaCl, which potentially may cause nanoparticle aggregation and iron leakage. 60 mg of the nanoparticles were added to 20ml of 0.9% (w/w) NaCl solution and incubated at 37◦C in a mildly shaking water bath. Particle size was measured after 0, 18, 24 and 48h using LLS, and iron leakage was determined by measuring the amount of iron in the supernatant after 48h using ICP-MS. 3.3.9 In vitro release study 5 mg of the IO loaded PLGA-mPEG nanoparticles were placed in each centrifuge tube and then 10ml of fresh PBS (phosphorus buffered solution) at pH=7.4 was added. The tubes were put into a 37oC orbital shaker bath and shaken horizontally at 120 times per minute. The tubes were removed from the shaker bath at predetermined time intervals and centrifuged at 10500 rpm at 18oC for 15 minutes. Then 9 ml of the supernatant was collected for the release analysis. After that, 9 ml of fresh PBS was refilled into the tubes. The nanoparticle pellet was re-suspended in 10 ml 28 PBS and returned to the shaker bath. The amount of iron released from the IO loaded PLGA-mPEG nanoparticles was measured using ICP-MS. 3.4 MR Characterization of the nanoparticles 3.4.1 In vitro MR Imaging In vitro r1, r2 and r2* relaxivities of the IOs and the IO loaded PLGA-mPEG nanoparticles suspended in water were measured. MR images of the nanoparticles were obtained using a Siemens Symphony 1.5 Tesla scanner with a head coil. MR imaging was carried out with different concentrations of the IOs and the IO loaded PLGA-mPEG nanoparticles from 0 mM to 0.5 mM. The spin echo sequence was used. The imaging parameters are flip angle = 90○, number of excitations (NEX) = 1, field of view (FOV) = 180mm and slice thickness = 5mm. The values for TR and TE of the IOs and the IO-loaded PLGA-mPEG nanoparticles to obtain r1, r2 and r2* relaxivities were given in Table 3.1. Table 3.1 The TE and TR parameters for measuring relaxivities of the IOs and IO loaded PLGA-mPEG nanoparticles. r1 r2 r2* IOs 25 ≤ TR≤ 200ms, 9 ≤ TE≤ 360 ms, 5≤ TE≤ 60ms, TE = 9ms TR = 2400 ms TR= 2400ms IO loaded PLGA- 25 ≤ TR≤ 20 ≤ TE≤ 160ms, 5 ≤ TE≤ 60ms, mPEG nanoparticles 6400ms, TR = 1600 ms TR = 1600ms TE = 12ms 29 3.4.2 Ex vivo MR Imaging This study was performed according to a protocol conformed to the animal care legislation and approved by Institutional Animal Care and Use Committee (IACUC), National University of Singapore. Male Sprague Dawley rats (200~250 g) were used. An amount of IO loaded PLGAmPEG nanoparticles equivalent to 3.69 mg Fe/kg body weight was intravenously injected as an aqueous dispersion (0.922 mg Fe/ml) over 300 s into the rat under anaesthesia. Another rat injected with equivalent volume of saline was used as control. The rats were dissected and sacrificed under anaesthesia one hour after the injection. Organs were imaged by MRI. 3.5 Biodistribution Male Sprague Dawley rats (200~250 g) were used. An amount of IO loaded PLGAmPEG nanoparticles equivalent to 1.87 mg Fe/kg body weight was intravenously injected as an aqueous dispersion (1.87 mg Fe/ml) over 300 s into the rat under anaesthesia. An equivalent concentration of IOs was injected into another rat to be used as comparison. For control, saline was injected into the rat. One hour after the injection, the rat is sacrificed. The blood vessels are flushed with saline before the dissection. The rat was dissected to obtain the liver, spleen, kidney, and brain. After that, the removed organs are washed with saline and dried with gauze. The organs 30 were then weighed to obtain the wet weight before freeze-drying. The dried organs were ground into powder, and weighed. Finally, 200mg of each type of organ powder were microwave digested and sent for ICP analysis to determine the iron content in the organs. 31 Chapter 4 Physicochemical Characterization In this project, Resovist®, a commercial MRI contrast agent, purchased from Schering AG, was used as the IOs to be encapsulated. According to the product phamplet, 1 ml of Resovist contained 28 mg of iron in the form of ferucarbotran. It was a stable, aqueous solution of SPIOs coated with carboxydextran in an approximate ratio of 1:1.1 (w/w). As the IOs play a significant role in influencing the properties of the IO loaded PLGA-mPEG nanoparticles, it is necessary to do a characterization study on them as well. The information gathered also served as a comparison when evaluating the IO loaded PLGA-mPEG nanoparticles. 4.1 Crystalline structure and surface chemistry XRD result presented in Figure 4.1 identifies the IOs to be magnetite (Fe3O4) as all the major peaks correspond to the spinel Fe3O4 phase. Further investigation of XPS in the Fe 2p (atomic orbital 2p of iron) region confirms the presence of Fe2+ and Fe3+ ions, as shown by Peak 1 ( Fe3+) and Peak 2 ( Fe2+) in Figure 4.2 32 (511) (422) 40 (440) (400) (220) Internisty (a.u) 30 50 2θ (deg) 60 70 Figure 4.1 Peaks in XRD patterns of the IO correspond to spinel Fe3O4 phase peaks. 3+ Intensity (a.u.) 1. Fe 2+ 2. Fe 2. 1. 720 718 716 714 712 710 708 706 704 702 700 Binding energy (eV) Figure 4.2 Fe 2p XPS of the IO showing Fe3+ and Fe2+ peaks. 33 4.2 Size Distribution and Iron loading Factors, such as particle size and surface property of nanoparticles, could influence how long nanoparticles remain in circulation, how the nanoparticles interact with cells and their ability to penetrate drug barriers such as the BBB and gastrointestinal tract. TEM image presented in Figure 4.3(a) shows that the iron cores of the IOs (the dark dots) are very small, approximately 5 nm. According to the literature, the hydrodynamic size of Resovist® is approximately 60 nm [55]. The difference observed here is due to the dextran coating on the iron cores. We also used TEM to study encapsulation and location of IOs in nanoparticles. The TEM image in Figure 4.3(b) confirms that IOs (dark domains) are encapsulated in the polymer matrix of PLGA-mPEG nanoparticles. 34 Figure 4.3 TEM images of (a) the IOs (bar = 20 nm) and (b) the IO loaded PLGA-mPEG nanoparticles (bar = 50 nm). The hydrodynamic diameter and polydispersity of the nanoparticles can be obtained using LLS. Polydispersity is a quantitative measure of the uniformity of the nanoparticles. Generally, a uniform distribution is pursued. The amount of iron incorporated and average hydrodynamic size of both IOs and IO loaded PLGAmPEG nanoparticles were summarized in Table 4.1. The actual size distribution of the IO loaded nanoparticles was ±12.5 nm, as shown in Figure 4.4. As the size distribution was fairly narrow, the size of the IO loaded nanoparticles could be considered to be quite uniform. Table 4.1 Properties of the IOs and IO loaded PLGA-mPEG nanoparticles. Sample Fe content (%) Average hydrodynamic diameter (nm) IOs 22.02 ±2.31(n=5) 60 [35] IO loaded PLGA-mPEG 1.37±0.02 (n=5) 233.0 nanoparticles 35 IO loaded PLGA-mPEG nanoparticles 100 Intensity (a.u.) 80 60 40 20 0 150 200 250 300 350 Hydrodynamic diameter (nm) Figure 4.4 Particle size distribution of IO loaded PLGA-mPEG nanoparticles. 4.3 Surface Morphology Surface morphology of IO loaded nanoparticles gives indication if any unencapsulated IOs present in the system as it may affect both magnetic property and release kinetics of IOs from the nanoparticles. FESEM image in Figure 4.5 shows that these nanoparticles are spherical and have relatively uniform size. No free IOs could be observed on the surface of these nanoparticles. 36 Figure 4.5 FESEM images of the IO loaded PLGA-mPEG nanoparticles (bar = 1 µm). 4.4 Surface Charge Surface charge determines whether the nanoparticles will agglomerate in blood, their adhesion and interaction with the negatively charged cell membranes. It also affects the stability of the nanoparticles [56]. Surface charge is indicated by the zeta potential. The greater the zeta potential, the more stable the nanoparticles suspension will be because of the electrostatic repulsion. The zeta potential of the IO loaded PLGA-mPEG nanoparticles was negative as shown in Table 4.2. This can be attributed to the presence of ionized carboxyl groups and oxygen atoms of the ester groups on the surface of the nanoparticles [57]. 37 Table 4.2 Zeta potential of the IO loaded NPs 4.5 Sample Zeta potential (mV) PLGA-mPEG + IO -28.45 ± 0.99 Stability The stability of the IO loaded PLGA-mPEG nanoparticles was studied by measuring changes in the particle size and the level of iron leakage. There was no significant change (less than 5%) in the size of the nanoparticles when exposed to NaCl solution at 37◦C during the period of the study as seen in Figure 4.6. Therefore, it can be deduced that no aggregation had occurred in 48 hrs and the PLGA-mPEG nanoparticles were resistant to electrolytes. This stability of PLGA-mPEG nanoparticles is a result of the steric stabilization provided by mPEG molecules [58]. After 48 hours, the iron leakage was measured. Only 0.4% of the encapsulated IO leaked out. Thus, the formulations exhibited good stability in presence of 0.9% NaCl.. 38 500 Mean diameter (nm) 400 300 200 100 0 0 10 20 30 40 50 Time (h) Figure 4.6 Stability of 233 nm IO loaded PLGA-mPEG nanoparticles in NaCl solution at 37◦C. 4.6 In vitro release profile From the in vitro release profile shown in Figure 4.7, it can be observed that the rate of release of IOs from the IO loaded PLGA-mPEG nanoparticles for the first 4 days was pretty constant, and then it increased a little (from day 5 to 6) before slowing down again. At the end of 9 days, about 20% of IOs were released. Subsequently, there was very little IOs release. In fact after 31 days, only 21.3% (SD = 1.8%) of IOs were released (Data is not shown in Figure 4.7). Hence, it can be deduced that iron leakage from the IO loaded PLGA-mPEG nanoparticles was very slow. Since the nanoparticles were non-toxic, they hold the potential to be used for prolonging MR imaging provided that the nanoparticles were not cleared from the body during the scan time. 39 30 IO loaded PLGA-mPEG nanoparticles in vitro release in PBS Cumulative release (%) 25 20 15 10 5 0 0 1 2 3 4 5 6 7 8 9 10 Time (day) Figure 4.7 In vitro release profile of the IO loaded PLGA-mPEG nanoparticles in PBS at 37◦C. 4.7 Summary We have shown that the IOs used for encapsulation are Fe3O4 through XRD and XPS. The IO loaded PLGA-mPEG nanoparticles are 233 ± 12.5 nm in diameter, have negative surface charge, and their Fe loading is about 1.37%. Successful encapsulation of the IOs inside the PLGA-mPEG matrix is verified by TEM and XPS. Through FESEM pictures, we can observe that the IO loaded PLGA-mPEG nanoparticles are spherical. It is also demonstrated that the IO loaded PLGA-mPEG nanoparticles are stable with little changes in size and insignificant amount of iron leakage after proplonged exposure to osmotic NaCl solution. Only a small amount of 40 IOs were released from the IO loaded PLGA-mPEG nanoparticles after 31 days. So far, the IO loaded PLGA-mPEG nanoparticles have exhibited physicochemical properties suitable for MRI applications. 41 Chapter 5 Magnetization properties The aim of this part of the project is to investigate the magnetic properties of the IOs and IO loaded PLGA-mPEG nanoparticles so as to access their feasibility as contrast agent for MRI and their efficacy as compared to the IOs that are currently being used. 5.1 Characteristics of superparamagnetic materials To obtain more information regarding the magnetic properties of the IOs and IO loaded PLGA-mPEG nanoparticles, the field dependence of the magnetization at a constant temperature, and specifically, the characteristics of the hysteresis cycle are evaluated. This is shown in Figure 5.1 for both IOs and IO loaded PLGA-mPEG nanoparticles. It can be observed that the two types of material generally display similar magnetic behaviour. They show characteristic of superparamagnetic particles with zero hysteresis cycle, and no coercive field and remanent magnetization. Hence, it can be deduced that encapsulation retains the superparamagnetism of the IOs. 42 100 Magnetization (emu/g) IO loaded PLGA-mPEG nanoparticles IOs 50 Hysteresis loops at 300K 0 -50 -100 -4000 -2000 0 2000 4000 Field (Oe) Figure 5.1 Magnetization curve for IOs and IO loaded PLGA-mPEG nanoparticles at 300K. 5.2 Magnetization – temperature dependence Figure 5.2 illustrates the influence of temperature on the magnetization for IOs for an applied field 20 kOe. Basically, a parabolic decrease of magnetization with temperature is observed. This is because as temperature rises, the increase in thermal motion interferes with the order produced by the molecular field which is responsible for the parallel orientation of the magnetic moments of a domain [13]. Figure 5.2 also reveals that the differences in magnetization between IOs and IO loaded PLGAmPEG nanoparticles became greater as the temperature increased. At 300 K, the saturation magnetization of IO loaded PLGA-mPEG nanoparticles was 83.5 emu/g while that of IOs was 72.9 emu/g. 43 It is known that magnetization of the IOs is directly correlated to their size: the larger the size the stronger the magnetization [59]. This could be the possible reason for the increased Ms of the IOs after encapsulation in this study, as IO agglomeration might occur during the formulation process. Actually, change of magnetic properties of the IOs after polymer encapsulation has been reported previously, but a decreased Ms was observed [13]. This could be due to the encapsulation of a single magnetic nucleus instead of a cluster of IOs, as TEM images taken showed a single iron core in Magnetization (emu/g) the polymer matrix. 100 95 90 85 80 75 70 65 60 55 IOs IO loaded PLGA-mPEG nanoparticles 0 100 200 300 400 Temperature (K) Figure 5.2 Magnetization as a function of temperature for the IOs and the IO loaded PLGA-mPEG nanoparticles (applied field 20 kOe). 44 5.3 Blocking temperature TB The divergence between the susceptibility in a ZFC process and in a FC process is another typical feature of superparamagnetic materials. It arises from the anisotropy barrier blocking of the magnetization orientation in the nanoparticles cooled with a ZFC process [60], thus demonstrating that superparamagnetism is indeed preserved after encapsulation. It can be observed in Figure 5.3 that as temperature increases, the ZFC magnetization increases and reaches a peak, where the temperature is known as TB. This is defined as the temperature at which the nanoparticle moments do not relax during the time scale of the measurement [59]. It is an important parameter in the study of a magnetic particle system as the IOs would exhibit superparamagnetic properties above TB [61]. Below TB, there is random orientation of the easy axes among the nanoparticles, the net susceptibility can be taken to be zero as the applied field is too small to overcome the magnetic anisotropy. Above TB, there is sufficient thermal energy to overcome the anisotropy and the nanoparticles are aligned according to the applied field. Therefore, TB marks the transition between the ferromagnetic and superparamagnetic states. After encapsulation, the blocking temperature had increased from 187 K to 212 K. It is known that the blocking temperature increases with IO particle size, as a greater energy for a larger particle size is required to overcome the anisotropy barrier. Again, this implies that there may be an agglomeration of IOs in the PLGA-mPEG matrix after encapsulation. 45 -2 χ (emu/(g.Oe)*10 ) (a) 0.36 0.34 0.32 0.30 0.28 0.26 0.24 0.22 0.20 0.18 0.16 0.14 0.12 0.10 187K ZFC FC without polymerized 0 100 200 300 400 Temperature (K) (b) (b) 212K 0.55 -2 χ (emu/(g.Oe)*10 ) 0.60 0.50 0.45 0.40 After polymerized 0.35 ZFC FC 0.30 0.25 0.20 0 100 200 300 Temperature (K) 400 Figure 5.3 Blocking temperatures of (a) IO and (b) IO loaded PLGA-mPEG nanoparticles. 46 5.4 Summary From the above experimental results, it can be seen that both the IOs and IO loaded PLGA-mPEG nanoparticles exhibit superparamagnetic behaviours. In addition, the IO loaded PLGA-mPEG nanoparticles have larger saturation magnetization and higher blocking temperature than the IOs. Both phenomena are known to occur when there is an increase in size of IOs, thus this can imply that there is an agglomeration of IOs inside the polymer matrix. Regardless of the mechanism behind the changes observed, incorporations of IOs into PLGA-mPEG nanoparticles have indeed altered the magnetization of the IOs, this may in turn affect the MR contrast effects of the IO loaded PLGA-mPEG nanoparticles. 47 Chapter 6 In vitro MR studies Since the magnetization measurements have shown that the IO loaded PLGA-mPEG nanoparticles had greater saturation magnetization than the commercial IOs, it warrants MRI experiments to be carried out to assess if this enhancement in magnetization can be correlated with higher MR relaxivities. 6.1 Relaxivity plots To obtain the r1, r2 and r2* relaxivities of the nanoparticles using MRI, images have to be taken over a range of TR and TE values. The relaxivity plots of IOs, and IO loaded PLGA-mPEG nanoparticles are presented in Figure 6.1 and the summary of their relaxivities are summarised in Table 6.1. There was little change observed with r1. However, the IO loaded PLGA-mPEG nanoparticles developed were about twice more efficient than the IOs based on their in vitro r2 and r2* data. 48 (a) IOloaded PLGA-mPEG nanoparticles IOs 6.0 -1 1/T1 Relaxation Rate (s ) 6.5 5.5 r1 = 11.4315 5.0 4.5 4.0 3.5 3.0 r1 = 7.47191 2.5 2.0 1.5 1.0 0.5 0.0 0.0 0.1 0.2 0.3 0.4 0.5 Concentration of Fe (mM) (b) -1 1/T2 Relaxation Rate (s ) 180 IO loaded PLGA-mPEG nanoparticles IOs 160 140 120 100 80 r2 = 532.73265 60 r2 = 282.36198 40 20 0 -20 0.0 0.1 0.2 0.3 0.4 0.5 Concentration of Fe (mM) 49 -1 1/T2* Relaxation Rate (s ) (c) 300 IOloaded PLGA-mPEG nanoparticles IOs 250 r2*=537.5239 200 150 100 r2*=266.46273 50 0 0.0 0.1 0.2 0.3 0.4 0.5 Concentration of Fe (mM) Figure 6.1 (a) r1, (b) r2 and (c) r2* relaxativities of the IOs and the IO loaded PLGA-mPEG nanoparticles. Table 6.1 r1, r2 and r2* relaxivities of the IOs and the IO loaded PLGA-mPEG nanoparticles. Sample r1 ( mM IOs 11.4 282.4 266.5 IO loaded PLGA-mPEG 7.5 532.7 537.5 −1 ⋅ s −1 ) r2 ( mM −1 ⋅ s −1 ) r2* ( mM −1 ⋅ s −1 ) nanoparticles 50 6.2 Qualitative analysis For qualitative analysis, a comparison of in vitro MR images of the IOs and IO loaded PLGA-mPEG nanoparticles suspended in water was conducted. From Figure 6.2, it can be observed that at TE=7 ms the IO loaded PLGA-mPEG nanoparticles produced darker images for all different concentrations of iron. This demonstrates that our IO loaded PLGA-mPEG nanoparticles could achieve greater a contrast effect than commercial IOs. Previously, Kim et al [62] had conducted in vitro MRI imaging of their SPIO developed to illustrate their contrast enhancement which was comparable to Resovist®. Figure 6.2 Comparison of IOs and IO loaded PLGA-mPEG nanoparticles at TE=7 ms 51 6.3 Investigations on encapsulation effects To further verify the effect of IO particle encapsulation on r2 and r2* relaxivity, MRI was carried out on blank PLGA-mPEG nanoparticles, IOs and mixtures of IOs with different concentrations of blank PLGA-mPEG nanoparticles. Figure 6.3 shows that blank PLGA-mPEG nanoparticles do not enhance the proton relaxation rate as their relaxation rates for different concentration are almost constant. In addition, the mixtures of blank PLGA-mPEG nanoparticles with 0.2 mM IOs had no effect on r2 and r2* relaxivities of IOs. Their relaxation rates were similar to that of 0.2mM IOs. This result confirms that it is the encapsulation of IOs with PLGA-mPEG that leads to the contrast enhancement, and pure physical mixing of IOs with blank nanoparticles cannot produce such an effect. 52 (a) 70 water 0.2mM IOs Blank PLGA-mPEG nanoparticles Blank PLGA-mPEG nanoparticles + 0.2mM IOs Relaxation Rate R2 (1/s) 60 50 40 30 20 10 0 0 28 140 280 Concentration of blank NPs (mg/L) (b) 100 water 0.2mM IOs Blank PLGA-mPEG nanoparticles Blank PLGA-mPEG nanoparticles + 0.2mM IOs Relaxation Rate R2* (1/s) 90 80 70 60 50 40 30 20 10 0 0 0 28 140 280 Concentration of blank NPs (mg/L) Figure 6.3 Relaxation rate (a) R2 and (b) R2* of blank PLGA-mPEG nanoparticles, IOs, and mixtures of them with different concentrations of blank PLGA-mPEG nanoparticles. 53 6.4 Theories behind relaxivity enhancement As mentioned earlier, the R2 relaxation rates increases with particle sizes in the motional averaging regime and the maximum R2 relaxation rate is achieved in SDR. Thus, it can be deduced that both the IOs and IO loaded PLGA-mPEG nanoparticles are either in the MAR or SDR. Gillis et al has reported that the maximum R2* relaxation rate that can be achieved by IOs when simulated at equatorial field of 1 kG ( ∆ωr = 2.36 × 107 rad / s ) and volume fraction of v = 5 × 10 −6 is equal to 160 s-1 [36]. Making use of their simulated results, we can estimate the maximum R2* relaxation rate for IOs which is actually magnetite. Given the magnetite volumic mass (5100kg/m3) and its chemical composition (Fe3O4), the volume fraction of magnetite for 1mM of iron is v = 15.2 × 10 −6 . By substituting v = 1.52 × 10 −6 and ∆ωr = 3.07 × 107 rad / s (equatorial field of magnetite is 1.3 kG) into (5), the maximum R2* relaxation rate is approximately 630 s-1. Since relaxivity is defined as the relaxation rate for 1mM of iron, the maximum r2* relaxivity is 630 mM −1 ⋅ s −1 . The r2* relaxivity achieved by our IO loaded PLGA-mPEG nanoparticles falls within this upper theoretical bound, thus proving that our results is reasonable. A similar work carried out by Pouliquen et al [18] had reported a decrease MR relaxivities after encapsulation. Since they did not carry out any r2* measurements, one explanation could be that their results were not accounted for by the MAR or SDR. Another possible reason for the decreased relaxivities could be due to the increased distance 54 between the IOs in the polymer matrix and the water molecules at the surface of the polymer matrix. 6.5 Summary There are several prominent phenomenons that are observed. Firstly, the r2 and r2* relaxivities of the IO loaded PLGA-mPEG nanoparticles approximately doubled those of the IOs after encapsulation. Secondly, the r2 and r2* relaxivities of IO loaded PLGA-mPEG nanoparticles could be considered as equal. Last but not least, the MR images have also proved that with the same concentration of iron present, the IO loaded PLGA-mPEG nanoparticles gave darker images. Therefore, the IO loaded PLGA-mPEG nanoparticles show great potential to serve as a better contrast agent for clinical MR imaging. 55 Chapter 7 Animal studies The desirable R2 and R2*-relaxation enhancing properties of the IO loaded PLGAmPEG nanoparticles warrant animal studies of their efficacy as a MRI contrast agent. However, before doing that, it will be useful to know the biodistribution of the injected IOs and IO loaded PLGA-mPEG nanoparticles. The biodistribution data can provide an idea of specificity to the organ sites to be used for achieving the best MR image. In addition, the results obtained can be used to investigate the effects of encapsulation of IOs on the biodistribution. In this section, the results of the biodistribution and ex vivo MRI are presented and discussed. 7.1 Biodistribution studies The distribution of iron in the various organs (heart, kidney, liver, spleen and brain) of the rats (control, IOs, and IO loaded PLGA-mPEG nanoparticles) was shown in Figure 7.1. From the graph, it can be observed that most of the IOs and IO loaded PLGA-mPEG nanoparticles were found in the liver. This is expected because majority of particles injected intravenously are lost to the reticulo-endothelial system (RES), they are taken up mainly by macrophages in the liver after opsonization by proteins in the bloodstream. The degree of opsonization depends on the size and surface properties of particles. The liver and spleen usually take up particles between 50-500 nm which coincide with the size of the IOs and IO loaded PLGA-mPEG 56 nanoparticles. This natural tendency enables passive targeting to RES. Thus to avoid targeting to RES, particles have to overcome opsonization and thus uptake by RES. For effective targeting to other organs, ligands such as folate [48] and lectins [49] may have to be attached to the surface of the nanoparticles. Figure 7.1 Biodistribution of iron in various organs (1 hr after injection) 7.2 Ex vivo MRI The IO loaded PLGA-mPEG nanoparticles were injected into the tail vein of male Sprague Dawley rats. The liver of the rat injected with IO loaded PLGA-mPEG nanoparticles in Figure 7.2 was shown darker compared to the control, thus 57 demonstrating the efficacy of these nanoparticles in enhancing the in vivo proton relaxation rate. Ferrofluid encapsulated PLGA particles had been investigated as a MRI contrast agent in the kidney of rabbit by other researchers, and the enhanced contrast of MRI image was reported after the injection of the composite nanoparticles [2]. Figure 7.2 MR imaging of the livers of the rats (upper is the control; bottom is that of the rat injected with IO loaded PLGA-mPEG nanoparticles). 7.3 Summary The biodistribution shows that the distribution of IO loaded PLGA-mPEG nanoparticles does not differ from that of the IOs. Both of them were found to accumulate in the liver which is expected as the liver is responsible for removal of foreign particles in the body. In order to evade the RES and reach the other organs like the brain, the IOs and IO loaded PLGA-mPEG nanoparticles may require surface modification and/or size reduction for successful penetration of the blood-brain 58 barrier. The ex vivo MR image of the rats’ livers demonstrates that the IO loaded PLGA-mPEG nanoparticles were indeed effective at reducing the signal intensity and produce darker image. Therefore, they can be used as MRI contrast agents. 59 Chapter 8 Conclusions and Recommendations 8.1 Conclusions In this project, IO loaded PLGA-mPEG nanoparticles were developed for use in MRI as an alterative to IOs because there is a need to develop special contrast agents that increase the MRI signal intensity for future applications such as imaging of specific molecular targets to allow for earlier recognition and characterization of disease, earlier and direct evaluation of treatment outcomes, and a deeper understanding of disease development. After thorough examinations of the physical and chemical properties, the IOs are found to be Fe3O4, the iron cores were about 5 nm as seen in TEM picture and the overall size (including the dextran coating) was about 60 nm. While the IO loaded PLGA-mPEG nanoparticles were approximately 233 nm, spherical, and negatively charged. They also exhibited size and storage stability as there was negligible changes in size and insignificant amount of iron leakage from the nanoparticles after 48 hours of exposure to osmotic agent NaCl. Only a small amount of IOs released from the IO loaded PLGA-mPEG nanoparticles even after 31 days. Thus, the nanoparticle formulation exhibited suitable properties for its MRI applications. 60 Investigations of the magnetic properties of the IOs and IO loaded PLGA-mPEG nanoparticles show that both the IOs and IO loaded PLGA-mPEG nanoparticles display superparamagnetic property such as zero hysteresis loops above blocking temperature. It is also observed that after encapsulation, Ms increased from 72.9 emu/g to 83.5 emu/g at 300 K and TB increased from 187 K to 212 K. This results is in agreement with the report in the literature [59, 61]. Therefore, the increased saturation magnetization and blocking temperature of IO loaded PLGA-mPEG nanoparticles can be due to agglomeration of IOs inside the polymer matrix. Since the objective is to develop the IO loaded PLGA-mPEG nanoparticles as MR contrast agents, it is important to assess their MR properties. Relaxivity of a contrast agent is the key factor in evaluating its effectiveness, therefore in vitro MR studies were conducted to determine the r1, r2 and r2* relaxivities of the IOs and IO loaded PLGA-mPEG nanoparticles. After encapsulation, there was insignificant change in r1 relaxivity while r2 relaxivity increased from 282.4 mM −1 s −1 to 532.7 mM −1 s −1 and the r2* relaxivity increased from 266.5 mM −1 s −1 to 537.5 mM −1 s −1 . IOs act as contrast agents in MRI by decreasing the signal intensity of images, hence only the r2 and r2* relaxivities are of importance. Since the r2 and r2* relaxivities almost doubled after encapsulation and in vitro MR images of IO loaded PLGA-mPEG nanoparticles were darker than that IOs, it can be said that the IO loaded PLGA-mPEG nanoparticles were more effective as MR contrast agent than the IOs. The increase in relaxivities may be explained by the MAR and SDR. Within the MAR, the relaxivity increases as size of IOs increases up to a maximum. The maximum relaxivity is 61 determined by the SDR. From theoretical calculation, the maximum relaxivity for IOs which is around 630 mM −1 s −1 has not been reached, thus the increase in relaxivity after encapsulation is reasonable. As in vitro results may not truly reflect the situations when used in animals, it is necessary to conduct animal experiments to show that the IO loaded PLGA-mPEG nanoparticles can act as effective MR contrast agents in animals. Prior to ex vivo imaging, biodistribution data was collected to determine the amount of IOs and IO loaded PLGA-mPEG nanoparticles accumulated in some of the major organs (heart, kidney, liver, spleen and brain). The results showed that the IOs and IO loaded PLGA-mPEG nanoparticles were found in the kidney, liver and spleen. In fact, liver had the highest concentration of both agents. This is expected because the macrophages in the liver are known to ingest particles of sizes 50 to 500 nm. The particles were practically not existing in the heart and brain. The ex vivo MR image of rats’ liver illustrated that the IO loaded PLGA-mPEG nanoparticles could produce a darker image of the liver. In conclusion, the objective of this project has been met since encapsulation had increased the relaxation rates of the IOs, making the IO loaded PLGA-mPEG nanoparticles being more effective as MR contrast agents than the commercial IOs (Resovist®). 62 8.2 Recommendations What has been done so far is just the foundation; there are several ways in each this project can be taken further in order to achieve better results. In this section, I shall give some suggestions on further developments. 1. Dynamic MRI may be carried out to determine how long the IO loaded PLGAmPEG nanoparticles stay in the body before they are excreted. 2. Surface of the nanoparticles can be modified to cross physiological barrier like the BBB so that images of other organs like the brain may be able to be imaged by MRI. To allow MR imaging of the brain, the surface of the nanoparticles may be coated with surfactant like tween 80 which can aid in penetrating the blood-brain barrier. The size of nanoparticles may have to be reduced by modifying the encapsulation method. 3. It may be desirable to develop uncoated IOs which have higher saturation magnetization than Resovist® so as to further amplify the MRI signal. 4. Ligands may be attached to the surface of nanoparticles to allow active targeting. An example will be attaching folate ligands so that the nanoparticles can be specifically targeted to cells like breast tumor cells where folate receptors are found in high concentrations. 5. Drugs may be incorporated into the polymer matrix so that the nanoparticles can act as both a therapeutic and diagnostic agent. 63 References [1] Weissleder R, Mahmood U. Molecular Imaging. Radiology 2001; 219:316333. [2] Lee SJ, Jeong JR, Shin SC, Huh YM, Song HT, Suh JS, Chang YH, Jeon BS, Kim JD. Intracellular translocation of superparamagnetic iron oxide nanoparticles [3] Lanza GM, Winter PM, Caruthers SD, Morawski AM, Schmieder AH, Crowder KC, Wickline SA. Magnetic resonance molecular imaging with nanoparticles. J Nucl Cardiol 2004; 11(6): 733-743. [4] Keevil SF. Magnetic resonance imaging in medicine. Phys Educ 2001; 476485. [5] Chatterjee J, Haik Y, Chen CJ. Size dependent magnetic properties of iron oxide nanoparticles. J Magn Magn Mater 2003; 257(1): 113-118. [6] Jordan A, Scholz R, Maier-Hauff K, Johannsen M, Wust P, Nadobny J, Schirra H, Deger S, Loening S, Lanksch W, Felix R. Presentation of a new magnetic field therapy system for the treatment of human solid tumours with magnetic fluid hyperemia. J Magn Magn Mater 2001; 225: 118-126. [7] Chatterjee J, Haik Y, Chen CJ. Size dependent magnetic properties of iron oxide nanoparticles. J Magn Magn Mater 2003; 257(1): 113-118. [8] Tartaj P, Morales MP, Veintemillas-Verdaguer S, Gonzalez-Carreno T, Serna CJ. The preparation of magnetic nanoparticles for application in biomedicine. J 64 Phys D: Appl Phys 2003; 36: R182-97. [9] Moghimi SM, Hunter ACH, Murray JC. Long-circulating and target-specific nanoparticles: theory to practice. Pharm Rev 2001; 53: 283-318. [10] Chambon C, Clement O, Blanche AL, Schouman-Claeys E, Frija G. Superparamagnetic iron oxides as positive MR contrast agents: in vitro and in vivo evidence. Magn Reson Imaging 1993; 11: 509-519. [11] Müller RH, Maassen S,Weyhers H, Specht F, Lucks JS. Cytotoxicity of magnetite-loaded polylactide/glycolide particles and solid lipid nanoparticles. Int J Pharm 1996; 138: 85–94. [12] Okassa LN, Marchais L, Douziech-Eyrolles S, Cohen-Jonathan M, Souc´e P, Dubois I, Chourpa. Development and characterization of sub-micron, poly(d,llactide-co-glycolide) particles loaded with magnetite/maghemite nanoparticles. Int J Pharm 2005; 302: 187-196. [13] Gomez-Lopera SA., Plaza RC, Delgado AV. Synthesis and characterization of spherical magnetite/biodegradable polymer composite particles. J Colloid Interface Sci 2001; 240: 40–47. [14] Lee SJ, Jeong JR, Shina SC, Kim JC, Changa YH, Chang YM, Kim JD. Nanoparticles of magnetic ferric oxides encapsulated with poly(D,L latide-coglycolide) and their applications to magnetic resonance imaging contrast agent. J Magn Magn Mater 2004; 272-276: 2432-2433. [15] Dresco PA, Zaitsev VS, Gambino RJ, Chu B. Preparation and properties of magnetite and polymer magnetite nanoparticles. Langmuir 1999; 15: 19451951. 65 [16] Pich A, Bhattacharya S, Ghosh A, Adler H-JP. Composite magnetic particles: 2. Encapsulation of iron oxide by surfactant-free emulsion polymerization. Polymer 2005; 46(13): 4596-4603. [17] Zheng WM, Gao F, Gu HC. Magnetic polymer nanospheres with high and uniform magnetite content. J Magn Magn Mater 2005; 288: 403-410. [18] Pouliquen D, Perdrisot R, Ermias A, Akoka S, Le Jeune JJ. Superparamagnetic iron oxide nanoparticles as liver MRI contrast agent: Contribution of microencapsulation to improved biodistribution. Magn Reson Imaging 1989; 7: 619-627. [19] Salata OV. Applications of nanoparticles in biology and medicine. J Nanobiotechnol 2004; 2(1): 3. [20] Feng SS, Chien S, Chemotherapeutic engineering: Application and further development of chemical engineering principles for chemotherapy of cancer and other diseases. Chem Eng Sci 2003; 58: 4087 – 4114. [21] Chen Y, Dalwadi G, Benson HAE. Drug Delivery across the Blood-Brain Barrier. Curr Drug Deliv 2004; 1(4): 1-16. [22] Zambaux MF, Bonneaux F, Gref R, Maincent P, Dellacherie E, Alonso MJ, Labrude P, Vigneron C. Influence of experimental parameters on the characteristics of poly(lactic acid) nanoparticles prepared by double emulsion method. J Control Release 1998; 50: 31–40. [23] Mu L, Feng SS. Fabrication, characterization and in vitro release of paclitaxel loaded poly(lactic-co-glycolic acid) (PLGA) nanospheres prepared by the spray dry technique with phospholipids/cholesterol as additives. J Control Release 2001; 76: 239–254. 66 [24] Jaime N, Delgado JN, & William AR. Wilson & Gisvold’s textbook of organic, medicinal and pharmaceutical chemistry, 1998 (10th ed.). Philadelphia: Lippincott-Raven. [25] Murakami H, Yoshino H, Mizobe M., Kobayashi M, Takeuchi H, Kawashima, Y. Preparation of poly(,-lactide-co-glycolide) latex for surface modifying material by a double coacervation method. Proc Int Symp Control Release Bioactive Mater 1996; 23: 361–362. [26] Kompella UB, Koushik K. Preparation of drug delivery systems using supercritical fluid technology. crc Crit Rev Ther Drug Carrier Sys 2001; 18(2): 173–199. [27] Couvreur P, Kante B, Roland M, Goit P, Bauduin P, Speiser P. Polycyanoacrylate nanocapsules as potential lysosomotropic carriers: Preparation, morphology and sorptive properties. J Pharm Pharmacol 1979; 31: 331–332. [28] Gupta AK, Gupta M. Synthesis and surface engineering of iron oxide nanoparticles for biomedical applications. Biomaterials 2005; 26: 3995-4021. [29] Roch A, Muller RN. Proc 11th Annu Meet Soc Magn Reson Med, Works in Prog 1992; 1447. [30] Roch A, Muller RN, Gillis P. J Chem Phys 1999; 110: 5403. [31] Roch A, Muller RN, Gillis P. Magn Reson Imaging 2001; 14: 94. [32] Laurent S, Nicotra C, Gossuin Y, Roch A, Ouakssim A, Vander Elst L, Cornant M, Soleil, Muller RN. Influence of the length of the coating molecules on the nuclear magnetic relaxivity of superparamagnetic colloids. Phys Stat Sol 67 (c) 2004; 1(12): 3644-3650. [33] Brooks RA, Moiny F, Gillis P. On T2-shortening by weakly magnetized particles: the chemical exchange model. Magn Reson Med 2001; 45: 10141020. [34] Kiselev VG, Posse S. Analytical model of susceptibility-induced MR signal dephasing: effect of diffusion in a microvascular network. Magn Reson Med 1999; 41: 499-509. [35] Jensen JH, Chandra R. Strong field behaviour of the NMR signal from magnetically heterogeneous tissues. Magn Reson Med 2000; 43: 226-236. [36] Gillis P, Moiny F, Brooks RA. On T2-shortening by strongly magnetized spheres: a partial refocusing model. Magn Reson Med 2002; 47: 257-263. [37] Ngaboni Okassa L, Marchais H, Douziech-Eyrolles L, Cohen-Jonathan S, Souce M, Dubois P, Chourpa I. Development and characterization of submicron poly(D,L-lactide-co-glycolide) particles loaded with magnetite/maghemite nanoparticles. Int J Pharm 2005; 302: 187-196. [38] Jeong JR, Lee SJ, Kim JD, Shin SC. Magnetic properties of Fe3O4 nanoparticles encapsulated with poly(D,L lactide-co-glycolide). IEEE Trans on Magnetics 2004; 40: 4: 3015-3017. [39] Chatterjee J, Haik Y, Chen CJ. Polyethylene magnetic nanoparticle: A new magnetic material for biomedical applications. J Magn Magn Mater 2002; 246: 382-391. [40] Zhitomirsky I, Niewczas M, Petric A. Electrodeposition of hybrid organicinorganic films containing iron oxide. Mater Lett 2003; 57: 1045-1050. 68 [41] Mauduit J, Bukh N, Vert M. Gentamycin/poly(lactic acid) blends aimed at sustained release local antibiotic therapy administered per-operatively. I. The case of gentamycin base and gentamycin sulfate in poly(DL-lactic acid) oligomers. J Control Release 1993; 23: 209-220. [42] Mauduit J, Bukh N, Vert M. Gentamycin/poly (lactic acid) blends aimed at sustained release local antibiotic therapy administered per-operatively. II. The case of gentamycin sulfate in high molecular weight poly (DL-lactic acid) and poly (L-lactic acid). J Control Release 1993; 23: 221-230. [43] Mauduit J, Bukh N, Vert M. Gentamycin/poly (lactic acid) blends aimed at sustained release local antibiotic therapy administered per-operatively. III. The case of gentamycin sulfate in films prepared from high and low molecular weight poly (DL-lactic acids) J Control Release 1993; 25: 43-49. [44] Sah HK, Chien YW. Effects of H+ liberated from hydrolytic cleavage of polyester microcapsules on their permeability and degradability. J Pharm Sci 1995; 84(11): 1353-1359. [45] Neuberger T, Schopf B, Hofmann H, Hofmann M, Rechenberg B. Superparamagnetic nanoparticles for biomedical applications: Possibilities and limitations of a new drug delivery system. J Magn Magn Mater 2005; 293: 483-496. [46] Mareda H. The enhanced permeability and retention (EPR) effect in tumour vasculature: the key role of tumor-selective macromolecular drug targeting. Adv Enzyme Regul 2001; 41: 189-207. [47] Sahoo SK, Sawa T, Fang J, Tanaka S, Miyamoto Y, Akaike T, Maeda H. 69 Pegylated zinc protoporphrin: a water-soluble heme oxygenase inhibitor with tumor targeting capacity. Bioconjug Chem 2002; 13: 1031-1038. [48] Aronov O, Horowitz AT, Gabizon A, Gibson D. Folate-Targeted PEG as a Potential Carrier for Carboplatin Analogs. Synthesis and in Vitro Studies. Bioconjug Chem 2003; 3: 563-74. [49] Bies C, Lehr CM, Woodley JF. Lectin-mediated drug targeting: history and applications. Adv Drug Deliv Rev 2004; 56: 425-35. [50] Sun WQ, Xie CS, Wu HF, Hu Y. Specific role of polysorbate 80 coating on the targeting of nanoparticles to the brain. Biomaterials 2004; 25: 3065-3071. [51] Kreuter J. Application of nanoparticles for the delivery of drugs to the brain. Int Congress Series 2005; 1277: 85– 94. [52] Kreuter J, Ramge P, Petrov V, Hamm S, Gelperina SE, Engelhardt B, Alyautdin R, Briesen HV, Begley SJ. Direct Evidence that Polysorbate-80Coated Poly(Butylcyanoacrylate) Nanoparticles Deliver Drugs to the CNS via Specific Mechanisms Requiring Prior Binding of Drug to the Nanoparticles. Pharm Res 2003; 20: 3: 409-416. [53] Alyautdin RN, Petrov VE, Langer K, Berthold A, Kharkevich, Kreuter J. Delivery of Loperamide across Blood-Brain Barrier with Polysorbate 80Coated Polybutylcyanoacrylate Nanoparticles. Pharm Res 1997; 14: 3: 325328. [54] Krachler M, Radner H, Irgolic KJ. Microwave digestion methods for the determination of trace elements in brain and liver samples by inductively coupled plasma mass spectrometry. Fresenius J Anal Chem 1996; 355: 120- 70 128. [55] Kirchin MA, Runge VM. Contrast Agents for Magnetic Resonance Imaging: Safety Update. Topics in Magn Reson Imaging, Contrast-Enhanced Magn Reson Imaging 2003; 14(5): 426-435. [56] A. M. Clarence and P. Neogi. Interfacial phenomena: equilibrium and dynamic effects. Marcel Dekker 1985; volume 17. [57] S. Stolnik, M. C. Garnett, M. C. Davies, L. Illum, M. Bousta, M. Vert, and S. S. Davis. The colloidal properties of surfactant-free biodegradable nanospheres from poly(-malic acid-co-benzyl malate)s and poly(lactic acid-co-glycolide). Colloids Surf A 1995; 97: 235-245. [58] Avgoustakis K, Beletsi A, Panagi Z, Klepetsanis P, Evangelatos G, Ithakissios DS. Effect of copolymer composition on the physicochemical characteristics, in vitro stability, and biodistribution PLGA–mPEG nanoparticles. Int J Pharm 2003; 259: 115–127. [59] Lin CR, Chiang RK, Wang JS, Sung TW. Magnetic properties of monodisperse iron oxide nanoparticles. J Appl Phys 2006; 99: 1-3. [60] Liu C, Rondinone AJ, Zhang ZJ. Synthesis of magnetic spinel ferrite CoFe2O4 nanoparticles from ferric salt and characterization of the size-dependent superparamagnetic properties. Pure Appl Chem 2000; 72: 37-45. [61] Gupta AK, Gupta M. Synthesis and surface engineering of iron oxide nanoparticles for biomedical applications. Biomaterials 2005; 26: 3995-4021. [62] Kim EH, Lee HS, Kwak BK, Kim BK. Synthesis of ferrofluid with magnetic nanoparticles by sonochemical method for MRI contrast agent. J Magn Magn 71 Mater 2005; 289: 328-330. 72 [...]... manipulates R1 of the surrounding molecules to increase the total signal In recent years, superparamagnetic iron oxides (SPIOs) that enhance R2 of the surrounding medium to produce signal voids on magnetic resonance images have been developed [4] Iron oxides (IOs) are the most-studied materials for magnetic targeting because of their favorable magnetic properties and high biocompatibility Superparamagnetic. .. out by imaging the organs of rats injected with IO loaded PLGA-mPEG nanoparticles Biodistribution of IO loaded PLGA-mPEG nanoparticles in rats were studied as well 6 1.3 Organization of thesis The thesis consists of (i) thorough literature review; (ii) description of materials and methods used in the novel formulation of biodegradable IO loaded PLGA-mPEG nanoparticles; (iii) results and discussions of. .. results are summarized, and some suggestions for future directions of this research are given 7 Chapter 2 Literature Review 2.1 Nanoparticles of Biodegradable Polymers 2.1.1 Basic information of Biodegradable Polymers Recently, there has been increased interest in developing long-circulating nanoparticles as a drug carrier The studies using polymeric biodegradable nanoparticles to encapsulate anti-tumor... of 120-200 nm in diameter, in order to decrease clearance by the reticuloendothelial system (RES) Nanoparticles used for drug delivery to the brain are generally the diameters of 60 – 400 nm Efforts have been made to modify the surface of nanoparticles to increase their systemic circulation time, by either physical adsorption of a hydrophilic polymer on the particle surface or chemical 9 grafting of. .. the particle surface or chemical 9 grafting of polymer chains onto particles To date, the most successful longcirculating biologically stable nanoparticles have been coated with PEG [21] 2.1.2 Manufacture techniques of nanoparticles There are many ways to manufacture the nanoparticles, for instance, dispersion of the preformed polymers or by polymerization of monomers [20] Some other more commonly used... half-life and biodistribution The first group are termed superparamagnetic iron oxides (SPIOs) where nanoparticles have a size greater than 50 nm (coating included) and the second type termed ultrasmall superparamagnetic iron oxides (USPIOs) where nanoparticles are smaller than 50 nm Both types of particles are commercially available Some examples of SPIOs are Lumirem®, silicon-coated particles with 300... (b) r2 and (c) r2* relaxativities of the IOs and the IO loaded 49 PLGA-mPEG nanoparticles Figure 6.2 Comparison of IOs and IO loaded PLGA-mPEG nanoparticles at TE=7 ms 51 Figure 6.3 Relaxation rate (a) R2 and (b) R2* of blank PLGA-mPEG nanoparticles, IOs, and mixtures of them with different concentrations of blank PLGA-mPEG nanoparticles 53 Figure 7.1 Biodistribution of iron in various organs (1 hr after... polymerization and interfacial polymerization Emulsion polymerization builds up a chain of polymers from single monomers When the monomer-contained organic phase and aqueous phase are brought together by mechanical force, interfacial polymerization will take place Couvreur et al [27] reported the production of nanoparticles of about 200 nm diameter by polymerizing mechanically the dispersed methyl or ethyl... SPIOs Superparamagnetic iron oxides SQUID superconducting quantum interference device TB blocking temperature TE time to echo TEM transmission electron microscopy TR time of repetition USPIOs ultrasmall superparamagnetic iron oxides w/o/w water in oil in water xii XPS X-ray photoelectron spectroscopy XRD X-ray diffraction ZFC Zero field cooling xiii Chapter 1 Introduction 1.1 Background Magnetic resonance. .. particles by coating a magnetic nucleus (magnetite) with a biodegradable PLA polymer, but they found these composite particles had decreased saturation magnetism Lee et al [14] prepared ferrofluidic PLGA nanoparticles and suggested that a decrease in particle size may increase the magnetic susceptibility of nanoparticles as a result of the increase in packing density or volume fraction of the nanoparticles ... some suggestions for future directions of this research are given Chapter Literature Review 2.1 Nanoparticles of Biodegradable Polymers 2.1.1 Basic information of Biodegradable Polymers Recently,... of iron loading (% w/w) was calculated as the ratio of the mass of iron (mg) that can be detected using ICP analysis to the sum of the mass of iron (mg) and the mass of polymer (mg) 3.3.7 Magnetic. .. are termed superparamagnetic iron oxides (SPIOs) where nanoparticles have a size greater than 50 nm (coating included) and the second type termed ultrasmall superparamagnetic iron oxides (USPIOs)