Incorporation of Magnetic Nanoparticles and Paramagnetic

Incorporation of magnetic nanoparticles and paramagnetic molecules into human red blood cells for use as proton relaxivity enhancers for blood pool magnetic resonance imaging.

Mounir Ibrahim (MSc)

This thesis is presented in partial fulfilment of the requirements for the degree of the Doctor of Philosophy of The University of Western Australia

School of Physics 2013

Abstract

Chelated gadolinium compounds and magnetic nanoparticles have been developed for use as magnetic resonance imaging (MRI) contrast agents. However, both chelated gadolinium compounds and magnetic particles are rapidly cleared from the blood stream, gadolinium via renal excretion and magnetic nanoparticles by the reticulo-endothelial system. In order to extend the lifetime of the contrast agents in the blood stream, we have encapsulated contrast agents within red blood cells (RBCs) as a potential strategy for slowing their clearance from the blood pool. A key strategy for loading RBCs with contrast agent is to incubate them in the presence of the contrast agents under hypo-osmolar conditions. Under these conditions the cells swell and holes open up in the cell membranes allowing particles to pass into or out from the cell. The RBCs can be re-sealed by bringing the osmolarity of the medium back up to physiological values.

Human RBCs were loaded with magnetic nanoparticles and gadoteric acid by two different methods. The methods comprised either hypo-osmolar incubation or a hypo-osmolar pulse in the presence of gadoteric acid or magnetic nanoparticles. The efficacy of the resulting RBCs as contrast agents for MRI was assessed by measuring the cell specific longitudinal and transverse proton relaxivities in a magnetic field of magnitude 1.4T.

Human RBCs were loaded with four different magnetic iron oxide nanoparticle systems (three different polymer coated systems and one uncoated system (-Fe2O3)) using the two techniques. Polymers used comprised poly(ethylene oxide) (PEO), poly(acrylic acid) (diblock), and propylene oxide-b-ethylene oxide (PPO-b-PEO) (triblock). The aim of this study was to compare the effectiveness of magnetic resonance contrast enhancement by RBCs loaded with magnetic nanoparticles by the two different methods. Here we also tested the hypothesis that only a fraction of the RBCs in a sample open pores to enable iron oxide nanoparticles loading with each of the techniques. The optimum osmolarity for particle loading was found to be close to 200 mOsm, a condition at which the greatest fraction of particles in the incubating medium enter the cells. The maximum size of aggregate that enters the cells under

Abstract these conditions was found to be approximately 120 nm. Osmolarities of 150 mOsm or lower cause excessive cell destruction. At osmolarities of 200 mOsm and above, and with the osmotic pulse method, cell survival was typically about 70%.

While use of uncoated magnetic nanoparticles was useful for studies of particle size limitations on cell loading, the inability to effectively wash extraneous particles from the cells makes quantitative interpretation of proton relaxivity enhancement problematical. Extraneous polymer-coated particles were easily washed from cells preparations facilitating quantitative interpretation of proton relaxometry measurements. For the polymer coated nanoparticle systems it was found that the osmotic pulse method of loading resulted in higher cell-specific transverse proton relaxivities (r2). Cell-specific transverse proton relaxivity up to 31 times higher than native red blood cells were achieved (using the triblock copolymer coated iron oxide nanoparticles introduced with the osmotic pulse technique).

Human RBCs were also loaded with gadoteric acid by either hypo-osmolar incubation or a hypo-osmolar pulse in the presence of gadoteric acid. The aim of this study was to compare the effectiveness of magnetic resonance contrast enhancement by RBCs loaded with gadoteric acid by these two different methods. The resulting enhancements in proton relaxivities of cell suspensions in 1.4 T were measured and the effect of incubation osmolarity on the resulting cell specific proton relaxivity was also studied. The osmotic pulse method was found to yield the greatest cell-specific relaxivity enhancements (71-fold for longitudinal relaxivity and 39-fold for transverse relaxivity). The spatial distribution of the gadolinium within the cells was studied using energy filtered transmission electron microscopy to generate gadolinium M-edge jump ratio images. All surviving cells exposed to gadoteric acid under hypo-osmolar conditions showed enhanced (relative to control cells) and generally uniform intensity within the cells in gadolinium jump ratio images suggesting all cells are susceptible to loading and that the loading is generally spatially uniform within each cell. There was some evidence for a small amount of precipitation or aggregation of gadolinium within some cells prepared by the hypo- osmolar incubation method.

Table of Contents

Table of contents

List of Figures………………………………………...... 9 List of Abbreviations………...... 15 Statement of candidate contribution ...... 16 Acknowledgements ...... 17

Chapter 1: Introduction 1.1 Magnetic Resonance Imaging development history ...... 18 1.2 Magnetic Resonance Imaging (Introduction) ...... 19 1.2.1 MRI scanner ...... 21 1.3 Role of contrast agents ...... 22 1.4 Limitations and side effects of current agents ...... 23 1.5 Human blood function and red blood cells (RBCs) ...... 24 1.5.1 Blood function ...... 24 1.5.2 Red blood cells ...... 25 1.6 The concept of using red blood cells as a “Trojan Horse” for contrast agents ...... 27 1.6.1 Equation for manipulating osmolarity (diffusion and osmosis) ...... 27 1.6.2 Manipulating red blood cells with osmotic pressure...... 31 1.6.3 Loading red blood cells: pore size and membrane reversibility ...... 32 1.6.4 Previous methods of drug loading into red blood cells ...... 32 1.7 Aims of this study ...... 34

Chapter 2: Theoretical Principles of MRI Contrast Agents 2.1 Basic magnetic resonance imaging theory ...... 36 2.1.1 Proton magnetic relaxation ...... 39

2.1.1.1 Longitudinal (or spin-lattice) relaxation time T1 ...... 39

2.1.1.2 The transverse relaxation times T2 and T2* ...... 40

2.1.1.3 Relationship between T2 and T2* ...... 42 2.2 Contrast agent mechanisms ...... 42 2.2.1 Ferromagnetism...... 44 2.2.2 Paramagnetism ...... 47 2.2.3 Superparamagnetism ...... 48

Table of Contents

2.3 Mechanisms of longitudinal Proton Relaxation Enhancement by paramagnetic materials ...... 49 2.4 Mechanisms of Transverse Proton Relaxtion Enhancement by Magnetized Particles ...... 50 2.4.1 Static dephasing regime ...... 51 2.4.2 Motional averaging regime ...... 52 2.4.3 Echo limited regime ...... 52

Chapter 3: Materials and Equipments 3.1 Magnetic Nanoparticles ...... 54

3.1.1 Polymer coated Fe3O4 nanoparticles ...... 54

3.1.2 -Fe2O3 nanoparticles from Sigma Aldrich ...... 54 3.1.2.1 Nanoparticle size fractionation ...... 55 3.2 Gadoteric acid...... 55 3.3 Transmission Electron Microscope (TEM) /EELS ...... 55 3.3.1 JEOL 2100F Transmission Electron Microscope (TEM) ...... 56 3.3.1.1 Electron gun ...... 58 3.3.1.2 Lens system ...... 58 3.3.1.3 Condenser lenses ...... 58 3.3.1.4 Objective lens ...... 58 3.3.1.5 Projector lenses ...... 59 3.3.1.6 Apertures ...... 58 3.3.1.7 Stigmators and deflectors ...... 59 3.4 Zetasizer Nano-ZS, Malvern Instruments ...... 60 3.5 Superconducting Quantum Interference Device (SQUID) Magnetometry ...... 62 3.6 The Minispec nuclear magnetic resonance (NMR) ...... 62 3.7 Scanning Probe Microscopy ...... 63 3.7.1 Dimension 3100 AFM ...... 65 3.8 Optical microscope ...... 65 3.9 Red blood cells (RBCs) ...... 65 3.10 Hemocytometer (Counting chamber) ...... 66 3.11 Loading RBCs with contrast agents ...... 66

Table of Contents

Chapter 4: Loading Erythrocytes with Maghemite Nanoparticles via Osmotic Pressure Induced Cell Membrane Pores 4.1 Introduction ...... 68 4.2 Methods and materials ...... 69 4.2.1 Preparation of Stock Solutions and Cell Suspensions ...... 69

4.2.1.1 Maghemite (γ-Fe2O3) nanoparticles ...... 69 4.2.1.2 Human Red Blood Cells (RBCs) ...... 69

4.2.2 Methods of Loading Cells with Maghemite (γ-Fe2O3) nanoparticles ...... 69 4.2.3 Measurement and Characterization of Cell Suspensions ...... 70 4.2.3.1 Transmission Electron Microscopy...... 70 4.2.3.1 Freeware ImageJ (NIH) measurements ...... 71 4.2.3.1 Inductively coupled plasma mass spectrometry (ICP-MS)...... 71 4.3 Results ...... 72

4.3.1 TEM and DLS of the -Fe2O3 nanoparticle ...... 72 4.3.2 Superconducting Quantum Interference Device (SQUID) Magnetometry ...... 73 4.3.3 TEM of RBCs ...... 74

4.3.4 The concentrations of -Fe2O3 iron inside the cells ...... 76 4.3.5 The distributions of particle/aggregate sizes inside and outside of the cells ...... 77 4.4 Discussion ...... 79 4.5 Conclusion ...... 81

Chapter 5 : Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods 5.1 Introduction ...... 82 5.2 Materials and Methods ...... 83 5.2.1 Preparation of Stock Solutions and Cell Suspensions ...... 83 5.2.1.1 Iron oxide nanoparticles ...... 83 5.2.1.2 Human Red Blood Cells (RBCs) ...... 85 5.2.2 Methods of Loading Cells with Magnetic Nanoparticles ...... 86 5.2.2.1 Hypo-osmolar Incubation Method (Osmotic incubation): ...... 86 5.2.2.2 Osmotic Pulse Method ...... 87 5.2.3 Measurement and Characterization of Cell Suspensions ...... 88 5.2.3.1 Cell counting ...... 88 5.2.3.2 Cell Size and Morphology ...... 88

Table of Contents

5.2.3.3 Dynamic light scattering (DLS) measurement Cell counting ...... 89 5.2.3.4 Proton relaxometry ...... 90 5.2.3.5 Calculation of Apparent Intracellular Proton Relaxation Rates and Loading Efficiencies...... 90 5.2.3.6 Transmission Electron Microscopy ...... 91 5.3 Results ...... 92 5.3.1 TEM and DLS of nanoparticles...... 92 5.3.2 Proton Relaxometry ...... 96 5.3.2 Apparent loading efficiency ...... 100 5.3.4 TEM of RBCs ...... 102 5.3.5 Nanoparticles size distribution inside cells ...... 108 5.3.6 The median projected cell area ...... 110 5.4 Discussion ...... 112 5.5 Conclusion ...... 114

Chapter 6: Enhancement of the Cell Specific Proton Relaxivities of Human Red Blood Cells via Loading with Gadoteric Acid 6.1 Introduction ...... 115 6.2 Materials and Methods ...... 116 6.2.1 Preparation of Stock Solutions and Cell Suspensions ...... 116 6.2.1.1 Gadoteric Acid Solution...... 116 6.2.1.2 Human Red Blood Cells ...... 116 6.2.1.3 Hemoglobin Solution ...... 116 6.2.2 Methods of Loading Cells with Gadolinium...... 117 6.2.2.1 Hypo-osmolar Incubation Method ...... 117 6.2.2.1 Osmotic Pulse Method ...... 117 6.2.3 Measurement and Characterization of Cell Suspension ...... 118 6.2.3.1 Cell Counting ...... 118 6.2.3.2 Cell Size and Morphology ...... 118 6.2.3.3 Proton relaxometry ...... 118 6.2.3.4 Transmission Electron Microscopy ...... 119 6.3 Results ...... 120 6.3.1 Proton Relaxometry ...... 120 6.3.2 Rates of survival of RBCs ...... 125

Table of Contents

6.3.3 The median projected cell areas ...... 126 6.3.4 TEM of RBCs ...... 127 6.3.5 Magnitudes of the gadolinium jump ratio image intensity ...... 129 6.4 Discussion and Conclusion ...... 129

Chapter 7: Atomic Force Microscopy and Magnetic Force Microscopy of iron oxide nanoparticles and Red blood cells loaded with iron oxide nanoparticles 7.1 Introduction ...... 133 7.1.1 Atomic force imaging modes ...... 136 7.1.2 Magnetic force microscopy (MFM) ...... 137 7.2 Methods and Materials ...... 139 7.2.1 Preparation of Stock Solutions and Cell Suspensions ...... 139 7.2.1.1 Iron oxide nanoparticles ...... 139 7.2.1.2 Human Red Blood Cells ...... 139 7.2.1.3 Methods of Loading RBCs with iron oxide nanoparticles ...... 139 7.2.2 Measurement and Characterization ...... 140 7.2.2.1 AFM scanning ...... 140 7.2.1.2 MFM scanning ...... 140 7.2.1.3 Samples preparations ...... 141 7.3 Results ...... 142

7.3.1 MFM of γ-Fe2O3 nanoparticles sample ...... 142

7.3.2 MFM with a magnet under the sample of -Fe2O3 nanoparticles sample ...... 144 7.3.3 MFM of RBCs section ...... 146 7.3.4 AFM of RBCs at different osmolarity ...... 149 7.3.5 AFM of RBCs in liquid ...... 149 7.4 Discussion ...... 150 7.5 Conclusion ...... 151

Chapter 8: Conclusion and future work 8.1 Discussion ...... 153 8.2 Speculation about the relationship between the pores that open up in RBCs membranes and the diffusion time of the water, nanoparticles, gadoteric acid and hemoglobin ...... 158 8.2.1 Friction and Diffusion ...... 158

Table of Contents

8.2.1.1 Diffusion time of the different components into and out from the cells during the loading period ...... 159 8.2.2 Our point of view of the loading process ...... 160 8.2.2.1 Consideration of rates of diffusion of different sized entities with regard to loading of RBCs by osmotic pressure differences ...... 160 8.3 Key conclusion ...... 164 8.3.1 Future directions ...... 165 8.4.1 Magnetic fractionations...... 165 8.4.2 There dimension images ( 3D images) ...... 165 8.4.3 In vivo studies (mice and rats) … ...... 165 8.4.4 Toxicology ...... 166

Bibliography ...... 167 Appendices ...... 173

List of Figures

Figure 1.1: The application of an external magnetic field, B0, causes the magnetic moments of the protons to process about the field direction ...... 20

Figure 1.2: An MRI image showing a defect in the blood-brain barrier after astroke ...... 21

Figure 1.3a: Magnetic Resonance Imaging Scanner cutaway ...... 21

Figure1.3b: Cardiac angiograph ...... 22

Figure 1.4: Atomic Force Microscopy image of Red blood cells ...... 25

Figure 1.5: The structure of haemoglobin: adult haemoglobin is a tetramerichemeprotein [α (2):β(2)] found in erythrocytes ...... 26

Figure 1.6: Chemical potential (µ) vers concentration (C) ...... 30

Figure 1.7: The diffusion of liquid between zone 1 and zone 2 via a semi-permeable membrane ...... 30

Figure 1.8: Red blood cells at different osmotic pressures ...... 32

Figure 2.1: A proton rotating about its own axis ...... 38

Figure 2.2: Proton spin states rotating on their own axis and wobbling about the axis of B0 ...... 38

Figure 2.3: Relaxation of the longitudinal component of magnetization after the application of a radio frequency pulse where, Mz(0) is the longitudinal magnetization immediately after the excitation pulse, M0 is the block saturation magnetization at equilibrium ...... 40

Figure 2.4: The relaxation of the transverse component of magnetization after the application of a radio frequency pulse. Where Mxy (t) is the transvers magnetization at time - t after the pulse and Mxy (0) is the initial maximum magnetization immediately after the pulses...... 41

Figure 2.5: Ferromagnetic materials domains: (a) the domains are randomly oriented (b) the domains are forced into alignment by an external magnetic field ...... 45

Figure 2.6: Closure domains at the end regions of sample ...... 46

Figure 2.7: Hysteresis loop in ferromagnetic, paramagnetic and superparamagnetic samples...……...... 46

Figure 2.8: Schematic representation of the alignment of magnetic moments in a paramagnetic sample (a) in the absence of an external magnetic field the magnetic moments are randomly oriented. (b) an applied magnetic field causes the magnetic dipoles to re-orient according to the magnetic field (H) ...... 48

List of Figures

Figure 3. 1: Schematic representation of the main parts of the TEM ...... 57

Figure 3.2: Schematic representation of the electron beams transmitted through a specimen...... 60

Figure 3.3: The Zetasizer Nano- ZS...... 61

Figure 3.4: The Minispec NMR Analyser ...... 63

Figure 3.5: AFM Dimension 3100 ...... 64

Figure 3.6: Schematic illustration of RBCs preparation ...... 65

Figure 3.7:Optical microscopy image of counting chamber square……………………… .66

Figure 4.1: TEM of γ-Fe2O3 nanoparticles cast onto grid from aqueous suspension ...... 72

Figure 4.2: (a) Magnetization vs applied magnetic field for the γ-Fe2O3 nanoparticles and (b) magnetization vs applied field behavior close to zero field ...... 73

Figure 4.3: (a) Transmission electron micrograph of γ-Fe2O3 nanoparticles embedded in resin and (b) iron map corresponding to region imaged in (a) generated from electron energy loss imaging ...... 74

Figure 4.4: (a) TEM of unstained human red blood cell after incubation with γ-Fe2O3 nanoparticles at 200 mOsm and (b) iron map corresponding to region imaged in (a) generated from electron energy loss imaging ...... 74

Figure 4.5: (a) TEM of unstained control human red blood cell after incubation at 200 mOsm with no γ-Fe2O3 nanoparticles and (b) iron map corresponding to region imaged in(a) generated from electron energy loss imaging ...... 75

Figure 4.6: (a) TEM of unstained human red blood cell after incubation with γ-Fe2O3 nanoparticles at 290 mOsm and (b) iron map corresponding to region imaged in (a) generated from electron energy loss imaging ...... 75

Figure 4.8: Distribution of particle/aggregate sizes inside (black) and outside (gray) cells after incubation with γ-Fe2O3 nanoparticles at the indicated osmolarity ...... 78

Figure 5.1: Polymer binding to the surface of Fe3O4 ...... 83

Figure 5.2a: Hypo-Osmolar incubation procedure ...... 87

Figure 5.2b: Osmotic pulse procedure ...... 88

Figure 5.3: RBCs suspension on a hemocytometer ...... 89

List of Figures

Figure 5.4: TEM images of the four different magnetic iron oxide nanoparticles systems . 92

Figure 5.5: Dynamic Light Scattering (DLS) of the four different magnetic iron oxide nanoparticle systems ...... 95 Figure 5.6a: RBC specific longitudinal proton relaxivity (r1) for RBCs exposed to the four different magnetic iron oxide nanoparticle systems under osmotic pulse and hypo- osmolar incubation conditions at 200mOsm together with values for RBCs submitted to the same conditions without exposure to iron oxide nanoparticles (Control RBCs)……... 96

Figure 5.6b: Ratios of RBC specific longitudinal relaxivity for treated cells to RBC specific longitudinal relaxivity for native cells ...... 97

Figure 5.7a: RBC specific transverse proton relaxivity (r2) for RBCs exposed to the four different magnetic iron oxide nanoparticle systems under osmotic pulse and hypo- osmolar incubation conditions at 200 mOsm together with values for RBCs submitted to the same conditions without exposure to iron oxide nanoparticles (Control RBCs)…….. 98

Figure 5.7b: Ratios of RBC specific transverse relaxivity for treated cells to RBC specific transverse relaxivity for native cells ...... 99

Figure 5.8: Transverse to longitudinal proton relaxation rate ratios (R2/R1) of RBCs loaded with the four different magnetic iron oxide nanoparticle systems under osmotic pulse and hypo-osmolar incubation conditions at 200mOsm together with the R2/R1 ratios for the stock nanoparticle suspensions...... 100

Figure 5.9: Apparent loading efficiencies calculated from the transverse and longitudinal proton relaxation rates (R2 and R1) of RBCs loaded with the four different magnetic iron oxide nanoparticles systems under osmotic pulse and osmotic incubation conditions ...... 101

Figure 5.10: TEM of unstained human red blood cell after incubation with γ-Fe2O3 nanoparticles by osmotic incubation method at 200 mOsm and (b) iron map corresponding to region imaged in (a) generated from electron energy loss imaging ...... 102

Figure 5.11: TEM of unstained human red blood cell after incubation by osmotic pulse method with γ-Fe2O3 nanoparticles at and (b) iron map corresponding to region imaged in (a) generated from electron energy loss imagingom electron energy loss imaging ...... 102

Figure 5.12: (a) TEM of unstained control human red blood cell after incubation at 200 mOsm with no Fe2O3 nanoparticles and (b) iron map corresponding to region imaged in (a) generated from electron energy loss imagingom electron energy loss imaging...... 103

Figure 5.13: TEM of unstained human red blood cell after incubation with diblock PEO-PAA polymer coate system nanoparticles by osmotic pulse incubation method (b) iron map corresponding to region imaged in (a) generated from electron energy loss imaging …...... 104

Figure 5.14: TEM of unstained human red blood cell after incubation with diblock PEO-PAA polymer coated system nanoparticles at 200 mOsm and (b) iron map corresponding to region imaged in (a) generated from electron energy loss imaging ...... 104

List of Figures

Figure 15.15: TEM of unstained human red blood cell after incubation with PEO polymer coated system nanoparticles by osmotic incubation method at 200 mOsm and (b) iron mapcorresponding to region imaged in (a) generated from electron energy loss imaging ...... 105

Figure 5.16: TEM of unstained human red blood cell after incubation with PEO polymer coated system nanoparticles by osmotic pulse incubation method (b) iron map corresponding to region imaged in (a) generated from electron energy loss imaging ...... 105

Figure 5.17: TEM of unstained human red blood cell after incubation with Triblock polymer coated system nanoparticles by osmotic pulse incubation method(b) iron map corresponding to region imaged in (a) generated from electron energy loss imaging ...... 106

Figure 5.18: TEM of unstained human red blood cell after incubation with Triblock polymer coated system nanoparticles by osmotic incubation method at 200 mOsm. (b) iron map corresponding to region imaged in (a) generated from electron energy loss imaging ...... 107

Figure 5.19: The size distribution of the four different magnetic iron oxide nanoparticle systems loaded into RBCs...... 108,109

Figure 5.20: Median projected cell area (horizontal lines in middle of boxes) with interquartile ranges (boxes) and maximum and minimum values (whiskers) for native RBCs and RBCs prepared by osmotic incubation and osmotic pulse methods ...... 111

Figure 6.1a: RBC-specific longitudinal proton relaxivities (r1) for RBCs incubated at 200 mOsm in the presence of gadoteric acid (Gd-RBC) and in the absence of gadoteric acid (Con-RBC) and for native RBCs. The data for the Gd-RBC were acquired at different incubation times ranging from 1 to 18 hours but no correlation between incubation time and relaxivity was found ...... 120

Figure 6.1b: RBC-specific transverse proton relaxivities (r2) for RBCs incubated at 200 mOsm in the presence of gadoteric acid (Gd-RBC) and in the absence of gadoteric acid (Con-RBC) and for native RBCs. The data for the Gd-RBC were acquired at different incubation times ranging from 1 to 18 hours but no correlation between incubation time and relaxivity was found ...... 121

Figure 6.2a: RBC-specific longitudinal proton relaxivity (r1) for RBCs exposed to gadoteric acid under osmotic pulse and hypo-osmolar incubation conditions (RBCs loaded with gadolinium) together with values for RBCs submitted to the same conditions without exposure to gadoteric acid (Con RBCs). Bars indicate the values obtained for a single preparation ...... 122

Figure 6.2b: RBC-specific transverse proton relaxivity (r2) for RBCs exposed to gadoteric acid under osmotic pulse and hypo-osmolar incubation conditions (RBCs loaded with gadolinium) together with values for RBCs submitted to the same conditions without exposure to gadoteric acid (Con RBCs). Bars indicate the values obtained for a single preparation ...... 122

List of Figures

Figure 6.3a: Ratios of RBC-specific longitudinal relaxivity for treated cells to RBC- specific longitudinal relaxivity for native cells. Bars indicate the values obtained for a single preparation ...... 123

Figure 6.3b: Ratios of RBC-specific transverse relaxivity for treated cells to RBC- specific transverse relaxivity for native cells. Bars indicate the values obtained for a single preparation ...... 124

Figure 6.4:Transverse to longitudinal proton relaxation rate ratios (R2/R1) of RBCs loaded with gadoteric acid under different conditions together with the R2/R1 ratio for gadoteric acid solution. Bars indicate the values obtained for a single preparation ...... 125

Figure 6.5: Rates of survival of RBCs after different processing protocols. Bars indicate the values obtained for a single preparation ...... 125

Figure 6.6: Median projected cell areas (horizontal lines in middle of boxes) with interquartile ranges (boxes) and maximum and minimum values (whiskers) for native RBCs and RBCs prepared by the osmotic incubation and osmotic pulse methods...... 126

Figure 6.7: (a) TEM image of unstained human RBC after incubation with gadoteric acid at 200 mOsm and (b) gadolinium jump ratio image corresponding to the region imaged in (a) generated from electron energy loss imaging ...... 127

Figure 6.8: (a) TEM image of unstained control human RBC after incubation at 200 mOsm with no gadoteric acid. (b) Jump ratio image corresponding to the region imaged in (a) generated from electron energy loss imaging ...... 127

Figure 6.9: Magnitudes of the gadolinium jump ratio image intensity differences between cells and resin background ...... 129

Figure 7.1: Standard tip mounted on a cantilever ...... 135

Figure 7.2: The basic principle of an AFM ...... 135

Figure 7.3: Phasing mode imaging...... 137

Figure 7.4: Interleave mode scanning , the scanner performs a main trace and retrace, then lifts the tip to the lift scan height and performs the interleave trace and retrace ...... 137

Figure 7.5: The MFM interleave scan and the magnetic of the sample surface…………..138

Figure 7.6: MFM scanning with magnetized tip only ...... 140

Figure 7.7: MFM scanning with a magnet under the samples ...... 141

Figure 7.8: 7.8a A non-contact AFM image and 7.8b,7.8 c ,7.8d, 7.8e and 7.8f the corresponding phase images of -Fe2O3 nanoparticles generated by the magnetized tip at different lift height. Field of view 5 μm × 5 μm ...... 142

List of Figures

Figure 7.9: 7.9a a non-contact AFM image and 7.9b,7.9c ,7.9d and 7.9e the corresponding phase images of -Fe2O3 nanoparticles generated by a permanent magnet installed underthe sample at different lift height. Field of view 10 μm × 10 ...... 144,145

Figure 7.10: 7.10a a non-contact AFM image of RBC section loaded with -Fe2O3 nanoparticles and 7.10b,7.10 c ,7.10d , 7.10e,7.10f ,7.10g.7.10h.7.10j and 7.10i the corresponding phaseimages of the RBC section loaded with -Fe2O3 nanoparticles generated by a permanent magnet installed under the sample at different lift height: 30,50,100,150,200 and 250 nm respectively . Field of view 7 μm × 7 μm ...... 146,147,148

Figure 7.11: RBCs shape versus osmotic pressure ...... 149

Figure 7.124a: Height Image of a RBC in liquid ...... 150

Figure 7.124b: Height Image of a RBC in liquid ...... 150

Figure 8.1: A schematic of the proposed loading mechanism ...... 163

List of abbreviations

List of Abbreviations

AFM :Atomic force microscopy DMSO: Dimethyl sulfoxide EELS : Electron energy loss spectroscopy EFTEM : energy filter transmission electron microscopy Hb : Hemoglobin ICP-MS : Inductivity coupled plasma mass spectroscopy MFM: Magnetic force microscopy mOsm : milliosmoles MRI : Magnetic resonance imaging MRS : Magnetic resonance spectroscopy NMR : Nuclear magnetic resonance RBC : Red blood cell RES : Reticulo-endothelial system SQUID : Superconducting quantum interference device TEM : Transmission electron microscopy WBC : White blood cell Mh : Maghemite Mt: Magnetic

Statement

Statement of candidate contribution

Chapter Title Contribution

M.Ibrahim 70 % Loading Erythrocytes with Maghemite L.Wee 2 % Nanoparticles via Osmotic Pressure Induced M.Saunders 8 % 4 Cell Membrane Pores R. Woodward 5 % T. G. St. Pierre 15 %

M.Ibrahim 70 % L.Wee 1 % M.Saunders 4 % Loading Erythrocytes with Magnetic R. Woodward 4 % 5 Nanoparticles: a Comparison of Osmotic Pulse M. House 5 % and Hypo-Osmolar Incubation Methods J. Murphy 2 % T. G. St. Pierre 6 % T. Becker 1 % J. Riffle 2 % S. Balasubramaniam 5 %

M.Ibrahim 70 % L.Wee 1 % 6 Enhancement of the Cell Specific Proton M.Saunders 4 % Relaxivities of Human Red R. Woodward 4 % Blood Cells via Loading with Gadoteric Acid M. House 5 % J. Murphy 2 % T. G. St. Pierre 14%

7 Atomic Force Microscopy and Magnetic Force Microscopy scanting of iron oxide M.Ibrahim 80 % nanoparticles and Red blood cells loaded with T. Becker 20 % iron oxide nanoparticles

Acknowledgements

This thesis would not have been possible without the guidance and the help of several individuals who in one way or another contributed and extended their valuable assistance in the preparation and completion of this thesis.

First I would like to express my deepest gratitude to my supervisor, Prof. Tim G. St. Pierre, for his excellent guidance, great efforts to explain things clearly and simply, patience, and providing me with an excellent atmosphere for doing research.

I would like to thank my Co-supervisor, Dr. Leonard Wee, who was abundantly helpful and offered invaluable assistance, support and guidance.

I thank Dr. Robert C. Woodward who has always been available for excellent advice and guidance on several aspects of the project. I would also like to thank Assoc/Prof. Michael House for his many contributions and advice, especially with regard to proton relaxometry measurements. I also thank Sharavanan Balasubramaniam for preparation of the synthetic polymers and their linking to the magnetic nanoparticles used in Chapter 5.

Special thanks to Dr. Ralph James for many useful discussions and suggestions.

I would like express my appreciation to Dr. Thomas Becker of the Nanochemistry Research Institute, Curtin University of Technology who helped with the atomic force microscopy and gave me the opportunity to work in his lab.

Thank you to Prof. Martin Saunders of the Centre for Microscopy, Characterisation and Analysis for your guidance and the help with the transmission electron microscopy.

Finally I would like to thank all other members of the Biomagnetics research group at UWA. It has been a pleasure to work with you all.

Chapter 1 General Introduction

1.1 Magnetic Resonance Imaging development history

The history of magnetic resonance imaging (MRI) can be traced to 1944 when Isidor Isaac Rabi won the Nobel Prize in Physics for his development of the resonance method for recording the magnetic properties of atomic nuclei. Rabi discovered that the spins of atomic nuclei could be studied by the combination of radio waves and a magnetic field [1- 4]. In 1952, Edward Purcell and Felix Bloch received the Nobel Prize in Physics for their independent discoveries of nuclear magnetic resonance (NMR) [5] [6]. In 1970, the theory of nuclear magnetic resonance was used by Paul Lauterbur, a chemistry professor at the State University of New York and Peter Mansfield, a physics professor at the University of Nottingham, to develop a new diagnostic technique called magnetic resonance imaging. In 2003, Lauterbur and Mansfield were awarded the Nobel Prize in Physiology or Medicine for their discoveries concerning MRI. Raymond Damadian has since disputed with Nobel Prize organizers over the invention of MRI claiming that he has the right to be mentioned in the Nobel prize [7]. In 1980, Damadian built the first commercial MRI by using his patented voxel by voxel method of imaging [8, 9]. He put together the first MRI scanner with the assistance of two postgraduate students, Michael Goldsmith and Larry Minkoff [10, 11].

In the early 1990s Seiji Ogawa found that MRI properties change when blood oxygen levels change. Ogawa realized he could use this contrast in blood oxygen level response to map images of brain activity during an MRI scan [12, 13]. This method is called functional magnetic resonance imaging (fMRI) and is now widely used in biology, neurobiology, psychology, neurology and other branches of research. It is also used to diagnose the physiological basis of mental illnesses and organic brain dysfunction in clinical medicine. Although Ogawa’s discovery was developed in the 1990’s, the the underlying biophysics relating the magnetic state of haemoglobin to its oxygenation state had been mentioned in 1930 by Linus Pauling [14, 15]

Chapter 1 General Introduction

1.2 Magnetic Resonance Imaging (Introduction)

Nuclear magnetic resonance (NMR) has become a well-established technique for analysis in physics and chemistry. Increasingly, there are applications of NMR-based procedures in medicine and biology such as MRI and Magnetic Resonance Spectroscopy (MRS). These medical applications exploit the different relaxation rates or chemical shifts of the nuclear magnetic dipole of specific atoms within organic molecules. MRI, now an indispensable tool for clinical diagnosis, reveals fine details of anatomy without using any ionizing radiation or radioactive tracers. MRI works with radio waves and is a painless and harmless way of looking inside the human body [16].

MRI uses the combination of a strong magnetic field and radio frequency excitation to generate three-dimensional images [17, 18]. Most of the clinical MRI studies performed today are essentially studies of anatomy, but MRI is such a flexible technique that it is possible to measure the physiological functioning of tissue [19-21]. In 1970, Raymond Damadian showed that the nuclear magnetic relaxation times of tissues and tumours differed, thus motivating scientists to consider magnetic resonance for the detection of disease [22]. Cancerous tissue emits response signals that last much longer than non- cancerous tissue. He would subsequently find that the response times of other kinds of diseased tissue, normally called "relaxation times", also vary dramatically.

There are generally two kinds of relaxation times that can be detected: they are known as longitudinal relaxation time and transverse relaxation time. The relaxation rate has been shown in numerous experiments to correlate strongly with the presence of materials with large magnetic moments. St Pierre et al. (2004) [23] showed that a single spin echo sequence is useful to evaluate iron overload in tissue. St Pierre and coworkers (2005) showed that transverse relaxation rate (R2) measurements by magnetic resonance imaging could be used to measure iron concentration in tissue [24].

The human body is composed primarily of diamagnetic materials which means that whole body imaging is possible [25]. MRI is based on imaging mostly water and to a lesser extent, fat, which is abundant in physiological tissue. Water molecules are composed of 19

Chapter 1 General Introduction hydrogen and oxygen. The hydrogen atom consists in turn of a negatively charged electron and a nucleus, which contains a single positively charged proton. The MRI signal is generated from this nucleus [17].

The proton behaves like a spinning magnet with a north and south pole which is referred to as a “spin”. Proton spins are randomly oriented in the absence of an external magnetic field

[26]. In the presence of an external magnetic field (B0) the spins align with (parallel) or against (antiparallel) the direction of the magnetic field (Figure 1.1). Overall, spins tend to align with a magnetic field in very small number, providing a small net magnetisation which is the basis of the MRI signal [27]. The net difference between the parallel and the antiparallel spins represents the spins that will create a detectable magnetic resonance signal [10, 11]. The most visually dominant component in an MRI scanner is the imaging magnet, which produces the B0 field [28]. The imaging magnet is an electromagnet made of several kilometres of superconducting wire formed into a coil. MRI takes advantage of the natural magnetic properties of human body tissue. A group of contrast media called MRI contrast agents is commonly used to enhance the visibility of internal body structure in MRI (Figure 1.2).

Figure 1.1: The application of an external magnetic field, B0, causes the magnetic moments of the protons to precess about the field direction.

Chapter 1 General Introduction

Figure 1.2: An MRI image showing a defect in the blood-brain barrier after a stroke. T1weighted images, left image without contrast medium - right image with contrast medium administered [28].

1.2.1 MRI scanner

The MRI scanner consists of a strong magnetic field (B0) in the longitudinal direction which is produced by a superconducting electromagnet. After activation, superconducting electromagnets require no current supply. For imaging, the more uniform the static field B0 is, the less the resonance frequency varies with position.

Figure 1.3a: Magnetic Resonance Imaging scanner cutaway [29]

Chapter 1 General Introduction

The most common field strength used for clinical scanners is 1.5 T although 3T and higher field are becoming available. Preclinical scanners can have fields over 10 T [30]. Gradient coils produce a spatial variation of the longitudinal magnetic field with respect to position along three orthogonal axes. A second coil arrangement produces a field transverse to the static field (B1) and perpendicular to the B0 field, which sends radio waves through the body. This coil is tuned to the resonant frequency (RF) of magnetization. The

B1 field is used to excite the magnetisation from its longitudinal orientation along B0 to a transverse orientation. According to Faraday’s law of induction, changing of the flux in the coils induces a voltage in the transverse coils. The received signal is in the same frequency band as the excitation pulses and often the same coil or coils used to generate the B1 field can be used to receive the signal in the transverse direction [31]. In MRI, the signal is just a current or voltage that is induced by an oscillating magnetic field. These signals are collected to create three-dimensional images of what is inside the body [32, 33]. The main cost of the MRI scanner is the coil of wire. To keep the coil of wire superconducting and with zero resistance, it has to be very cold. Liquid helium is used to cool the coil wire [31] which is four degrees above absolute zero. Usually an MRI scanner contains several thousand litres of liquid helium.

1.3 Role of contrast agents

Figure1.3b: Cardiac angiograph [34]

Angiography is a medical imaging technique used to visualize the inside, or lumen, of blood vessels and organs of the body. Angiography is also commonly performed to identify vessel narrowing in patients. Contrast agent is added to the blood to make it visible on images. For example this technique is used to look at the left side of the heart and at the 22

Chapter 1 General Introduction arterial system; or the jugular or femoral vein, to look at the right side of the heart and at the venous system.

Contrast agents are widely used in clinical practice with about 30% of MRI scans being contrast enhanced [35]. In clinical MRI studies, contrast agents are often used to enhance the contrast between different tissues by altering the local proton relaxation times [36]. In other words, the contrast between different tissues (healthy against diseased) can be increased. Usually, when contrast agents are injected intravascularly to establish the difference between normal and abnormal tissue there is a difference in the uptake of the contrast media in the various tissues. Therefore, the more the uptake of contrast media in a tissue, the more the tissue details will appear. This process is called enhancement. Paramagnetic gadolinium chelates and superparamagnetic nanoparticles have been developed for use as MRI contrast agents [37].

Recent developments in this area include the use of contrast agents with large magnetic moments. In this study, I used superparamagnetic nanoparticles and gadoteric acid as potential contrast agents. In addition, nanoparticles of Fe2O3 and Fe3O4 contrast agents can exhibit a property known as superparamagnetism. This occurs at room temperature when they are below a certain size, which is determined by the magneto-crystalline anisotropy of the particles. Gadoteric acid (C16H25GdN4O8, Dotarem, Guerbert, France) contrast agent was used because of its paramagnetic properties. The contrast agent magnetic properties have the advantage of high magnetic susceptibilities which result in large proton relaxivities when the particles are suspended in aqueous media. If such materials can be developed with sufficient specificity, they are potentially useful for highlighting structures that are otherwise hard to detect, such as small tumours, cardiac failure and brain functions.

1.4 Limitations and side effects of current agents

There are several limitations in the use of current contrast agents such as side effects, toxicity, and macrophage clearance. The side effects of the contrast agents can be reduced by injecting them at minimum dosage to enhance the contrast between different tissues. In addition, the optimal contrast agent should create enhanced resolution at minimum dosage 23

Chapter 1 General Introduction so that unwanted side effects are minimized or avoided. The contrast medium should be excreted from the body without any interference with endogenous species so that it does not accumulate in the body.

Chelated-gadolinium compounds are rapidly cleared from the blood stream via the kidney with an elimination half-life of approximately 2 hours [38, 39]. The chelated gadolinium contrast medium has been reported safe compared with the iodinated contrast medium [40, 41]. However, there are case reports of acute renal failure after gadolinium contrast medium has been administered to patients at a dose of 0.1- 0.44 mmol/kg [42]. Sam et al. confirmed in 2003 that gadolinium injected into a patient should be less than 0.44 mmol/kg. They also reported that gadolinium contrast medium is 20 times safer than iodinated agents [43]. There is no record of aggressive peritoneal dialysis prescriptions despite the very poor peritoneal dialysis clearance of gadolinium [44].

Magnetic particles are rapidly cleared from the blood stream via the reticulo-endothelial system (RES). Large nanoparticles above 200nm are quickly eliminated by the liver. To reduce this problem, nanoparticles are coated with a hydrophilic layer when their size is below 200nm [45]. The kidneys functions to eliminate small nanoparticles sizes especially the particle size ranging from 10-20nm [46].

Contrast agents still require development in a number of areas. First, non-toxic nanoparticles must be manufactured with precisely controlled sizes and chemical composition. Second, any interaction between the nanoparticles with healthy cells and tumour cells must be understood. A still poorly understood area is the toxicology of superparamagnetic particles when introduced into the human body. Third, in order to be a potential marker for tumours, the nanoparticles should have an affinity towards tumour cells. An understanding of how these nanoparticles can bind to tumour cells will be necessary to enable the effective design and synthesis of suitable materials.

1.5 Human blood function and red blood cells (RBCs)

1.5.1 Blood function 24

Chapter 1 General Introduction

Blood is a complex liquid that performs a number of critical functions. It transports:  oxygen from the lungs to all cells of the body;  carbon dioxide from cells to the lungs;  nutrients from the digestive organs to cells;  waste products from cells to the kidneys, lungs and sweat glands; and  enzymes to various cells. Blood also regulates body pH through buffer and amino acids. It plays a role in the regulation of normal body temperature because it contains a large volume of water (an efficient heat absorber and coolant). It regulates the water content of cells, mainly through dissolved sodium ions. It prevents body fluid loss through the clotting mechanism. Special combat-unit cells in blood protect the body from toxins and foreign microbes. Blood leaves the heart via the arteries that branch repeatedly until they become capillaries. Oxygen (O2) and nutrients diffuse across capillary walls and enter tissues. Carbon dioxide (CO2) and waste products move from tissues into the blood. Oxygen-deficient blood leaves the capillaries and flows through veins to the heart. The blood flows to the lungs where it releases CO2 and picks up O2. The oxygen-rich blood then returns to the heart.

1.5.2 Red blood cells

(a) (b)

Figure 1.4: a and b Atomic Force Microscopy (AFM) images of RBCs.

Blood is the body’s only fluid tissue. It is composed of liquid plasma and formed elements that include erythrocytes (red blood cells - RBCs), leukocytes (white blood cells), and thrombocytes (platelets). RBCs are biconcave discs about 7.7 µm in diameter. They have

Chapter 1 General Introduction a surface area of approximately 140 m2, a volume of approximately 100 m3[47] and are without nuclei (Figure 1.4 a and b). RBCs lack a nucleus therefore they can neither reproduce nor carry out extensive metabolic activities. The plasma membrane is selectively permeable and consists of proteins and lipids. The membrane encloses cytoplasm and a red pigment called haemoglobin. Haemoglobin molecules (Figure 1.5) constitute about 33 percent of RBC weight and are responsible for the red colour of blood. The haemoglobin molecule consists of a protein called globin and pigment called haem or heme, each haem containing one iron atom. The overall structure of haemoglobin comprises 4 globins and 4 haem groups [48-50]. Normal blood concentrations of haemoglobin are 14 to 20 mg/mL in infants, 12 to 15 mg/mL in adult females and 14 to 16.5 mg/mL in adult males.

Figure 1.5: The structure of haemoglobin: adult haemoglobin is a tetrameric hemeprotein [α(2):β(2)] found in erythrocytes [51].

One RBC contains approximately 280 million molecules of haemoglobin and is dedicated to respiratory gas transport. The biconcave shape of an RBC gives a much greater surface area than a sphere or cube, therefore providing a relatively large surface area for the diffusion of gas molecules to pass through its membrane where it combines with haemoglobin. Haemoglobin reversibly binds with oxygen and most oxygen in the blood is bound to haemoglobin. Hemoglobin is composed of the protein globin, made up of two alpha and two beta chains, each bound to a heme group. Each heme group bears an atom

Chapter 1 General Introduction of iron, which can bind to one oxygen molecule. Each hemoglobin molecule can transport four molecules of oxygen.

The cell membrane of an RBC becomes fragile and the cell is non-functional in about 120 days. A healthy male has about 5.4 million RBCs per cubic millimetre (mm3) of blood, and a healthy female has about 4.8 million. RBCs do not divide but are renewed by cells in bone marrow. In adults the production of RBCs takes place in the red bone marrow in the spongy bone of the cranium, ribs, sternum, the bodies of vertebrae, and the proximal epiphyses of the humerus and femur [48-50].

1.6 The concept of using red blood cells as a “Trojan Horse” for contrast agents

Chelated gadolinium compounds and magnetic nanoparticles have been developed for use as MRI contrast agents. However, both of these are rapidly cleared from the blood stream: gadolinium via renal excretion and magnetic nanoparticles by the reticulo-endothelial system (RES). In order to extend the lifetime of the contrast agent in the blood stream, attempts have been made to encapsulate contrast agents within RBCs so that their clearance is slowed. A key strategy for loading RBCs with contrast agents is to incubate them in the presence of the contrast agent under hypoosmolar conditions, inducing pores in the cell membrane. The RBCs can be re-sealed by bringing the osmolarity of the medium back up to physiological values [52]. It is essential to control the osmotic pressure during this loading process.

1.6.1 Equation for manipulating osmolarity (diffusion and osmosis)

A difference in concentration leads to a flow of matter from a region of high concentration to a region of low concentration. This movement or flux (J) is given by Fick’s law of diffusion. Osmosis is a special case of diffusion. Variations in the concentrations of dissolved ion species will produce diffusion as the system moves towards equilibrium. This can occur either through bulk liquid or a permeable barrier. If a semi-permeable barrier is used allowing only the water to pass through, osmosis occurs as the system tends to dilute the trapped species by a net flow of water into the high concentration region. 27

Chapter 1 General Introduction

There are five recognized mechanisms for transport across cell membranes: solubility- diffusion, pore-diffusion, coupled carrier, active transport, exocytosis and endocytosis. Osmosis is a special case of diffusion. Here the solvent diffusion is a driven process between two phases separated by a semi-permeable membrane which allows water, but not large molecules, to pass through. A difference in concentration of solute cannot be achieved by the transport only of these molecules. We must return to the thermodynamic basis to define the osmotic pressure equation. The van 't Hoff formula of osmotic pressure, which is identical to the formula of ideal gas pressure, is a direct outcome of the second law of thermodynamics. The formula is derived by applying a closed cycle (reversible and isothermal).

First law of thermodynamics The conservation of energy in an isolated system is expressed thermodynamically as:

Where ΔU: is the change in internal energy of the system. ΔQ: the net heat into the system. ΔW is the work performed by the system on its surroundings

Second law of thermodynamics For the case of ideal reversible heat flow at constant temperature T, the corresponding change in the local entropy S given by:

Combining (1.1) and (1.2)

Chapter 1 General Introduction

Gibbs Free Energy Gibbs Free Energy (G) is defined in terms of the other thermodynamic functions pressure (P) and temperature (T):

(1.4)

At constant T and P (1.5) ΔW is the work performed by the system on its surroundings equal to the mechanical work

(Wm =PΔV) plus Electrical (Wel) and Chemical Potential (Wc)

Where:

Wm is the mechanical work performed by the system at pressure P with change in volume ΔV.

Wel is the work need to move a charge Δq (coulomb) across an electrical potential (joule per coulomb).

Wc is the work need to move Δn moles of matter across a chemical potential µ (Joule per mole). Combining (1.6) and (1.10) Experimentally μ is found to be related to concentration (c) according to:

Where R is the Universal Gas Constant.

Chapter 1 General Introduction

Figure 1.6 Chemical potential (µ) vers concentration (C)[53]. Integrating (1.12) gives

Where: μo is a constant at standard T and P. The difference in chemical potential between two regions 1 and 2 can then be written:

Where c1 is the initial molarity of solution and c2 is the final molarity of solution after equilibration.

Figure 1.7: The diffusion of liquid between zone 1 and zone 2 via a semi-permeable membrane.

Chapter 1 General Introduction

When the system is not at constant T and P we must use a more general expression:

̅ ̅ where ̅ is the partial molar volume and ̅ the partial molar entropy. At constant T and P we cannot achieve equilibrium (Δμ = 0). If we only have T held constant then at equilibrium there will be a pressure difference across the membrane.

̅ This pressure difference is called the osmotic pressure. This pressure will act to drive water (the permeable species) through the membrane in a sense acting to reduce the concentration by diluting the higher side. An approximation useful in very dilute situations comes from the linearity of the ln function near x = 1. [53]

This approximation is referred to as van‘t Hoff's Formula.

1.6.2 Manipulating red blood cells with osmotic pressure.

All biological membranes have a property known as semi-permeability which is the ability to allow water and other small molecules to pass through them. Semi-permeable membranes do not to allow the passage of large molecules that are present in water. Since water can flow freely in both directions, but large molecules cannot, water passes slowly from a weak solution on one side of a semi-permeable membrane to a more concentrated solution on the other. The more concentrated solution can be said to draw the water with a less concentrated solution. This process is called osmosis and is driven by osmotic pressure. The concentration of particles (osmoles) per unit weight (kg) of solvent is known as the osmolarity of the solution. In medicine it is more commonly expressed as milliosmoles per kilogram of water (mOsm / kg H2O). Blood in the human body is approximately 290 mOsm/kg an osmolarity referred to as an isotonic solution. RBCs have the same solute (or particle) concentration as isotonic solution as they gain and lose water. This creates equilibrium. A solution is also called hypotonic when it contains less solute than the RBC. Here, the solvent rushes into the RBC. Solution is hypertonic when it

Chapter 1 General Introduction contains solute at a higher concentration than within the RBC. Hypotonic and hypertonic media cause changes in the size and shape of RBCs (see figure below) [54].

Figure 1.8: Red blood cells at different osmotic pressures. 1.6.3 Loading red blood cells: pore size and membrane reversibility

In this thesis my key strategy for loading RBCs with contrast agent is to incubate them in the presence of the contrast agent under hypo-osmolar conditions. This involves inducing pores in the RBC membrane by varying the osmotic pressure. The RBCs can be re-sealed by bringing the osmolarity of the medium back up to physiological values [52, 55-58].

1.6.4 Previous methods of drug loading into red blood cells

Several methods have been used to incorporate drugs into RBCs. Most of them are osmosis-based systems. I will list below the methods most frequently used.

 Hypotonic hemolysis

RBCs are incubated in hypotonic solution in the presence of drugs. This causes RBCs to swell and pores to open in the cell membrane. The cells can maintain their integrity down to a tonicity of 150mOsm/Kg and below this point the cells lyse. The RBCs are re-sealed by bringing the osmolarity of the medium back up to physiological values - 290-300 mOsm/kg [52, 59, 60].

 Hypotonic dilution

Chapter 1 General Introduction

Packed RBCs are diluted with a 2-20 volume of an aqueous solution of drugs. The solution tonicity is then controlled by adding a hypertonic buffer. The final suspension is centrifuged and then incubated in an isotonic buffer [61-63].

 Red Cell Loader

In 1998 Magnani et al. designed a machine called a Red Cell Loader with a capacity of 50mL of solution (cells in suspension). This machine’s process is based on two hypotonic dilutions of washed RBCs followed by an isotonic re-seal of the cells [64].

 Hypotonic dialysis

RBCs in isotonic solution are placed in a conventional dialysis tube immersed in a 10-20 volume of a hypotonic buffer. The drugs are added to the isotonic buffer inside the dialysis bag. The tonicity of the dialysis tube is restored by directly adding a calculated amount of a hypotonic solution to the surrounding medium or by replacing it with a large volume of isotonic buffer [65, 66].

 Hypotonic preswelling

This method was developed for drug loading by Rechstiner in 1975 [67]. It is based on the principle of controlled RBCs swelling in a hypotonic buffer solution by creating centrifugal force at low g. This is followed by the addition of drugs until the lysis point of the RBCs is detected and then the mixture is centrifuged. The RBCs are resealed by adding a calculated amount of hypertonic solution. The cell suspension is incubated at 37C [61, 67-69] .

 Isotonic lysis

This method is also called osmotic pulse. Here we obtain hypotonic RBC buffer by incubating RBCs in a solution with high membrane permeability such as dimethyl sulfoxide (DMSO). This is followed by adding drugs which cause a very high flow of the

Chapter 1 General Introduction solution into the RBCs, maintaining equilibrium. After adding an isotonic buffer solution at 37C and centrifugation, the RBCs re-seal. [57, 70, 71].

 Chemical perturbation of the membrane

RBCs are exposed to certain chemicals to increase the membrane’s permeability. This method was first mentioned by Deuticke et al. in 1973 [72-74].

 Electroencapsulation

This method was first mentioned in 1973 by Zimmerman. Here electrical pulse is used to incorporate molecules into the RBCs [75].

 Entrapment by endocytosis

This method was reported by Schrier et al. in 1975 [76]. Endocytosis involves the addition of one volume of washed and packed erythrocytes, to nine volumes of buffer containing 2.5 mM ATP, 2.5 mM MgCl2, and1mM CaCl2, followed by incubation for 2 minutes s at room temperature [77, 78].

 Loading by electric cell fusion

This method involves the initial loading of drug molecules into erythrocyte ghosts followed by adhesion of these cells to target cells. The fusion is accentuated by the application of an electric pulse, which causes the release of an entrapped molecule [79, 80].

1.7 Aims of this study

The principle aim of this study is to demonstrate that iron oxide nanoparticles and gadoteric acid can be loaded into RBCs in a concentration sufficient to obtain a strong enhancement of signal in MRI. Encapsulating magnetic nanoparticles within red blood cells is one strategy for extending the lifetime of MRI contrast agents in the bloodstream. Both the 34

Chapter 1 General Introduction gadolinium chelates and magnetic particles are rapidly cleared from the blood stream: gadolinium via renal excretion and magnetic nanoparticles by the reticulo-endothelial system (RES). In order to extend the lifetime of magnetic particles in the blood stream, attempts have been made to encapsulate magnetic particles within RBCs so that macrophage clearance is slowed. A convenient method for loading RBCs with magnetic nanoparticles involves inducing pores in the RBC membrane by varying the osmotic pressure [52, 57].

This study has two broad directions relating to the loading and uptake of superparamagnetic iron oxide nanoparticles and paramagnetic gadolinium chelates by RBCs. In the first part of the study, I introduce size-controlled iron oxide nanoparticles produced by chemical synthesis and uncoated nanoparticles into RBCs to examine the uptake and binding of the nanoparticles. Here, we compare the effectiveness of magnetic resonance contrast enhancement by RBCs loaded with magnetic nanoparticles using two different methods. We investigate the optimal conditions for loading RBCs using these techniques and identify the limit of particle (or aggregate) size that can be incorporated through the pores by using transmission electron microscopy (TEM) and iron mapping by energy filtered transmission electron microscopy (EFTEM) [81].

The second part of the study assessed the performance of two different methods of loading RBCs with gadoteric acid. The first method involved incubating the RBCs in the presence of gadoteric acid under hypoosmolar conditions. The RBCs were subsequently re-sealed by bringing the osmolarity of the medium back up to physiological values (300 mOsm) [52]. The second method involved preparing the RBCs in a hypoosmolar solution of DMSO which diffuses into the cells, thus preventing an osmotic pressure difference across the cell membrane. The cells were then rapidly flushed with a slightly hypoosmolar solution containing the gadoteric acid to induce an osmotic pressure pulse, temporarily opening pores in the cell membrane [57]. Previous studies have investigated the enhancement of the relaxivity of gadolinium-loaded RBCs [57]. However, it has not been clear until now whether all RBCs load with gadolinium during the preparation process. Here we tested the hypothesis that only a fraction of the RBCs in a sample open their pores to enable gadolinium loading with each of the techniques. 35

Chapter 2 Theoretical Principles of Magnetic Resonance Imaging Contrast Agents

2.1 Basic magnetic resonance imaging theory

In 1946 Felix Bloch demonstrated that electromagnetic fields are created by any spinning, charged particles such as hydrogen nuclei. The MRI signal is generated from the hydrogen nucleus [17]. It is based on imaging mostly water and to a lesser extent fat, which is abundant in physiological tissue.

The spin of the hydrogen nucleus is possible at +1/2 or -1/2. In the absence of an external magnetic field (B0), the energy of these two spin states is equal. A proton rotating about its own axis generates a magnetic moment (mp) (Figure 2.1). The magnetic moment of each -27 hydrogen nucleus within the body is equal to 14.106 X10 J/T [82]. The magnetic moments of the hydrogen nuclei within the body are randomly oriented and cancel out in the absence of an external magnetic field. The sum of all these moments is zero and the bulk magnetisation is zero [26, 82].

In the presence of an external magnetic field (B0), the magnetic moments of the +1/2 spins align with the direction of the magnetic field (parallel). The magnetic moment of the -1/2 spins align against it (antiparallel) (Figure 2.2). Therefore the energy level of the +1/2 spin state is lower than the energy level of -1/2 spin state. Overall, spins tend to align with the magnetic field such that there is a small difference in number between the +1/2 spins and - 1/2 spins. This provides a net magnetic moment (µ) along the field direction which is the basis of the MRI signal [27, 82]. The net magnetic moment (µ) is given by the formula:

Where N is the number of protons, µH is the nuclear magnetic moment of the proton, E is the energy difference between the two spin states, k is Boltzmann’s constant and T is temperature.

Moreover, in the presence of an external magnetic field (B0), the proton spins of the nuclei rotate on their own axes and wobble (or precess) about the axis of the B0 (Figure 2.3). The

Chapter 2 Theoretical Principles of Magnetic Resonance Imaging Contrast Agents precession frequency of the magnetic moment of the proton spins around the external field

B0 is given by the Larmor equation [27, 82]:

where ω0 (Hz) is the angular precessional frequency of the proton, ϒH (Hz/Tesla) is the proton gyromagnetic ratio, B0 is the external magnetic field experienced by the proton (Tesla) e is the charge, m is the mass of the proton, g is a factor that includes the effects of the spin of the nucleons as well their orbital angular momentum and the coupling between the two.

The net number of the parallel (+1/2) versus the antiparallel (-1/2) spins represent the spins that will create a detectable magnetic resonance signal [10, 11]. In order to excite the low energy level spin (+1/2) to the higher energy level (-1/2) a radio frequency (RF) pulse is applied at the Larmor frequency. The RF pulse is an electromagnetic wave transmitted in a three dimensional (x,y,z) coordinate system. If an RF pulse is applied along the x axis perpendicular to B0, it creates a magnetic field - B1, which is much weaker than the external magnetic field B0. The magnetic field B1 oscillates at frequency along the x axis which causes the proton spins of the nuclei ( spins) to precess around it with a frequency ω1:

Since the frequency ω1 is proportional to the magnetic field B1 and since B0 >> B1, then

ω0 >> ω1. Here the proton spins of the nuclei ( spin) precess around B0 with frequency

ωo and around B1 with frequency ω1. This proton precessing causes a spiral motion of the net magnetization vector from the z axis into the x-y plane. Resonance occurs when the frequency of the proton precession is equal to the frequency (ωRF) of the RF pulse, i.e.

Chapter 2 Theoretical Principles of Magnetic Resonance Imaging Contrast Agents

. If the precessional frequency ( ) of the spins is not equal to the RF frequency the system will not resonate. The proton moments will not turn over into the x-y plane and no energy will be added to the protons.

To conclude, initially the protons precess about B0 along the z axis, but they are out of phase. When an RF pulse is applied, where this creates the magnetic field B1.

Here the protons also tend to precess around field B1 and will then be in phase. This creates transverse magnetization and simultaneously the B1 field causes a spiral downward motion of the proton magnetization vector as I described above.

Figure 2.1: A proton rotating about its own axis.

Figure 2.2: Proton spin states rotating on their own axis and wobbling about the axis of B0.

Chapter 2 Theoretical Principles of Magnetic Resonance Imaging Contrast Agents

2.1.1 Proton magnetic relaxation

Proton magnetic relaxation is the process of the proton spins relaxing back to their equilibrium state when the RF excitation is removed. In this case the proton spins will tend to realign with the external magnetic field direction - B0 and the net magnetic moment (µ) will go back to the original equilibrium value as given by equation (2.1). Here we have two different types of proton relaxation - longitudinal relaxation and transverse relaxation.

Longitudinal relaxation is when the relaxation is in the applied magnetic field (B0) direction. Transverse relaxation is when the relaxation is perpendicular to the applied magnetic field (B0) in the rotating frame of reference [27].

2.1.1.1 Longitudinal (or spin-lattice) relaxation time T1

T1 is called the longitudinal relaxation time because it is the time it takes the proton spins to re-align along the direction of the external magnetic field B0 after the RF pulse is turned off.

T1, also known as “spin lattice relaxation time”, refers to the time it takes the proton spins to go back to a state of equilibrium by giving back the energy they gained from the RF pulse to the surrounding lattice. When an RF pulse is applied at the Larmor frequency, it causes spin excitation from the low spin energy state (+1/2) to the higher energy spin state (-1/2). The bulk magnetization vector steadily decreases when the number of the +1/2 spins decrease. The bulk magnetization will be oriented in the opposite direction of the magnetic field when the number of -1/2 spins is higher than the number of +1/2 spins. The duration of the excitation pulse controls the direction of the bulk magnetization vector, ranging from a small angle to 180° [82]. When the applied external RF is turned off, the proton spins relax back to their equilibrium energy state and at the same time, the proton spins dephase with each other. The proton spin net magnetization along the direction of the external magnetic field B0 slowly recovers. The characteristic relaxation time it takes the net magnetization along the z axis to recover or to go back to the original equilibrium value is called T1. The change in longitudinal magnetization as a function of time (Mz (t)) after a 180° excitation pulse is shown in figure 2.3:

Chapter 2 Theoretical Principles of Magnetic Resonance Imaging Contrast Agents

Figure 2.3 : Relaxation of the longitudinal component of magnetization after the application of a radio frequency pulse where, Mz(0) is the longitudinal magnetization immediately after the excitation pulse, M0 is the magnetization at equilibrium (taken from [82]).

2.1.1.2 The transverse relaxation times T2 and T2*

* T2 is called the transverse relaxation time. T2* is the time constant that characterises the rate at which the transverse component of the magnetization vector (in the x-y plane) decays to zero. The mechanisms that cause the dephasing of spins and hence transverse magnetization can be categorized into those that are reversible (by application of appropriate field gradient pulses or RF pulses) and those that are irreversible. The contribution to the transverse relaxation time of irreversible processes is known as T2. At equilibrium, the bulk magnetization vector is oriented only along the direction of the external magnetic field B0 (z axis). When an external RF excitation pulse is applied, the bulk magnetization vector of the proton spins reorients from the z direction through an angle that generates a transverse component of magnetization vector in the x-y plane. During the application of the RF pulse the proton spins tend to start precessing in phase with each other around the direction of B0. When the RF excitation is removed, the spin moments then tend to dephase with each other owing to two mechanisms: spin-spin interactions and inhomogeneities in the externally applied magnetic field. Inhomogeneities in the externally applied magnetic field cause proton spins in different locations to precess at different frequencies because each spin is exposed to slightly different magnetic field

Chapter 2 Theoretical Principles of Magnetic Resonance Imaging Contrast Agents strength. Hence the precessing spins will dephase over time. The most common spin-spin interaction in MRI is the interaction of the magnetic moment of the proton in one hydrogen atom in a water molecule with the magnetic moment of the proton in the other hydrogen atom of the molecule. The magnetic field generated by each proton effectively perturbs the field applied externally and hence contributes to dephasing of the proton spins. The magnitude of this effect is dependent on the rate at which the water molecule tumbles since rapid tumbling can average out the perturbing fields and reduce the dephasing rate. This dephasing of the magnetic moments (spins) is observable at the macroscopic level as decay in the transverse magnetization signal, as shown in Figure 2.4.

Figure 2.4: The relaxation of the transverse component of magnetization after the application of a RF excitation pulse. Mxy (t) is the transverse magnetization at time - t after the pulse and Mxy (0) is the initial maximum magnetization immediately after the pulse (taken from [82]).

Chapter 2 Theoretical Principles of Magnetic Resonance Imaging Contrast Agents

2.1.1.3 Relationship between T2 and T2*

The transverse relaxation time T2* can be expressed in terms of a transverse relaxation rate

(R2* = 1/T2*). Similarly the transverse relaxation time T2 can be expressed in terms of a transverse relxation rate (R2 = 1/T2).

The free induction decay rate R2* is then given by:

Where is the contribution to the free induction decay rate, R2*, from reversible processes.

2.2 Contrast agent mechanisms

Molecules or nanoparticles that display high saturation magnetization and high magnetic susceptibility are of great interest in medical applications and have potential to be used as MRI contrast agents.

In MRI the two most commonly exploited contrast mechanisms are: longitudinal relaxation enhancement contrast (R1 = 1/T1) and transverse relaxation enhancement contrast (R2 =

1/T2). With contrast agents like gadolinium (Gd) chelates where the relaxation enhancement is predominantly through R1 enhancement, the contrast agent generally results in brighter or higher intensity regions. This is often referred to as positive contrast enhancement. For other contrast agents such as iron oxide superparamagnetic nanoparticles, where R2 enhancement is the dominant mechanism, the contrast agent generally results in darker or lower intensity regions. This is often referred to as negative contrast.

In the research presented in this thesis I used paramagnetic Gd chelates and superparamagnetic iron oxide nanoparticles. Therefore I describe below the phenomena of

Chapter 2 Theoretical Principles of Magnetic Resonance Imaging Contrast Agents ferromagnetism, paramagnetism and superparamagnetism for the benefit of readers unfamiliar with these concepts.

The magnetic induction (B) in free space is given by:

If we place a magnetic sample in this magnetic field it becomes magnetized. The magnetic induction inside the sample is given by:

Where µ0 is the permeability of free space, H is the externally applied magnetic field intensity and M is the sample magnetization vector.

In many situations, M is proportional to H and can be described mathematically by:

where is the magnetic susceptibility of the medium (or particle).

The magnetization (M) of a sample is equal to the magnetic moment created in the sample by the external magnetic field divided by the volume of the sample. Materials can be classified depending on their magnetic susceptibility: for example, paramagnetic materials have positive χ. In other words, the magnetization vector is parallel to the applied magnetic field (H). On the other hand, all diamagnetic material has negative χ and, as such, magnetize in a direction opposite to the applied magnetic field [83].

Chapter 2 Theoretical Principles of Magnetic Resonance Imaging Contrast Agents

2.2.1 Ferromagnetism

Only a few materials are ferromagnetic and nearly all of them are metals such as Fe, Co, Ni, Gd and Dy. The magnetic susceptibility (χ) of ferromagnetic materials is very large. Ferromagnetic materials have regions of spontaneous magnetization known as magnetic domains. The magnetic domains in ferromagnetic materials are usually randomly oriented in zero applied magnetic field. When an external magnetic field is applied, the domains are forced into alignment (Figure 2.5). In the absence of an externally applied magnetic field, this kind of material can retain some magnetization owing to incomplete randomization of magnetic domains after removal of an applied field. Usually interaction between dipole moments will result in some antiparallel alignments [53] but ferromagnetism describes a situation where the interactions between atomic dipoles are quantum mechanical in nature and are strong enough to align all of the atomic moments even in zero applied field. These interactions are known as magnetic exchange interactions. The magnetic domains within the material then exhibit non-zero magnetisations in zero applied fields.

Ferromagnetic materials are endowed with their particular magnetic properties owing to the presence of magnetic domains. On application of a magnetic field, domains with magnetization in the direction of the applied field grow while domains with magnetization opposed to the applied field tend to shrink [53]. The magnetization of a ferromagnetic material saturates when a strong external magnetic field is applied. Very strong applied fields can create a single magnetic domain. On removal of an applied magnetic field, the domain wall can become stuck or pinned and then the domain’s structure will not return to its original state after removing the external magnetic field [53]. This phenomenon is known as magnetic hysteresis.

Closer examination of the domain structure in zero applied field reveals the presence of small transverse domains near the end of the sample. These are called closure domains (Figure 2.6). Despite their size, they can close the magnetic loop between two adjacent domains resulting in a further decrease in magnetostatic energy.

Chapter 2 Theoretical Principles of Magnetic Resonance Imaging Contrast Agents

Ferromagnetic samples are often found to be demagnetized. In order to magnetize them, an external magnetic field is applied to the sample. The process of this magnetisation is shown in Figure 2.7. The magnetization (M) increases slowly at first, then more rapidly as the field (H) is increased until it reaches the saturation point (MS). When the magnetic field decreases, the new curve does not retrace the original curve. In addition to this, when the field reduces to zero, some magnetization remains; a phenomenon called remanent magnetization (MR).

At low temperature, ferromagnetic materials exhibit spontaneous magnetization (atomic dipole moments aligned) in zero field. At higher temperature spontaneous magnetisation vanishes and the material is paramagnetic. Here the magnetization reduces when the temperature is augmented. The material shows paramagnetic behaviour when temperature is higher than Curie temperature; at this point, there is no net magnetic moment or atomic dipole alignment in zero field. As the temperature decreases in a ferromagnetic sample the magnetic susceptibilities (χ) increase until it reaches a critical point at At this point χ is infinite.

H=0 H

(a) (b) Figure 2.5: Ferromagnetic materials domains: (a) the domains are randomly oriented (b) the domains are forced into alignment by an external magnetic field.

Chapter 2 Theoretical Principles of Magnetic Resonance Imaging Contrast Agents

Figure 2.6 : Closure domains at the end regions of sample [83]

Figure 2.7: M vs. H loops in ferromagnetic, paramagnetic and superparamagnetic samples.

Chapter 2 Theoretical Principles of Magnetic Resonance Imaging Contrast Agents

2.2.2 Paramagnetism

Most metals with a magnetic susceptibility χ greater than 0 are paramagnetic. In paramagnetic samples where there is the absence of an external magnetic field, the orientation of the atomic magnetic moments is random and the net magnetic moment is zero, as shown in Figure 2.8a. An applied magnetic field causes the magnetic dipoles to re- orient according to the field as shown in Figure 2.8b. Here the magnetic moment gain by the sample is generally much smaller than the maximum possible magnetic moment when standard laboratory strength or clinical MRI scanner strength fields are applied. For example, a sample containing an N atoms in perfect alignment should have a magnetic moment equal to N.µ, where µ is the magnetic dipole of each atom. In fact the sample in this instance gains a magnetic moment smaller than the maximum possible magnetic moment (N.µ) because of only partial alignment of dipoles. Here, thermal agitation has a very important effect on the degree of magnetic moment alignment.

In 1895 Pierre Curie discovered experimentally that the magnetization (M) of a paramagnetic sample is directly proportional to the applied magnetic field and inversely proportional to the temperature in which the sample is placed (for relatively small magnetic field strengths). The equation, known as “Curie’s law” is given by:

where M is the resulting magnetization, χ is the magnetic susceptibility, H is the magnetic field applied to the sample, T is absolute temperature and C is a material-specific Curie constant [83]. At high magnetic field and low temperature Curie’s law is not valid because when the saturation of magnetization is reached, no further alignment can be accrued. Increasing the magnetic field at this point will not increase the degree of alignment because

Chapter 2 Theoretical Principles of Magnetic Resonance Imaging Contrast Agents all the dipole moments are already fully aligned. Therefore Curie's law is applicable at low magnetization but not where the magnetization is approaching saturation. The magnetic susceptibilities (χ) for paramagnetic materials are much less than one.

The process of the magnetization of paramagnetic samples is shown in Figure 2.7: the magnetization (M) increases slowly as the magnetic field (H) is increased until it reaches the saturation point MS. When the magnetic field decreases the new curve retraces the original curve. In addition, when the field reduces to zero the magnetization goes back to zero.

(a) (b) Figure 2.8 : Schematic representation of the alignment of magnetic moments in a paramagnetic sample (a) in the absence of an external magnetic field the magnetic moments are randomly oriented. (b) An applied magnetic field causes the magnetic dipoles to re-orient according to the magnetic field (H).

2.2.3 Superparamagnetism

Superparamagnetism is a form of magnetism which appears in ferromagnetic or ferrimagnetic nanoparticles. Usually, above the Curie temperature, ferromagnetic or ferrimagnetic materials undergo transition into a paramagnetic state by heating. This rule is not applicable in superparamagnetism because superparamagnetism occurs below the Curie temperature of the materials. Here the formation multiple magnetic domains within a particle become energetically unfavourable when the nanoparticle size is below a certain critical size which forces the nanoparticles to have a single magnetic domain. In 1947 Kittel [84] confirmed that, below a rough estimate of critical size - 10 -100nm for a spherical sample of a common ferromagnetic material, the particle will form a single magnetic domain only. 48

Chapter 2 Theoretical Principles of Magnetic Resonance Imaging Contrast Agents

At very low temperatures, the single domain magnetization is stable and its orientation with respect to the crystal axes in zero applied magnetic field is determined by a magnetic anisotropy energy which is originates from the crystal structure and to a lesser extent the shape of the crystal. As the temperature is raised, thermal energy can become sufficient to cause the magnetization of the particle to fluctuate in direction about the low energy orientation and above a critical temperature (called the blocking temperature) the thermal energy is sufficient to overcome the magnetic anisotropy energy and cause the magnetization to randomly fluctuate in all directions. Magnetic nanoparticles at temperatures greater than the blocking temperature are described as being superparamagnetic.

An ensemble of superparamagnetic nanoparticles retains no magnetization in the absence of an external magnetic field. External applied magnetic field cause the magnetic dipoles of superparamagnets to reorient according to the field which results in a magnetization much higher than that of paramagnetic substances.

The process of the magnetization of superparamagnetic samples is shown in Figure 2.7. The magnetization (M) of the paramagnetic sample increases rapidly as the field (H) increases until it reaches the saturation point MS. In addition to this, the magnetization for superparamagnetic materials reduces to zero when the externally applied magnetic field goes back to zero.

2.3 Mechanisms of longitudinal Proton Relaxation Enhancement by paramagnetic materials

R1 (=1/T1) contrast agents work by increasing the longitudinal relaxation rate of water protons near them by enhancing the stimulation of spin-lattice relaxation processes.

Paramagnetic ions such as Gd and Mn are usually used to increase the R1 contrast.

Chapter 2 Theoretical Principles of Magnetic Resonance Imaging Contrast Agents

The theory for longitudinal proton relaxation enhancement, based primarily on the effect of chelated paramagnetic ions such as gadoteric acid (C16H25GdN4O8 ), is relatively well developed.

The relaxation mechanisms in tissues depend on the interaction of the water molecules with macro molecules via bound states. Varying magnetic fields caused by the relative thermal motion between the water and macromolecules induce exchange of spin between the protons in water and the protons of the macro molecule. These varying fields are enhanced with paramagnetic contrast agents in place of the macromolecules. Here the exchange mechanism is accelerated. Paramagnetic atoms have a number of unpaired electrons and therefore electron spin. The magnetic field of an electron spin is 657 times stronger than the field of a proton spin therefore the interacting field is much stronger. In the case of the gadolinium compounds, the ion is Gd3+ and has seven unpaired spins in its valence shell. For the spins to be effective as proton relaxation enhancers, they ideally have precession frquencies that match the Larmor frequencies of the protons. Gd ions are toxic and cannot be used without being bound to chelates.

The paramagnetic contrast agents reduce T1 to a degree depending on the concentration.

From theory it follows that the influence of the relaxivity r1 of a contrast agent with concentration C on the T1 values of tissues can be written

1/T1 (observed) =1/ T1(tissue) + r1.C

2.4 Mechanisms of Transverse Proton Relaxation Enhancement by Magnetized Particles

The mechanisms by which magnetized particles (such as superparamagnetic or ferromagnetic particles) within aqueous media or tissues can enhance transverse proton relaxivity can be characterized into three regimes known as the static dephasing regime, the motional averaging regime, and the echo limited regime.

Chapter 2 Theoretical Principles of Magnetic Resonance Imaging Contrast Agents

2.4.1 Static dephasing regime

The static dephasing regime is applicable when all water molecules are stationary at a fixed

point around the nanoparticles. This assumption is useful when .

Where τD is the translational diffusion time around the magnetized particle,

Ra is the magnetized particle radius, D is the water diffusion coefficient, and Δω is the difference in angular precession frequency between the local field experienced by a proton at the equatorial line of the particle surface and at the pole:

Where µ0 is the magnetic permeability of free space, M is the particle magnetization and γ is the proton gyromagnetic ratio [82, 85].

The static dephasing model assumes the protons are effectively stationary. In this regime the transverse relaxation time depends only on the volume and the magnetization of the particle. During the measurement time the water molecule does not have enough time to diffuse a significant distance around the nanoparticles, therefore the nanoparticles must be large enough relative to typical diffusion distances for this regime to apply. Dephasing of protons in this regime is primarily due to the different precession frequencies of the different protons in different parts of space owing to the variations in magnetic field strength in space caused by the presence of the magnetized particles.

Chapter 2 Theoretical Principles of Magnetic Resonance Imaging Contrast Agents

2.4.2 Motional averaging regime

This model is valid only when: τD << 1/∆ [82]. This regime is valid where the nanoparticle size is sufficiently small (below about 30nm for pure water) allowing the molecules to diffuse around the nanoparticles in a random path. By following this random path, the proton experiences a range of magnetic field strengths. In this regime the rate of the precession will vary from faster rates when the proton is near the pole of the nanoparticles to slower rates at the equator. This variation takes place because the precession rate is proportional to magnetic field. Here the proton will have its precession rate averaged as a result of its motion. Dephasing in this regime takes longer which results in a longer transverse relaxation times (T2). Here T2 is in inversely proportional to the square of the particle size [82].

2.4.3 Echo limited regime

This regime is applicable where there are large nanoparticles with a diameter size above a few hundred nanometres (for pure water). The echo limited model is very similar to the static dephasing model with one exception, the proton transverse relaxation rate R2 is also dependent upon the echo time that is used in measurements. In this regime T2* remains the same as that for the static dephasing case.

The inter-echo time used is very important in this model. In this regime the transverse relaxation time depends on the volume, the magnetization of the particle and the inter-echo time used during the measurement. This model is useful when the refocusing pulses become effective. The echo limited model becomes applicable where the refocusing pulses become effective at recovering the transverse magnetisation lost due to the magnetic contrast agent. The effectiveness is a qualitative measure of the magnitude of the transverse magnetization that is recovered after a 1 0 refocusing pulse [82]. Here the diffusion of water molecules is proportional to the echo time. Therefore water molecules will not diffuse very far from the nanoparticles if the echo time is short. In this case a large percentage of the transverse magnetization will be recovered by the 1 0 refocusing pulse.

Chapter 2 Theoretical Principles of Magnetic Resonance Imaging Contrast Agents

In contrast, if the echo time is large enough to allow a high percentage of the water molecules to diffuse around the nanoparticles, there will be an irreversible loss of transverse magnetization. Therefore, for a large echo time (2τcp) when 2τcp > τD the refocusing pulse is not as effective in this case T2* and T2 are equivalent. Where τD is the translational diffusion time around the magnetized particle. When 2τcp < τD partial refocussing of the transverse magnetisation leads to an increase in T2 compared to T2*. [82].

Chapter 3 Materials and Instrumentation

Several instrumental techniques and materials are used in each of the subsequent chapters of this thesis. Here I describe the common materials used and their methods of preparation together with the instrumentation that was used to characterise the materials.

3.1 Magnetic Nanoparticles

In order to satisfy criteria of low toxicity, magnetite (Fe3O4) and maghemite (-Fe2O3) nanoparticles have been chosen to serve as the magnetic nuclei of the particles for this study. In general, iron oxide nanoparticles have been extensively studied, well characterised in many aspects, with a toxicity that has been demonstrated to be quite low and well tolerated in the human body [86].

3.1.1 Polymer coated Fe3O4 nanoparticles

We received a range of synthetic polymer coated nanoparticles being synthesised as part of an ARC Discovery project carried out in the School of Physics in collaboration with the

Chemistry Department of Virginia Tech (Virginia, USA). Magnetite (Fe3O4) nanoparticles were coated with three different polymer systems. The magnetite nanoparticles used in all three complexes were synthesized via reductive thermolysis of iron acetylacetonate in benzyl alcohol adapted from the method of Pinna et al 2005, and Miles et al 2009 [87, 88]. The magnetite for each sample of the polymer-coated particles had a particle diameter of ~8 nm and a standard deviation of 2.5 nm.

3.1.2 -Fe2O3 nanoparticles from Sigma Aldrich

Maghemite (-Fe2O3) nanoparticles were obtained from Sigma Aldrich. These nanoparticles have very broad size distributions up to 800 nm. Therefore, a method was developed to fractionate these nanoparticles in order to obtain narrower size distributions from 10 nm to 200 nm.

Chapter 3 Materials and Instrumentation

3.1.2.1 Nanoparticle size fractionation

Fifty mg of iron oxide nanoparticles (-Fe2O3) were added to eight separate tubes. Each tube was filled with 10 mL of distilled water. The tubes were shaken well on Vortex mixer (Scientifica, Model VELP) for 2 minutes then sonicated for 15 minutes using an ultrasonic homogenizer (Biologics, Model 3000) at ambient temperature. They were then centrifuged (BOECO, Model C-28A) for 30 minutes at 3000 g. This centrifugation was equivalent to the effect of ambient gravity force for 2 months. In this thesis I used the nanoparticles in the supernatant phase to incubate the RBCs in the presence of magnetic nanoparticles. The size distribution of the nanoparticles in the stock suspension was measured using dynamic light scattering (DLS) (Zetasizer Nano-ZS, Malvern Instruments) and transmission electron microscopy (TEM) using a JEOL 2100F microscope. The TEM images demonstrate that the majority of particles are roughly spherical with the smallest particles being approximately 10 nm in size. The TEM also showed the presence of aggregated particles. DLS measurements on the suspensions indicated particle sizes ranging from 30 to 300 nm demonstrating that stable aggregates were present in the aqueous suspension as well as being observed in the dry state under TEM.

3.2 Gadoteric acid

Gadoteric acid (C16H25GdN4O8 (279.32 mg/mL)) was obtained from Guerbet (France). The osmolarity of the stock gadoteric acid was 1350 mOsm.kg-1. In order to produce solutions with 300 mOsm osmolarity, 3.5 mL of distilled water was added to 1 mL of C16H25GdN4O8 stock solution. The solution was vigorously shaken for 2 minutes using a vortex mixer (VELP Scientifica).

3.3 Transmission Electron Microscope (TEM) /EELS

The transmission electron microscope (TEM) can obtain a resolution approaching 0.1 nm. In addition to providing this increased resolution, the TEM provides the opportunity to carry out a wide range of experiments to investigate the structural, compositional, and

Chapter 3 Materials and Instrumentation chemical properties of the sample with atomic resolution. The JEOL 2100F TEM has been optimised for biological imaging and analysis.

3.3.1 JEOL 2100F Transmission Electron Microscope (TEM)

The JEOL 2100F was used for analyzing iron oxide nanoparticles inside red blood cells (RBCs) and to determine the nanoparticle sizes. The JEOL 2100F can be operated with an accelerating voltage: 80,100,120,160 and 200 keV. It has an 11 Mpix camera capable of recording both images and diffraction data. Element distribution images can be acquired with a Gatan Tridiem energy filter for a wide range of elements, including important biological elements such as C, N, O, P, and S from a field of view as large as 12 micron. Also this TEM contains software for automated electron tomography data acquisition. The JEOL 2100 TEM is made up of an electron gun, electromagnetic lenses, apertures; detectors, stigmators, deflectors and a sample stage.

Chapter 3 Materials and Instrumentation

Figure 3.1: Schematic representation of the main parts of the TEM[89]. 57

Chapter 3 Materials and Instrumentation

3.3.1.1 Electron gun

In the JEOL 2100F TEM, the electron gun comprises a LaB6 thermionic emission electron source with accelerating voltage: 80, 100, 120, 160 and 200 kV.

3.3.1.2 Lens system

Since electrons are deflected in a magnetic field, the lenses in TEM are electromagnetic lenses. Magnetic fields are generated when a current is applied in the coil. The magnetic field goes out into the vacuum creating the lens field that is used for focussing the electron beam. The focussing can be controlled by adjusting the current through the coils.

3.3.1.3 Condenser lenses

There are two kinds of condenser lenses in the TEM. The first, C1, controls the size reduction of the electron beam. The second, C2, controls the intensity of the electron beam by changing the strength of the condenser C2.

3.3.1.4 Objective lens

The objective lens provides magnification between 20 to 50 times. It has a key effect on the final image.

3.3.1.5 Projector lenses

Projector lenses are used to project the image onto the microscope viewing screen.

3.3.1.6 Apertures

The TEM has three sets of apertures: condenser apertures, objectives apertures and selected area apertures. An aperture is a hole in a metal foil that can be used to select a specific set

Chapter 3 Materials and Instrumentation of electrons by allowing them to pass through the hole, while blocking the remaining electrons with the metal foil.

3.3.1.7 Stigmators and deflectors

Stigmators are used to compensate for asymmetry in the magnetic fields from the main lenses. Deflectors are used to adjust the position of the beam to assist with the optical alignment of the microscope. The JEOL 2100F TEM operates with an imaging technique whereby a beam of electrons is transmitted through a specimen. An image is formed, magnified and directed to appear on a layer of photographic film which is then projected onto a fluorescent screen. There are many advantages in using the JEOL 2100F TEM. These are: increased resolution, depth of field, depth of focus and secondary signals from the specimen. On the other hand there are some limitations such as: sample thickness and sample damage. The TEM can obtain a resolution approaching 0.1 nm which is perfect to study the size of nanoparticles. Also, the resolution of a TEM is limited by the wavelength of the illuminating wave. The resolving power of TEM is given by dmin ( is the wavelength of incident illumination,  is the aperture angle) From this equation it can be concluded that the resolution limit of TEM is a few tenths of a nanometre, which is sufficient to study nanoparticle size and the size of other particles.. Secondly, it can be seen from the equation above that dmin in TEM is inversely proportional to . Therefore the depth of field can be increased by decreasing the aperture angle  For example, if a thick sample is being analysed, the depth field can be increased to a greater value than the thickness of the sample. Thirdly, the JEOL 2100 TEM has a depth of focus ranging from a few metres to tens of thousands of meters. In other words there is no need to adjust the focus to take images. If the image is in focus it will remain in view. This large range plays an important role in terms of taking images. Finally, another important advantage of using the JEOL 2100 TEM is that electrons are one type of ionizing radiation which produces a wide range of secondary signals from the specimen and some of these are summarized in the figure below.

Chapter 3 Materials and Instrumentation

Incident high KV Beam BSE X-Ray S E Auger Visible Light Electron

Specimen

Bremsstrahlumg X_Rays Elastically scattered Inleastically Electrons Scattered Electron Direct Beam

Figure 3.2: Schematic representation of the electron beams transmitted through a specimen.

The JOEL 2100 also presents some limitations. Samples have to be thin enough to permit a beam of electrons to transmit through them. This means many materials require extensive sample preparation to produce a sample thin enough to be electron transparent. There is a potential that the sample may be damaged by electron beam, particularly in the case of biological material[89].

3.4 Zetasizer Nano-ZS, Malvern Instruments

The nanoparticle size distribution for the nanoparticles was measured using both TEM and a Zetasizer Nano-ZS (green badge) model ZEN3500, Malvern Instruments.

Chapter 3 Materials and Instrumentation

This instrument is fitted with a 532nm ‘green’ laser. It is used for measurement types such as particle size, molecular weight and zeta potential in a liquid medium over a wide range of concentrations. Zetasizer Nano-ZS, Malvern Instruments uses the theory of dynamic light scattering (DLS) to measure the size of small particles such as nanoparticles. The size of the nanoparticle measured is the diameter of the sphere that diffuses at the same speed as the particle being measured. The Zetasizer system determines the size by first measuring the Brownian motion of the particles in a sample using dynamic light scattering (DLS) and then interpreting a size from this by analysing the autocorrelation of the fluctuating light intensity in the time domain from the suspension. The particles in a liquid move about randomly and their speed of movement is used to determine the size of the particle. It is known that small particles move quickly in a liquid and large particles move slowly. Dynamic light scattering is also known as photon correlation spectroscopy or quasi-elastic light scattering. When light hits small particles the light scatters in all directions (Rayleigh scattering) so long as the particles are small compared to the wavelength (below 250 nm). The Zetasizer Nano-ZS can also be used to probe the behaviour of complex fluids such as concentrated polymer solutions[90].

Figure 3.3: The Zetasizer Nano- ZS [91].

Chapter 3 Materials and Instrumentation

3.5 Superconducting Quantum Interference Device (SQUID) Magnetometry

Magnetic susceptibility measurements were made using a Quantum Design MPMS-7 SQUID-based magnetic susceptometer. The MPMS-7 allows measurements with sample temperatures from 2K to 800K on encapsulated samples in fields up to 70 kOe with a precision of 0.1 kOe. DC-magnetic susceptibility and magnetisation can be measured on powder samples or single crystals. The SQUID input coils are connected to a system of super-conduction detection coils. The measurement is carried out while the sample moves through this coil system. When the sample moves inside the detection coil, the movement of the magnetized sample creates an electric current in the detection circuit. The variation of the current in the detection coil produces corresponding variation in the SQUID output voltage, which is proportional to the change in the magnetic field [92].

3.6 The Minispec nuclear magnetic resonance (NMR) analyser (proton relaxometry)

The Minispec NMR analyser is a time-domain NMR (TD-NMR) instrument used to measure the spin-lattice relaxation time constant (T1) and spin-spin relaxation time constant

(T2) of samples containing water or fat. The T1 and T2 data for this thesis have been collected using a Bruker mq-60 MHz Minispec NMR Analyzer (Bruker, Ettlingen, Germany) operating at a magnetic field of 1.4 T. Samples were placed in an aluminium block within a water bath at 37.5 °C for at least 15 minutes prior to relaxation rate measurements. Proton transverse relaxation times (T2) were obtained from fitting a monoexponential decay curve to signal data generated by a Carr-Purcell-Meiboom-Gill (CPMG) spin-echo pulse sequence with an echo spacing of 1 ms and a repetition time of 5 s. Longitudinal relaxation times (T1) were obtained from fitting a monoexponential recovery curve to signal data generated with an inversion recovery (IR) pulse sequence using 10 logarithmically spaced inversion times between 10 and 10000 ms. R1 (R1=1/T1) and R2 (R2=1/T2 ) data are reported in this thesis. 500 µL of sample solution or suspension was used for relaxometry [93].

Chapter 3 Materials and Instrumentation

Figure 3.4: The Minispec NMR Analyser.

3.7 Scanning Probe Microscopy

In a scanning probe microscope, a very sharp tip/probe is scanned across the sample surface. In an atomic force microscope (AFM) and magnetic force microscope (MFM) the probe is mounted on a small leaf spring and the motion of the cantilever is used to generate an image. Scanning probe microscopes generate three dimensional images and can be operated in either air or solutions. The basic principle of a scanning probe microscope is to scan a very sharp tip across the surface of a sample. The scanning motion is usually performed with a piezo tube consisting of two parts, one for the movement in the x-y-plane and the other one for z-movements. In case of scanning tunneling microscopy (STM) applications, the tip can be a thin Pt wire, which is etched on one end to obtain the sharp tip. In AFM applications, a sharp tip is attached to a flexible spring (the cantilever). A laserbeam is directed on the back of the cantilever and then reflected onto a 4-segment

Chapter 3 Materials and Instrumentation photodetector, which allows detection of the motion of the cantilever. Depending on the application and the imaging mode, different parameters are used as feedback signals. As the tip is scanned in the x-y-plane, the feedback signal is used to maintain a constant tip surface distance, which is achieved by the z-piezo, which moves the tip up and down, following the surface profile. With the ability to position the tip precisely within sub-nanometer precision, three dimensional images of the sample surface with magnifactions up to 10 million times can be taken. The AFM/MFM images for this theses have been collected using a Dimension 3100 AFM [94]. This instrument can be operated in different modes such as contact mode, tapping mode, phase imaging, lateral force mode, force imaging, magnetic force microscopy (MFM) and electric techniques (CAFM, Tuna). The details of the use of the instrument in magnetic force mode will be discussed in more detail in Chapter 7.

Figure 3.5: AFM Dimension 3100.

Chapter 3 Materials and Instrumentation

3.8 Optical microscope

A Nikon Optiphot 2 Trinocular Microscope with darkfield illumination was used to count the RBCs. A ProgRes C10 plus (Firewire and 3 Megapixel CCD sensor) was attached to this microscope to record the images. This microscope can provide images with: 12 V ,100W Illuminator, CFUW 10x eyepieces; plan 4x, 40 and 100x height resolution objective and a darkfield condenser 1.2-1.4.

3.9 Red blood cells (RBCs)

Figure 3.6: Schematic illustration of RBCs preparation.

Blood is the body’s only fluid tissue. It is composed of liquid plasma and formed elements include: erythrocytes or red blood cells (RBCs), leukocytes or white blood cells (WBCs) and platelets. Blood was collected into heparinised tubes as shown in figure 3.5 and centrifuged at 3000 g at 4°C for 10 minutes to separate the RBCs from the plasma. The RBCs were then washed three times with phosphate buffered saline (PBS, 300 mOsm, pH 8). Suspensions of RBCs at 300 mOsm were prepared by suspending approximately 2 mL of packed RBCs in 10 mL of PBS.

3.10 Hemocytometer (Counting chamber)

A Bright-line hemocytometer was used for counting RBCs in vitro. The hemocytometer consists of a chamber with volume of 0.1µL, with a cell depth 0.1mm and has rulings that 65

Chapter 3 Materials and Instrumentation cover 9 square mm. The central square mm is ruled into 25 groups of 16 small squares, each group separated by triple lines as shown in figure 3.6. The volume of each of the 16 small squares is 0.00025 cubic mm.

Figure 3.7: optical microscopy image of counting chamber square in the hemocytometer.

3.11 Loading RBCs with contrast agents

In this thesis I investigate the optimal conditions for loading RBCs with magnetic resonance contrast enhancing materials using two different approaches and identify the limit of particle (or aggregate) size that can be incorporated into the cells through pores induced in their membranes. The first approach of loading the RBCs with magnetic contrast agent comprised incubation of the RBCs in the presence of magnetic contrast agent under hypo-osmolar conditions which cause the RBCs to swell and pores to open up in the cell membranes (Figure 3.7). The RBCs were subsequently re-sealed by bringing the osmolarity of the medium back up to physiological values [52] .The second approach involved preparing the RBCs in a hyperosmolar solution of dimethyl sulfoxide (DMSO) which can diffuse into the cells thus preventing an osmotic pressure difference across the

Chapter 3 Materials and Instrumentation cell membrane.The cells were then rapidly flushed with a slightly hypo-osmolar solution containing the magnetic nanoparticles to induce an osmotic pressure pulse temporarily opening pores in the cell membrane [57]. Full details of these techniques will be given in the relevant Chapters (Chapters, 4, 5, and 6).

Chapter 4 Loading Erythrocytes with Maghemite Nanoparticles via Osmotic Pressure Induced Cell Membrane Pores

4.1 Introduction

In clinical magnetic resonance imaging (MRI) studies, contrast agents are often used to enhance the contrast between different tissues by altering the local proton relaxation times [36].

Recent developments in this area include the use of nanoparticles with large magnetic moments as contrast agents. Such nanoparticles exhibit a property known as superparamagnetism at room temperature when they are below a certain size that depends on the magneto-crystalline anisotropy of the particles. Their superparamagnetic properties have the advantage of high magnetic susceptibilities which result in large proton relaxivities when the particles are suspended in aqueous media.

Magnetic particles are cleared from the blood stream via ingestion by macrophages. In order to extend the lifetime of magnetic particles in the blood stream, attempts have been made to encapsulate magnetic particles within red blood cells (RBCs) so that macrophage clearance is slowed. A convenient method for loading RBCs with magnetic nanoparticles involves inducing pores in the RBC membrane by varying the osmotic pressure. A key strategy for loading RBCs with magnetic particles is to incubate them in the presence of the particles under hypo-osmolar conditions [55, 56]. The RBCs can then be re-sealed by bringing the osmolarity of the medium back up to physiological values. Here, we investigate the optimal conditions for loading RBCs by this technique and identify the limit of particle (or aggregate) size that can be incorporated through the pores.

Chapter 4 Loading Erythrocytes with Maghemite Nanoparticles via Osmotic Pressure Induced Cell Membrane Pores

4.2 Methods and Materials

4.2.1 Preparation of Stock Solutions and Cell Suspensions

4.2.1.1 Maghemite (γ-Fe2O3) nanoparticles

Maghemite (γ-Fe2O3) nanoparticles were obtained from Sigma Aldrich. Magnetometry measurements at 300 K were made using a Quantum Design MPMS-7 SQUID-based magnetic susceptometer. The particles were found to exhibit a wide range of aggregate sizes in aqueous suspension with larger aggregates falling out of suspension over time. In order to produce a stable suspension of particles, 50 mg of -Fe2O3 were mixed with 10 mL of distilled water. The suspension was vigorously shaken for two minutes using a vortex mixer and was then sonicated for 15 minutes using an ultrasonic homogenizer (Biologics, Model 3000). The suspension was then centrifuged for 30 minutes at 3000 g. The supernatant was decanted and used as the stock -Fe2O3 suspension for the remaining steps. The size distribution of the nanoparticles in the stock suspension was measured using dynamic light scattering (DLS) (Zetasizer Nano-ZS, Malvern Instruments). Human blood (approx 8 mL) was collected into heparinised tubes and was centrifuged at 3000 g at 4°C for 10 minutes to separate the RBCs from the plasma.

4.2.1.2 Human Red Blood Cells (RBCs)

The RBCs were washed three times with phosphate buffered saline (PBS, 300 mOsm, pH 8). A stock suspension of RBCs at 300 mOsm was prepared by suspending approximately 2 mL of packed RBCs in 10 mL of PBS.

4.2.2 Methods of Loading Cells with Maghemite (γ-Fe2O3) nanoparticles

The following protocol was used to incubate the RBCs in the presence of a stable suspension of -Fe2O3 nanoparticles at a given osmolarity. X mL (where X is 0.13, 0.4, 1.0,

2.0 or 4.0) of the -Fe2O3 suspension were added to Y mL (where Y is 4, 2, 2, 2 or 2) of the

Chapter 4 Loading Erythrocytes with Maghemite Nanoparticles via Osmotic Pressure Induced Cell Membrane Pores stock RBC suspension at 300 mOsm to produce a suspension of RBCs and magnetic nanoparticles at 290, 250, 200, 150 or 100 mOsm respectively. The RBCs were then incubated with -Fe2O3 nanoparticles at 4°C for 12 hours with gentle shaking using a vortex mixer. After the incubation period, the osmolarity of the suspension was brought back up to 300 mOsm by adding Z mL (where Z is 0.4, 1.2, 3.0, 6.0 or 12.0) of PBS at 400 mOsm to the suspension of cells and nanoparticles.

4.2.3 Measurement and Characterization of Cell Suspensions

4.2.3.1 Transmission Electron Microscopy

The resulting cell suspensions were prepared for transmission electron microscopy (TEM) in the following way. The suspensions were centrifuged and the cell pellet was re- suspended in PBS at a volume ratio of 1:1. The suspension was then fixed with glutaraldehyde at a volume ratio of 1:20. After fixation, the cells were serially dehydrated for 40 seconds each time in consecutive solutions of 50, 70, 95 and 100% ethanol. The sample was dried by washing twice in 100% acetone. After dehydration, the sample was infiltrated with Spurrs resin for 3 minutes each in consecutive solutions of acetone and Spurrs resin at volume ratios of 3:1, 1:2 and 1:3, respectively. The sample was bonded in 100% Spurrs resin followed by polymerization overnight at 70°C. An ultramicrotome was used to cut 100-nm thick sections from the resin block. The sections were then mounted on carbon-coated copper grids. A series of control samples was prepared using the above protocol using distilled water in place of the stock -Fe2O3 suspension. TEM (JEOL-2100, Philips) was used to record bright-field images of these sections. Energy-filtered TEM (EFTEM) was conducted using an electron energy-filter system (Gatan, Tridiem). Elemental maps for iron were acquired by the conventional three-window method [95] using the Fe M-edge. Background (pre-edge) images were acquired at 45 eV and 50 eV, with the signal (post-edge) image acquired at 59 eV. All images were acquired with an energy window (slit width) of 5 eV. These conditions were optimized to provide good signal-to-noise ratio with suitable background removal. All TEM images were acquired at 120 keV at room temperature. Particle size distributions within the RBCs and outside the

Chapter 4 Loading Erythrocytes with Maghemite Nanoparticles via Osmotic Pressure Induced Cell Membrane Pores

RBCs were measured by analysing at least 20 iron maps, each with a field of view of 5 × 5 μm.

4.2.3.2 Freeware ImageJ (NIH) measurements

Freeware ImageJ (NIH) was used to measure the intracellular and extracellular areas in each field of view, as well as the intracellular and extracellular size distribution for particles/aggregates specifically containing iron. Each particle/aggregate was fitted with an ellipse and its effective size d was calculated as 2√(ab), where a and b are the semi-minor and semi-major axes of the fitted ellipse, respectively. The volume of each particle or 3 cluster was estimated as (1/6)πd . Concentrations of -Fe2O3 trapped within the cells were then estimated from the ratio of total cell to nanoparticle volume in the imaged 100-nm thick sections by assuming that each particle/aggregate had the density of bulk -Fe2O3.

4.2.3.3 Inductively coupled plasma mass spectrometry (ICP-MS).

In order to determine the iron concentration in the stock -Fe2O3 suspension, 1 mL of stock

-Fe2O3 was added to 10 mL of concentrated HNO3 and heated at 95°C to reduce the volume to 1 mL in total. The digestion was then made up to 30 mL with dilute (1%) HNO3. The iron concentration in the digested solution was measured by ICP-MS.

Chapter 4 Loading Erythrocytes with Maghemite Nanoparticles via Osmotic Pressure Induced Cell Membrane Pores

4.3 Results

4.3.1 TEM and DLS of the -Fe2O3 nanoparticle

Figure 4.1 TEM of γ-Fe2O3 nanoparticles cast onto grid from aqueous suspension.

Figure 4.1 shows a TEM image of -Fe2O3 nanoparticle aggregates prepared directly from a suspension in distilled water by casting them on the TEM grid. This figure demonstrates that the majority of particles are roughly spherical with the smallest particles being approximately 10 nm in size. DLS measurements on the suspensions indicated particle sizes ranging from 30 to 300 nm demonstrating that stable aggregates were present in the aqueous suspension as well as being observed in the dry state under TEM.

Chapter 4 Loading Erythrocytes with Maghemite Nanoparticles via Osmotic Pressure Induced Cell Membrane Pores

4.3.2 Superconducting Quantum Interference Device (SQUID) Magnetometry

A magnetic hysteresis curve of the -Fe2O3 nanoparticles at 300 K is presented in Figure

4.2. The magnetometry measurements show that the -Fe2O3 nanoparticles mainly exhibit superparamagnetic behavior but with some remanent magnetization at zero field (Figure

4.2(b)). The saturation magnetization for the -Fe2O3 is 66 emu/g, a value slightly lower than the published value for bulk -Fe2O3 (76 emu/g) [96].

(a) (b)

Figure 4.2: (a) Magnetization vs. applied magnetic field for the γ-Fe2O3 nanoparticles and (b) magnetization vs. applied field behavior close to zero field.

Chapter 4 Loading Erythrocytes with Maghemite Nanoparticles via Osmotic Pressure Induced Cell Membrane Pores

4.3.3 TEM of RBCs

(a) (b)

Figure 4.3: (a) Transmission electron micrograph of γ-Fe2O3 nanoparticles embedded in resin and (b) iron map corresponding to region imaged in (a) generated from electron energy loss imaging.

(a) (b)

Chapter 4 Loading Erythrocytes with Maghemite Nanoparticles via Osmotic Pressure Induced Cell Membrane Pores

(a) (b)

Figure 4.3(a) is a TEM micrograph of the -Fe2O3 nanoparticles embedded in resin showing the presence of a large aggregate. Figure 4.3(b) shows the EELS iron map corresponding to the same region imaged in Figure 4.3(a). Figure 4.4(a) shows a TEM 75

Chapter 4 Loading Erythrocytes with Maghemite Nanoparticles via Osmotic Pressure Induced Cell Membrane Pores image of an unstained section through a red blood cell after incubation with the -Fe2O3 nanoparticles at 200 mOsm with its corresponding iron map image in Figure 4.4(b). The image indicates that nanoparticles can be found inside the cell. In comparison, a TEM and iron-map image of an unstained section through a control RBC are shown in Figure 4.5. While electron dense particles can be seen in the TEM image, the iron map indicates the absence of -Fe2O3 nanoparticles. Similar TEM and iron-map images of RBCs loaded at

150 and 250 mOsm showed low uptake of -Fe2O3 into the RBCs. No iron-containing particles were found inside the RBCs when loaded at 290 mOsm (see Figure 4.6).

4.3.4 The concentrations of -Fe2O3 iron inside the cells

The concentrations of -Fe2O3 iron found inside the cells determined from analysis of the

TEM iron map images together with the expected average concentrations of -Fe2O3 iron in the final cell suspensions determined from the ICP measured concentration of iron in the stock nanoparticle suspension are shown for each preparation in Figure 4.7. The concentration of -Fe2O3 iron in the cells incubated at 290 mOsm was found to be zero.

Cells incubated at 250 mOsm appeared to have intracellular -Fe2O3 iron concentrations slightly less than the predicted average -Fe2O3 iron concentration. Cells incubated at 200 mOsm were found to have intracellular -Fe2O3 iron concentrations somewhat higher than the predicted average -Fe2O3 iron concentration. For the cells incubated at 150 mOsm, only a few were found intact. Although the intracellular -Fe2O3 iron concentration was found to be less than the predicted average -Fe2O3 iron concentration, the error on the measurement could be somewhat larger than that calculated because of the small number of cells measureable. At 100 mOsm, no intact cells were found and hence no intracellular -

Fe2O3 iron concentration is reported.

Chapter 4 Loading Erythrocytes with Maghemite Nanoparticles via Osmotic Pressure Induced Cell Membrane Pores

Figure 4.7: Estimations of concentrations of γ-Fe2O3 Fe inside cells determined from TEM together with predicted average concentration of γ-Fe2O3 Fe in final cell suspension determined from ICP measurement of stock γ-Fe2O3 suspension. Error bars on measurements inside cells are derived from the standard deviation of values determined from the twenty fields of view analyzed. Error bars on the average Fe concentration values are determined from the analytical error on the ICP measurement method.

4.3.5 The distributions of particle/aggregate sizes inside and outside of the cells

The distributions of particle/aggregate sizes found inside and outside of the cells is shown in Figure 4.8. The largest clusters found inside the cell were approximately 120 nm (found in the cells incubated at 200 mOsm

Chapter 4 Loading Erythrocytes with Maghemite Nanoparticles via Osmotic Pressure Induced Cell Membrane Pores

Figure 4.8: Distribution of particle/aggregate sizes inside (black) and outside (gray) cells after incubation with γ-Fe2O3 nanoparticles at the indicated osmolarity.

Chapter 4 Loading Erythrocytes with Maghemite Nanoparticles via Osmotic Pressure Induced Cell Membrane Pores

4.4 Discussion

Previous studies of the incorporation of magnetic nanoparticles into RBCs have shown that it is possible to load particles into the cells by inducing pore formation through osmotic pressure [55, 56]. In these previous studies, based on osmotic incubation methods, there were no systematic studies of the effect of different osmotic pressures on the resulting loadings. Furthermore, very few details were given with respect to the methodology of loading, probably because of commercial sensitivities. In the case of Brahler and colleagues [56], it was claimed that an osmolarity of 20 mOsm was used. However, this is clearly incorrect because there is no cell survival whatsoever at 20 mOsm. In the case of Antonelli and coworkers [97] , no osmolarity was given at all. Hence, the current systematic study of the effect of osmolarity on the nanoparticle loading of the cells builds on these earlier studies and enables determination of the optimum osmolarity for loading and the limits on the particle sizes that can be loaded. Brahler and colleagues [56] showed a single conventional TEM image of a loaded cell indicating the presence of particles. However, no elemental mapping was used to confirm the composition of the particles. Antonelli and colleagues [97] also showed some TEM images but with no elemental mapping. As seen in this study, electron dense particles can be seen that are not iron oxide nanoparticles. They may be artefacts of the preparation procedures. Hence the elemental mapping shown in this study builds further on these earlier studies. This study aimed to find the optimum osmolarity for loading and the limits on the particle sizes that can be incorporated within the cells.While the fundamental particle sizes of the -Fe2O3 used in this study were generally less than 100 nm, both the TEM and DLS indicated that particles were mostly found in aggregates ranging up to a few hundred nanometers in size. The suppressed saturation magnetization compared with the literature value for bulk -Fe2O3 [96] is consistent with the TEM observation that the fundamental particle sizes range down to approximately 10 nm. The mixture of superparamagnetic and magnetically blocked behavior also indicates a large range of particle sizes.

The observation that no -Fe2O3 particles were found within cells incubated with nanoparticles at 290 mOsm (Figures 4.6, 4.7, and 4.8) while -Fe2O3 particles were found

Chapter 4 Loading Erythrocytes with Maghemite Nanoparticles via Osmotic Pressure Induced Cell Membrane Pores within cells incubated at lower osmolarities (Figures 4.4, 4.7, and 4.8) strongly suggests that the mechanism by which particles have entered the cells is through osmotic pressure induced pores in the cell membrane.

Figures 4.7 and 4.8 suggest that the optimum incubation condition for loading RBCs with nanoparticles is at 200 mOsm. At 200 mOsm, the ratio of the -Fe2O3 iron concentration within the cells to the average concentration in the cell suspension is greatest (Figure 4.7). The data suggest that the concentration within the cells is greater than that outside the cells. While there could be a trapping of particles within cells while the pores are still open, resulting in a greater concentration inside than out, the most likely explanation of this observation is that there is a systematic error in our calculation of intracellular concentrations. There may be a systematic error in the section thickness. Also, our calculations have assumed that the aggregates of nanoparticles have the density of -Fe2O3, whereas the aggregates must have a density somewhat less than that for -Fe2O3 because of the interstitial spaces between the nanoparticles. Hence it is unlikely that the intracellular concentration exceeds the extracellular concentration in the 200 mOsm preparation. Nevertheless, the error on the density estimate is likely to be systematic across the different loading conditions and as such the data strongly indicate that there is a more efficient loading of nanoparticles at 200 mOsm compared with the other incubation conditions. Figure 8 also suggests a more efficient loading of particles since a wider range of aggregate sizes is observed inside the cells. The increased range of aggregate sizes observed in cells incubated at 200 mOsm compared with those incubated at 250 mOsm may indicate larger pores forming under the greater osmotic pressure. The cut-off of approximately 120-nm for the aggregate size distribution for cells incubated at 200 mOsm suggests that the pore sizes range up to at least this size. Interestingly, it has been found that the maximum pore sizes induced in RBC membranes by electroporation are also approximately 120 nm [98]. The aggregate size distribution found in the cells incubated at 150 mOsm is not as broad as that found for the cells incubated at 200 mOsm. This observation may reflect the fact that very few cells survived the incubation process at 150 mOsm. Possibly only those with smaller pore sizes survived. Alternatively, the very low number of cells observed may result in the

Chapter 4 Loading Erythrocytes with Maghemite Nanoparticles via Osmotic Pressure Induced Cell Membrane Pores observed distribution being biased towards the individual cells measured because of a lack of statistical averaging.

4.5 Conclusion

Magnetic nanoparticles can be loaded into RBCs by temporarily incubating the cells with the nanoparticles under reduced osmolarity conditions. The optimum osmolarity for particle loading is close to 200 mOsm, a condition at which the greatest fraction of particles in the incubating medium enter the cells. The maximum size of aggregate that enters the cells under these conditions is approximately 120 nm. Osmolarities of 150 mOsm or lower cause excessive cell destruction. It is possible to load nanoparticles at 250 mOsm but at this osmolalrity only a few nanoparticles were observed inside the cells. RBCs loaded with -

Fe2O3 nanoparticles by this method are potential candidates for MRI contrast agents with an extended residence time in the blood stream.

Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods

5.1 Introduction

Magnetic nanoparticles have been developed for use as magnetic resonance imaging (MRI) contrast agents. Their superparamagnetic properties have the advantage of high magnetic susceptibilities which result in high proton relaxivities when the particles are suspended in aqueous media [36]. Magnetic particles are cleared from the blood stream via ingestion by macrophages. In order to extend the lifetime of the particles in the blood stream, attempts have been made to encapsulate magnetic particles within erythrocytes, also known as red blood cells (RBCs), so that macrophage clearance is slowed [99].

In this study we compare the effectiveness of magnetic resonance contrast enhancement by RBCs loaded with magnetic nanoparticles by two different methods. The first method (Figure 5.2) involved incubating the RBCs in the presence of four different magnetic iron oxide nanoparticle systems (three different polymer coated systems and one uncoated system) under hypo-osmolar conditions. The RBCs were subsequently re-sealed by bringing the osmolarity of the medium back up to physiological values [52]. The second method (Figure 5.3) involved the preparation of RBCs in a hyper-osmolar solution of dimethyl sulfoxide (DMSO) which diffuses into the cells thus preventing an osmotic pressure difference across the cell membrane. Suspensions of RBCs prepared in this way were then rapidly flushed with a hypo-osmolar solution containing magnetic iron oxide nanoparticles. The rapid dilution of the DMSO outside the RBCs causes a temporary osmotic pressure difference across the cell membrane which becomes balanced again once the DMSO internal to the RBCs is able to diffuse from the cells. This temporary pressure difference, or “osmotic pulse”, temporarily opens pores in the red cell membranes potentially allowing nanoparticles to diffuse into the cells [57]. The osmotic pulse method was used with four different magnetic nanoparticle systems. The nanoparticle systems comprised uncoated iron oxide particles and three sets of iron oxide particles coated with different stabilizing polymers. Previous studies have investigated the enhancement of the proton relaxivity of RBCs loaded with iron oxide nanoparticles [56]. Here we test the relative efficacy of two

Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods different methods of loading RBCs with magnetic nanoparticles and compare the effect of different stabilizing polymer coatings of the efficacy.

5.2 Materials and Methods

5.2.1 Preparation of Stock Solutions and Cell Suspensions

5.2.1.1 Iron oxide nanoparticles

A) Uncoated Maghemite

Maghemite (-Fe2O3) nanoparticles were obtained from Sigma Aldrich. The particles were found to exhibit a wide range of aggregate sizes in aqueous suspension with larger aggregates sedimenting over time. In order to produce a stable suspension of particles, 50 mg of -Fe2O3 were mixed with 10 mL of distilled water. The suspension was vigorously shaken for two minutes using a vortex mixer and was then sonicated for 15 minutes using an ultrasonic homogenizer (Biologics, Model 3000). The suspension was then centrifuged for 30 minutes at 3000 g. The supernatant was decanted and used as the stock -Fe2O3 suspension for the remaining steps. The transverse and longitudinal proton relaxivities for the stock solution, measured at 37 °C and a field strength of 1.4T, were 236.2 and 10.3 s-1 mM Fe-1 respectively

B) Polymer stabilized magnetite

Figure 5.1: Polymer binding to the surface of Fe3O4. 83

Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods

Three different polymer stabilized magnetite (Fe3O4) nanoparticles were used for loading the erythrocytes. The magnetite nanoparticles used in all three complexes were synthesized via reductive thermolysis of iron acetylacetonate in benzyl alcohol adapted from the methods of Pinna et al (2005) [88] and Miles et al (2009) [87].

 PEO Stabilized Magnetite

A 2850 g mole-1 poly(ethylene oxide) stabilizer was strongly bound to the magnetite via a tris(ammonium phosphonate) binding group [100, 101]. The PEO stabilizer was prepared using 3-hydroxypropyltrivinylsilane as an initiator for polymerization of ethylene oxide, catalyzed by a coordination catalyst (Zn3[Co(CN)6]2), to produce a polymer with a molecular weight distribution of 1.2. Zwitterionic ammonium phosphonate groups were then added across the vinylsilyl sites on the initiator end of the polymer utilizing a procedure described previously [101] and these functional endgroups were bound to the magnetite particles. The PEO stabilizer generates a predominantly charge neutral hydrophilic coating on the magnetite particle. The composition of the complexed magnetite was approximately

50% polymer with the remainder being Fe3O4. The transverse and longitudinal proton relaxivities for the complex, measured at 37 °C and a field strength of 1.4T, were 45.4 and 12.7 s-1mM Fe-1 respectively.

 Diblock PEO-PAA Stabilized Magnetite

The second polymer stabilizer used was a diblock copolymer of poly(ethylene oxide) (PEO) (2000 g/mole) and poly(acrylic acid) (PAA) (12,800 g/mole). A PEO oligomer with one non-functional methoxy endgroup and one hydroxyl endgroup was reacted with bromoisobutyryl bromide to yield a bromoalkyl functional PEO. This functional PEO was then utilized to polymerize t-butyl acrylate in a controlled free radical polymerization (ATRP). Finally, the t-butyl groups were removed to afford the PEO-PAA diblock copolymer. The composition of the complexed magnetite was approximately 67% polymer with the remainder being Fe3O4. The transverse and longitudinal proton relaxivities for the

Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods complex, measured at 37 °C and a field strength of 1.4T, were 46.6 and 9.5 s-1 mM Fe-1 respectively.

 Triblock Amphiphilic Copolymer Stabilized Magnetite

The third stabilizer consisted of a block copolymer of poly(ethylene oxide) (PEO) and poly(propylene oxide) (PPO). The polymer consists of a 1300 g mole-1 PEO block bound to the magnetite via a tris (ammonium phosphonate) binding group at one end. Attached to this initial PEO is a 2900 g mole-1 PPO block followed by a 3000 g mole-1 PEO block. In brief, the initial PEO block was prepared using the 3-hydroxypropyltrivinylsilane as an initiator, as described above. The PEO was then used as a macroinitiator for coordination- catalyzed polymerization of the PPO block, then the PEO-PPO was utilized as a macroinitator for the final PEO polymerization using conventional base catalysis. The PEO- PPO-PEO triblock stabilizer generates a neutral magnetite-polymer complex with amphiphilic characteristics due to the hydrophilic and hydrophobic PEO and PPO blocks respectively. The composition of the complex was approximately 50% polymer with the remainder being Fe3O4. The transverse and longitudinal proton relaxivities for the complex, measured at 37 °C and field strength of 1.4T, were 105.6 and 11.0 s-1 mMFe-1 respectively.

The iron concentrations in the stock suspensions of the three different magnetic iron oxide nanoparticle polymer coated systems in water were 0.511, 0.255 and 0.629 mg Fe/mL for the PEO-coated, diblock-coated and triblock-coated nanoparticle systems respectively. The concentration of iron in the stock suspensions of the uncoated nanoparticles was 0.047 mg Fe/mL.

5.2.1.2 Human Red Blood Cells (RBCs)

Human blood (approx. 8 mL) was collected into heparinised tubes and centrifuged at 3000 g at 4°C for 10 minutes to separate the RBCs from the plasma. The RBCs were washed three times with phosphate buffered saline (PBS, 300 mOsm, pH 8). A stock suspension of RBCs

Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods at 300 mOsm was prepared by suspending approximately 2 mL of packed RBCs in 10 mL of PBS.

5.2.2 Methods of Loading Cells with Magnetic Nanoparticles

5.2.2.1 Hypo-osmolar Incubation Method (Osmotic incubation):

The following protocol was used to incubate the RBCs in the presence of the four different magnetic iron oxide nanoparticle systems (the three different polymer coated systems and one uncoated system) at a given osmolarity. 1 mL of iron oxide nanoparticle suspension was added to 2mL of the stock RBC suspension at 300 mOsm to produce a suspension of RBCs and magnetic nanoparticles at 200 mOsm. The RBCs were then incubated with the iron oxide nanoparticles at 4°C for 12 hours with gentle shaking using a vortex mixer. After the incubation period, the osmolarity of the suspension was brought back up to 300 mOsm by adding 3mL of PBS at 400 mOsm to the suspension of cells and nanoparticles [52]. The cell suspension was incubated at 37C for 1 hour and was then washed with PBS 5 times before returning to 4C for storage before measurements or further preparation. A schematic of the procedure is shown in Figure 5.2a. A series of control samples was prepared using the above protocol using distilled water in place of the four different magnetic iron oxide nanoparticle systems.

Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods

Figure 5.2a: Hypo-Osmolar incubation procedure.

5.2.2.2 Osmotic Pulse Method (Osmotic pulse)

1 mL of 2M dimethyl sulfoxide (DMSO, formula (CH3)2SO) was added to 2 mL of RBCs suspended at 300 mOsm to produce a suspension of RBCs and DMSO. The RBCs were then incubated with DMSO for less than 1 min. This cell suspension was then centrifuged for 5 min at 3000 g. The supernatant was discarded and the RBC pellet was re-suspended in 3 mL of iron oxide nanoparticle suspension for 1 min. The moment the iron oxide nanoparticle suspension is added, there is a short osmotic pulse caused by the sudden dilution of the extracellular DMSO [57]. The resulting cell suspension was centrifuged for 5 min at 3000 g. The supernatant of this solution was also discarded. In order to help the cell membrane pores close, 1 mL of dextran (1.32 mM) was added to the RBC pellet. The cell suspension was incubated at 37C for 1 hour and was then washed with PBS 5 times before returning to 4C for storage before measurements or further preparation. A schematic of the procedure is shown in Figure 5.2b. Control samples were prepared using the above protocol with distilled water in place of the four different magnetic iron oxide nanoparticle systems.

Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods

Figure 5.2b: Osmotic pulse procedure.

5.2.3 Measurement and Characterization of Cell Suspensions

5.2.3.1 Cell Counting

Concentrations of RBCs in suspension were measured using a hemocytometer (model - BrightLine, Hausser Scientific, Horsham, USA). Digital optical microscope images of the hemocytometer cells were captured then analysed using ImageJ (v 1.42q, National Institute of Health) to count the cells. Cell counting was used to calculate cell specific proton relaxivities and percentages of cells that survived the preparation procedures.

5.2.3.2 Cell Size and Morphology The RBCs suspension was placed in the hemocytometer.The cell depth of the hemocytometer is 0.1mm; each square 0.2mm separated by triple line is ruled into 16 small squares. Digital optical microscope images of the hemocytometer cells were captured then the projection area was calculated by using ImageJ (v 1.42q, National Institute of Health)

Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods

The digital images acquired from the optical microscope as shown in figure 5.3 were used to assess the morphology of the RBCs. The effect of loading of the magnetic nanoparticles into the RBCs on cell size was assessed by measuring the distribution of projected cell areas in the digital images of the cells in the hemocytometer. The number of cells measured for each preparation ranged from 87 to 186.

Figurer 5.3: RBCs suspension on a hemocytometer

5.2.3.3 Dynamic light scattering (DLS) measurement:

The nanoparticle size distribution of the stock suspensions of the four different magnetic iron oxide nanoparticle systems (three different polymer coated systems and one uncoated system) were measured using Zetasizer Nano-ZS (green badge) model ZEN3500, Malvern Instruments.

Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods

5.2.3.4 Proton relaxometry

Proton relaxation rate measurements were made on a Bruker mq-60 MHz Minispec NMR Analyzer (Bruker, Ettlingen, Germany) operating at a magnetic field of 1.4 T. Samples were placed in an aluminium block within a water bath at 37.5 °C for at least 15 minutes prior to proton relaxation rate measurements. Proton transverse relaxation times (T2) were obtained from fitting a monoexponential decay curve to signal data generated by a Carr- Purcell-Meiboom-Gill (CPMG) spin-echo pulse sequence with an echo spacing of 1 ms and a repetition time of 5 s. Longitudinal relaxation times (T1) were obtained from fitting a monoexponential recovery curve to signal data generated with an inversion recovery (IR) pulse sequence using 10 logarithmically spaced inversion times between 10 and 10000 ms.

R1 (R1=1/T1) and R2 (R2=1/T2 ) data are reported in this chapter. 500 µL of cell suspension was used for each measurement. For some investigations, 500 µL of supernatant from the first wash of the cells was taken separately for proton relaxometry. The cell specific longitudinal (r1) and transverse (r2) proton relaxivities were calculated by dividing R1 and

R2 by the number of the RBCs per mL.

5.2.3.5 Calculation of Apparent Intracellular Proton Relaxation Rates and Loading Efficiencies.

From the measured R1 and R2 of the cell suspensions we calculated the apparent R1 and R2 of the intracellular medium using the following formula:

Where the R (solution) is the is the R1 or R2 of the suspending solution without cells and is the volume fraction of cells in suspension given by

Where Cell conc. is the number of blood cells per ml of suspension and Vblood cell is the average volume of a blood cell taken as 9 x 10-11 mL.

Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods

The loading efficiencies were calculated by dividing the apparent R1 and R2 of the intracellular medium by the R1 and R2 of the incubating solutions including any dilutions to adjust the osmolarity, respectively. A value of 1 for loading efficiency corresponds to an apparent relaxation rate for the intracellular medium that is the same as the incubating solution.

5.2.3.6 Transmission Electron Microscopy (TEM)

The nanoparticle size distribution of the stock suspensions of the four different magnetic iron oxide nanoparticle systems (three different polymer coated systems and one uncoated system) were measured using TEM. Each sample was prepared by drying a drop of nanoparticle suspension on a carbon-coated copper grid.

The cell suspensions were prepared for transmission electron microscopy (TEM) in the same method showing in chapter 4 section 4.2.3. Particle size distributions within the RBCs and outside the RBCs were measured by analysing at least 20 iron maps, each with a field of view of 5 × 5 μm. Freeware ImageJ (v 1.42q, National Institute of Health) was used to measure the intracellular size distribution for particles/aggregates specifically containing iron. The intracellular areas in each field of view, as well as the intracellular size distribution for particles/aggregates specifically containing iron were measured. Each particle/aggregate was fitted with an ellipse and its effective size d was calculated as 2√(ab),where a and b are the semi-minor and semi-major axes of the fitted ellipse, respectively.

Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods

5.3 Results

5.3.1 TEM and DLS of nanoparticles

Figure 5.4a: TEM of Fe3O4 coated with PEO

Figure: 5.4b TEM of Fe3O4 coated with Diblock

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Figure 5.4 c:TEM of Fe3O4 coated with Triblock.

Figure 5.4d: TEM of uncoated Fe2O3. Figure 5.4: TEM images of the four different magnetic iron oxide nanoparticles systems.

Figure 5.4a shows a TEM image of iron oxide nanoparticles coated with poly(ethylene oxide) (PEO) prepared directly by casting them on the TEM grid. This figure demonstrates that the majority of particles are roughly spherical with particle size ranging from 3 - 24

Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods nm. DLS measurements on the suspensions indicated the particles Z weighted average size is 15 nm with PDI at 1.28.

Figure 5.4b shows a TEM image of iron oxide nanoparticles coated with the diblock copolymer of poly(ethylene oxide) (PEO) and poly(acrylic acid) (PAA). This figure demonstrates that the majority of particles are roughly spherical with particle size ranging from 5-25 nm. DLS measurements on the suspensions indicated that the particles Z weighted average size is 30 nm with PDI at 1.3

Figure 5.4c shows a TEM image of iron oxide nanoparticles coated with the triblock copolymer of poly(ethylene oxide) (PEO) and poly(propylene oxide) (PPO) prepared directly by casting them on the TEM grid. This figure demonstrates that the majority of particles are roughly spherical with particle size ranging from 4-30 nm. DLS measurements on the suspensions indicated that the particles Z weighted average size is 50nm with PDI at 1.28.

Figure 5.4d shows a TEM image of the uncoated γ-Fe2O3 nanoparticles prepared directly from a suspension in distilled water by casting them on the TEM grid. This figure demonstrates that the majority of particles are roughly spherical with the smallest particles being approximately 10 nm in size. It also shows that the particles aggregated. DLS measurements on the suspensions indicated that the particles Z weighted average size is 87nm with PDI at 1.3 demonstrating that stable aggregates were present in the aqueous suspension as well as being observed in the dry state under TEM . Figure 5.5 shows intensity-weighted particles size distributions derived from the DLS measurements for each type of nanoparticles suspension.

Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods

Figure 5.5: Dynamic Light Scattering (DLS) of the four different magnetic iron oxide nanoparticle systems.

Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods

5.3.2 Proton Relaxometry

Figure 5.6a: RBC specific longitudinal proton relaxivity (r1) for RBCs exposed to the four different magnetic iron oxide nanoparticle systems under osmotic pulse and osmotic incubation conditions at 200mOsm together with values for RBCs submitted to the same conditions without exposure to iron oxide nanoparticles (Control RBCs).

Figure 5.6a shows the cell specific longitudinal proton relaxivities (r1) for RBCs loaded with the four different magnetic iron oxide nanoparticle systems (three different polymer coated systems and one uncoated system) and control RBCs. The greatest relaxivity was obtained for the RBCs loaded with Fe3O4 coated with diblock copolymer using the osmotic pulse method. The next greatest relaxivity was obtained with osmotic pulse method for the

RBCs loaded with Fe3O4 coated with triblock copolymer. Only modest rises in relaxivity were observed for RBCs incubated with Fe3O4 coated with PEO by the two loading methods. The r1 of the RBCs loaded by osmotic incubation with the uncoated Fe2O3 was approximately the same as that for the RBCs loaded with uncoated Fe2O3 by osmotic pulse.

The r1 for RBC controls incubated at 200 mOsm was approximately the same for both preparation methods. 96

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Figure 5.6b: Ratios of RBC specific longitudinal relaxivity for treated cells to RBC specific longitudinal relaxivity for native cells.

Figure 5.6b shows the ratios of r1 for RBCs loaded with the four different magnetic iron oxide nanoparticle systems to r1 for native RBCs. The figure indicates that loading the RBCs by the osmotic pulse method with the diblock, triblock, and PEO coated, and the uncoated nanoparticles yielded 10.9, 6.8, 2.6 and 2.2-fold increases in longitudinal proton relaxation rate respectively when compared with native cells. The hypo-osmolar incubation method at 200 mOsm yielded r1 relaxivity increases less than 2-fold for all the three different polymer coated systems, and an increase of 2.6-fold for when using the uncoated particles.

Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods

Figure 5.7a: RBC specific transverse proton relaxivity (r2) for RBCs exposed to the four different magnetic iron oxide nanoparticle systems under osmotic pulse and osmotic incubation conditions at 200 mOsm together with values for RBCs submitted to the same conditions without exposure to iron oxide nanoparticles (Control RBCs).

Figure 5.7a shows the cell specific transverse proton relaxivities (r2) for RBCs loaded with the four different magnetic iron oxide nanoparticle systems by the two different methods together with relaxivities for the control RBCs. The greatest relaxivity was obtained for the RBCs loaded with iron oxide nanoparticles coated with the triblock copolymer using the osmotic pulse method. The next greatest relaxivity was obtained with osmotic incubation at 200 mOsm for the RBCs loaded with uncoated iron oxide nanoparticles followed by RBCs loaded with nanoparticles coated with the diblock copolymer using the osmotic pulse method. Only modest rises in relaxivity were found for RBCs loaded with the PEO, diblock and triblock coated nanoparticles using the hypo-osmolar incubation method at 200 mOsm.

Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods

Figure 5.7b: Ratios of RBC specific transverse relaxivity for treated cells to RBC specific transverse relaxivity for native cells.

Figure 5.7b shows the ratios of r2 for nanoparticles loaded RBCs to r2 for native RBCs for the four different magnetic iron oxide nanoparticle systems and the two different loading methods. The figure indicates that the RBCs loaded with three iron oxide nanoparticle systems coated with PEO, diblock, triblock copolymer and the uncoated systems by osmotic pulse method yielded 5, 22.6, 31.8 and 10.2-fold increases in transverse relaxivity respectively. Using the hypo-osmolar incubation method at 200 mOsm, increases in transverse relaxivity of 1.3, 1.7, 3.8 and 27.7-fold were observed respectively.

Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods

Figure 5.8: Transverse to longitudinal proton relaxation rate ratios (R2/R1) of RBCs loaded with the four different magnetic iron oxide nanoparticles systems under osmotic pulse and osmotic incubation conditions at 200mOsm together with the R2/R1 ratios for the stock nanoparticle suspensions.

Figure 5.8 shows the R2/R1 ratios for suspensions of RBCs loaded with the four different magnetic iron oxide nanoparticle systems by the two different methods together with the

R2/R1 ratios for the stock nanoparticle suspensions. All of the RBCs suspensions have

R2/R1 ratios elevated above 1 and were of a similar magnitude to the stock suspensions.

5.3.3 Apparent loading efficiency

Figure 5.9 shows the apparent loading efficiency based on both the apparent intracellular R1 and R2 for RBCs. The figure indicates that the loading efficiency of RBCs calculated from

R2 is generally higher than the loading efficiency calculated from R1. The loading

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Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods efficiencies for the polymer coated particles were higher using the osmotic pulse method. While for the uncoated particles the loading efficiencies were higher in the osmotic incubation method.

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5.3.4 TEM of RBCs

(a) (b)

Figure 5.10: (a)TEM of unstained human red blood cell after incubation with γ-Fe2O3 nanoparticles by osmotic incubation method at 200 mOsm and (b) iron map corresponding to region imaged in (a) generated from electron energy loss imaging.

(a) (b) Figure 5.11: (a)TEM of unstained human red blood cell after incubation by osmotic pulse method with γ-Fe2O3 nanoparticles at and (b) iron map corresponding to region imaged in (a) generated from electron energy loss imaging.

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(a) (b) Figure 5.12: (a) TEM of unstained control human red blood cell after incubation at 200 mOsm with no Fe2O3 nanoparticles and (b) iron map corresponding to region imaged in (a) generated from electron energy loss imaging.

Figure 5.10 is a transmission electron micrograph through an unstained section of a RBC after its incubation with uncoated -Fe2O3 nanoparticles by the osmotic incubation method at 200 mOsm. The corresponding iron map image is shown in Figure 5.10b. The iron map image shows the presence of large iron oxide nanoparticle aggregates both inside and outside the cell. Similar images were observed for cells incubated with -Fe2O3 nanoparticles by the osmotic pulse method (Figure 5.11). For comparison, TEM and iron- map images of an unstained section through a control RBC are shown in Figure 5.12. While electron dense particles can be seen in the TEM image, the iron map indicates the absence of iron-containing magnetic nanoparticles. Figure 5.13 shows a transmission electron micrograph of an unstained section of a RBC after its preparation with iron oxide nanoparticles coated with diblock copolymer (PEO- PAA) using the osmotic pulse method. The corresponding iron map image is shown in Figure 5.13b. All the electron dense features seen in Figure 5.13(a) appear to have higher intensities in the iron map in Figure 5.13(b) suggesting that a large number of iron oxide nanoparticles loaded into the RBC. However, most other RBC sections did not show densities of particles as high as this particular cell shown in Figure 5.13. Example images

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Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods for cells prepared with iron oxide nanoparticles coated with diblock copolymer (PEO-PAA) using the osmotic incubation method are shown in (Figure 5.14). Only a few iron oxide clusters are observed in the RBC sections.

(a) (b) Figure 5.13: (a)TEM of unstained human red blood cell after incubation with diblock PEO- PAA polymer coated system nanoparticles by osmotic pulse incubation method (b) iron map corresponding to region imaged in (a) generated from electron energy loss imaging.

(a) (b) Figure 5.14: (a)TEM of unstained human red blood cell after incubation with diblock PEO- PAA polymer coated system nanoparticles by osmotic incubation method at 200 mOsm and

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(b) iron map corresponding to region imaged in (a) generated from electron energy loss imaging.

(a) (b) Figure 5.15: (a)TEM of unstained human red blood cell after incubation with PEO polymer coated system nanoparticles by osmotic incubation method at 200 mOsm and (b) iron map corresponding to region imaged in (a) generated from electron energy loss imaging.

(a) (b) Figure 5.16: (a)TEM of unstained human red blood cell after incubation with PEO polymer coated system nanoparticles by osmotic pulse incubation method(b) iron map corresponding to region imaged in (a) generated from electron energy loss imaging.

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Figure 5.15 is a transmission electron micrograph of an unstained section through a RBC after its incubation with iron oxide nanoparticles coated with PEO polymer using the osmotic incubation method. The corresponding iron map image is shown in Figure 5.15b. Some of the electron dense features seen in Figure 5.15(a) appear to have higher intensities in the iron map in Figure 5.15(b) suggesting that some iron oxide nanoparticles loaded into the RBCs. Similar images (Figure 5.16) were observed for cells prepared with iron oxide nanoparticles coated with polymer PEO using the osmotic pulse method. Figure 5.17 is a transmission electron micrograph of an unstained section through a RBC after its incubation with iron oxide nanoparticle coated with triblock copolymer (PEO-PPO- PEO) by osmotic pulse method. The corresponding iron map image is shown in Figure 5.17(b). The electron dense features seen in Figure 5.17(a) appear to have higher intensities in the iron map in Figure 5.17(b) suggesting that some iron oxide nanoparticles loaded into the RBCs. Similar images (Figure 5.18) were observed for cells prepared with iron oxide nanoparticles coated with triblock polymer using the osmotic incubation method.

(a) (b) Figure 5.17 : (a)TEM of unstained human red blood cell after incubation with triblock polymer coated system nanoparticles by osmotic pulse incubation method(b) iron map corresponding to region imaged in (a) generated from electron energy loss imaging.

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(a) (b) Figure 5.18 : (a)TEM of unstained human red blood cell after incubation with triblock polymer coated system nanoparticles by osmotic incubation method at 200 mOsm. (b) iron map corresponding to region imaged in (a) generated from electron energy loss imaging.

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5.3.5 Nanoparticles size distribution inside cells

Figure 5.19 a: The size distribution of the two different magnetic (Mt) iron oxide nanoparticle coated with PEO and diblock copolymer systems loaded into RBCs together with size distribution of the stock nanoparticle suspensions. 108

Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods

Figure 5.19 b :The size distribution of the magnetic (Mt) iron oxide nanoparticles coated with triblock copolymer system and uncoated Maghemite (Mh) iron oxide nanoparticles loaded into RBCs together with size distribution of the stock nanoparticle suspensions.

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Figure 5.19a and 5.19b show the distributions of high intensity particle/aggregate sizes found in the iron map images inside the RBCs. The distributions were approximately log-normal in character. The largest clusters were found in RBCs loaded with uncoated nanoparticles by the osmotic incubation method at 200 mOsm approximately (up to 120 nm). The osmotic incubation method tended to result in size distributions shifted to larger particle sizes than those found for the osmotic pulse method. The aggregates within cells also tended to have particle size distributions shifted to higher sizes than those for the stock particle suspensions.

The median projected cell area

Figure 5.20 shows the median projected cell area with interquartile ranges (boxes) and maximum and minimum values for native RBCs and RBCs prepared by osmotic incubation and osmotic pulse methods. Figure 5.20 demonstrates that nanoparticle-loaded cells had significantly smaller mean projected areas in the optical microscope images compared with native cells. The osmotic pulse technique resulted in cells that were significantly smaller than corresponding cells prepared by the osmotic incubation method.

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Figure 5.20 Median projected cell area (horizontal lines in middle of boxes) with interquartile ranges (boxes) and maximum and minimum values (whiskers) for native RBCs and RBCs prepared by osmotic incubation and osmotic pulse methods

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5.4 Discussion

Previous studies of the incorporation of magnetic nanoparticles into RBCs, including those in Chapter 4, have shown that it is possible to load particles into the cells by inducing pore formation through osmotic pressure [52, 56, 97]. These previous studies used only one method for loading particles (either citrate coated [56] or carboxydextran coated and silica coated[97]), with no elemental mapping studies, no particle size distribution analysis, and no determinations of loading efficiencies. This study aimed to compare the effectiveness of magnetic resonance contrast enhancement by RBCs loaded with four different magnetic iron oxide nanoparticle systems (three different polymer coated systems and one uncoated system) by two different methods.

The data in Figures 5.6 and 5.7 indicate that the longitudinal and transverse proton relaxivities of RBCs can be enhanced by incubation in the presence any of the four different magnetic iron oxide nanoparticle systems under hypo-osmolar incubation at 200 mOsm and osmotic pulse conditions. For the polymer coated nanoparticle systems it was found that the osmotic pulse method of loading resulted in higher longitudinal (r1) and transverse (r2) cell specific proton relaxivities, respectively, while for the uncoated nanoparticle system the hypo-osmolar incubation method was found to yield higher cell specific proton relaxivities. These observations are qualitatively consistent with the hypothesis that during the osmotic pulse incubation the pore in the cell membrane does not open to the maximum size in the very short time of the pulse and therefore only smaller nanoparticles can easily enter into the RBCs, while during the hypo-osmolar incubation at 200 mOsm the pore size can open to a maximum of approximately 120 nm [52, 102] thus allowing larger clusters (as found in the uncoated particles) to enter the cells.

The apparent loading efficiencies for the RBCs loaded with uncoated nanoparticles by the osmotic incubation method are very high (Figure 5.9). However the TEM images of the loaded cells (Figure 5.10) showed that there was a significant fraction of the iron oxide particles outside the cells. This suggests that the washing sequence to remove excess particles following loading was ineffective. It appears that the uncoated particles were 112

Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods colloidally unstable in the saline solution used for washing. This leads to the formation of aggregates in the wash, and probably incubation, solutions. These aggregates spun down with the cells in the centrifuge and hence cannot be effectively removed. A similar behavior was observed for the pulse method (Figure 5.11) although in this case the initial removal of excess particles was more effective as there is no saline used in the first step and hence the aggregate formation occurs only in later washes. The apparent intracellular medium proton relaxation rates for the uncoated nanoparticles are therefore likely to be higher than would be observed for samples with effective washing and hence can not be relied upon for interpretation of loading efficiency.

The polymer coated samples are completely stable in saline solutions and so washing of cells following loading with these particles was always effective. No sign of magnetic nanoparticles on the outside of cells was observed for any of the polymer coated particles. Hence the apparent loading efficiencies calculated for these particles likely reflects the true loading efficiencies. As such, it appears that the diblock copolymer, in particular, dramatically improves loading efficiency. The mechanism by which this occurs may be related to the interaction of the polymer with the cell membranes either through electrostatic or entropic forces. Further research is required to elucidate this mechanism.

Generally the TEM images showed dispersed clusters of nanoparticles within the cells at rather low number concentrations. Often no particles were observed within a particular section through a cell. A notable exception was the TEM and iron map image of a RBC loaded with diblock polymer coated magnetic nanoparticles using the osmotic pulse method (Figure 5.13). While the RBCs prepared with diblock polymer coated nanoparticles by the osmotic pulse method did give some of the highest cell specific proton relaxivities, they were not sufficiently higher than other preparations to account for the very high density of particles seen in Figure 5.13. It is of note that the RBC section in Figure 5.13 is rather small suggesting that it is a section through an extremity of the cell volume. Hence we speculate that during the sample preparation procedure for TEM analysis, the centrifugation step may cause any iron oxide nanoparticles within the cell to be concentrated in the lower extremity of the cell. Such phenomena would also explain the apparent accumulation of 113

Chapter 5 Loading Erythrocytes with Magnetic Nanoparticles: a Comparison of Osmotic Pulse and Hypo-Osmolar Incubation Methods particles on the interior of RBC surface figure 5.13b. Figure 5.13 could then be explained in terms of a rare but feasible chance of sectioning through the lower extremity of the cell. As such, quantitative use of number densities of nanoparticle aggregates within cells for comparison with cell specific proton relaxivities should be treated with caution since densities of particles may be somewhat higher than generally indicated in the TEM images.

The cell specific transverse relaxivity for the RBCs loaded with uncoated nanoparticles by the osmotic incubation method appeared to be comparatively high (Figures 5.7a and 5.7b), it should be noted that because of the lack of stabilizing coating, it was not possible to effectively wash the cells after preparation because (unlike the polymer coated particles) nanoparticle clusters were spun down together with the cells during the washing process. As such, nanoparticle clusters are present in the suspending medium of the cells as well as within the cells as seen in Figure 5.10 and would contribute to the measured transverse relaxation rates hence confounding the calculation of cell specific relaxivities. The cell specific transverse relaxivities presented in Figure 5.7 for the uncoated particles are therefore likely to be higher than would be observed for samples with effective washing (as seen for the polymer coated nanoparticle preparations).

The smaller cell sizes observed for cells prepared by the osmotic pulse method when compared to the corresponding cells prepared by the osmotic incubation method (Figure 5.20) may be due to a larger outflux of hemoglobin during the period of lowered osmolarity (as observed in our study loading RBCs with gadolinium [52] – see Chapter 6).

5.5 Conclusion

Enhancement of cell specific proton relaxivities by loading RBCs with magnetic nanopartciles was most efficient with the diblock and triblock polymer coated nanoparticles loaded by the osmotic pulse technique. Proton transverse relaxation rates up to 31 times higher than native red blood cells were achieved (using the triblock copolymer coated iron oxide nanoparticles introduced with the osmotic pulse technique).

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6.1 Introduction

Chelated gadolinium compounds have been used as contrast agents in medical magnetic resonance imaging (MRI) because of their paramagnetic properties which enhance the contrast between different tissues by altering the local proton relaxation times [36]. The effects on MRI are increased longitudinal and transverse proton relaxation rates when the chelated gadolinium compounds are suspended in the blood stream. However, the chelated gadolinium compounds rapidly clear from the blood stream via renal excretion [103, 104]. The compounds also diffuse into extra-cellular spaces in tissues making use of the agents to measure perfusion or other blood pool measurements problematic [57].

In order to extend the lifetime of the contrast agent in the blood stream and prevent it diffusing into tissues, it has been proposed and demonstrated that encapsulating gadoteric acid (C16H25GdN4O8) within red blood cells (RBCs) will slow clearance by the kidneys [57]. A key strategy for loading RBCs with gadoteric acid is to incubate them in the presence of the gadoteric acid under hypo-osmolar conditions in order that pores open up in the cell membrane to allow the gadolinium chelates to enter the cell. The RBCs can be re- sealed by bringing the osmolarity of the medium back up to physiological values.

The aim of this study was to assess the performance of two different methods of loading RBCs with gadoteric acid. The first method involved incubating the RBCs in the presence of gadoteric acid under hypo-osmolar conditions. The RBCs were subsequently re-sealed by bringing the osmolarity of the medium back up to physiological values (300 mOsm) [52]. The second method involved preparing the RBCs in a hyper-osmolar solution of dimethyl sulfoxide (DMSO) which diffuses into the cells thus preventing an osmotic pressure difference across the cell membrane. The cells were then rapidly flushed with a slightly hypo-osmolar solution containing the gadoteric acid to induce an osmotic pressure pulse temporarily opening pores in the cell membrane [57]. While previous studies have investigated the enhancement of the relaxivity of gadolinium loaded RBCs [57], it has not been clear whether all RBCs load with gadolinium during the preparation process or whether it is only older cells with more fragile membranes that form

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6.2 Methods and Materials

6.2.1 Preparation of Stock Solutions and Cell Suspensions

6.2.1.1 Gadoteric Acid Solution

We obtained clinical grade gadoteric acid (C16H25GdN4O8 (279.32 mg/mL)) from Guerbet (France). In order to produce solutions with an osmolarity of 300 mOsm, 3.5 mL of distilled water was added to 1 mL of the as received C16H25GdN4O8 solution. Using a vortex mixer, the solution was vigorously shaken for two minutes.

6.2.1.2 Human Red Blood Cells (RBCs)

6.2.1.3 Hemoglobin Solution

Hemoglobin solution was prepared by lysing 1mL of the stock RBC suspension. RBCs were lysed by sonication using an ultrasonic homogenizer (model - Biologics 3000) followed by centrifugation at 3000 g at 4°C for 10 mins to remove cell debris. A series of hemoglobin solutions with different concentrations was obtained by diluting the stock hemoglobin solution with PBS.

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6.2.2 Methods of Loading Cells with Gadolinium

6.2.2.1 Hypo-osmolar Incubation Method

RBCs were incubated in the presence of gadoteric acid using the following procedure. X mL of distilled water (where X is 2, 1, 0.4 or 0) was added to a solution made up from 1 mL of gadoteric acid at 300 mOsm and 1 mL of RBC suspension at 300 mOsm to produce a suspension of RBCs and gadoteric acid at 150, 200, 250 or 300 mOsm respectively. The RBCs were then incubated with gadoteric acid at 4°C for 2 hours (or a specified number of hours up to 24 in a study of the effect of incubation time on the loading method) with gentle shaking in a vortex mixer. After the incubation period, the osmolarity of the suspension was brought back up to 300 mOsm by adding Z mL of PBS at 400 mOsm (where Z is 6, 3, 1.2 or 0) to the suspension of cells and gadoteric acid. The final suspension was incubated at 37C for 1 hour and was then washed with PBS 5 times before returning to 4C for storage before measurements or further preparation. Control samples were prepared with the above protocol using PBS in place of gadoteric acid.

6.2.2.2 Osmotic Pulse Method

1 mL of 2M dimethyl sulfoxide (DMSO, formula (CH3)2SO) was added to 2 mL of RBCs suspended at 300 mOsm to produce a suspension of RBCs and DMSO. The RBCs were then incubated with DMSO for less than 1 min. This cell suspension was then centrifuged for 5 min at 3000 g. The supernatant was discarded and the RBC pellet was re-suspended in 3 mL of gadoteric acid at 300 mOsm for 1 min. The moment the gadoteric acid solution is added, there is a short osmotic pulse caused by the sudden dilution of the extracellular DMSO [57]. The resulting cell suspension was centrifuged for 5 min at 3000 g. The supernatant of this solution was also discarded. In order to help the cell membrane pores close, 1 mL of dextran (1.32 mM) was added to the RBC pellet. The cell suspension was incubated at 37C for 1 hour and was then washed with PBS 5 times before returning to 4C for storage before measurements or further preparation. Control samples were prepared using the above protocol with PBS in place of gadoteric acid.

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6.2.3 Measurement and Characterization of Cell Suspension

6.2.3.1 Cell Counting

Concentrations of RBCs in suspensions were measured using a hemocytometer (model - BrightLine, Hausser Scientific, Horsham , USA). Digital optical microscope images of the hemocytometer cells were captured then analysed using ImageJ (v 1.42q, National Institute of Health) to count the cells.

6.2.3.2 Cell Size and Morphology

The digital images acquired from the optical microscope were used to assess the morphology of the RBCs. The effect of loading of gadolinium into the RBCs on cell size was assessed by measuring the distribution of projected cell areas in the digital images of the cells in the hemocytometer. The number of cells measured for each preparation ranged from 88 to 229.

6.2.3.3 Proton relaxometry

Proton relaxation rate measurements were made on a Bruker mq-60 MHz Minispec NMR Analyzer (Bruker, Ettlingen, Germany) operating at a magnetic field of 1.4 T. Cell suspensions were prepared for relaxometry with a haematocrit of approximately 50%. Samples were placed in an aluminium block within a water bath at 37.5 °C for at least 15 minutes prior to relaxometry measurements. Proton transverse relaxation times (T2) were obtained from fitting a monoexponential decay curve to signal data generated by a Carr- Purcell-Meiboom-Gill (CPMG) spin-echo pulse sequence with an echo spacing of 1 ms and a repetition time of 5s. Longitudinal relaxation times (T1) were obtained from fitting a monoexponential recovery curve to signal data generated with an inversion recovery (IR) pulse sequence using 10 logarithmically spaced inversion times between 10 and 10000 ms.

R1 (R1=1/T1) and R2 (R2=1/T2 ) data are reported in this paper. 500 µL of cell suspension was used for each relaxometry measurement. For some investigations, 500 µL of

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Chapter 6 Enhancement of the Cell Specific Proton Relaxivities of Human Red Blood Cells via Loading with Gadoteric Acid supernatant from the first wash of the cells was taken separately for proton relaxometry measurement. Standard curves of R1 and R2 proton relaxation rates were also made using 500 µL samples of haemoglobin solutions at different dilutions ranging from iron concentrations of 2mM to 10 mM. Cell specific proton relaxivities were calculated for each cell suspension by dividing the measured relaxation rate (R1 or R2) by the number of cells per unit volume.

6.2.3.4 Transmission Electron Microscopy

Cell suspensions were prepared for transmission electron microscopy (TEM) in the following way. The suspensions were centrifuged and the cell pellet was re-suspended in PBS at a volume ratio of 1:1. The suspension was then fixed with glutaraldehyde at a volume ratio of 1:20. After fixation, the cells were serially dehydrated for 40 seconds each time in consecutive solutions of 50, 70, 95 and 100% ethanol. The sample was dried by washing it twice in 100% acetone. After dehydration, the sample was infiltrated with Spurrs resin for 3 minutes each in consecutive solutions of acetone and Spurrs resin at volume ratios of 3:1, 1:2 and 1:3 respectively. The sample was bonded in 100% Spurrs resin followed by polymerization overnight at 70°C. An ultramicrotome was used to cut 100-nm thick sections from the resin block. The sections were then mounted on carbon- coated copper grids.

Gadolinium jump ratio images were generated by acquiring energy-selected images just below (pre-edge) and above (post-edge) the gadolinium M-edge using a Gatan Tridiem energy filter at 200 kV. The pre-edge image was obtained at 130 eV and the post-edge image at 155 eV with both images acquired using a slit-width of 10 eV. The jump ratio image was obtained by dividing the post-edge image by the pre-edge image [81]. Jump ratio images were used to determine the spatial distribution of gadolinium in the samples since gadolinium-rich regions of the sample will result in a higher intensity in the jump ratio image.

Gadolinium M-edge jump ratio images of thin sections through resin embedded red blood

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Chapter 6 Enhancement of the Cell Specific Proton Relaxivities of Human Red Blood Cells via Loading with Gadoteric Acid cells generally showed a uniform intensity inside the cell (see later). The mean fractional intensity difference between the intensity within the cells and the intensity in the surrounding resin was measured for each sample of cells by digital image analysis. Multiple regions of interest (n  30) were selected that avoided any intensity hot spots (i.e. apparent gadolinium precipitates or clusters) both inside the cells and within the resin background. Regions of interest were selected from multiple gadolinium M-edge jump ratio fields of view (5 x 5 m) from each sample of cells.

6.3 Results

6.3.1 Proton Relaxometry

Figure 6.1a: RBC-specific longitudinal proton relaxivities (r1) for RBCs incubated at 200 mOsm in the presence of gadoteric acid (Gd-RBC) and in the absence of gadoteric acid (Con- RBC) and for native RBCs. The data for the Gd-RBC were acquired at different incubation times ranging from 1 to 18 hours but no correlation between incubation time and relaxivity was found.

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Figure 6.1b: RBC-specific transverse proton relaxivities (r2) for RBCs incubated at 200 mOsm in the presence of gadoteric acid (Gd-RBC) and in the absence of gadoteric acid (Con- RBC) and for native RBCs. The data for the Gd-RBC were acquired at different incubation times ranging from 1 to 18 hours but no correlation between incubation time and relaxivity was found.

No systematic change of either the longitudinal or transverse proton relaxivities of the RBCs with variation in the incubation time (over the range 1 to 18 hours) using the hypo- osmolar incubation method was found. Figures 6.1a and 6.1b show the longitudinal (r1) and transverse (r2) proton relaxivities of RBCs incubated with gadoteric acid (Gd-RBC) at 200 mOsm using the hypo-osmolar incubation method, control RBCs incubated at 200 mOsm, and native RBCs. The r1 and r2 values for the control and native RBCs are similar and much lower than those for the RBCs incubated with gadolinium at 200 mOsm. The mean r1 and mean r2 for the native RBCs are, however, slightly but significantly higher

(p<0.0001) than the mean r1 and r2 values for the control RBCs. No systematic change of either the longitudinal or transverse proton relaxation rates of the first wash of the control RBCs with variations in incubation time was found. By comparing the relaxivities of the hemoglobin solution with the proton relaxation rates of the first wash of the control RBCs, it was estimated that approximately 20% of the hemoglobin is released from the cells during the hypo-osmolar incubation.

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Figures 6.2a and 6.2b show the longitudinal (r1) and transverse (r2) proton relaxivities, respectively, for gadolinium-loaded and control RBCs prepared under different conditions. The greatest longitudinal and transverse relaxivities were obtained for the RBCs loaded with gadolinium using the osmotic pulse method followed by those for RBCs loaded with gadolinium by the osmotic incubation method at 150 mOsm. There were only modest rises in relaxivity for RBCs incubated at 200 and 250 mOsm (Figure 6.2a, 6.2b). The r1 and r2 values for RBCs incubated at 300 mOsm were approximately the same as those for all of the control RBC suspensions.

Figure 6.3a: Ratios of RBC-specific longitudinal relaxivity for treated cells to RBC-specific longitudinal relaxivity for native cells. Bars indicate the values obtained for a single preparation.

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Figure 6.3b: Ratios of RBC-specific transverse relaxivity for treated cells to RBC-specific transverse relaxivity for native cells. Bars indicate the values obtained for a single preparation.

Figures 6.3a and 6.3b show the ratios of r1 and r2 for gadolinium-loaded cells to r1 and r2 for native RBCs. The osmotic pulse method yields a 71-fold increase in longitudinal proton relaxation rate and a 39-fold increase in transverse proton relaxation rate compared with native cells. Incubation with gadolinium at 150 mOsm using the hypo-osmolar method produces a 58-fold increase in r1 and a 16-fold increase in r2, while the incubation with gadolinium at 300 mOsm yields a ratio of 0.8 for r1 and 0.7 for r2 indicating a slight decrease in relaxivities compared to native RBCs.

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Figure 6.4: Transverse to longitudinal proton relaxation rate ratios (R2/R1) of RBCs loaded with gadoteric acid under different conditions together with the R2/R1 ratio for gadoteric acid solution. Bars indicate the values obtained for a single preparation.

The R2/R1 ratios for suspensions of RBCs loaded with gadolinium by the different methods compared with the R2/R1 ratio for gadoteric acid solution is shown in Figure 6.4. While the gadoteric acid solution has an R2/R1 ratio of approximately 1, all of the gadolinium-loaded

RBC suspensions have R2/R1 ratios elevated above 1.

6.3.2 Rates of survival of RBCs

Figure 6.5: Rates of survival of RBCs after different processing protocols. Bars indicate the values obtained for a single preparation. 125

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6.3.3 The median projected cell area

Figure 6.5 shows the rates at which RBCs survived the gadolinium loading process. Survival rates were about 75% for most methods with the exception of the 150 mOsm incubation method where survival dropped to only 37%. The optical microscope images of the RBCs indicated that all surviving cells had the typical morphology of native RBCs. However, all populations of surviving cells had significantly smaller mean projected areas in the optical microscope images (Figure 6.6) compared with native cells. The osmotic pulse technique resulted in cells that were significantly smaller than all cells prepared by the osmotic incubation method apart from those prepared at 150 mOsm.

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6.3.4 TEM of RBCs

(a) (b) Figure 6.7 : (a) TEM image of unstained human RBC after incubation with gadoteric acid at 200 mOsm and (b) gadolinium jump ratio image corresponding to the region imaged in (a) generated from electron energy loss imaging.

(a) (b) Figure 6.8: (a) TEM image of unstained control human RBC after incubation at 200 mOsm with no gadoteric acid. (b) Jump ratio image corresponding to the region imaged in (a) generated from electron energy loss imaging.

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A transmission electron micrograph through an unstained (i.e. no heavy metal staining) section of a RBC after its incubation with gadoteric acid at 200 mOsm is shown in Figure 6.7a. The corresponding gadolinium jump ratio image is shown in Figure 6.7b. The gadolinium jump ratio image shows a generally uniform signal within the cell that is higher in intensity than the surrounding resin. However, some of the electron dense features seen in Figure 6.7a appear to have somewhat higher intensities in the gadolinium jump ratio image in Figure 6.7b suggesting some aggregation or precipitation of gadolinium. Similar images were observed for cells incubated with gadolinium at 250 and 150 mOsm. The largest clusters found inside cells were approximately 57 nm (found in the cells incubated at 150 and 200 mOsm). Images for cells loaded with gadolinium by the osmotic pulse method were also similar with regard to the higher intensity within the cells relative to the resin background but showed no high intensity aggregates in the cells.

By comparison, a transmission electron micrograph and a gadolinium jump ratio image through an unstained section of a control RBC are shown in Figure 6.8. The signal intensity within the cell in the gadolinium jump ratio image is uniform and slightly lower than that observed in the surrounding resin. Similar images were observed for the RBCs incubated with gadolinium at 300 mOsm.

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6.3.5 Magnitudes of the gadolinium jump ratio image intensity

Figure 6.9: Magnitudes of the gadolinium jump ratio image intensity differences between cells and resin background.

Figure 6.9 shows the magnitudes of the intensity differences between the resin background and the internal regions of the cells (omitting any aggregates) in the Gadolinium jump ratio images. Cells exposed to gadolinium at low osmolarities showed higher intensities within the cell than outside the cell while cells exposed to gadolinium at 300 mOsm and control cells showed lower intensities within the cell than outside the cell.

6.4 Discussion and Conclusion

The data in Figures 6.1a and 6.1b indicate that the longitudinal and transverse proton relaxivities of RBCs can be enhanced by incubation in the presence of gadoteric acid under hypo-osmolar conditions. The reproducibility of the relaxivity enhancement with repeated 129

Chapter 6 Enhancement of the Cell Specific Proton Relaxivities of Human Red Blood Cells via Loading with Gadoteric Acid preparations is characterized by a coefficient of variation of approximately 38% for both r1 and r2. The very slightly higher relaxivities of the native RBCs relative to the control RBCs (Figures 6.1a and 6.1b) are presumably because of hemoglobin leakage from the control cells during the hypo-osmolar incubation.

Figures 6.2 and 6.3 suggest that exposure of RBCs to gadoteric acid at 300 mOsm does not result in any gadoteric acid entering the cells since no enhancement of the proton relaxivities of the RBCs is observed. As the osmolarity of the incubating medium is lowered, the degree of proton relaxation enhancement increases suggesting that greater cellular loadings of gadoteric acid are achieved at lower osmolarities. These observations are qualitatively consistent with the hypothesis that lowering the osmolarity of the RBC suspension results in pores opening up in the cell membranes thus allowing gadoteric acid to diffuse into the cells. However, it is not clear why higher loadings of gadoteric acid are achieved with lower osmolarities. One possible hypothesis is that the loading is related to the volume of hemoglobin released from the RBCs during the osmotic incubation. It appears that the majority of hemoglobin is not released from the cells even though pores are opened in the cell membranes. This conclusion is drawn from the observation that the washes from the control cells contain only about 20% of the total hemoglobin from the cells (when incubated at 200 mOsm). As such, the hemoglobin release may occur only during the initial drop in osmolarity where temporary osmolarity differences between the inside and outside of the cell may cause osmotic pressure induced flows ejecting hemoglobin. This concept is consistent with hemoglobin molecules in RBCs tending to have mutually attractive interactions [105]. The hypothesis that the gadoteric acid load is related to the fraction of hemoglobin ejected from the cell during the lowering of the osmolarity may also explain why the RBCs exposed to gadoteric acid during an osmotic pulse are observed to have the highest proton relaxivities and hence presumably the highest gadoteric acid loadings. The osmotic pulse method is likely to result in greater peak osmolarity differences between the inside and outside of the cells owing to the pulsed nature of the procedure and hence greater osmotic pressure induced flows out from the cells might be expected. As such, it may be expected that a greater fraction of the hemoglobin load is ejected from the cell compared to the slower osmolarity drops caused by the osmotic

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Chapter 6 Enhancement of the Cell Specific Proton Relaxivities of Human Red Blood Cells via Loading with Gadoteric Acid incubation method.

The longitudinal proton relaxivity enhancement of the RBCs prepared by the osmotic pulse method in this study (71-fold) was almost identical to that found by Johnson et al [57] (73- fold) using the same method. However, the transverse proton relaxivity enhancement of the RBCs prepared by the osmotic pulse method in this study (39-fold) was somewhat less than that achieved by Johnson et al (83-fold) [57]. This difference in R2 enhancement may be related to the variable and shorter echo spacing employed by these workers for measuring

R2. In particular, the very short echo spacing (0.1 ms) used for some of their measurements would increase R2 relative to our constant and longer echo spacing of 1 ms.

The observations that the R2/R1 ratios for all of the gadolinium-loaded RBCs were elevated above the ratio for gadoteric acid solution (Figure 6.4) suggests non-uniform distribution of the gadoteric acid within the cell suspensions. While some degree of enhancement of the

R2/R1 ratio is expected because of the compartmentalization of gadoteric acid within the cells, there may also be a contribution from precipitation of gadolinium into aggregates as seen in the transmission electron microscope images of the cells incubated with gadoteric acid under hypo-osmolar conditions (Figure 6.7) assuming that the aggregates are not an artifact of the sample preparation procedure for TEM.

The rate of cell survival for the 200 mOsm incubation method was 77% suggesting that the majority of the hemoglobin found in the washings from the cells (which accounted for approximately 20% of the hemoglobin in the cells at the start of the procedure) was due to completely lysed cells. This observation suggests that the volume of hemoglobin ejected from cells that survived the procedure was a very small fraction (perhaps only 1 or 2%) of the initial cell hemoglobin content and could explain why the apparent gadoteric acid loading at 200 mOsm was relatively small. Furthermore, the observation that the gadolinium loaded cells remain red colored is consistent with a significant amount of hemoglobin being retained within them. Although hemoglobin loss in the other preparations in this study was not measured, Johnson et al [57] measured a drop of about 15% in mean cell hemoglobin content on using the osmotic pulse method. Such a change in

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Chapter 6 Enhancement of the Cell Specific Proton Relaxivities of Human Red Blood Cells via Loading with Gadoteric Acid mean cell hemoglobin is somewhat greater than could be accounted for in our 200 mOsm incubation method and as such is consistent with the hypothesis that the osmotic pulse method results in a larger outflux of hemoglobin compared with the osmotic incubation methods and hence creates more intracellular space for gadoteric acid loading. It is noteworthy that, of all the preparation methods, the osmotic pulse method results in the lowest cell sizes (Figure 6.6).

The observation that all gadolinium jump ratio images of cells that demonstrated enhanced proton relaxivities exhibited higher internal signal than the resin background while those that did not demonstrate enhanced proton relaxivities (the control RBCs and RBCs incubated in the presence of gadoteric acid at 300 mOsm) exhibited lower internal signal than the resin background indicates that the gadolinium jump ratio signal within the cells is indicative of the presence or absence of uniformly distributed gadolinium. Furthermore, the observation that none of the imaged cells exposed to gadoteric acid under hypo-osmolar conditions yielded gadolinium jump ratio signal intensities less than the resin background suggests that the osmotic pulse and osmotic incubation methods result in all cells being loaded with gadoteric acid rather than only a fraction of the cells. Although there is not an obvious strong trend relating the intra-cell intensities in the gadolinium jump ratio images with the relaxivities of the cells, it should be noted that the jump ratio images are semi quantitative because slight variations in the thickness of sample sections may perturb the intensities.

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7.1 Introduction

Part of my research has focused on loading human red blood cells (RBCs) with iron oxide nanoparticles. The atomic force microscope (AFM) has been used in this work to image iron oxide nanoparticles and human RBCs loaded with nanoparticles to examine where the introduced nanoparticles are located in the cells. AFM is particularly useful because of its capability to generate precise topographic images of the sample by scanning the surface with a nanometer–scale probe, with lateral resolution around 1 nm and vertical resolution as low as 0.1 nm under controlled conditions [106]. The AFM technique also enables images of RBCs to be obtained in air and liquid media. In this chapter I will discuss some of my work on iron oxide nanoparticles and human RBCs loaded with these nanoparticles. I will discuss the different modes of AFM image acquisition useful for biophysical studies and propose an interpretation of the AFM images. I will also discuss magnetic force microscopy (MFM) as a natural extension of AFM with regards to imaging RBCs that have been loaded with iron oxide nanoparticles.

The first aim of this study was to determine where the nanoparticles are located in cells by comparing the AFM and the MFM images. In AFM scanning mode images are formed from the signal collected from the surface of the sample only. While the MFM scanning technique can probe deeper below the surface of the cells. Therefore we thought that it will be possible to observe if the nanoparticles are inside the cells or on the surface.

The second aim of this research was to demonstrate that RBC shape changes according to the level of osmotic pressure and that reversal of osmotic pressure causes cell recovery. We hypothesised that swelling of the cell causes openings to appear in the cell membrane which allow iron oxide nanoparticles to enter. Our second hypothesis was that the opening and closing of the membrane is reversible. However there have been no definitive studies to date.

I faced several difficulties with these two aims as described here which I tried hard to resolve but given the time restraints of my candidature I was not able to bring to a

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Chapter 7 Atomic Force Microscopy and Magnetic Force Microscopy of iron oxide nanoparticles and red blood cells loaded with iron oxide nanoparticles successful conclusion. I faced several problems in terms of scanning in liquid using glutaraldehyde. Furthermore, scanning in air did not answer the questions mentioned above due crystallising of the medium onto the cells. Therefore I will present in this chapter the techniques that I attempted and the results that I have achieved as an archive for future researchers who may attempt to improve on the methods

In 1929 Schmalz invented the Profiler using an optical lever arm to monitor the motion of a sharp probe mounted at the end of a cantilever. Russell Young and colleagues between 1965 and 1971 created the topografiner which combined the detection of tunneling current with a scanning device [107]. In 1980 Gerd Binnig and Heinrich Rohrer developed the modern scanning probe microscope (SPM) [108]. They earned the Nobel prize in physics in 1986 for their design of the scanning tunneling microscope [109]. This was followed by the development of the AFM by Gerd Binning, Calvin Quate and Cristoph Gerber[110].

The basic principle of an AFM is to scan a very sharp tip across the surface of a sample. The tip is mounted on a cantilever and the motion of the cantilever is used to generate an image. The scanning motion is usually performed with a piezo tube consisting of two parts, one for the movement in the x-y plane and the other for z movements. The atomic force microscope (AFM) uses a probe which is mounted on a small leaf spring (the cantilever) [111]. In AFM applications, a sharp tip is attached to the cantilever as shown in figure 7.1. A laser beam is directed onto the back of the cantilever and then reflected onto a 4-segment photodetector, which detects the motion of the cantilever. Depending on the application and the imaging mode, different parameters are used as feedback signals. As the tip is scanned in the x-y plane, the feedback signal is used to maintain a constant tip surface force which is achieved by moving the cantilever via the z-piezo, which moves the tip up and down following the surface profile as shown in figure 7.2. With the ability to position the tip within sub-nanometer precision, three -dimensional images of the sample surface with magnifications up to 10 million times can be taken. AFM can be operated in air or liquid solutions [106, 110, 111].

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Figure 7. 1: Standard tip mounted on a cantilever[112].

Figure 7.2: The basic principle of an AFM [113]

AFM can be operated in different modes based on interactions between the tip and the sample. In imaging operation, the feedback parameters generaly are set up to ensure a contact tip-sample force and the image is composed of the x,y and z coordinates of the scan.

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7.1.1 Atomic force imaging modes a. Contact mode

In contact mode, relatively soft cantilevers are used with a spring constant typically between 0.01 and 50 N/m [114]. The tip is in permanent contact with the sample surface and the deflection of the cantilever acts as a feedback signal. A contact force of the tip on the surface is maintained by keeping a constant cantilever deflection [111, 115]. Contact mode images can be taken in air, gas or even in liquids. Contact mode imaging allows high scan rates and can be used to obtain high resolution images of the surface lattice of samples, for example, crystals. However, due to the permanent contact between the tip and surface, lateral forces present during the scanning motion of the tip can lead to deformations of the surface of samples and or damage to tips resulting in loss of resolution over time. b. Tapping mode

In tapping mode imaging, the cantilever oscillates near its resonant frequency. The amplitude of this oscillation can be used as a feedback signal. The tip is only in contact with the sample at the lowest point during the oscillation cycle lightly taping on the surface. In tapping mode there are nearly no lateral forces present and therefore it is possible to image very soft and loosely held samples. Imaging is possible in air, gas and liquids. In order to overcome the adhesion between the tip and the sample, tapping mode imaging in air is performed with cantilevers with a higher spring constant than the ones used in liquid. Compared to contact mode, tapping mode imaging is often done with lower scan rates [115, 116]. c. Phase mode imaging

In phase mode imaging, the phase shift of the oscillating cantilever relative to the driving signal is measured. This phase shift can be correlated with specific material properties that

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Chapter 7 Atomic Force Microscopy and Magnetic Force Microscopy of iron oxide nanoparticles and red blood cells loaded with iron oxide nanoparticles affect the tip or sample interaction. The phase shift can be used to differentiate areas on a sample with such differing properties as friction, adhesion, and viscoelasticity as shown in figure 7.3. The techniques are used simultaneously with tapping mode, so topography is measured at the same time [115, 117].

Figure 7.3 : Phasing mode imaging[113].

7.1.2 Magnetic force microscopy (MFM)

The magnetic force microscopy (MFM) is similar to AFM imaging technique but in this case the force sensing probe is magnetic. MFM scanning is accomplished in a two-step process called lift mode or interleave mode. Firstly the tip is scanned over the sample to measure the local surface topography. This information is saved in the scanner memory as the standard scan. Secondly the tip is lifted at a constant height above the sample surface and a second scan is performed as shown in figure 7.4. This called the interleave scan.

Figure 7.4: Interleave mode scanning , the scanner performs a main trace and retrace, then lifts the tip to the lift scan height and performs the interleave trace and retrace[113]. 137

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In MFM scanning the tip is either composed or more often coated with magnetic materials. In interleave scan mode the magnetic field gradient causes a force on the tip that is measured to image the magnetic structure of the sample. Here the cantilever responds to interactions between the tip and the stray fields above the sample [118]. In standard scan the distance between the tip and the sample is around 1nm. Here the short range forces between the sample and the surface dominate such as Van der Waals attraction forces. The probe tip tracks the sample topography as in normal contact AFM. In interleave mode the tip is lifted away from the sample at certain height where the short range force between the sample and the surface are negligible. At this point the long range forces (magnetic force) dominate and the cantilevers responds to the magnetic forces only. This induces changes in the cantilevers resonance frequency or phase and the image obtained by the MFM result only from the convolution of the magnetic properties of both sample and the tip. It is very important to maintain a reasonable distance between the tip and the sample. If the distance between the tip and the sample is quite large then the magnetic signals are weaker and this decreases the MFM phase signal and reduces image resolution. Conversely, if the distance between the tip and the sample is too small, the phase images obtained by the MFM may result from the convolution of the long rang forces and the short range forces and hence the MFM images will contain topography artefacts. The magnetic force between the sample and tip is given by:

Where m is the magnetic moment of the tip as shown in figure 7.5, H is the stray magnetic field from the sample.

Figure 7.5: The MFM interleave scan and the magnetic of the sample surface. 138

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Because the magnetic stray field from the sample affects the magnetized state of the tip and vice versa, in most cases it is difficult to obtain quantitative information from the MFM measurement. To interpret the information quantitatively, the configuration of the tip must be known. MFM has a typical resolution of 30 nm [106] although resolutions as high as 10 nm are attainable [119].

7.2 Methods and materials

7.2.1 Preparation of Stock Solutions and Cell Suspensions

7.2.1.1 Iron Oxide Nanoparticles

Maghemite (-Fe2O3) nanoparticles were obtained from Sigma Aldrich. The stock -Fe2O3 suspension was prepared as described in Chapter 3

7.2.1.2 Human Red Blood Cells

Human red blood cells (RBCs): Human blood (approx.8mL) was collected into heparinised tubes and centrifuged at 3000 g at 4°C for 10 minutes to separate the RBCs from the plasma. The RBCs were washed three times with phosphate buffered saline (PBS, 300 mOsm, pH 8). A stock suspension of RBCs at 300 mOsm was prepared by suspending approximately 2 mL of packed RBCs in 10 mL of PBS.

7.2.1.3 Methods of Loading RBCs with Iron Oxide Nanoparticles

RBCs were loaded with -Fe2O3 nanoparticles at different osmolarity (as described in Chapter 4) under hypo-osmolar conditions which caused the RBCs to swell and pores to open up in the cell membranes. The RBCs were subsequently re-sealed by bringing the osmolarity of the medium back up to physiological values [52].

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7.2.2 Measurement and Characterization

7.2.2.1 AFM scanning

A Dimension 3010 (Digital Instruments) AFM was operated in contact and tapping mode to scan the RBCs dried onto pieces of mica at different osmolarities. All scans were performed at an ambient temperature.

7.2.2.2 MFM scanning

A magnetic tip (model MESP) was magnetized in a one Tesla external field using an electromagnet. In this study a non-magnetic probe holder for the AFM Dimension 3010 (Digital Instruments) was used to hold the magnetized tip. Interleave mode scans at different heights were used to ensure that the cantilever responses were predominately due to the magnetic forces. All scans were performed at ambient temperature. In this chapter the MFM images we obtain by two different scanning techniques. The first method used a magnetized tip and the phase images obtained by the MFM resulting only from magnetic field induced by the tip into the sample as shown in figure 7.6. In the second method a permanent magnet was placed under the sample to induce a large magnetic field in the sample as shown in figure 7.7.The phase image obtained by the MFM results from the convolution of the magnetic field induced by the tip and the permanent magnet.

Figure 7.6: MFM scanning with magnetized tip only.

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Figure 7.7: MFM scanning with a magnet under the samples.

7.2.2.3 Sample preparations

 -Fe2O3 nanoparticles: -Fe2O3 nanoparticles were prepared from a suspension in

distilled water as described in Chapter 4. A drop of the -Fe2O3 nanoparticles was dried onto a piece of freshly cleaved mica.

 RBCs on mica: RBCs incubated with -Fe2O3 nanoparticles at 200 mOsm were dried onto pieces of freshly cleaved mica.

 RBCs sections: Resin embedded human RBCs incubated with -Fe2O3 nanoparticles were prepared for the AFM as described in Chapter Four. An ultramicrotome was used to cut 100-nm thick sections from the resin block. The sections were then mounted on carbon-coated copper TEM grids.

 RBCs in phosphate buffered saline (PBS): Here freshly cleaved mica was plunged into a solution of poly-L-lysine for 2 hours. The RBCs were fixed by glutaraldehyde to make sure that the RBCs maintained their same shape when they were placed onto the mica. The mica and the microscope coverslips were rinsed with 1% aqueous glutaraldehyde solution and washed with PBS.

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7.3 Results

7.3.1 AFM and MFM of -Fe2O3 nanoparticles sample

Figure 7.8a : Topographic image of -Fe2O3 Figure 7.8b : Phase image of -Fe2O3 nanoparticles nanoparticles, lift at 20 nm (magnetized tip)

Figure 7.8c : Phase image of -Fe2O3 Figure 7.8d : Phase image of -Fe2O3 nanoparticles, lift at 30 nm nanoparticles, lift at 50 nm (magnetized tip) (magnetized tip)

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Figure 7.8f : Phase image of -Fe2O3 nanoparticles, lift at 100 nm (magnetized tip)

Figure 7.8a a non-contact AFM image and 7.8b,7.8 c ,7.8d, 7.8e and 7.8f the corresponding phase images of -Fe2O3 nanoparticles generated by the magnetized tip at different lift heights. Field of view 5 m × 5 m.

Figure 7.8a shows a topographic tapping mode AFM image of -Fe2O3 nanoparticles. 7.8b, 7.8c, 7.8d, 7.8e and 7.8f show the phase image corresponding to the region imaged in Figure 7.8a generated in MFM imaging mode at lift heights of 20, 30, 50 and 100 nm respectively. Here the phase image contrast was generated by the magnetic field induced in the sample by the magnetized tip itself. By comparing, the topographic image and the phase images it is possible to identify very weak magnetic signal from the nanoparticles in the phase image. However the signal is only slightly higher than that observed in the background. The best phase image resolution was obtained at lift height 20 nm. These images demonstrate that when the lift height increases the phase image signal decreases. When the lift height reaches 100 nm the magnetic signals intensity from the nanoparticles decreases to zero. The fact that the image persists to heights up to 30 to 50 nm indicates that the image acquired at 20nm is a magnetic signal and not an artefact of close contact with the sample. The higher lift height images enable the more magnetic parts of the topography image to be identified.

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7.3.2 MFM with a permanent magnet under the sample of -Fe2O3 nanoparticles

Figure 7.9a : Height image of -Fe2O3 Figure 7.9b : Phase image of -Fe2O3 nanoparticles nanoparticles, lift at 50 nm (magnet under the sample) ( magnet under the sample)

Figure 7.9c : Phase image of -Fe2O3 Figure 7.9d : Phase image of -Fe2O3 nanoparticles, lift at 100nm nanoparticles, lift at 150nm (magnet under the sample) (magnet under the sample)

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Figure 7.9e : Phase image of -Fe2O3 nanoparticles Figure 7.9f : Phase image of -Fe2O3 , lift at 200nm nanoparticles, lift at 8 µm (magnet under the sample) (magnet under the sample)

Figure 7.9a a non-contact AFM image and 7.9b,7.9c ,7.9d and 7.9e the corresponding phase images of -Fe2O3 nanoparticles generated by a permanent magnet installed under the sample at different lift heights. Field of view 10 m × 10 m.

Figure 7.9a shows a topographic tapping mode AFM image of -Fe2O3 nanoparticles. 7.9b, 7.9c, 7.9d and 7.9e show the phase image corresponding to the region imaged in Figure 7.9a generated from MFM imaging mode at lift heights of 50, 100, 150 and 200 nm respectively. Here the magnetic moment of the -Fe2O3 nanoparticles sample was increased by placing a permanent magnet under the sample. By comparing the topographic image and the phase images we can see that the magnetic signal intensity from the nanoparticles in the phase image is much higher than that observed in the surrounding background. The best phase image resolution was obtained at lift height 50 nm. These images demonstrate that as the lift height increases the magnitude of the phase image contrast decreases and the resolution of the image also decreases. At 200 nm lift height the phase image shows only weak magnetic signals from the -Fe2O3 nanoparticles. At a lift height of 8 µm no magnetic information is observed. The fact that the image persists to heights up to 200nm indicates that the image acquired at 50nm is a magnetic signal and not an artefact of close contact with the sample. The higher lift height images enable the more magnetic parts of the topography image to be identified. The fact that the lift heights at which the image is detectable are much higher than the case where no magnet was applied indicate that the magnet has increased the magnetization of the particles in the sample.

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7.3.3 MFM of RBC sections

RBC

Figure 7.10a: Height image of a RBC section after Figure 7.10b : Phase image of a RBC section

its incubation with -Fe2O3 nanoparticles at 200 after its incubation with -Fe2O3 nanoparticles at mOsm 200 mOsm (Lift at 30 nm)

Figure 7.10c : Phase image of a RBC section after Figure 7.10d : Phase image of a RBC section after

its incubation with -Fe2O3 nanoparticles at 200 its incubation with -Fe2O3 nanoparticles at 200 mOsm (Lift at 40 nm) mOsm (Lift at 50 nm)

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Figure 7.10e : Phase image of a RBC section after Figure 7.10f : Phase image of a RBC section after

its incubation with -Fe2O3 nanoparticles at 200 its incubation with -Fe2O3 nanoparticles at 200 mOsm (Lift at 60 nm) mOsm (Lift at 80 nm)

Figure 7.10g : Phase image of a RBC section after Figure 7.10h : Phase image of a RBC section after

its incubation with -Fe2O3 nanoparticles at 200 its incubation with -Fe2O3 nanoparticles at 200 mOsm (Lift at 100 nm) mOsm (Lift at 150 nm)

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Figure 7.10j : Phase image of a RBC section after Figure 7.10i : Phase image of a RBC section after

its incubation with -Fe2O3 nanoparticles at 200 its incubation with -Fe2O3 nanoparticles at 200 mOsm (Lift at 200 nm) mOsm (Lift at 250 nm)

Figure 7.10a, a non-contact AFM image of RBC section loaded with -Fe2O3 nanoparticles and 7.10b, 7.10c, 7.10d, 7.10e, 7.10f , 7.10g, 7.10h, 7.10i and 7.10j the corresponding phase images of the sections of RBCs loaded with -Fe2O3 nanoparticles generated by a permanent magnet installed under the sample at different lift height: 30, 50, 100, 150, 200 and 250 nm respectively. Field of view 7 m × 7 m.

An AFM topography image, using tapping mode, of a section of a RBC after its incubation with -Fe2O3 nanoparticles at 200 mOsm is shown in figure 7.10a. The corresponding phase images (magnetic images) at lift heights of 30, 40, 50, 60, 80, 100, 120, 150, 200 and 250 nm are shown in Figure 7.10b, 7.10c, 7.10d, 7.10e, 7.10f, 7.10g, 7.10h, 7.10i and 7.10j respectively. The phase images show a magnetic signal generated from the -Fe2O3 nanoparticles located inside and outside the RBC. However, when the lift height increases the magnetic signals become weaker. Here the phase images were generated from the magnetic field induced into the sample by a permanent magnet installed under the sample.

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7.3.4 AFM of RBCs at different osmolarities

Figure 7.11: RBCs shape versus osmotic pressure.

Figure 7.11 shows topographic tapping mode AFM images of RBCs, prepared at different osmotic pressures ranging from 160 to 290 mOsm, dried onto pieces of freshly cleaved mica. The aim of this measurement was to find a correlation between the RBC size and the osmotic pressure but unfortunately due to salt crystallization, drying and artifacts there was no consistent change in RBCs shape and size with osmolarity. In order to avoid salt crystallization, drying and artefacts we attempted to scan RBCs in liquid.

7.3.5 AFM of RBCs in liquid

Figure 7.12a and 7.12b show topography obtained using tapping mode AFM of a stock suspension of RBCs in PBS at 300 mOsm and pH 8. Here The RBCs were scanned in liquid to avoid the salt crystallization problems. Unfortunately the images are dominated by tip like artefacts. It was assumed that these artefacts were due to the cells movement during imaging process.

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Figure 7.12a: Height image of a RBC in liquid. Figure 7.12b Height image of a RBC in liquid.

7.4 Discussion

In this study the aim was to compare the images of RBCs generated by the AFM and MFM scanning mode and then determine where the nanoparticles are located in cells. I faced some difficulties in this aim such as crystallizing of the salt in the medium and scanning in liquid using glutaraldedyde.

During this study I manged to compare the magnetic images generated by the magnetic field induced into the -Fe2O3 nanoparticles sample by the magnetized tip and by a permanent magnet installed under the sample. The magnetic field generated by the permanent magnet under sample is much greater than the magnetic field generated by the magnetized tip. Therefore the phase images created by permanent magnet under sample have better resolution than the phase images created by the magnetized tip alone at the same lift height as shown in figures 7.8 and 7.9. When the sample–tip distance increased the short range force becomes negligible and the magnetic force become dominate. Here the resolution of the magnetic images created by permanent magnet are much better than the magnetic image resolution created by the magnetized tip alone as shown in figures 7.8

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Chapter 7 Atomic Force Microscopy and Magnetic Force Microscopy of iron oxide nanoparticles and red blood cells loaded with iron oxide nanoparticles and 7.9. In addition, creating a strong magnetic field in the sample allows the MFM user to record pure magnetic images without any short range forces effects as shown in figure 7.8. Previous studies of the incorporation of magnetic nanoparticles into RBCs have shown that it is possible to load particles into the cells by inducing pore formation through osmotic pressure [52, 56]. The second aim of this study was to record a magnetic signal from γ-

Fe2O3 nanoparticles loaded into the RBCs. The MFM images of the RBCs section (phase images figure 7.10 ) indicated that a significant number of particles were inside the RBCs. This result was also found in our previous studies using TEM [52]. Here it is obvious as the lift height increase the magnetic signal resolution decrease as shown in figure 7.10.

To get more specific results we used the AFM to scan the RBCs at different osmotic pressures. The scanning was done in tapping mode and the RBCs were dried onto pieces of mica. Unfortunately this method was not perfect to measure the RBCs size at different osmolarities as shown in figure 7.11 due to the crystallization of the salt from the medium and because when the RBCs were dried their shape changed due to other effects.

AFM scanning of RBCs in PBS was not successful due to motion of the RBCs caused by the probing tip as shown in figure 7.12. There is the possibility of scanning RBCs in PBS if we improve the scanning condition such as finding a method to make the RBCs stick onto a surface without adversely affecting their shape.

7.5 Conclusion

In MFM scanning mode the best qualities phase images were generated using an external magnetic field to magnetize the magnetic nanoparticles in the sample. The best resolution was obtained with relatively small lift heights. This was very obvious when MFM scanning were done on γ-Fe2O3 nanoparticles sample in the presence of a strong and weak magnetic field.

In terms of RBCs sample preparation for AFM scaning , sections of resin embedded RBCs incubated with γ-Fe2O3 nanoparticles were the best sample to record a magnetic signal

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Chapter 7 Atomic Force Microscopy and Magnetic Force Microscopy of iron oxide nanoparticles and red blood cells loaded with iron oxide nanoparticles from the γ-Fe2O3 nanoparticles loaded into the cells. RBCs dried onto pieces of freshly cleaved mica were not good targets to scan RBCs due to the crystallizing of the salt from the medium and due to the RBCs changing their shape during the drying porcess. RBC imaging in PBS may be possible if a method to fix cells onto a surface can be perfected.

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Chapter 8 Conclusion and Future Work

In this thesis I study the loading and uptake of superparamagnetic iron oxide nanoparticles and gadolinium chelates by red blood cells (RBCs) under hypo-osmolar conditions. The location of the iron oxide nanoparticles and the gadoteric acid inside the RBCs was confirmed by using energy filtered transmission electron microscopy (EFTEM). The size distribution of the particulate contrast agents inside the RBCs was also measured. The maximum size of aggregate of particles that entered the cells was found to be approximately 120 nm. All surviving cells exposed to gadoteric acid and iron oxide nanoparticles by the two loading methods used in this thesis had enhanced proton relaxivity demonstrating their potential to act as magnetic resonance contrast enhancing agents for the blood pool. However, this enhancement was demonstrated in vitro, therefore in vivo studies are now required.

8.1 Discussion

In Chapter Four I introduced iron oxide nanoparticles, produced by chemical synthesis, into RBCs. This was to examine the uptake and capture of the nanoparticles at different osmolarities that range from 100 to 290 mOsm before being returned to an osmolarity of 300 mOsm. Chapter Four also investigated the optimum osmolarity for loading magnetic nanoparticles into RBCs and the limits on the particle sizes that can be incorporated within the cells. As the fundamental particle size of the magnetic particles used in this study was generally less than 100 nm, these measurements were conducted by transmission electron microscopy (TEM). Previous studies of the incorporation of magnetic nanoparticles into RBCs have shown that it is possible to load particles into the cells by inducing pore formation through osmotic pressure [56, 97]. Studies by Antonelli et al. and Brahler et al. [56, 97] examined a very particular osmolarity but in both studies the loading method was not clear and information on osmolarity was either missing or clearly incorrect.

In this study, RBCs were incubated for 12 hours with iron oxide nanoparticles with a broad range of particle and aggregate sizes (from 10 to 600 nm) and at different osmolarities ranging from 100 to 290 mOsm before being returned to an osmolarity of 300 mOsm. In addition, iron oxide nanoparticles were loaded into the RBCs by an osmotic pulse technique which resulted in significant MRI contrast enhancement. TEM images were used 153

Chapter 8 Conclusion and Future Work to measure the nanoparticle size distributions within the cells. However, the TEM images demonstrated that no magnetic nanoparticles were found within cells incubated with nanoparticles at 290 mOsm consistent with the hypothesis that it is osmotic pressure differences across the cell membrane that cause pores to open up in the membrane. The TEM and iron mapping by electron energy loss spectroscopy show no iron oxide nanoparticles inside the RBC sections at 290 mOsm. Here we also calculated the concentrations of nanoparticles trapped within the cells by using TEM and iron mapping by electron energy loss spectroscopy. An osmolarity of 200 mOsm was found to be the optimal condition for loading the cells with nanoparticles. At this osmolarity, it was shown that the concentration of particles within the cells relative to the average concentration in the suspension is maximized. For this reason we loaded the RBCs with polymer coated iron oxide nanoparticles by using the osmotic incubation method at 200 mOsm only (Chapter 5). In order to compare the efficiency of this method with the other loading technique, we loaded the RBCs with iron nanoparticles by the osmotic pulse loading technique. For the polymer coated nanoparticles systems it was found that the osmotic pulse method of loading resulted in higher longitudinal (r1) and transverse (r2) cell specific proton relaxivities, while for the uncoated nanoparticle system the hypo-osmolar incubation method was found to yield higher cell specific proton relaxivities. These observations are qualitatively consistent with the hypothesis that during the osmotic pulse incubation the pores in the cell membrane do not open to their maximum size in the very short time of the pulse and therefore only smaller nanoparticles can easily enter into the RBCs, while during the hypo-osmolar incubation at 200 mOsm the pore size can open to a maximum of approximately 120 nm [52, 102] thus allowing larger clusters (as found in the uncoated particles) to enter the cells. However, as discussed later, there may be another explanation for this observation. Generally the TEM images showed dispersed clusters of nanoparticles within the cells at rather low number concentrations.

In Chapter Five, as mentioned above, I compare the effectiveness of magnetic resonance contrast enhancement in RBCs loaded with four different magnetic iron oxide nanoparticles systems (three different polymer coated systems and one uncoated system) using the two different loading techniques. The observed MRI contrast enhancement from the uncoated iron oxide nanoparticles system was not a reliable measure of efficacy of loading due to the 154

Chapter 8 Conclusion and Future Work difficulty of washing away the excess unloaded nanoparticles from the cell suspension. Conversely, the results from the coated iron oxide nanoparticles were more reliable because of the amenability of the uncontained particles to be effectively washed from the cell suspension. This phenomenon was confirmed by the TEM images. The entire polymer coated nanoparticle TEM images showed no trace of nanoparticles outside of the cells (Chapter 5) while the TEM images of the uncoated nanoparticles showed the existence of nanoparticles outside of the cells (Chapter 4). To conclude, to achieve a more specific result it is important to used coated iron oxide nanoparticles to ensure that the excess of the iron oxide nanoparticles can be easily washed.

In Chapter Six, RBCs were loaded with gadoteric acid using the two methods. Here the gadoteric acid (C16H25GdN4O8) was used as the active MR contrast enhancing cell load. As we were not sure if these molecules would behave in a similar fashion to the nanoparticles, the loadings by the osmotic incubation were carried out at 150, 200, 150 and 300 mOsm. The osmotic pulse method was found to yield the greatest cell-specific relaxivity enhancements (71-fold for longitudinal relaxivity and 39-fold for transverse relaxivity). Johnson et al. [57] have investigated the enhancement of the proton relaxivity of gadolinium loaded RBCs but it was not clear whether all RBCs had loaded with gadolinium during the preparation process or whether it was only some, perhaps older, cells with more fragile membranes that form pores that admit the contrast agent. Here we also tested the hypothesis that only a fraction of the RBCs in a sample open pores to enable gadolinium loading with each of the techniques. The observation that none of the imaged cells exposed to gadoteric acid under hypo-osmolar conditions yielded gadolinium jump ratio signal intensities less than the resin background suggests that the osmotic pulse and osmotic incubation methods result in all cells being loaded with gadoteric acid rather than only a fraction of the cells.

The Gd M-edge jump ratio image intensities within control RBCs were less than the image intensities in the surrounding resin, resulting in a negative contrast. Cells incubated with Gd all showed generally uniform positive contrast with respect to the resin background except for those incubated at 300 mOsm indicating once again that the osmotic pressure difference across the cell membrane is the key factor that initiates access of contrast agent 155

Chapter 8 Conclusion and Future Work into the cells. Small numbers of high contrast nanoparticles were observed in the cells suggesting that some of the Gd had clustered. The RBCs prepared at the higher osmolarities had larger clusters (median ~45nm) compared with those prepared at 200 mOsm and 150 mOsm (median sizes ~20nm and 10 nm respectively). Cells prepared with the osmotic pulse method also showed positive contrast but no clustering of Gd. These results suggest that Gd is able to enter and be retained by the cells via both the hypo- osmolar incubation and osmotic pulse methods. Since 300 mOsm is very close to physiological osmolarity (approximately 290 mOsm) the conditions are not hypo-osmolar and explain why the Gd M-edge jump ratio image RBC contrast remains negative when the

RBCs are incubated with gadoteric acid at that osmolarity. The r1 enhancement was greatest for the cells prepared by the osmotic pulse method despite these cells showing only modest Gd M-edge jump ratio image contrast. It is possible that r1 enhancement is suppressed in the RBCs prepared by the incubation method because of the clustering of the Gd. The RBCs projected areas in the optical microscope images in Chapter Five and Six show that, of all the preparation methods, the osmotic pulse method results in the lowest cell sizes. This suggested that the RBCs loaded with contrast agents by osmotic incubation method might have longer half-life into the blood stream than the RBCs loaded by the osmotic pulse technique since the spleen is able to preferentially select damaged cells for removal from the blood stream. Future studies will be required to test this hypothesis.

In Chapter six it is observed that the haemoglobin release may occur only during the initial drop in osmolarity where temporary osmolarity differences between the inside and outside of the cell may cause osmotic pressure induced flows ejecting hemoglobin. This concept is consistent with hemoglobin molecules in RBCs tending to have mutually attractive interactions. The hypothesis that the gadoteric acid load is related to the fraction of hemoglobin ejected from the cell during the lowering of the osmolarity may also explain why the RBCs exposed to gadoteric acid during an osmotic pulse are observed to have the highest proton relaxivities and hence presumably the highest gadoteric acid loadings. The osmotic pulse method is likely to result in greater peak osmolarity differences between the inside and outside of the cells owing to the pulsed nature of the procedure and the high concentrations of DMSO used and hence greater osmotic pressure induced flows out from 156

Chapter 8 Conclusion and Future Work the cells might be expected. As such, it may be expected that a greater fraction of the hemoglobin load is ejected from the cell compared to the slower osmolarity drops caused by the osmotic incubation method.

In Chapter Six it is not clear why higher loadings of gadoteric acid are achieved with lower osmolarities. One possible hypothesis is that the loading is related to the volume of hemoglobin released from the RBCs during the osmotic incubation. Therefore, more studies are required in this field.

Chapter Seven shows the possibility of using magnetic force microscopy (MFM) as a natural extension of atomic force microscopy (AFM) to probe the nature of iron oxide nanoparticle loading of RBCs. In MFM scanning mode, the best quality phase images were generated using an external magnetic field to magnetize the magnetic nanoparticles in the sample. In this chapter MFM was used to record images generated by the magnetic fields induced by the magnetic nanoparticles and the RBCs loaded with magnetic nanoparticles. The magnetic field generated by the permanent magnet under the sample is much greater than the magnetic field generated by the magnetized tip and is thus able to generate greater magnetizations in the magnetic nanoparticles. The greater particle magnetizations caused by the strong uniform field from the permanent magnet resulted in the phase images having better resolution than the phase images created by the magnetized tip alone at the same lift height. Furthermore, particles were detectable at greater lift heights (and presumably greater depths into the RBC) when the permanent magnet was used to provide a uniform magnetizing field.

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8.2 Speculation about the relationship between the pores that open up in RBCs membranes and the diffusion time of the water, nanoparticles, gadoteric acid and hemoglobin.

In the following paragraph, I draw together the various observations made in the different chapters if this thesis and, together with some considerations of the relative diffusion rates of different sized particles in aqueous suspension, speculate on the mechanisms by which magnetic nanoparticles and gadoteric acid molecules may enter RBCs under the different hypo-osmolar conditions encountered.

8.2.1 Friction and Diffusion

In the context of hydrodymamics, friction is the force resisting the relative motion of fluid layers. Fluid friction is the friction between layers within a viscous fluid that are moving relative to each other. The viscous friction coefficient ζ (kg s-1) is given by Stokes formula [120]

Where a: is the radius of the particle. η: is a constant called the viscosity of the fluid

The diffusion constant (D (m2 s-1) ) of particles in fluid is given by the Einstein relation [120]

Where

kB: is Boltzmann constant. T: is absolute temperature

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Combining (8.1) and (8.2)

8.2.1.1 Diffusion time of the different components into and out from the cells during the loading period.

With the assumption that RBCs have a spherical shape with radius 3.5µm, η equal to the viscosity of water. I also assume that the nanoparticles, hemoglobin, hydrated ions and gadoteric acid molecules have a spherical shape. The characteristic time, t, for a particle to diffuse the distance of the order of the radius (R1) of a RBC is approximated by

From this relation, we can see that the characteristic time, t, that it takes a particle to diffuse a certain distance in a given fluid at a given temperature is proportional to the radius of the particle, a. Table 8.1 shows calculations of the characteristic times it would take different ions, molecules, and nanoparticles to diffuse a distance approximately equal to the radius of a RBC in water at 27C. The calculations assume that the particles are spherical.

Particle Radius (nm) D (m2s-1) t (ms) Water molecule 0.15 1.46E-09 1 Hydrated ion 0.30 7.32E-10 3 Hemoglobin 3.1 7.09E-11 29 Nanoparticle 20 1.10E-11 186 Nanoparticle 30 7.32E-12 279 Nanoparticle 50 4.39E-12 465 Gadoteric acid 0.50 4.39E-10 7

Table: 8.1 Size, self diffusion coefficients (D) and characteristic diffusion times (t) of hydrated ions, water molecules, hemoglobin, gadoteric acid and iron oxide nanoparticles to diffuse a distance of the order of the radius of a RBC in a medium with a viscosity of pure water at 300K.

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8.2.2 Our point of view of the loading process

We hypothesise that swelling of the cell causes openings to appear in the cell membrane which allow small nanoparticles to enter between the hemoglobin (Hb) molecules. However, large nanoparticles have less chance to enter into the RBCs during the loading procedure for many reasons. For example, it is known that Hb molecules occupy about one third of the internal volume of the RBC. This high density makes it very difficult for large nanoparticles to diffuse between the Hb molecules.

8.2.2.1 Consideration of rates of diffusion of different sized entities with regard to loading of RBCs by osmotic pressure differences

If the osmolarity within a RBC is greater than the osmolarity external to the cell, there is an osmotic pressure difference between the inside and outside of the cell. It is worth noting that the usual osmolarity inside a RBC is approximately 290 mOsm and that it can be shown that the haemoglobin molecules contribution to the osmolarity is negligible i.e. the osmolarity is predominantly determined by the concentration of hydrated ions which are much smaller than haemoglobin molecules. Since the cell membrane acts as a semi- permeable membrane, allowing water molecules to pass freely while restricting most other naturally occurring molecules and ions, the difference in osmolarity causes a net flux of water molecules into the cell causing the cell to expand. The evidence presented in this thesis (chapter 4, 5 and 6), suggests that the expansion of the cell results in large pores or holes opening up in the cell membrane. In Chapter Four we showed that these holes could be as large as up to 120 nm in size under appropriate conditions.

The opening up of large pores would most likely facilitate the passage of molecules such as haemoglobin and hydrated ions out form the cell as well as allowing passage of molecules and particles external to the cell into the cell since the sizes of these ions, molecules, and particles are generally smaller than the size of the holes being opened. The nature of the holes still requires further research. For example, the holes may be static or dynamic in nature. Holes may open up at a particular place in the membrane and stay open until the cell membrane shrinks again (static). Alternatively, it could be imagined that while the cell is 160

Chapter 8 Conclusion and Future Work expanded, holes may be in a constant state of opening and closing at different points around the cell (dynamic). Regardless of whether the holes are static or dynamic, we can infer certain differences in behaviour between the different particulate species involved during the osmotic pressure based loading of RBCs described in this thesis. We note that the characteristic time for a spherical particle to diffuse a certain distance through a fluid medium with a given viscosity at a given temperature is proportional to the radius of the particle (see Table 8.1). As such, a hydrated Na+ or Cl- ion (each with radii of approximately 0.3 nm) will diffuse about 10 times further in a given time than a haemoglobin molecule (with radius approximately 3nm). Hence, we can envisage that once large pores open up in the cell membrane, small hydrated ions, which are the dominant contributors to the osmolarity within the cell, will be able to diffuse out from the cell and reduce the osmotic pressure difference, allowing the cell shrink back to close to its original size under membrane tension forces before a significant number of haemoglobin molecules have had time to diffuse out from the cell. This picture could explain why we observed relatively little haemoglobin escaping from the cells in our loading experiments. Furthermore, depletion forces between the haemoglobin molecules are likely to further slow down their exit from the RBCs. Even in the case of the osmotic incubation technique, it may be that the large pores are only open for very short periods of time – the time it takes for the small hydrated ions to diffuse out from the cell – a timescale that we can estimate to be of the order of 3ms using Stokes formula and the Einstein relation (see Table 8.1).

How does this picture fit with the observed difference in loading efficiency between the osmotic incubation and osmotic pulse methods? The loading efficiency was found to be greater for the osmotic pulse method than the osmotic incubation method. Possible reasons for the greater efficiency with the osmotic pulse method could include (a) larger pores and (b) pores open for a longer period of time. Interestingly, we found no evidence for larger pores being formed with the osmotic pulse method. If any difference was found, it was that the osmotic incubation method may have resulted in larger pores. In principle, the osmotic pressure difference across the cell membrane was much higher for the osmotic pulse method since the osmolarity differences achieved with the DMSO reach 900 compared with a difference of just 100 or so with the osmotic incubation method. Thus, the lack of 161

Chapter 8 Conclusion and Future Work evidence for larger pores may be related to the maximum size of pore that a RBC can sustain before being destroyed. The possibility that the pores remain open longer using the osmotic pulse method is very likely. DMSO molecules are about 0.3nm radius (unhydrated) and so are likely to be larger than hydrated ions which are about the same size 0.3 nm when hydrated. As such we expect them to diffuse out of the cell more slowly once the concentration gradient is formed during the osmotic pulse technique. Since the concentration of DMSO used is so much greater than the naturally occurring ions within the cell, the timescale for equilibrating the osmolarity between the inside and outside of the cell will be greater than that for the osmotic incubation method. Together, these two factors suggest that the pores will be open for a longer period of time with the osmotic pulse technique than the osmotic incubation technique which could explain why we see better loading efficiency with the former. Such a scenario would also explain why more Hb is released when using the osmotic pulse technique.

How does this picture fit with what we observe for the diffusion of gadoteric acid and nanoparticles into cells? Gadoteric acid is a little larger than DMSO (approximately twice the size) but a lot smaller than haemoglobin molecules. As such we expect it to diffuse slightly slower than DMSO and hydrated ions but faster than haemoglobin. The similarity in size to the DMSO implies that gadoteric acid could diffuse into the cell significantly on the timescale of the pores being open. The key limitation to the gadoteric loading will be the remaining haemoglobin molecules which will move very little distance on the timescale of the pores being open and which occupy about one third of the volume of the RBC. The observation that magnetic nanoparticles of the dimension of several tens of nanometers can be observed to load into the RBCs is not easily explained by this model of relative diffusion. The timescales of diffusion will be (a) much longer than the timescale of pore opening and (b) there is very little space between the haemoglobin molecules for the nanoparticles to diffuse into. As such, the observation that particles do enter the cells suggests that there may be other mechanisms besides diffusion at play. The involvement of other mechanisms is also consistent with our observation of very different loading efficiencies of the same iron oxide nanoparticles coated with different polymers. The polymers could very likely play a role in modifying electrostatic or entropic interactions between the cell and the particles. The PEO-PAA coated particles appeared to have a 162

Chapter 8 Conclusion and Future Work particularly enhanced loading efficiency relative to the other particles. It is not clear at this stage why a PEO-PAA coating should enhance the efficiency with which particles enter the cell. The PAA polymer is expected to be associated with negative charges and the cell surface of a RBC is also negatively charged so a naïve first impression would be that the PAA would hinder particles approaching the cell. However, the polymer may be interacting with the charge double layer surround each RBC thus causing a higher than average concentration of nanoparticles immediately adjacent to cell membranes. In this way, the likelihood of a nanoparticle being able to diffuse into a cell would be greatly increased because of the effectively much higher concentration of particles near the cell membrane. Since the nanoparticles would diffuse into the RBC during the period when it has expanded, there would be a margin within the cell with a lower Hb concentration thus providing space for the nanoparticles (as illustrated in Figure 8.1). Further research is clearly required in order to test these hypotheses and elucidate the mechanism by which the PEO-PAA enhances the loading of magnetite nanoparticles into the cells.

Figure 8.1 a schematic of the proposed loading mechanism. First, The RBCs swell with water induced by the osmotic pressure and the pores are formed. Second, small particles such as gadoteric acid or hydrated ions may diffuse between the Hb molecules into or out from the cell. Only large nanoparticles that are initially very close to the cell membrane can diffuse into the RBC and occupy the space between the Hb molecules and the cell membrane. Third, when the RBCs shrink the large particles are trapped between the Hb molecules and the cell membrane. 163

Chapter 8 Conclusion and Future Work

8.3 Key conclusion

An osmolarity of 200 mOsm was found to be the optimal condition for the loading of red blood cells (RBCs) with nanoparticles using the osmotic incubation technique. At this osmolarity, it was shown that the concentration of particles within the cells relative to the average concentration in the suspension is maximized. At 200 mOsm, the maximum size aggregate of particles that entered the RBCs was approximately 120 nm suggesting that the maximum pore size generated by the hypo-osmolar conditions is of the order of 120 nm. For the polymer coated nanoparticles systems it was found that the osmotic pulse method of loading resulted in higher proton transverse relaxation rates (r2) while for the uncoated nanoparticles system, the results are not so clear because of the problem of not being able to wash the cells efficiently. Proton transverse relaxation rates up to 31 times higher than native red blood cells were achieved (using the triblock copolymer coated iron oxide nanoparticles introduced with the osmotic pulse technique).

Loading RBCs with gadoteric acid (C16H25GdN4O8) resulted in the enhancement of proton relaxivities in RBCs suspensions. The osmotic pulse method was found to yield the greatest cell-specific relaxivity enhancements (71 fold for longitudinal relaxivity and 39 fold for transverse relaxivity). All surviving cells exposed to gadoteric acid under hypo- osmolar conditions showed enhanced (relative to control cells) and generally uniform intensity within the cells in gadolinium jump ratio images, suggesting all cells are susceptible to loading and that the loading is generally spatially uniform within each cell. There was some evidence for a small amount of precipitation or aggregation of gadolinium within some cells prepared by the hypo-osmolar incubation method.

It is possible to use magnetic force microscopy (MFM) as a natural extension of the atomic force microscopy (AFM) to elucidate details of the nature of magnetic nanoparticle inclusions in RBCs. In the MFM scanning mode the best quality phase images were generated using an external magnetic field to magnetize the magnetic nanoparticles in the sample.

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8.4 Future directions

8.4.1 Magnetic fractionation

Further to this study we propose to use magnetic fractionation on the loaded RBCs. The aim of this study would be to demonstrate whether all RBCs load with iron oxide nanoparticles or whether only a fraction of cells load. Fractionation experiments would also give information on the distribution of cell loadings. We hypothesize that the swelling of the cell causes openings to appear in the cell membrane which allows iron oxide nanoparticles to enter the cells. The magnetic fractionation methods described by Karl et al. 2010 [121] will determine the fraction of RBCs loaded with iron oxide nanoparticles

8.4.2 Three dimensional transmission electron microscopy imaging

Measuring nanoparticle dispersions and in vitro cellular uptake by electron microscopy, the uptake of nanoparticles by human RBCs has been examined by quantitative TEM imaging to determine the size distribution of the nanoparticles in RBCs and the iron concentration. The future work in this area is to extend this work into 3-D using the Gatan 3-View system in the transmission electron microscope (TEM) to serial section resin-embedded cells exposed to nanoparticles. By combining the information from TEM imaging and serial sectioning with that provided by other techniques (such as optical imaging), a more detailed picture of the three dimensional distribution of nanoparticles within the RBCs will be possible. Such information would elucidate whether nanoparticles are driven to one end of the cell during centrifugation processes for example. Without answering this question, the data from standard TEM images remain ambiguous.

8.4.3 In vivo studies (mice and rats)

We propose to conduct in vivo studies of RBCs loaded with iron oxide nanoparticles and gadoteric acid in order to measure their contrast enhancing ability in vivo as well as characteristic lifetimes within the blood stream. We suggest loading the contrast agent into mouse or rat RBCs by collecting 1 mL of their blood using heparinised tubes. This will be

165

Chapter 8 Conclusion and Future Work followed by loading the animal RBCs with contrast agents using the two loading methods as described in Chapters Five and Six. This will be followed by re-injecting the loaded cells into the animal’s blood stream. The animal will be examined by a MRI scanner that will become available at UWA later this year (Bruker: 9.4 Tesla, 30 cm Bore). This scanner will allow us to study the contrast enhancement of contrast agents and their half life time in the animal blood stream.

8.4.4 Toxicology

Toxicity is determined by the combination of the biological half-life of the compounds and any potential breakdown products [82]. Both the chelated gadolinium compounds and magnetic particles are rapidly cleared from the blood stream, gadolinium via renal excretion and magnetic nanoparticles by the reticulo-endothelial system (RES) [82]. However, when those contrast agents are loaded into the RBCs, the rate of clearance depends on the half-life of the loaded RBCs into the blood stream and the clearance pathways may be different from those observed for contrast enhancing agents delivered to the plasma compartment of blood. Therefore in vivo studies are necessary to determine the half-life of RBCs loaded into the blood stream and the biodistribution of the agents at different time points after initial delivery into the blood stream.

We obtained gadoteric acid (C16H25GdN4O8) from Guerbet (France). Gadoteric acid is approved for use with all ages ranging from newborns to the elderly. It can not be used for at-risk patients including cardiovascular patients or patients with renal impairment. It remains to be seen whether the toxicity of gadoteric acid is changed when delivered to the body in RBC carriers.

Magnetic nanoparticles at relatively low concentrations can generate significant improvement in MRI contrast and image quality. However the toxicology of superparamagnetic nanoparticles is still poorly understood when introduced into the human body [82]. Again, the toxic behavior of the particles when delivered within RBC carriers may be different from that observed when delivered directly into the plasma compartment of the blood. 166

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Appendices

Appendices The following appendices are journal articles published by the candidate over the course of the candidature. They are submitted as additional evidence for candidature for the degree of Doctor of Philosophy.

Appendix1-Mounir Ibrahim, Leonard Wee, Robert C. Woodward, Martin Saunders and Tim G. St. Pierre. Loading Erythrocytes with Maghemite Nanoparticles via Osmotic Pressure Induced Cell Membrane Pores. Am. Inst. Phys. Conf. Proc. Series, Vol 1311, 8th International Conference on the Scientific and Clinical Applications of Magnetic Carriers, pp 375-381.

Appendix 2 -Mounir Ibrahim, Leonard Wee, Michael J. House, Robert C. Woodward, Martin Saunders, John Murphy, and Tim G. St. Pierre. Enhancement of the cell Specific Proton Relaxivities of Human Red Blood Cells via Loading with Gadoteric Acid. IEEE Transactions on Magnetics, 2013. 49(1): p. 414-420.

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