UNIVERSITY OF CINCINNATI

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1 UNIVERSITY OF CINCINNATI Date: I,, hereby submit this work as part of the requirements for the degree of: in: It is entitled: This work and its defense approved by: Chair:

2 In Vivo MR Microscopy of Tumor- Targeted Liposome Combining USPIO and Saposin-C A dissertation submitted the Division of Research and Advanced Studies of the University of Cincinnati in partial fulfillment of the requirements for the degree of DOCTOR OF PHILOSOPHY (Ph.D.) in the Department of Biomedical Engineering of the College of Engineering by Vinod Kaimal M.S., University of Cincinnati, 2002 B.E., University of Pune, India, 1999 Committee: Dr. Scott K. Holland, Chair Dr. Xiaoyang Qi Dr. William Ball Dr. Christy K. Holland Dr. Jing-huei Lee 2007

3 Abstract In recent years, molecular imaging, defined as the visualization of biologic process in vivo at the molecular level, has gained prominence in the detection and monitoring of cancerous tissue. Several targets for the molecular imaging of cancer have been identified but in order to produce selective contrast enhancement, target specific contrast agent delivery methods and optimized imaging techniques need to be developed. This dissertation describes a molecular imaging approach that uses liposomes loaded with Ultrasmall SuperParamagnetic Iron Oxide (USPIO) nanoparticles targeted to tumor cells. Liposomes made from dioleylphosphatidylserine, (DOPS) are being used as carriers for Saposin-C, a fusogenic protein, currently being investigated for its effect in inducing cell death in a variety of human cancer cells. We developed an efficient method for loading Saposin-C-DOPS liposomes with USPIO. Further, we tested the uptake and MR detectability of these contrast-laden liposomes, referred to as Sc-DOPS-IO, in vitro in tumor cell cultures. Finally, we evaluated the feasibility of detecting saposin-c- DOPS-USPIO liposomes using MR imaging in vivo. The Sc-DOPS-IO liposomes encapsulated an average of 94.8 ±12.8 µg Fe /ml when a 1:30 molar ratio of saposin-c to DOPS was used with 1mM DOPS and exhibited a transverse relaxivity (r2) of ±4.9 mm -1 s -1. Uptake in tumor cell cultures was found to be proportional to both time of exposure and initial concentration of Sc-DOPS- IO in growth medium. The cells were MR detectable against a uniform agarose background when average encapsulation was greater that 1pg Fe/cell, using high iii

4 resolution T2* weighted MR imaging. In vivo experiments indicate accumulation of Sc- DOPS-IO in tumor xenografts in mice and T2 mapping of the xenografts show a significant drop in mean tumor T2 about 4 hours post injection of Sc-DOPS-IO. These results demonstrate the feasibility of using MR imaging for quantitative estimation of delivery and uptake of targeted drugs and early detection of tumors using targeted contrast agents. The methods presented here are promising for translation to applications in human subjects for clinical trials of new chemotherapy drugs. iv

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6 Acknowledgements I would like to thank my advisor, Dr. Scott Holland, for his support, guidance and patience during the course of my dissertation research. Dr. Holland allowed me the opportunity to pursue my research interests and ensured that I had access to every resource that I needed. His profound knowledge and insights enabled me to overcome many challenging problems and I consider myself very fortunate to have him for a mentor. I would also like to thank Dr. Xiaoyang Qi for his guidance and for giving me the opportunity to work in his lab. I would like to thank the other members of my dissertation committee, Dr. Christy Holland, Dr. William Ball and Dr. Jing-huei Lee. Their suggestions and feedback during our meetings were invaluable. I owe special thanks to Dr. Christy Holland for encouraging me to pursue a doctorate in medical imaging. I am grateful to the Department of Biomedical Engineering at the University of Cincinnati and the Imaging Research Center at Cincinnati Children s Hospital for their support. I have greatly benefited from my interactions with individuals from both these institutions. I would like to acknowledge and thank Scott Dunn, Ron Pratt, Ph.D., Bernard Dardzinzki, Ph.D., Diana Lindquist, Ph.D., Vince Schmithorst, Ph.D., Zhengtao Chu, MD, Doug Richardson Ph.D., Istvan Pirko, MD and Alexei Bogdanov, Ph.D. They have vi

7 contributed to this dissertation by sharing their knowledge and experience and have helped me tremendously with some difficult problems. I have fond memories of my days in graduate school at UC thanks to my friends here in Cincinnati and elsewhere. I am thankful for their friendship and support. I d also like to thank my uncle Bali, a UC alumnus, who inspired me to chase the American dream! Finally, I d like to thank my parents Dr. K.G. Kaimal and Mrs. Usha Kaimal, and my brothers Vikram and Vivek. Their unwavering support in all my endeavors and their confidence in me have helped me get to this point. vii

8 TABLE OF CONTENTS TABLE OF CONTENTS...8 Chapter 1 Introduction Magnetic Resonance Imaging for Molecular Imaging Motivation MRI and Cancer Present Trends and Challenges in Molecular Imaging of Cancer Overview of dissertation Chapter 2 Background and Review of Relevant Topics Magnetic Resonance Imaging Brief History MR: A Quantum mechanical description MR: A classical approach Contrast in MR Imaging MR Microscopy MR Imaging Parameters and Image Optimization Molecular and Cellular Imaging using MRI Contrast Agents Superparamagnetic Contrast Agents Cellular and Molecular Imaging Using MRI In vivo labeling of cells with iron oxide contrast Issues for detection of iron oxide contrast using MRI Summary Chapter 3 In Vitro Experiments Overview of research goals The protein Saposin C and the lipid DOPS Mechanism of uptake in tumor cells Preliminary Experiments Encapsulation of USPIO in DOPS-Sapsoin-C liposomes Method Analysis of liposome solution Dynamic Light Scattering Determining concentration of iron Average values of iron concentration and R2 values in Sc-DOPS-IO In vitro cell experiments Preparation Effect of exposure time on cell cultures Effect of initial concentration of ScDOPS-IO on cell cultures Fixing cells in agarose Imaging Results

9 3.6.1 Statistical analysis of In vitro results Discussion and Conclusions Chapter 4 In vivo Experiments General Guiding Principles Specific Aim Imaging Setup Injection of ScDOPS-IO Pilot Experiments Animal Model Imaging Injection of ScDOPS-IO Image post processing and analysis ICP-AES Analysis Imaging experiments in mice bearing human neuroblastoma xenografts Experiments in mice bearing human pancreatic tumor xenografts Study design Imaging Image post processing and analysis Results Statistical analysis of In vivo results Discussion Chapter 5 Conclusion A Summary and Discussion of the Major Findings Background Encapsulation of USPIO in DOPS liposome In vitro cell experiments In vivo experiments Discussion and Future research Conclusion BIBLIOGRAPHY Appendix Additional data, Graphs, Methods I. Relaxation times of USPIO in saline and gelatin II. ICP-AES Procedure III. Freeze-fracture and electron microscopy of Saposin-C DOPS liposomes IV. Fluorescence imaging in mouse tumor model Appendix CCHMC Animal Imaging Protocol Appendix Abstracts and Conference Proceedings

10 GLOSSARY OF ABBREVIATIONS ADC: Apparent Diffusion coefficient CNR: Contrast-to-Noise Ratio CTL Computed tomography DOPS: Di-oleyl-Phosphatidylserine DOT: Diffuse Optical tomography EMF: electromotive force FID: Free Induction Decay FIESTA: FLASH: Fast Low Angle Shot FOV: Field of View HUVEC: ICP-AES: Inductively coupled plasma-atomic emission spectroscopy LUV: Large Unilamellar Vesicles MION: Mono-crystalline iron oxide nanoparticles MRI: Magnetic Resonance Imaging NEX: Number of Excitations NMR: Nuclear Magnetic Resonance PEG: Polyethylene glycol PET: Positron Emission Tomography PS: Phosphatidylserine RES: Reticulo-endothelial system REV: Reverse-phase encapsulation vesicles RF: Radio Frequency ScDOPS-IO: Saposin-C-DOPS- USPIO liposomes ( See also: SCPSPL)

11 SCPSPL: Saposin-C Phosphatidylserine Proteoliposome SI: Signal Intensity SNR: Signal to Noise ratio SPECT: Single Photon Emission Computed Tomography SPIO: Superparamagnetic Iron Oxide SUV: Small Unilamellar Vesicles USPIO: Ultrasmall Superparamagnetic Iron Oxide VEGF: Vascular Endothelial Growth Factor 11

12 Chapter 1 Introduction The past several years have seen rapid growth in the field of molecular imaging. Molecular imaging can be defined as the characterization, typically in vivo, of biologic processes at a cellular and molecular level. A related term, cellular imaging, can be defined as the visualization of targeted cells in vivo. In some respects, molecular imaging is not a new discipline; nuclear imaging approaches such as positron emission tomography (PET) and single photon emission computer tomography (SPECT) have used molecular imaging concepts for over a decade to visualize biodistribution of labeled compounds. In these imaging techniques, however, the image contrast is governed by the local concentration of the radiolabeled reporter compound rather than the anatomical features. Recently, the merging of concepts in molecular biology with non-invasive imaging technologies has lead to some novel approaches in the field of molecular and cellular imaging, one of which is described in this dissertation. 1.1 Magnetic Resonance Imaging for Molecular Imaging Ideally, a molecular imaging technique is characterized by high sensitivity, minimal background signal and high temporal and spatial resolution. No imaging modality exists that has all these features; existing imaging modalities have their strengths and weaknesses (Figure 1-1). The highest sensitivity is provided by nuclear (PET and SPECT) and optical

13 Figure 1-1: Comparison of various imaging modalities with regard to spatial resolution and sensitivity. Adapted from Molecular Imaging by Markus Rudin [1]. (fluorescence, bioluminescence) imaging techniques. These modalities allow the detection of tracers or probes at less than nanomolar concentrations. For small animal scanners, the spatial resolution is limited to about 1mm. The high-resolution modalities, MRI, X-ray and computed tomography (CT), are characterized by inferior sensitivity. Sensitivity can be improved by using contrast agents. For example, in CT, contrast enhancement is achieved by using electrodense ( radio-opaque) contrast agents containing iodine. Proton ( 1 H) MRI maps the weighted distribution of tissue (mostly water) protons. The signal intensity is governed by a number of weighting factors including the longitudinal and transverse relaxation times (T1 and T2/T2*), diffusion, perfusion, and magnetization exchange due to chemical interactions. (Please see Chapter 2 for a detailed discussion of MR imaging and microscopy). These parameters are tissue specific and form the basis of the high soft-tissue contrast provided by MRI. The principle limit on resolution in MRI is given by the 13

14 displacement of water or adipose molecules between excitation and detection due to incoherent motion i.e. diffusion. This is about 1.3µm assuming a 1ms diffusion time [2]. In practice, however, resolution is limited by instrumental parameters and by sensitivity. As voxel volume decreases, signal to noise ratio (SNR) will decrease. Typically, a 100 µm isotropic resolution can be achieved in vivo using high field strength scanners. The low sensitivity of MRI means that direct visualization of molecular targets is not possible without signal amplification. This is achieved by paramagnetic or superparamagnetic MR contrast agents, which exert their effects over significant volumes and therefore changing the relaxation properties of a large number of water molecules. Typically, in MRI, concentrations of the order of >10-6 M have to be achieved in order to exert detectable effects on water proton relaxation. Another challenge is the actual delivery of these contrast agents to a specific tissue or target. The target should be expressed in sufficient quantities, should preferably be specific to the disease and should be accessible to the imaging agent. Once the target is identified, a high affinity ligand, usually proteins, peptides or antibodies are then coupled with the imaging agent and/or a therapeutic agent for deliver in vivo. Developing target specific reporter moieties to deliver MR contrast agents for molecular and cellular imaging applications is an active area of research. Several recent approaches are also discussed in Chapter Motivation MRI and Cancer The replication of cells in our body is tightly regulated by complex mechanisms. Any breakdown of this cell cycle leading to uncontrolled proliferation is broadly termed as cancer. Cancer can also travel to other parts of the body (metastases) through the blood and lymphatic fluid. The Annual Report to the Nation on the Status of Cancer [3], by the NIH 14

15 estimates that there will be approximately 1.5 million new cases of cancer and over half a million deaths dues to cancer in MRI has been used for several years clinically for the detection of tumors. The wide variety of contrast mechanisms enables physicians and radiologists to probe several aspects of tumor physiology non-invasively[4]. Apart from anatomical and morphological information on tumors, MRI can provide information on vascular function such as perfusion, provide information about tissue density using diffusion weighted imaging and can also provide metabolic information using MR spectroscopy. However, visualization of molecular targets, receptors or processes related to tumor physiology using MRI has not been possible until recently, due to the relative insensitivity of MRI. Recent advances in the design of high relaxivity MR contrast agents and strategies for targeting these contrast agents to specific targets have enabled MRI to be used in the molecular imaging of cancer Present Trends and Challenges in Molecular Imaging of Cancer The growing popularity of MR imaging in preclinical cancer research has lead to the development of novel imaging probes. These probes could be designed with several goals in mind, including but not limited to the following: a) Tracking of cancer cells b) Imaging of receptors or targets that are expressed during tumor progression (eg. probes targeted to the α v β 3 integrin receptors expressed on the angiogenic endothelium of tumors) c) Imaging of drug delivery to tumors d) combined delivery of diagnostic and therapeutic agents. These advances have made MRI an invaluable tool in preclinical cancer research and drug discovery. However, currently, there is no standard technique available for the imaging of molecular targets using MRI. Several different strategies have been adopted by various groups, for coupling either gadolinium or iron oxide based contrast agents to a targeting agent. (See section 2.4). The focus of these studies is usually just imaging. In this 15

16 dissertation, we have combined a potential anti-cancer agent and an imaging contrast agent into a single complex, encapsulated in a liposomal delivery vector, enabling not just imaging of the molecular target, but MR detection of the delivery of the chemotherapeutic agent. Several issues need to be addressed in order to successfully accomplish MR detection of targeted chemotherapeutic and imaging agents. Primarily, the issue is sensitivity. As discussed in the previous section, the low inherent sensitivity of MRI necessitates the use of high relaxivity contrast agents to identify molecular targets. These contrast agents tend to be superparamagnetic iron oxide (SPIO) based, which effect maximum signal change in MR images per unit metal. Quantifying the negative signal change caused by SPIO can be difficult, so strategies have to be developed to quantify, for example, contrast uptake in cells. The contrast agents have to be reliably coupled with the targeting agent. Next, the imaging probe has to be delivered to the target site, overcoming vascular, interstitial and cell membrane barriers. It must be noted that each application for imaging a certain target would require unique strategies for probe development, delivery to target, imaging and quantification. Ideally, a bottom-up approach is preferred, where the development of the imaging probe is incorporated into the design of the targeting moiety. However, this dissertation takes a top-down approach, taking an existing targeting moiety and incorporating in it an MR contrast agent. The dissertation addresses each of the above mentioned challenges, from development of the imaging probe by coupling an SPIO based contrast agent with the targeting agent and delivery vector, to optimizing imaging strategies, to developing methods to quantify the concentration of the probe at the target site in both in vitro and in vivo models. Therefore, this dissertation is an attempt to take existing techniques in molecular imaging using MRI and apply to toward a novel application. The generality of the methodologies developed could be adapted for several similar applications in the preclinical molecular imaging using MRI. More importantly, a significant advantage of these novel strategies for 16

17 molecular imaging using MRI in preclinical research is the possibility of translation to the clinic and provides the primary motivation for this dissertation. 1.3 Overview of dissertation Each molecular imaging application requires a unique strategy depending on the disease, available targets for imaging and the method of delivery of the imaging agent. The broad goal of this dissertation is to develop and test an MR imaging probe that could be used to visualize the delivery of targeted nanoparticles to tumors in vivo. The nanoparticle of interest is a proteo-liposome complex, made from the interaction of the human protein Saposin-C with the lipid di-oleyl-phosphatidylserine (DOPS). Saposin-C is a small protein comprising of approximately 80 amino acids. At acidic ph, Saposin-C combines with DOPS, and the resulting complex can then be made into liposomes of nm using either sonication or extrusion techniques. Preliminary investigations by Qi et al indicate anti-cancer properties of the protein, when delivered to tumor cells by using the DOPS liposomal vector. Saposin-C is thought to cause tumor cell death by inducing apoptosis. The mechanism of this action is presently under investigation. In order to incorporate MR imaging in the evaluation of this potential anti-cancer agent, the governing hypotheses were: a) Ultrasmall Superparamagnetic Iron Oxide nanoparticles can be efficiently loaded into the Saposin-C-DOPS liposomes b) The uptake of USPIO laden Saposin-C-DOPS liposomes in tumor cells can be visualized both in vivo and in vitro using high resolution MR microscopy at 7T. 17

18 In order to test these hypotheses, the following specific aims of were proposed. These are listed below: SA 1: Characterize the relaxivity of USPIO in solution and in gelatin, as a model for relaxivity in tissue. SA 2: Develop a method to efficiently encapsulate USPIO particles in Saposin-C- DOPS liposomes. a) Remove unencapsulated USPIO and verify encapsulation using both MR and non-mr methods. b) Determine the relaxivity (r2) of the Saposin-C-DOPS-USPIO in solution. SA 3: Perform in vitro experiments to determine if the Saposin-C-DOPS-USPIO liposomes are taken up by tumor cells in sufficient quantities for MR detection. a) Characterize the uptake of contrast agent using MR techniques. b) Verify using non-mr techniques. SA 4: Determine if the Saposin-C-DOPS USPIO complex can be targeted to tumors in vivo and detected using high resolution MR imaging. a) Optimize MR imaging techniques to improve sensitivity to enable detection of Saposin-C-DOPS-USPIO nanoparticles in vivo b) Perform experiments in order to evaluate the delivery of Saposin- C-DOPS-USPIO liposomes to tumors in vivo. The following chapters are sequenced broadly in the same order as the specific aims listed above, except for specific aim 1, which is covered in Appendix 1. Chapter 2 provides an overview of fundamentals of MR imaging and discusses the applications and challenges of using MRI for molecular imaging. Recent literature in the area of contrast enhanced MRI 18

19 and molecular imaging is also discussed. In Chapter 3, the preparation of the Saposin-C- DOPS- USPIO nanoparticles and the in vitro experiments on tumor cells is described, corresponding to specific aims 1,2 and 3. Chapter 4 describes the in vivo experiments in a mouse tumor model, corresponding to specific aim 4. Chapter 5 concludes the dissertation with a discussion of the results, challenges and a look toward the future of MRI and molecular-imaging. 19

20 Chapter 2 Background and Review of Relevant Topics In this chapter, the fundamental concepts of magnetic resonance imaging are outlined. Beginning with a brief history of the development of NMR, the physical principles of generation of MR images and the various mechanisms for generating contrast are described. Further, the various factors that influence the resolution and signal to noise ratio of MR images are discussed. In the latter half of the chapter, the use of contrast agents for enhancing MR images for high resolution applications is discussed. An overview of the literature and recent advances in contrast enhanced MR imaging, targeted contrast agent delivery mechanisms and molecular imaging is provided. The chapter concludes with some background information on the use of liposomes for delivery of drug or contrast agents. 2.1 Magnetic Resonance Imaging Brief History Magnetic Resonance Imaging is a powerful, non-invasive imaging modality that has gained widespread use in both clinical and research settings. Major work in magnetic resonance began with Isaac Rabi [5], who developed a resonance method for recording magnetic properties of atomic nuclei. He was awarded the Nobel Prize in Physics in The principle of Nuclear magnetic Resonance was discovered independently by Felix Bloch

21 [6]and Edward Purcell [7] in 1946, for which they shared the Nobel prize in Physics. NMR has long been used in chemistry for studying the chemical composition and molecular structure of materials. In 1971 Raymond Damadian [8] showed that the nuclear magnetic relaxation times of tissues and tumors differed, thus motivating scientists to consider magnetic resonance for the detection of disease. Paul Lauterbur [9]acquired the first MR image using linear gradient fields in Paul Lauterbur and Sir Peter Mansfield [10, 11], who developed echo planar imaging and is also credited with conception of slice selective excitation and the first human MRI, shared the Nobel Prize in Medicine in Other milestones in the history of NMR include the discovery of spin echo phenomenon in the fifties by Erwin Hahn[12], development of high resolution NMR spectroscopy using Fourier methods, by Richard Ernst[13], and development of NMR spectroscopy for determining 3-D structure of biological macromolecules, by Kurt Wuthrich [14] MR: A Quantum mechanical description Atoms which exhibit the NMR phenomenon have nuclei with odd number of protons and/or odd number of neutrons and have an associated nuclear spin angular momentum. In biological specimens, hydrogen, ( 1 H) is the most abundant and also the most sensitive. These hydrogen nuclei are often referred to as spins and can be qualitatively visualized as spinning charged spheres that give rise to a small magnetic moment. Magnetic resonance imaging is based on the interaction of these spins, with an external magnetic field. From basic quantum mechanics spins have angular momentum characterized by the spin quantum number I. Spin angular momentum is a vector quantity given by: 21

22 S = hi where ħ = 1.05 x J.s (Plank s constant, h, divided by 2π). The magnitude of spin angular momentum is given by 2-1 S = h I( I + 1) 2-2 The spin angular momentum S has an associated magnetic moment µ, with the relation ì =! S =! hi 2-3 where γ is a constant unique to each nuclear species and is called the gyromagnetic ratio. In the absence of an external magnetic field, microscopic magnetic moments are randomly oriented and the net magnetization will be zero. However, in the presence of a strong external static magnetic field (B 0 ), usually directed along the z-direction, the magnetic moments align with the static magnetic field, hence a net magnetization in the direction of the B 0. Physically, some of the spins will point opposite ( anti-parallel) to the field direction (higher energy state) and a slightly greater number of the spins will point in the same direction as the field (lower energy state) and this is seen as a net magnetic field in the same direction as B 0. Mathematically, taking the vector nature of the angular momentum, the possible values of the z-components of magnetic moment in the presence of the magnetic field along z axis are µ z = "hm, where m = I,( I! 1),( I! 2),...,! I

23 For spin ½ particles, such as protons, there are two possible values for µ z, ± ½γħ (corresponding to a spin state parallel and anti-parallel to the B 0 field, respectively). The discrete magnetic moment leads to discrete energy values given by: E " ì B# = " µ Bz = "! mh = z B Total spin, gyromagnetic ratio, relative natural abundance and relative sensitivity of common nuclei [15]. 2-5 m = -1/2 + 1/2 0!E E ~ B 0 S z m = + 1/2-1/2 0 Figure 2-1: The two energy levels of a spin one-half system placed in a uniform magnetic field (Zeeman Effect). A transition from higher to lower energy states (vertical line) results in photon emission with energy E. 23

24 For a proton, m = ± 1/2, we have two energy states. This is an example of the Zeeman Effect[16], where atomic or nuclear magnetic moments in the presence of a B-field lead to atomic and nuclear level splitting (Figure 2-1). The magnitude of energy absorbed or released by the proton spin system, upon transition between the upper and lower energy states, can be found from Eq. 2-6: $ E = E( m = # 1 / 2) # E( m = + 1/ 2) = " h B = h! 0 0 where, " 0 =! B 0 is the frequency associated with the emission or absorption of the quantum of energy and is called the Larmor precession frequency. An arbitrary state, "!, of a nucleus may be written in terms of a linear combination of basis states labeled by m "! = # a m m m! 2-8 where a m represents the amplitude or spin population in each state. The result of a measurement, called the expectation value, of the z - component of spin (I z ) is given by: # $ I $ " z =! m a m 2 m where am 2 is the normalized probability of finding a single nucleus in the state m. In a real 2-9 system we have an ensemble of spins, each occupying a particular spin state and therefore 24

25 one evaluates the average value which in turn is represented by a sum over all subensembles, each with classical probability p ψ : " $ I $! = # p " $ I $! z $ $ z 2-10 For an ensemble of spin one-half particles, such as protons where I z = ± 1/2, Eq. 2-8 and the average expectation value, Eq. 2-10, can be rewritten, respectively, as: #! = a 1 + / 2! + a" " 1/ 2! / 2 1 1/ and 1 $ & 2 ) * I z * ( = a1/ 2 ' a' 2 % 1/ 2 2! # " This last equation states that the ensemble average expectation value of I z is determined by the difference in population between the upper and the lower energy levels. From Boltzmann statistics[17], at thermal equilibrium the population of the energy levels is a m 2 =! m exp ( mh" B / k T) exp 0 B ( mh" B / k T) where k B = 1.38 x J/K is the Boltzmann constant at temperature T. The ratio and difference of the population of the two energy states for nuclei with m = ± 1/2, provided k B T "" h! B0, can be shown to be, respectively: 0 B

26 a a ' 1/ 2 + 1/ = & h( B $ ' 0 exp % k BT #! " # 1 " h! B k B T and a 2 # a# + 1/ 2 1/ 2 2 " h! B 2k B 0 T 2-15 For an ensemble of N number of spin - 1/2 nuclei per unit volume, with an individual magnetic moment ± ħγ/2, the bulk magnetization of the ensemble in the direction of the B 0 - field is given by M 0 ' & h( N 2! # $ % " 2 k B B 0 T 2-16 Absorption or emission of a quantum of energy can produce change in the spin orientation or the energy state of the tissue nuclei. This energy comes from an electromagnetic radiation or radio-frequency (RF) pulse, denoted by B 1 (t). The nuclei will absorb energy only when the RF pulse frequency ω 1 matches the Larmor frequency of the protons, ω 0. The individual protons will now precess about the B 0 at the Larmor frequency, ω 0 = γb 0, and at the same time about the B 1 at a frequency ω 1 = γb 1. In addition, the application of an RF pulse tilts or rotates the net magnetization, M 0, away from the z-axis toward the x-y plane and a spiral motion of the net magnetization vector is observed as a result. 26

27 B 0 B 0 E 2 E 1!E ~ ~ RF Energy =!E Energy Released =!E Figure 2-2 Application of a 90 0 RF pulse results in pushing some of the spins to a higher state. As a result the two states will be approximately equally populated (right), hence zero net longitudinal magnetization and maximum transverse magnetization. However, after a few moments the spins will go back to their equilibrium state by emitting an RF pulse MR: A classical approach From a classical standpoint, the MR signal is derived from the interaction of the spins with three types of magnetic fields: 1) the main field, B 0, 2) The radiofrequency (rf) field B 1, and 3) the linear gradient fields G, as described below. In the absence of the external magnetic field B 0, the spins are randomly oriented. However, in the presence of B 0, the magnetic moment vectors are aligned in the direction of B 0 to create a net magnetic moment. By convention, this is often referred to as the longitudinal magnetization, along the z-direction. The spins also exhibit resonance at a well defined frequency called the Larmor frequency, ω. In a magnetic field B, ω is given by the previously defined equation 2-7, which can be restated as: f " = 2! B

28 where γ is the gyromagnetic ratio, a unique constant for each type of atom. For 1 H, γ/2π = MHz/Tesla. The net effect of the B 0 field is to polarize the sample, inducing a net magnetization pointed in the z-direction, with strength M 0. To obtain an MR signal, a radiofrequency pulse B 1, tuned to the resonant frequency of the spins is applied in the xy (transverse) plane. The RF pulse at the Larmor frequency effectively applies a torque on the magnetization vectors, inducing nutation by an angle dependant on the strength and duration of B 1. z z M M y! B 1 B 1 x x Figure 2-3: Precession of the magnetization vector about the z axis (left) and viewed from a rotating reference frame (a) (right) y Upon turning the excitation off, the tipped vectors, which are also precessing at the Larmor frequency, begin to relax back to equilibrium. From Faradays laws of induction, the rotating magnetization vectors induce an emf in an RF receiver coil oriented to detect changes of magnetization in the xy plane. This primary MRI signal is measured in the time domain as an oscillating, decaying emf, induced by the free precession of magnetization (Figure 2-4) is known as Free Induction Decay (FID). 28

29 MRI Signal Time Figure 2-4: Free Induction Decay (FID) signal following an RF-pulse. The time constant characterizing the growth of magnetization along the z-axis during relaxation is called T 1 or spin-lattice relaxation time while the time constant characterizing the decay of the magnetization component in the xy plane is called T 2 or spin-spin relaxation time. T 1 relaxation occurs due to magnetic field fluctuations at the Larmor frequency brought about by the random motions of molecules in the surrounding medium (lattice). These molecules in motion each have magnetic moments, and the movement of these moments leads to a magnetic noise that encompasses a broad frequency range including the Larmor frequency. Magnetic noise at the Larmor frequency will stimulate transition to the lower energy state. T 2 relaxation occurs via fluctuations of a magnetic field caused by the random motion of molecules resonating at the same frequency. Fluctuation in the individual proton spins leads to a loss of phase coherence in the xy-plane with no net loss of energy from the system. Spin-spin relaxation is additionally affected by dephasing arising from bulk 29

30 inhomogenieties in B 0. The behavior of the magnetization vector! M can be described by the Bloch Equation: # d M dt # = $ M" B ext 1 + ( M T 1 0! M z ) 1 ) z! T 2 ( M x ) x + M y ) y) 2-18 where x ), y ) and z ) are unit vectors in the x,y and z directions respectively, M 0 is the equilibrium magnetization arising from the B 0 field and B ext incorporates the various magnetic fields applied, including the RF field. RF _ ss B B 0 +G z z 0 B = B 0 + G z z _ = _B BW FT B 1 (t) G z Phase _ ss /2 B 0 0 Z 0 z Signal = FID _ z (a) Sample Points (b) Figure 2-5: (a) The net magnetic field in the presence of a slice select gradient along the z-direction (G z ) versus the position (z). The slope of the gradient and the bandwidth determines the slice thickness (Δz). Since the slice is offset by z 0 the center frequency of the RF-pulse is offset from the static Larmor frequency by γg z z 0. (b) The slice select gradient has two lobes: a positive (first lobe) and a negative (second lobe). The evolution of the phase of the spins during the slice select gradient is: at the beginning and end of the first lobe spins have a negative & positive maximum phase, respectively; while at the end of the second lobe the spins are rephased. The FID signal produced after the slice select gradient is also shown. Spatial localization in MR imaging is achieved by applying linear gradient magnetic fields. For example, if a gradient, G z is applied along the z direction, then the magnetic field 30

31 strength will vary depending on the z location. The frequency of the spins now becomes a function of their z location.( Fig 2-5). Assuming that a thin slice of the sample perpendicular to the z-axis is excited by turning on the z-gradient, the MR signal from that slice, after removal of the high (Larmor) frequency component (demodulation) is given by: t t & # & # s( t) = (( m( x, y)exp$ ' i* (( Gx () d) x! exp$ ' i* (( G y () d) y! dxdy x y % 0 " % 0 " OR 2-19 s( t) = "" x y m( x, y) e! i2# [ kx ( t) x+ k y ( t) y] dxdy 2-20 where m(x,y) is the distribution of transverse magnetization in the xy plane. Variation in the phase of magnetization is achieved by applying a linear field of gradient G y, called a phase encoding gradient. The phase accumulated after a time τ y, duration of G y, is $ ( x, y) = ( #( x, y) % # 0 )! = " G y!. After G y is turned off nuclei will y y y return to the resonance frequency determined by the main field, while they retain the y- dependent phase angle. Finally, in order to distinguish precessional frequencies from different spatial locations within the selected slice, a gradient in the x-direction (G x ), called frequency encoding gradient is applied. Due to this gradient the frequency of precession will change linearly with location (Fig. 2-6): 31

32 t = 0!/2 t = 0 y G x RF t G z G y x _ (x,y ) = _G x x G x Signal x Acquisition (a) (b) (c) ON T s Figure 2-6: The slice select gradient gives the same precession frequency in the slice (a). The application of the frequency encoding gradient will provide us a position dependent precessing frequency (b). As an example, a 2D Fourier transform imaging pulse sequence diagram and the resulting FID signal is shown (c). The arrow in the phase encoding gradient is to indicate a step wise increment of the gradient, starting with a positive maximum G y. The following signal equations state that the baseband signal (i.e. signal obtained after demodulation) is a planar integral of the transverse magnetization multiplied by a spatially dependant phase factor. If linear gradients are applied, the resultant phase modulation will also vary linearly with spatial position. On comparing the signal equation with the 2-D Fourier transform equation, it is seen that k x and k y correspond to the variables in the spatial frequency domain, referred to in MR literature as k-space. Thus the distribution of transverse magnetization in the xy plane can be obtained simply by taking the inverse Fourier transform of the k-space signal. [18]. t * & # k x ( t) = exp$ ' i* (( Gx () d) x! 2+ % 0 "

33 AND k y t * & # ( t) = exp$ ' i* (( G y () d) y! 2+ % 0 " Contrast in MR Imaging Using fat and water as an example, various contrast mechanisms can be described in detail. Fat has a shorter T 1 than water. A simple saturation recovery sequence, (Figure 2-8) which consists of a string of 90 pulses each separated by TR, produces a signal dependent on the size of M z prior to the 90 pulse. The signal intensity from water or fat is given by the equation: I( x, y) = K" ( x, y)[1! e! TR / T1( x, y) ] e! TE / T 2( x, y) 2-23 where K is the constant of proportionality, ρ is the spin density and I(x,y) is the intensity of the signal from either fat or water in the voxel at (x,y). In a spin-echo sequence, a 180 pulse is applied along the y axis (rotating frame of reference) after the 90 excitation pulse is applied along the x axis., which serves to reverse the decay due to magnetic field inhomogenieties (Figure 2-10). After the 180 flip, the spins refocus and produce an echo at time TE, the amplitude of which depends in the intrinsic T 2. A T2 weighted image can be obtained by using long TR to avoid T 1 contrast and setting TE ~T 2 to highlight T 2 differences. Water has a longer T 2, i.e., the spins take longer to dephase than fat. So water appears brighter on a T 2 weighed image compared to fat. From Figure 2-7 it is evident that fat produces a larger MR signal and will appear brighter in a T 1 weighed image. T 1 weighting is achieved by using short TE and TR~T 1. Then 33

34 the T 2 relaxation term exp(-te/t 2 ) ~1. To obtain T 2 weighted contrast while avoiding problems with dephasing, spin echo sequences can be used, where the readout at time TE will depend on intrinsic T 2 not T 2 *. (a) (b) Figure 2-7: (a)mechanism of T1 contrast and difference in recovery time of the magnetization vector in fat and water. (b)mechanism of T2 weighted contrast (Top-right) and difference in T2 in fat and water (bottom-right). 34

35 !/2!/2!/2!/2!/2!/2 RF TR TE RF TR G z FID G y G x Signal Echo Acquisition (a) ON TR (b) Time Figure 2-8: Pulse sequence diagram for saturation recovery imaging implemented in gradient echo mode (a) and T 1 recovery curves for short TR (partial recovery) (b). For long TR one can predict the recovery of T 1 to its plateau value and the resulting PD weighted images. Dephasing of the magnetization in the transverse plane caused by field inhomogenieties produces an additional suppression of the signal apart from the natural T 2 responses. This decay, characterized by the time constant T 2 * is always much faster than T 2 processes (Figure 2-9). Thus the decay time T2* represents a combination of external field induced (T2 ) and thermodynamic (T2) effects: 1 T * 2 1 = T T ' Figure 2-9: T2* decay is always faster than T2 processes. 35

36 The additional signal loss can be overcome by shimming (i.e. correcting the inhomogenieties in the main magnetic field) and by using a pulse sequence with a 180 refocusing pulse after the 90 excitation pulse (i.e. a spin-echo sequence). However, T2* mechanisms can also be effectively used to generate contrast in images, especially when local magnetic field susceptibility differences between tissues are present. Such susceptibility differences can also be induced in the tissue of interest using contrast agents. Gradient echo imaging methods, i.e. methods that do not use a refocusing pulse, can be used to highlight the T2* differences in tissue[18]. Since there is no 180 refocusing pulse, the signal decay follows the T2* envelope rather than the T2 envelope (Figure 2-11). To briefly summarize the acquisition process for a T2* weighted echo, the RF pulse is applied simultaneously with the slice select gradient. This RF pulse produces a rotation of the magnetization into the transverse plane, typically from 10 to 90. Then the phase encoding gradient is turned on and a dephasing frequency encoding gradient is applied at the same time as the phase encoding gradient so as to cause the spins to be in phase at the center of the acquisition period. This gradient is negative in sign from that of the frequency encoding gradient turned on during the acquisition of the signal. An echo is produced when the frequency encoding gradient is turned on because this gradient refocuses the dephasing which occurred from the dephasing gradient. The echo time (TE) is defined as the time between the start of the RF pulse and the maximum in the signal. The sequence is repeated every TR seconds. The phase encoding gradient is varied between a maximum and a minimum value for each repetition. Thus, in each acquisition, which takes time TR, one line of k-space data is acquired. 36

37 Such methods are ideal for imaging of tissues in which the T 2 * has been reduced by using contrast agents. Sequences such as FLASH (Fast Low Angle SHot) [19] are based on elimination or spoiling of transverse magnetization before the next rf pulse. The reconstructed signal is given by the relation: 1! e I( x, y) = # 0 ( x, y)sin" 1! e! TR / T1! TR / T1 e cos" *! TE / T 2 where I (x,y) is the image intensity and θ is the flip angle. 2-25!/2! x x z z z t = 0 t = 0 + t =! M 0 B 1 z t = _ !/2- Pulse Dephasing M 0 x x!-pulse Rephasing (a) x z t = 2 _ = TE y M 0 exp( -TE/T 2 ) 1 2 B y RF G z G y G x Echo Acquisition _ (b) _ ON Figure 2-10: (a) Spin-echo generation: the π/2-pulse flips the magnetization onto the y -axis and all spins are in-phase. However, as time goes on the spins begin to dephase because of precessional frequency differences. At time t = τ a π-pulse rotates all the spins about the x -axis and the spin vectors continue to precess with reduced phase differences in time till they rephase once again at time t = 2τ. (b) SE pulse sequence diagram: we note the slice select gradient is applied during both pulses. The first positive gradient in G x (called pre-readout gradient) gives a positive phase difference (maximum) that will stay the same till the π-pulse is applied, which will reverse the phase difference. Then it will remain constant until the read gradient (the second) is applied. During the read gradient, spins will begin going back in phase, reaching a zero phase difference at TE [18]. 37

38 RF TR TE Spins G z Variable spoiler G x (Read) G y Out of Phase In Phase Out of phase Echo G x Echo Acquisition ON ON (a) (b) Figure 2-11: (a) Gradient echo formation: instead of a π- pulse, a bipolar read gradient with a refocusing part twice the area of the dephasing part, is used to form an echo (called gradient echo) [20]. (b) Gradient spoiled GRE (SPGR) pulse sequence diagram: an additional gradient that varies from cycle to cycle is used to spoil the transverse magnetization from a previous cycle. 2.3 MR Microscopy MR microscopy refers to high resolution MR imaging, usually of the order of tens of microns for in vitro samples and 100 micron isotropic resolution or better for in vivo systems. Several factors determine what the attainable resolution is for a given MR experiment. The nuclei being imaged is usually hydrogen ( 1 H) contained primarily in water, which is abundant in biological systems. The more nuclei a given voxel contains, the higher the signal from the voxel. The MR signal is also dependant on the imaging parameters, pulse sequence, type of encoding etc. Different pulse sequences and encoding systems tailored to observe different phenomena like diffusion, perfusion or susceptibility artifacts will yield different resolutions. Finally, resolution also depends on the hardware i.e. the magnet field strength, the gradient system and the RF transmit and receive systems. The field strength of the scanner determines the polarization, as given in equation 2-6 and hence affects the SNR. At a higher resolution, the voxel size and hence the number of nuclear spins available is reduced. However, by increasing the field strength, the nuclear spin population differential between 38

39 the two energy states is increased as shown in Figure 2-1. SNR increases roughly linearly with magnetic field. Stronger gradients are required to achieve higher resolution images. This is depicted in Figure 2-12 for the slice selection gradient. Stronger, faster switching gradients enable finer spatial encoding of spins. frequency G z1 G z2!f!z1!z2 Position Figure 2-12: Plot of Larmor frequency vs position along the direction of the slice select gradient. The slope of the two lines G z1 and G z2 represent the strength of the gradient. For any given bandwidth f of the rf pulse, the stronger gradient produces a thinner slice MR Imaging Parameters and Image Optimization Although noise cannot be completely eliminated, there are ways to increase the signal to noise ratio (SNR). The SNR is defined as: SNR = Average Signal / Standard Deviation of the Noise 2-26 The major source of noise in MRI is thermal noise from the sample. This is due to the random thermal vibration of ions, electrons, etc. in the sample giving random transitions between energy states. Coils that are small in size and closer to the body will give improved 39

40 SNR as compared to large coils, because large coils cover a larger sensitive volume and hence larger random transitions are seen by it. Other sources of noise are thermal noise in the RF coil and electronics noise. RF coil noise is also reduced with the coil dimensions The SNR is dependent on voxel size, the number of measurements and receiver band width according to the expression: SNR = K! Voxel Size! Number of Measurements Receiver Band Width 2-27 where K is a proportionality constant that roughly varies linearly with B 0. For a 3D imaging the number of measurements is the product N x N y N z NEX, where N x,y,z are the number of k- space samples in the direction of each gradient and NEX is the number of excitations. In terms of the pixel sizes x, y, & z, the voxel size is x y z for 3D imaging and x y for 2D imaging. The k-space parameters are related to field of view (image size in cm) along the x, y and z-axes (FOV x,y,z ) by FOV x FOV z = = BW " G BW rec x " G rec z = = 1 " G! T x 1 "# G z z s = = 1! k 1! k z x,, FOV y = BW " G rec y,max = 1 "# G y y,max = 1! k y, and 2-28 where BW rec = the receiver band width, G y,max = the maximum amplitude of the phaseencoding gradient, T s = the sampling interval during frequency encoding, (the time between successive sampling points), and τ y,z = duration of time the gradients G y and G z 40

41 are on, respectively. The pixel sizes/resolutions in the image, given in terms of k space parameters, are! x FOV x y = = =,! y = =, and! N x 1 " G T x s 1 k x FOV N y 1 k y z = FOV N z z = 1 k z, 2-29 where T s = N x T s = acquisition or sampling time. Using Eqs and 2-29, the SNR equation (Eq.2-27) can be rewritten as FOV FOV FOV = Ts. N N N x y z SNR K " " " " NEX "! x y z 2-30 Note that for 2D imaging the quantity FOV z / N z should be replaced by the thickness of the slice. The SNR can be improved by increasing the NEX, the number of measurements. However, there is a tradeoff: Increasing the NEX increases the total scan time, Scan Time = TR! N! N NEX. y z! 2-31 Also, keeping FOV constant and decreasing N y improves the SNR but at the expense of resolution ( y = FOV y /N y ). We can keep pixel size y constant by increasing both FOV y and N y. In this case the SNR increases but the scan time increases as well. Narrower bandwidth (BW rec = 1/ T s = N x /T s ) will also increase the SNR, but a longer sampling time results in a longer TE and hence a decreased signal as a result of increased T 2 dephasing. In addition, at high resolutions, stronger, faster-switching gradients are required to overcome signal degradation from broadening induced by susceptibility effects and diffusion. MR spectral line width is defined as the full width half maximum (FWHM) of the Lorentzian function obtained as by taking the Fourier transform of the FID and given by: 41

42 1 FWHM =!T * Susceptibility mismatches create local field inhomogenieties, ( B 0 ) and increase line width, resulting in signal attenuation. These susceptibility effects can be overcome by using large gradients such that:! 0 B =<< G! r 2-33 Where B 0 is local variation in the magnetic field generated over a pixel of dimension r and G is the strength of the gradient along that direction, Finally, improving the sensitivity if RF coils improve the SNR, enabling higher resolution scans. The following equation shows the dependence of SNR on various parameters of the RF circuit: SNR ' ( 0 & B $ % i R 1 #! V " noise S 2-34 where ω 0 is the resonant frequency, B 1 /i is the coil sensitivity defined as the transverse magnetic field generated by the coil per unit volume, V s is the sample volume and R noise is the noise resistance from the sample and the coil. In conclusion, this section provides the reader with an overview of the principles of magnetic resonance imaging and the challenges of MR microscopy. Readers are referred to Magnetic Resonance Imaging by Mark Haacke [18] and Principles of Nuclear Magnetic Resonance Microscopy by Paul T Callaghan [21]for an in-depth treatment of these topics. 42

43 2.4 Molecular and Cellular Imaging using MRI Contrast Agents Inherent contrast in MR images can be improved with the use of paramagnetic contrast agents. Most contrast agents reduce both T 1 and T 2. Agents are classified as T 1 agents if they shorten T 1 more than T 2, and T 2 agents if they affect T 2 more than T 1. Positive contrast agents are those that produce an enhancement of the signal or brightening of the tissue of interest, while a negative contrast agent produces a loss of signal and darkening in the tissue of interest. The relaxation properties of contrast agents in solution are often characterized by relaxivity, defined as the ability to reduce relaxation rates R 1 =1/T 1, R 2 =1/T 2 and R 2 *=1/T 2 *. Relaxivities, r1, r2 and r2* are described by the concentration normalized relaxation times, usually in units of (moles second) Superparamagnetic Contrast Agents An atom with an unpaired electron has a non-vanishing permanent magnetic moment with an associated nonzero dipole magnetic field and is referred to as paramagnetic. The moments tend to align with an external magnetic field, producing a bulk magnetic moment. Ferromagnetic materials, on the other hand, have permanent domains of electron spin magnetic moments which combine to form strong macroscopic self-fields existing independently of external fields( e.g. Fe, Co, Ni). These domains are randomly distributed throughout the material. When they are aligned by the temporary application of an external field, a permanent macroscopic magnet is produced. The act of continuously subdividing ferromagnetic material produces particles that are each just one domain. These materials are 43

44 called Superparamagnetic and consist of individual domains of elements that have ferromagnetic properties in bulk. Their magnetic susceptibility is between that of ferromagnetic and paramagnetic materials. With no external field, their thermal motion leads to vanishing magnetizations. However, with an external field, the alignment of these domain particles produce a strong self-field, in a manner analogous to ferromagnetic materials. Superparamagnetic contrast agents are composed of a water insoluble crystalline magnetic core, usually magnetite (Fe 3 O 4 ) or maghemite (γ-fe 3 O 4 ). The mean core diameter ranges from 4 to 10 nm. This crystalline core is often surrounded by a layer of dextran or starch derivatives. The total size of the particle is expressed as the mean hydrated particle diameter. USPIO, Ultrasmall Superparamagnetic Iron Oxide nanoparticles, which usually have single crystal cores, have a mean hydrated particle diameter less than 50 nm. SPIOs on the other hand, have aggregated iron oxide crystal cores and have a diameter larger than 50nm Cellular and Molecular Imaging Using MRI Spatial resolution in MR imaging is defined by the strength of the three gradients and the size of the receiver coil, the field of view (FOV) and the matrix dimensions that cover the region of interest, as described in section 2.3. Steeper gradients enable acquisition of thinner slices and smaller fields of view but at the cost of less signal due to smaller voxels. However, using a higher magnetic field and smaller receiver coils, some improvement in signal to noise ratio (SNR) is observed. Weissleder et al [22, 23] define molecular imaging as the in vivo characterization and measurement of biologic process at the cellular and molecular level. Some of the challenges in this area include: 1) development and use of high affinity probes with reasonable 44

45 pharmacokinetics. 2) Overcoming vascular, interstitial, cell membrane or blood-brain barriers to deliver the probes to the desired site. 3) Optimized, fast and high resolution imaging techniques. In recent years, MR resolution has been pushed to the order of tens of microns, usually with the help of MR contrast agents like USPIO and chelated gadolinium. MR imaging at 100µm isotropic spatial resolution or better is generally considered MR microscopy[21]. In the following sections, a review of recent and relevant literature primarily in the area of MR contrast enhancement using iron-oxide based contrast agents is provided In vivo labeling of cells with iron oxide contrast Imaging using iron oxide contrast agents was first introduced for hepatic imaging[24]. In vivo studies [25] have shown that mononuclear phagocytic system (MPS) cells or macrophages internalize USPIO by absorptive endocytosis. When injected into the blood stream, the USPIO particles are taken up by the cells of the reticuloendothelial system (RES), a network of macrophage cells lining blood vessels whose function is to remove foreign substances. The size of the particles also has an effect on the clearance rate and hence halflife of the particles in blood. Depending of the half-life of the particles, they end up in the liver s Kupffer s cells, lymph nodes spleen and bone marrow. Tumor cells in the liver, lymph nodes and spleen do not have an effective reticuloendothelial system, so their relaxation rates are not altered by USPIO. This has been useful in the identification of malignant lymph nodes and liver tumors[26]. For improved uptake and internalization into the cytoplasm in non-phagocytic cells, optimized methods need to be developed for delivery of the USPIO particles to the desired site. Early attempts by Bulte et al[27, 28] and Josephson et al [29] used magnetoliposomes or 45

46 lectins to increase uptake in non-phagocytic cells. An improvement in labeling was achieved by Lewin et al[30] by linking the USPIO particles to the HIV tat peptide, which contains a membrane translocating signal that efficiently transports the USPIO in particles into cells. Dodd et al [31] have demonstrated the labeling of T-cells with SPIO particles causing (negative) contrast enhancement in labeled cells in gelatin. Another promising application of USPIO is in the monitoring of implanted stem cells in vivo. The implantation of cultured stem cells such as embryonic stem cells into the brain holds great promise in treating several pathological conditions. Hoehn et al [32] were able to monitor the migration of embryonic stem cells containing USPIO toward ischemic regions within the rat brain in vivo using MRI. The labeling of cells was achieved by a lipofection technique generally used for infusion of DNA into cell nuclei using lipofection reagent FuGENE. Figure 2-13: Lipofection technique used by Hoehn et al [32]. (Figure obtained from Franklin et al [33]were able to monitor in vivo the activity and migratory capacity of USPIO labeled oligodendrocytes implanted in the rat brain, for up to 7 days after implantation. Zhao et al [34, 35] have created a contrast agent that binds to apoptotic cells. They labeled the C2 domain of synaptotagmin I with SPIO nanoparticles and show that it is 46

47 possible to detect apoptotic cells both in vivo and in vitro using this novel contrast agent. Rausch et al [36] were able to detect USPIO enhancement in macrophages after ischemic brain damage cause by occlusion of the middle cerebral artery in the rat. Bulte et al [37] developed a versatile class of magnetic tags called magnetodendrimers which allow them to efficiently label mammalian cells including human neural stem cells Liposomes Liposomes are microscopic, fluid-filled vesicles whose walls are made of layers of phospholipids similar to the phospholipids that make up cell membranes. These spherical vesicles form when phospholipids are hydrated. In water under low shear conditions, the phospholipids arrange themselves in sheets, the molecules aligning side by side in like orientation, hydrophilic end up and hydrophobic end down. These sheets then join, hydrophobic end to hydrophobic end, to form a bilayer membrane which encloses some of the water in a phospholipid sphere[38]. Typically, several of these vesicles will form one inside the other in diminishing size, creating a multilamellar structure of concentric phospholipid spheres separated by layers of water. Utilizing ultra high-shear processing, liposomes which are of unilamellar structure, a single phospholipid bilayer sphere enclosing water can be produced. Besides being much smaller than multilamellar liposomes, these unilamellar liposomes are of uniform size, usually 200 nanometers or less in diameter. Liposomes have a long history in the study of biological membranes and they have become versatile tools in biology, biochemistry and medicine[39]. Liposomes are used to deliver vaccines, enzymes, or drugs (e.g., insulin and some cancer drugs) to the body[40, 41]. When used in the delivery of certain cancer drugs [42, 43], liposomes help to shield healthy 47

48 cells from the drugs' toxicity and prevent their concentration in vulnerable tissues (e.g., the kidneys, and liver), lessening or eliminating the common side effects of nausea, fatigue, and hair loss[44, 45]. Liposomes are especially effective in treating diseases that affect the phagocytes of the immune system because they tend to accumulate in the phagocytes, which recognize them as foreign invaders. They have also been used experimentally to carry normal genes into a cell in order to replace defective, disease-causing genes[46], a process referred to as lipofection. In some cases liposomes attach to cellular membranes and appear to fuse with them, releasing their contents into the cell. Sometimes they are taken up by the cell, and their phospholipids are incorporated into the cell membrane while the drug trapped inside is released. In the case of phagocytic cells, the liposomes are taken up, the phospholipid walls are acted upon by organelles called lysosomes, and the medication is released into the cytoplasm. Liposomes can be custom designed for almost any need by varying the lipid content, size, surface charge and method of preparation. The preparation of liposomes requires careful attention to ease of formulation, encapsulation efficiency and production capacity. Liposomal delivery systems are still largely experimental; the precise mechanisms of their action in the body are under study, as are ways in which to target them to specific diseased tissues Liposomes as Vectors for Contrast Agents Bulte et al [27] were successful in labeling peripheral blood mononuclear cells (PMBCs) by incubating them with large unilamellar vesicles containing encapsulated dextran magnetite particles. The liposomes were made from Egg yolk l-alpha-phosphatidylcholine, cholesterol and bovine phosphatidylserine and seven cycles of freeze thaw extrusion was 48

49 used to prepare the liposomes. Encapsulation of the dextran-magnetite particles in the liposome as well as their uptake into the PMBCs was confirmed by electron microscopy. Previous attempts have also been made to target liposomes to tissue in vivo. Knut- Egil et al [47-49] have developed and characterized paramagnetic ph sensitive liposomes for monitoring pathologic changes in ph using MRI. They use the ph sensitive system dipalmitoylphosphatidylethanolamine/ palmitic acid (DPPE/PA) combined with gadolinium diethylenetriamine pentaaceticacid bismethylamide (GdDTPA-BMA). They found that the incorporation of cholesterol increased the stability of DPPE/PA in blood but decreased the ph sensitivity. However, exchanging the PA with the double chained amphiphile dipalmitoylglycerosuccinate (DPSG) yielded liposomes with improved properties. Liposomal systems have been used as a dual drug delivery and imaging contrast agent vehicle by Viglianti et al [50]. The purpose of their study was to determine if MnSO 4 /doxorubicin loaded liposomes could be used for in vivo monitoring of liposome concentration distribution and drug release using MRI. They used a temperature sensitive liposome formulation and MnSO 4 release in tissue was found to significantly lower T1. The feasibility of monitoring drug uptake via the T1 shortening effect of MnSO 4 was shown in a murine tumor model and the thermally sensitive liposomes showed a clear pattern of accumulation at the periphery of the tumor Magnetoliposomes Bulte et al [51] note that two prerequisites have to be fulfilled for liposome based formulations to be useful in imaging application: the contrast agent should remain stable or bound to the liposome without dissociating in vivo. Secondly, since liposomes are also seen as foreign entities within the body, they are removed from blood circulation by macrophages. 49

50 It has been shown [52] that adding a small percentage of phospholipids that bear a polymeric poly-ethylene glycol (PEG) chain can prevent rapid uptake of the liposomes in vivo. Magnetoliposomes, in which the aqueous interior of the liposome was occupied by magnetic iron oxide, was synthesized by Cuyper et al [53]. In the classical encapsulation procedure, the material to be encapsulated is present during membrane formation. This method is more difficult for larger particles like iron oxide crystals. Several methods to improve the capture efficiency of iron oxide particles in liposomes have been described in the literature [51, 54-59]. The basic procedure is to mix lauric-acid stabilized magnetite/maghemite cores with the lipid in organic solvent, followed by removal of the organic solvent by nitrogen gas or vacuum drying. Aqueous phase is added followed by brief sonication resulting in the formation of vesicles containing the iron oxide cores. Other methods include detergent dialysis, reverse phase evaporation, extrusion and freeze-thawing [60]. However, the encapsulation of dextran coated iron oxide particles is more complicated due to the larger size of the particles and the tendency of these particles to aggregate in organic solvent. A novel approach to increase the efficiency of encapsulation of dextran coated iron oxide particles and to obtain homogeneous non-aggregated population of superparamagnetic liposomes was developed by Bogdanov et al [61]. They showed that a reversible covalent attachment of the dextran coated iron oxide particles to aminophospholipids during the preparation of water/lipid emulsion would increase entrapment of colloids in reverse-phase evaporation vesicles (REV). This was achieved by treating the mono-crystalline iron oxide (MION) particles with sodium periodate to generate aldehyde groups on the dextran surface. These groups are known to react with amino groups of aminophospholipids with the formation of covalent Schiff bonds. An efficiency of up to 49% Fe trapped in lipid was achieved. 50

51 Ryuta Saito et al [62]used liposomes as a vector for convection enhanced delivery (CED) of contrast into rat brain tumor models. They encapsulated the gadolinium based contrast agent gadodiamide in liposomes, which was then injected into rat brains. The distribution of the liposome-contrast complex in the rat brain tumor model was monitored in vivo using MR imaging. Minimum concentration of Gd necessary for detection was determined from in-vivo imaging experiments. In another set of experiments, they infused the drug Doxil, used to treat several cancers, in the liposome contrast complex, to determine optimum convection enhanced delivery parameters for complete targeted coverage of the tumor. Arbab et al [57, 59, 63] have used various concentrations of ferumoxides(fe)-ploy-l- Lysine complexes to magnetically label cells in-vitro. The negatively charged ferumoxide particles (Feridex IV; Berlex Labs Inc.) were complexed with the polycationic transfection agent poly-l-lysine. Several different types of cells, including human cervical carcinoma cells and human mesenchymal stem cells were labeled by endocytosis the FE-PLL complex. Sufficient iron was taken up by the cells that significant signal intensity loss was detected with as few as 1000 cells in a 1.5T MR scanner. Pauser et al [64] tested three different types of liposomes, REVs (reverse phase evaporation vesicles), SUVs (small unilamellar vesicles) and SUV-PEGs (small unilamellar vesicles stabilized with polyethylene glycol) as vectors for transporting iron oxide cores into tumor cells, in-vitro and in-vivo. They characterized the relaxivities of the contrast agent in liposomes at various concentrations at 0.46 Tesla and also estimated the encapsulation of iron in the liposomes to be less than 1% of initial iron concentration. In-vivo experiments were performed in rats implanted with CC531 adenocarcinoma in the liver. For all types of 51

52 liposomes, iron was found in the necrotic areas of the liver. In imaging experiments, the first maximum of signal intensity (SI) reduction was attributed to diffusion and perfusion of the free USPIO, while the SI reduction after 24 hours was attributed to internalization of the USPIO into cells after interaction of the liposomal vesicles with the tumor cell membranes. Among recent literature in iron oxide enhanced MR microscopy and cellular imaging, two articles, one on the detection threshold of single SPIO labeled cells by Heyn et al, [65] and the other on the use of micron sized iron oxide particles for cellular imaging by Shapiro et al [66, 67], are noteworthy. The detection threshold for single SPIO labeled cells and the effect of resolution and SNR for a balanced steady-state free precession (SSFP) [18, 68] sequence (3D-FIESTA) are described in [65]. An expression that predicts the minimum mass of SPIO required to detect a single cell against a uniform signal background is derived. Shapiro et al [66, 67] studied the MRI properties of single micron-sized iron oxide particles (MPIOs) with sizes ranging from 0.96 µm to 5.80 µm. The capacity of cells to endocytose these MPIOs was investigated and the MRI properties of the labeled cells were measured at 7 T and 11.4 T. Cells labeled with MPIOs contained on about 100 pg of iron on an average, which is approximately three times the concentration obtained with the best strategies to label nanometer size particles, i.e. USPIOs. An important aspect of targeted drug delivery of either drugs or imaging agents to tissue in vivo is the selectivity of the moiety. The first step is to identify the target. The target should be expressed in sufficient quantities, should be specific to a certain disease and should be accessible to the drug/imaging agent. Once a target has been identified, a high affinity ligand can be chosen for the target. These ligands include proteins, peptides and antibodies. 52

53 Figure 2-14: Above figure shows the principle of targeted contrast imaging. The imaging agent is coupled with the ligand using a carrier. The ligand is targeted to the receptor on the cell surface. The imaging label then usually taken up by the cell by receptor-mediated endocytosis. The drug or imaging agent is then conjugated with the ligand using one of many available approaches including including periodate oxidation/borohydride-reduction to form Schiff bonds[69], gluteraldehyde crosslinking[70], the biotin streptavidin system[71] and amine-sulfhydryl group linkage[72]. A review of lipid based nanoparticles for contrast enhanced MRI is available by Mulder et al [73]. Attractive molecular targets in cancer include markers for angiogenesis like αvβ3 integrins and Vascular Endothelial Growth Factor (VEGF) receptor and markers for apoptosis, including phosphatidylserine. Schmieder et al [74] target the αvβ3 integrins expressed on neovasculature in melanoma xenografts in a mouse model using Gadolinium based contrast particles conjugated to an αvβ3 integrin antagonist (a thiolated peptidomimetic vitronectin antagonist) and show that the MR signal 53

54 enhancement in the tumor is approximately 50% better compared to non-targeted nanoparticles. Mantyla et al [75] use biotinylated USPIO particles to target Scavidin expressing human umbilical vein endothelial cells (HUVEC) cultures in vitro. By coupling biotinylated USPIO with the novel fusion protein Scavidin consisting of macrophage scavenger receptor class A and avidin, and using T2 weighted MR imaging, they show scavidin expressing cells were capable of binding and mediating endocytosis of USPIO in vitro. In another study, van Tilborg et al [76] use an Annexin A-5 functionalized lipid based contrast agent to detect apoptosis in tumor cells. They use Gd-DTPA-BSA lipids incorporated into the lipid bilayer of pegylated liposomes. Multiple human recombinant annexin A5 molecules were covalently coupled to introduce specificity for apoptotic cells based on the affinity of annexin A5 for phosphatidylserine, which is expressed on the exoplasmic membrane of apoptotic cells. Work by Deans et al [77] shows yet another approach for visualizing cells using contrast enhance MRI. They induce MRI contrast in an in vivo system via expression of the transferrin receptor ( TfR) and ferrtin (FTH) proteins in a transfected mouse neural stem cell line. When grown in an iron rich medium, the transgenic cells accumulated significantly more iron than control cells and become MR visible. 2.5 Issues for detection of iron oxide contrast using MRI An important consideration for the detection of iron oxide in cells is the concentration required for detection. SPIO contrast agents, with magnetic moments more that 3 orders of magnitude greater than paramagnetic contrast agents, behave differently from paramagnetic contrast agents under compartmentalization. Compartmentalization causes a dramatic 54

55 increase in the R2* to R2 ratio compared to the values obtained when uniformly suspended in gel. This is explained by the static dephasing regime theory[78]. Figure 2-15: Prediction of the minimum Fe (pg) required for detection for various SNRs (above) and minimum fconcentration of Fe for detection plotted against SNR for various concentration. The minimum Fe for detection scales linearly with voxel volume decreases with increasing SNR following a hyperbolic relationship. (Figure reproduced from Heyn et al [65] with author s permission) Work by Heyn et al [65] shows the relationship between the minimum detectable concentration of iron compartmentalized in a cell and the voxel volume and SNR. As shown in Figure 2-15 (a), the minimum iron concentration required for detection ( m det c in pg) scales linearly with voxel volumes. This means that smaller voxel volumes (i.e. higher resolution) 55

56 enables detection of smaller quantities of iron. Similarly, higher SNR enables detection of smaller quantities of iron (Figure 2-15 (b)). The challenge is to simultaneously improve the SNR and resolution. Reducing the voxel size for higher resolution results in less signal and therefore reduces the SNR. This has to be compensated for by using higher field strengths (B 0 ), optimized receiver coils and by signal averaging. As shown in section 2.3, signal averaging provides diminishing returns on SNR since doubling the SNR requires quadrupling the imaging time. 2.6 Summary Several approaches to contrast enhanced MR imaging for various applications have been discussed in this chapter. Ultimately, each application or disease model requires a unique strategy for labeling and imaging, keeping in mind the goals and the usefulness of the data obtained. Visualizing the delivery of the Saposin-C-DOPS liposomes to tumors using MRI requires combining the existing delivery vehicle (Saposin-C-DOPS) and a suitable contrast agent (USPIO). In the next chapter, the method for encapsulation of MR contrast (USPIO) in liposomes is described. The uptake of these contrast laden liposomes in tumor cells is verified by MR techniques and other techniques. 56

57 Chapter 3 In Vitro Experiments This chapter begins with a brief review of some preliminary experiments. The remainder of the chapter is based on the manuscript submitted to Magnetic Resonance in Medicine entitled MR detection of targeted delivery of USPIO and Saposin-C to tumors using liposomes. The procedure for the encapsulation of USPIO in Saposin-C DOPS liposomes, experiments conducted on in vitro cell cultures and methods for quantifying the results obtained are described. 3.1 Overview of research goals The protein Saposin C and the lipid DOPS In this dissertation, we attempt to use MR imaging to detect the delivery of liposomes to deliver both particles of Ultrasmall Super-Paramagnetic Iron Oxide (USPIO) contrast agent and the protein Saposin-C to the tumor in a mouse model. Saposin-C and dioleoylphosphatidylserine (DOPS or PS) are two molecules that are naturally synthesized in humans. PS is a phospholipid component of cell membranes and is also commercially available (Avanti Polar Lipids, Alabaster, AL). Saposin-C is a small heat-stable, proteaseresistant and lipid-associated non-enzymatic glycoprotein containing approximately 80 amino acids. At acidic ph, Saposin-C binds to PS to form a proteoliposome compound (SCPSPLs) of size nm[79]. In preliminary studies by Qi et al at Cincinnati Children s Hospital, this complex has been shown be taken up specifically by tumor cells 57

58 through fusion with the cell membrane followed by endocytosis. Further, these preliminary data also indicate that the Saposin-C-DOPS liposomes induce apoptosis of a variety of human cancer cells, without affecting normal cells, the mechanism of which is currently under investigation Mechanism of uptake in tumor cells Neoplastic cells are metabolically more active and proliferate more rapidly than normal cells. Cancer cells produce significant amounts of acid including products of anaerobic glycolysis and carbon dioxide during aerobic respiration[80]. Hydrogen ions accumulate in the respective tissue, if a high glycolytic rate and a high lactic acid production coincide with an insufficient drainage by convective and/or diffusive transport. As a result, the mean of extracellular ph in solid tumors is lower than that in the normal tissues of origin. The mean ph values obtained in most human tumors were as low as 6.15 [80]. Most of the chemotherapeutic agents in clinical use are directed primarily at the cell nucleus. In contrast, some agents that primarily interfere with cellular membranes modulate membrane organization and fluidity, membrane lipid biosynthesis and metabolism. In the last two decades, cellular membranes have become novel targets for anticancer drugs [81]. Some tumors, squamous cell carcinomas and astrocytomas, have even lower ph values of less than 6.0. Such an acidic environment provides the basis for the cancer cell specificity seen with use of the Saposin C-DOPS complex. Two key factors are important to lead the specific tumor targeting of SCPSPLs. First, tumor blood vessels are leaky. Hypoxia is an epigenetic factor that stimulates expression and release of vascular endothelial growth factor (VEGF) from tumor cells. VEGF is known as vascular permeability factor. Thus, tumor vasculature is disorganized and leaky when compared to normal vessels. Unlike the tightly packed vascular endothelium of blood vessels in most normal tissues, there are gaps as large as nm between adjacent 58

59 endothelial cells in angiogenic blood vessels in tumors[82]. These gaps allow SCPSPLs pass through into tumor cells. Second, PS lipids have been found to be exposed on the surface of human tumor blood vessels and cells. More than 5 to 10 fold more PS is present on the surface of tumor cells than cells of normal tissues and vessels[83, 84]. SCPSPLs have strong affinity and interaction with PS lipids in the biological membranes. PS exposure on the outer leaflet of membranes will provide a favorable platform for SCPSPLs function. Once SCPSPLs have leaked out through tumor blood vessels, the strong affinity between SCPSPLs and PS in tumor cells and blood vessels increases the targeting specificity and retention of SCPSPLs in tumors. 3.2 Preliminary Experiments Several preliminary experiments were conducted to gather data and estimate the difficulties of the research project and establish the specific aims of the dissertation. These are listed below. 1. Measurement of relaxivity of USPIO in saline and in gelatin 2. Fluorescence tagging of USPIO particles with Flurescin (Green Fluorescent Protein) 3. Encapsulation of USPIO in liposomes by simple sonication. 4. Simulation of the FLASH signal equation with varying TE/TR/Flip Angle to determine the dependence of contrast on these parameters. These preliminary experiments helped determine which approaches would be suitable and which ones would have to be abandoned in trying to accomplish the specific aims of the research project. Measurement of relaxivity of USPIO in saline and gelatin helped in understanding the T2/T2* shortening effect of USPIO and in optimizing the imaging 59

60 sequences used for measuring the relaxivity of various phantoms. The measurements are included in Appendix I. In order to create a dual tagged agent, an attempt was made to label USPIO with the green fluorescent probe Fluorescin. However, this approach was later abandoned after developing a chemical coupling procedure for the encapsulation of USPIO in DOPS liposomes, as the fluorescent label would be quenched and rendered ineffective. Initially, attempts were made to encapsulate USPIO in DOPS liposomes using simple sonication techniques. Lipid in organic solvent was dried in a standard round bottom test tube using slow flow of N 2 gas. USPIO solution was then added io the dried lipid and the mixture was sonicated for 10 minutes in a bath type sonicator. The solution was passed through a con-a sepharose 4B column to separate the free USPIO from the liposomes. This procedure was found to be ineffective in encapsulating USPIO in the liposomes. This led to an exploration of other methods to increase encapsulation efficiency. The imaging approach that would provide the most sensitivity for detection of USPIO/SPIO contrast agents are T2* weighted gradient echo techniques. The spoiled gradient echo imaging sequence in the Bruker Biospin 7T small imaging scanner, known as FLASH (Fast Low Angle Shot) imaging protocols were tested on several imaging phantoms for detection of USPIO contrast. The FLASH signal equation was also simulated in MATLAB to study the effect of varying different parameters like TE, TR and flip angle. 3.3 Encapsulation of USPIO in DOPS-Sapsoin-C liposomes Method In order to encapsulate USPIO (Combidex, Advanced Magnetics, MA, size ~20-30 nm) efficiently in liposomes, a chemical binding method described by Bogdanov et al [61] was adapted. Briefly, the dextran coating on the USPIO particles was oxidized to generate aldehyde groups. Typically, 1ml of USPIO solution in PBS (20 mg Fe/ml) was oxidized 60

61 using mg Sodium Periodate for 30 minutes at a ph of 6.0. The reaction was stopped by addition of ethylene glycol (10 mm). The USPIO solution was then purified using a PD10 column (Amersham/ GE Healthcare Uppsala, Sweden) as follows: The PD10 column was first equilibrated with PBS by passing 5 volumes of PBS and then the column was centrifuged at 1000 g. The eluate was discarded and the solution containing oxidized USPIO particles and ethylene glycol was pipetted drop by drop to the center of the column. The column was again centrifuged at 1000 g and the eluate containing 99% USPIO solution free of sodium periodate and ethylene glycol was obtained. The surface aldehydes form a covalent Schiff bond at high ph with amines of lipids such as DOPS. After drying the lipid (1 mm) by evaporation of organic phase using slow flowing N2 gas, 200 µl of oxidized USPIO solution at ph 6 was added along with 0.2 mg of Saposin-C and vortexed. Note: While a higher molar ratio (1:3) of Saposin-C to DOPS is being used for testing the efficacy of Saposin-C as an anti-cancer agent, a molar ratio of 1:30 was used for the purpose of using Saposin-C as a targeting agent, without causing a large amount of cytotoxicity. This ratio was used in all standard preparations. Vortexing of the solution was followed by addition of 300 µl of Sodium Borate at ph 9. The mixture was vortexed again and then mixed with 2 ml of diethyl ether. Sonication for about 15 minutes, followed by removal of diethyl ether thorough flow of N 2 gas resulted in the formation of reverse evaporation vesicles. The vesicles were sized by passing through a 200 nm polycarbonate membrane (11 to 15 passes) using the Liposofast extruder (Avestin Inc., Ottawa, Canada). To remove the USPIO particles attached to the outer surface of the liposome, the solution was further dialyzed against a low ph (4.5) solution of sodium chloride (0.15 mm) and sodium citrate (0.15 mm) for several hours. Finally, the unencapsulated USPIO was removed by passing the solution through a Con-A sepahrose 4B column (Amersham Biosciences Corp., NJ.). A con-a sepharose column (approximately 0.5 ml per 1 ml of liposome solution) was equilibrated with a solution of 0.15 mm sodium 61

62 chloride and 0.15 mm sodium citrate at ph 6.0. The liposome solution was added in 250µl aliquots and the first 250 µl eluted from the column was discarded. The liposome solution was stabilized by adding 20 µl of 1M HEPES buffered to ph 8 (20 mm final concentration) to prevent aggregation. The liposome solution was passed through the Liposofast extruder using a 200 nm membrane in order to further re-size the liposomes. For convenience, the Saposin-C-DOPS-USPIO liposomes will be referred to as ScDOPS-IO. Figure 3-1: Above figure shows a Saposin-C DOPS- USPIO liposome in various stages of the encapsulation process. The USPIO particles in the exterior (brown circles) are removed by low ph dialysis and then by gel filtration on a con-a sepharose 4B column. 3.4 Analysis of liposome solution Dynamic Light Scattering The liposome size distribution was determined by dynamic light scattering experiments (Coulter N4 particle sizer). Figure 3-2 (a) shows the size distribution of DOPS liposomes and USPIO nanoparticles in solution after sonication for 10 minutes in a bath type sonicator. The smaller peak represents USPIO particles with a mean diameter of 30 nm and the larger peak represents DOPS liposomes of about 190 nm. Figure 3-2 (b) shows the size distribution of USPIO loaded liposomes (ScDOPS-IO) made using the chemical coupling method described in section 3.3, followed by purification 62

63 using a con-a sepharose 4B column and further sizing using the extruder. The liposomes have a mean diameter of about 230 nm. Figure 3-2: Liposome size distribution (a) USPIO particles ( mean diameter ~30 nm) and DOPS liposomes (mean diameter~190 nm) obtained by sonication.(b) Size distribution profile of liposomes containing UPSIO chemically coupled to the lipid, followed by purification of the solution to remove free USPIO Determining concentration of iron ICP-AES Inductively coupled plasma atomic emission spectroscopy (ICP-AES) was used to determine the concentration of iron in the vesicles. ICP-AES measurements were performed 63

64 at the department of Chemistry, University of Cincinnati. The detailed protocol followed is documented in Appendix 1 Section II. A standard preparation of ScDOPS-IO solution was found to encapsulate about 90 µg Fe/ml when Saposin-C was used in a 1:30 molar ratio. This preparation was used to determine R2 relaxivity of the SC-DOPS-IO solution by plotting serial dilutions of Sc-DOPS-IO solution against the corresponding R2 values obtained by MR relaxometry MR Relaxometry MR relaxometry was performed to determine the r2 relaxivity of the liposome solution. A standard T2 mapping protocol provided in Paravision was used with minor modifications. ( MSME_T2fit: TE=10ms to 200ms, TR=2000ms ) Figure 3-3: Figure shows an example of an MR image obtained using a T2 weighted spin-echo sequence ( TE/Tr= ms/2000 ms). Multiple spin echoes can be used to generate a series of images with different echo times and the T2 is calculated by fitting the image data with a monoexponential on a pixel-by-pixel basis. The vials contain liposome solutions of different concentrations. The r2 relaxivity of the ScDOPS-IO solution was calculated by taking serial dilutions and calculating R2 (=1/T2) vales of each vial. The concentration normalized relaxation rate i.e. r2 is determined by plotting the R2 values against the corresponding molar concentration of iron 64

65 and taking the slope of the best-fit line. The iron concentration was determined by ICP-AES (Ref: Section ) the r2 of the ScDOPS-IO solution is shown in Figure 3-2 and was calculated to be s -1 mm -1. R2 vs Concentration of Iron in Sc-DOPS-IO liposome solution R2( 1/s) y = x R 2 = [Fe] in mm Figure 3-4: Plot of R2 relaxation rates of various concentrations of Saposin-C-DOPS-USPIO solution vs the molar concentration of iron in the solution, as measured by ICP-AES. The relaxivity r2 is the slope of the line and is s -1 mm -1. Due to the variability in encapsulation induced by the above mentioned factors, a standard curve was generated by mixing known concentration of USPIO (iron oxide) and DOPS-Saposin C liposomes. Serial dilutions of this solution containing unbound and unencapsulated USPIO was imaged on the 7 T MR scanner and the R2 values where plotted to get a standard curve. To determine the iron content of a liposome solution, the iron in the liposomes was freed by adding about 10 µl of surfactant (Triton-X100) per ml and the ph was adjusted to 4.5 by addition of a 0.1M HCl in order to free USPIO bound to lipids by Schiff bonds. The solution used to make the standard curve also contained 10 µl of 65

66 surfactant and was adjusted to a ph of 4.5. The relaxation rate of a preparation of ScDOPS-IO solution, was calculated by MR relaxometry and the iron concentration was then determined from the standard curve (Figure 3 5) Average values of iron concentration and R2 values in Sc-DOPS-IO The iron concentration in a standard Sc-DOPS-IO preparation was 94.8 ±12.8 µg/ml when a 1:30 molar ratio of Saposin-C to DOPS was used with 1mM DOPS. Transverse relaxivity (r2) of the ScDOPS-IO was ±4.9 mm -1 s -1. R2 vs. concentration of iron y = x R 2 = R2=1/T concentration of Fe in microg/ml Figure 3-5: Above figure shows the standard curve generated by mixing known quantities USPIO with 1mM of Saposin-C-DOPS liposomes. This curve was used to estimate iron concentrations in Saposin-C-DOPS-USPIO liposome solutions by measuring the relaxation rate R2 after releasing the encapsulated iron using a surfactant. Iron concentrations for the standard curve were determined by ICP-AES Effect of Saposin C on iron encapsulation When compared to ScDOPS-IO liposomes, liposomes not containing iron showed a lower encapsulation of USPIO nanoparticles. An average of 36.33±4.5 µg Fe/m was found to be encapsulated in DOPS liposomes using same encapsulation protocol. 66

67 3.5 In vitro cell experiments Preparation Experiments were performed to determine the uptake of ScDOPS-IO liposomes in tumor cell cultures by incubating the cells with the contrast laden liposomes. The neuroblastoma tumor cell line CHNLA-20 available at Cincinnati Children s Hospital was used for both in-vitro studies and for implantation in mice for in vivo studies. Cells were grown in complete medium consisting of Iscove's modified Dulbecco's medium (IMDM, Bio Whittaker, Walkersville, MD) supplemented with 3 mm L-glutamine (Gemini Bioproducts, Inc., Calabasas, CA), insulin and transferrin 5 µg/ml each and 5 ng/ml of selenous acid (ITS Culture Supplement, Bio Whittaker, Walkersville, MD), and 20% fetal bovine serum at 37 C in a humidified 5% CO 2 atmosphere. Cell lines were subcultured by detaching with trypsin from culture plates Effect of exposure time on cell cultures Samples of neuroblastoma cells were prepared with approximately 500,000 cells per group in 1 ml IMDM growth medium. ScDOPS-IO liposomes were added to the samples at (final concentration of 18 µg Fe/ml ). Growth medium containing free USPIO and DOPS- USPIO liposomes were used as controls, in addition to cells not exposed to any contrast agent or liposomes. Cells were collected at various time points (24, 18 and 12 hours), washed, trypsinized and fixed in 1% agarose for imaging. A majority of cells exposed to ScDOPS-IO were killed and detached from the plate. These cells were also collected, washed and fixed and agarose. Cells exposed to only DOPS or DOPS-USPIO did not show any toxicity. 67

68 3.5.3 Effect of initial concentration of ScDOPS-IO on cell cultures In another experiment, neuroblastoma cells (5x10 5 per well) were incubated with various concentrations of ScDOPS-IO liposomes for 12 hours. 60 µl, 80 µl, 100 µl and 120 µl of ScDOPS-IO solution was added to cells (5x10 5 per well) in 1ml growth medium resulting in a final Iron concentration 5.1 µg, 6.6 µg, 8.1 µ and 9.6 µg respectively. The cells were collected, including those that were detached from the plate were also collected, washed and fixed in 1% agarose. All experiments were performed in triplicate Fixing cells in agarose In order to immobilize the cells for MR imaging, cells were fixed in 1% agarose. After washing and collecting the cells, they were re-suspended in 0.5 ml fresh growth medium. A 2% agarose solution in deionized water was prepared in a conical flask by microwave heating. Care was taken to ensure that no bubbles were present in the solution. First, 1 ml of 2% agarose was pipetted into standard 2 ml or 4 ml vials to serve as a base and allowed to set. After the base layer was ready, the cells in growth medium were pipetted into the vials. Then 0.5 ml of 2% agarose was carefully pipetted into the vial and allowed to mix with the growth medium containing cell. After ensuring that the agarose and growth medium were mixed, the mixture was allowed to set. Finally, a layer of 2% agarose was added to the top of the vial to seal off the layer containing the cells Imaging High resolution MR imaging of the cells was performed using a 7T Bruker Biospec scanner (Bruker BioSpin MRI GmbH, Ettlingen, Germany) equipped with the B-GA12 actively shielded 12 cm gradient subsystem delivering up to 400mT/m. Relaxometry and imaging experiments were performed using a 38 mm birdcage resonator. Gradient echo methods optimized for T2* weighting were used to visualize cells loaded with ScDOPS-IO. 68

69 A 3D FLASH imaging sequence with TR/TE/θ of 200 ms/35 ms/10 and a 320!320!64 matrix was used for a 3.2 cm!3.2 cm!0.64 cm FOV resulting in an isotropic 100 µm resolution. Relaxometry: Spin echo and gradient echo images were acquired with varying TE (Spin Echo: TE=15 to 50 ms, TR=4000 ms and Gradient Echo : TE=5 to 25 ms, TR=1000 ms). The natural logarithm of the signal intensity of cells fixed in agarose in each vial was determined over a 3D slab and plotted against TE. The negative slope of the best fit line determined by linear regression for each set of points is the corresponding relaxation rate (R2 or R2*). 3.6 Results Figure 3-6 shows a typical axial imaging of the vials containing cells fixed in 1% agarose using a high resolution 3D FLASH sequence. The black dots or signal voids represent voxels containing cells loaded with USPIO contrast. An increasing number of signal voids are observed in cells incubated with the ScDOPS-IO solution for longer periods. For reference, images of fixed cells that were exposed only to free USPIO in the growth medium and cells that were incubated in growth medium without any contrast are included. In order to quantitatively investigate these standard preparations, MR relaxometry was performed on the vials. There is a corresponding increase in the relaxation rates R2 and R2* of vials containing cells loaded with increasing amount of iron. This indicates that the longer the cells are exposed to the ScDOPS-IO liposomes, the more iron enters the cell by endocytosis. It must be noted that exposure of tumor cells for longer periods (>12 hrs) resulted in increased cytotoxicity. To cells that may have taken up iron but may have detached from the culture plate, the dead cells in the medium were also collected by centrifugation of the original medium followed by washing to remove liposomes. 69

70 Figure 3-6: A typical high resolution (100 µm) T2* weighted FLASH image of vials containing cells fixed in agarose. The vial in the center contains control cells, not loaded with any iron. The black dots in the other vials represent the voxels that have cells loaded with iron, which causes a loss of signal from those voxels. From Figure 3-7 and Figure 3-8, the R2 and R2* ( mean ± standard error of measurement) values ate plotted as shown in Figure 3-9 b and c. Also shown are the actual values of iron concentration as determined from ICP-AES in identical samples. For example, the R2 and R2* relaxation rates of cells in agarose were s -1 and s -1 respectively for cells exposed to ScDOPS-IO for 24 hours vs. the control cells, which had R2 and R2* relaxation rates of 7.84 s -1 and s -1 respectively. Relaxation rates and ICP measurements are also tabulated in Table 3-1. From ICP-AES measurements of the absolute Iron concentration in cell samples, the average Iron uptake per cell was also estimated. The average iron content in cells was about 3 pg/cell, 2.12 pg/cell and 1.48 pg/cell for cells incubated with ScDOPS-IO for 24, 18 and 12 hours respectively. 70

71 Natural log of signal intensity vs TE for spin echo sequence 24 Hrs ln(si) TE in seconds 18 Hrs 12 Hrs USPIO 24 Hrs Unloaded Cells Linear (24 Hrs) Linear (18 Hrs) Linear (12 Hrs) Linear (USPIO 24 Hrs) Linear (Unloaded Cells) Figure 3-7: ( Above)The natural log of signal intensities of the vials in spin echo images was plotted vs TE. The negative slope of the line gives the relaxation rate R2 of each vial. (Error bars are smaller than data point symbol) Natural log of signal intensity vs TE for gradient echo images ln(si) TE in seconds 24 Hrs 18Hrs 12 Hrs Unloaded cells USPIO cells Linear (24 Hrs) Linear (18Hrs) Linear (12 Hrs) Linear (Unloaded cells) Linear (USPIO cells) Figure 3-8( Above) The natural log of signal intensities of the vials in gradient echo images was plotted vs TE. The negative slope of the line gives the relaxation rate R2* of each vial. (Error bars are smaller than data point symbol) 71

72 a) 24 hrs 18 hrs 12 hrs USPIO Unloaded cells 24 hrs R2* values R2* values Hrs Hrs R2* in 1/s Hrs 12 Hrs USPIO 24 hrs Unloaded cells 0 b) c) Cell sample R2* in 1/s Hrs 12 Hrs Cell sample USPIO 24 hrs Unloaded cells hrs Iron conent (µg) hrs 12 hrs 24 hrs USPIO d) 0 Sample Figure 3-9: a) MR images of cells fixed in agarose. T2* weighted 3D-FLASH imaging on 7T. b) R2 relaxation rates of labeled cells fixed in agarose. c) R2* relaxation rates of labeled cells fixed in agarose. d) Plot of Iron concentration in cells as determined by ICP AES of 1 set of identical samples of cells obtained at the indicated time points. 72

73 Correlation of R2 and R2* with actual iron concentration in cells R2* R2 Linear (R2) Linear (R2*) R 2 = R 2 = R2 and R2* ( 1/ms) [Fe] in cells ( ICP) 0 Figure 3-10: The R2 and R2* values of the cells fixed in gelatin is plotted against the actual concentration of iron in cells (µg/2.5x10 5 cells) as determined by ICP-AES ( also tabulated in Table 3-1). Both R2 and R2* values show good correlation with the concentration of iron in cells. Table 3-1: Relaxation rates R2 and R2* and actual iron concentration in cells ICP-AES Sample R2( s -1 ) R2*(s -1 ) µg/ 2.5x10 5 cells Control Cells 7.8± ± USPIO 24 Hrs 7.8± ± ScDOPS-IO 12Hrs ScDOPS-IO 10.4± ± Hrs 12.9± ± ScDOPS-IO 24 Hrs 14.6± ±

74 Figure 3-11: a) Image strip showing spin echo ( top row) and gradient echo( bottom row) images of vials containing Iron oxide loaded cells fixed in agarose. Vials in columns 1,2,3 and 4 contain cells that were incubated with 5.1 µg, 6.6 µg, 8.1 µ and 9.6 µg of Fe respectively. b) Plot of R2 and R2* relaxation rates of cells fixed in agarose plotted against initial Iron concentration in growth medium. The mean values and standard deviation over 3 samples are shown. When the cells were loaded with the varying concentrations of liposomes for the same time period, the uptake was found to be proportional to the initial concentration in the growth medium, as expected. Cells exposed to lower concentration of ScDOPS-IO showed less uptake as compared to cells exposed to higher concentrations of ScDOPS-IO, as determined by measurements of R2 and R2* relaxation rates of the cells fixed in agarose. Spin echo and gradient echo image samples are shown in Figure 3-11(a) and the relaxation rates are plotted in Figure 3-11 (b) for each of the four vials. It can also be seen from the graph that R2* is much more sensitive to the presence of iron oxide compared to R2, indicated by the steeper slope of the line. 74

75 3.6.1 Statistical analysis of In vitro results Descriptive statistics were used to quantify the results of the various in vitro experiments. For liposome size measurements, the mean size of the liposomes for each preparation was reported. Since extrusion through a membrane of size 200 nm was used, the standard deviation of the diameter of the liposomes was less than 40 nm, as provided by the instrument (Coulter N4 particle sizer). Iron concentration in liposomes was reported as mean ± standard deviation when measured by relaxometry. ICP-AES measurements were used to validate the measurements and also for determining iron concentrations of solutions that were used for making standard curves. For example, the iron concentration in a standard Sc- DOPS-IO preparation was 94.8 ±12.8 µg/ml when a 1:30 molar ratio of Saposin-C to DOPS was used with 1mM DOPS. Transverse relaxivity (r2) of the ScDOPS-IO was ±4.9 mm -1 s -1. Best fit lines for determining iron concentrations has R 2 value greater than For analysis of MR images of vials containing cells fixed in agarose, mean and standard deviation of signal intensity was reported. Relaxometric measurements included mean relaxation rates R2 and R2* and the standard error in measurement, as determined by SigmaPlot from the best fit line. These relaxation rates R2 and R2* were also correlated with the actual concentration of iron in the cells, determined by ICP analysis. Correlation coefficients of R 2 =0.97 and R 2 =0.85 were obtained for R2 and R2* measurements respectively. 3.7 Discussion and Conclusions The objectives of performing in vitro experiments on tumor cell lines were the following: 1) To verify that the tumor cells take up the ScDOPS-IO 2) To determine if the cells that have taken up ScDOPS-IO could be detected using high resolution MR imaging. 3) To develop methods to quantify the uptake of ScDOPS-IO in the cells and 4) To verify using 75

76 non imaging methods the actual iron concentrations in the cells. It must be noted that while ScDOPS-IO caused cell death in tumor cells, the anti-cancer effect of Saposin-C was not measured in these experiments. Saposin-C is though to induce cell death by apoptosis, the mechanism of which is under investigation. Biological assays like TUNEL staining are better suited to detect and quantify apoptosis. In the experiments described in this chapter, the primary purpose of the Saposin-C protein was to serve as a targeting agent by enabling membrane fusion of the DOPS liposomes with the PS on the outer leaflet of tumor cells. In order to detect specific cells using MR imaging, the cells need to be loaded with contrast agents. The maximum sensitivity (i.e. maximum signal change per unit metal) is provided by iron oxide contrast agents and encapsulation within liposomes or cells further enhances the relaxivity of these agents, unlike Gd based contrast agents as discussed in Chapter 2. Even so, a significant amount of SPIO nanoparticles is required to be taken up by cells to become MR visible. To detect a single cell against a uniform background using image sequences tuned for maximum T2* contrast requires that the cell encapsulate at least 1pg of iron [67]. In in vitro experiments on the tumor cell lines, cells were routinely loaded with USPIO nanoparticles at levels of 3 pg/cell, enabling MR detection when fixed in agarose. In vitro experiments also confirmed that the uptake of ScDOPS-IO in cells is directly proportional to both incubation time and the initial concentration of the ScDOPS-IO in medium. As expected, measurements of T2* or R2* were more sensitive to the presence of iron oxide contrast compared to T2 or R2. Finally, an MR methodology to quantify the uptake of ScDOPS-IO in cells using relaxometric measurements on standard preparations of cells fixed in agarose was developed and confirmed by using ICP-AES to correlate the actual iron concentration in cells with the R2 and R2* values. 76

77 In conclusion, the experiments described in this chapter address Specific Aims 2 and 3, setting the stage for in vivo experiments to detect the delivery of ScDOPS-IO in a mouse tumor 77

78 Chapter 4 In vivo Experiments In this chapter, the animal experiments performed in a mouse tumor xenograft model are described. All animal experiments were performed in compliance with the guidelines set by the Institutional Animal Care and Use Committee (IACUC) at the Cincinnati Children s Hospital Research Foundation. The chapter is based partly on the manuscript submitted to Magnetic Resonance in Medicine titled MR detection of targeted delivery of USPIO and Saposin-C to tumors using liposomes. Experiments were conducted primarily to determine if the ScDOPS-IO liposomes would accumulate within a tumor on systemic delivery via tail vein injection and whether the uptake could be detected by MR imaging. The potential anticancer activity of Saposin-C was not examined in these experiments. Rather, the focus was to determine if MR imaging is a suitable modality to detect the delivery of ScDOPS-IO to the tumor. 4.1 General Guiding Principles Specific Aim The overall goal of the in vivo experiments on a mouse tumor xenograft model was to test the hypothesis that the ScDOPS-IO liposomes will be targeted to the tumor when injected into the blood via tail vein injection. Further, it was hypothesized that the accumulation of ScDOPS-IO in the tumor would be sufficient for MR detection using highly

79 sensitive, high resolution gradient echo techniques. The specific aim of these experiments is listed below: SA 4: Determine if the Saposin-C-DOPS USPIO complex can be targeted to tumors in vivo and detected using high resolution MR imaging. a) Optimize MR imaging techniques to improve sensitivity to enable detection of Saposin-C-DOPS-USPIO nanoparticles in vivo b) Perform in-vivo experiments in order to evaluate the delivery of Saposin-C- DOPS-USPIO liposomes to tumors in vivo Imaging Setup In order to find the best imaging techniques and optimize the imaging setup, several pilot experiments were conducted. Due to the difficulties in catheterizing a mouse tail vein for injection in the scanner, mice were imaged prior to injection and the injection was performed outside the scanner, followed by post-injection imaging. Therefore it was not possible to perform dynamic imaging in order to determine the pharmacokinetics of the Sc- DOPS-IO solution by tracking the signal intensity changes immediately after injection. To avoid the difficulties in co-registering pre and post-contrast images, a high resolution 3D imaging sequence was developed. By employing a 3D acquisition scheme, it was possible, in some cases, to match the pre and post contrast image locations very closely in retrospective image review. In order to ensure that maximum SNR was achieved, a cross (i.e. dual) coil imaging setup was used as shown in Figure 4-1. The animal was placed in a 72 mm birdcage quadrature (volume) coil. This coil was used only in transmit mode, to apply the excitation rf pulse. A surface coil was placed on the subcutaneous tumor and was used for receiving the MR signal. Using a surface coil matched to the tumor size improves the filling factor of the receiver coil, reducing receiver noise and improving detection sensitivity. The two coils were 79

80 Transmit Channels 1 and 2 Received signal To Active Decoupler Quadrature Volume coil (transmit) Surface coil (receive) Figure 4-1: Figure shows schematic of the cross coil setup. The surface coil is used only in receive mode. The quadrature volume coil is used only in transmit mode and the coils are actively decoupled. actively decoupled so that only the volume coil is active during transmission and only the surface coil is active during receiving. However, during initial setup for each imaging experiment, only the volume coil was used for adjusting the center frequency, the reference gains for the rf pulses and the digitizer filling, since only the volume coil is used for transmitting. However, in order to achieve a homogeneous B 0 field in the field of view (FOV), shimming of the magnetic field was done with the surface coil activated. While performing animal imaging experiments, it is necessary to anesthetize the animal and limit motion in order to avoid artifacts in the image. Isoflurane (0.5 to 2.5 %) mixed with medical air was delivered using a nose-cone constructed from a syringe tube to anesthetize the animal. The respiration rate was monitored and warm air was used to keep the animal at a constant 30 C in the bore of the magnet. The scans were not gated either using respiration or cardiac waveforms. By using a surface coil, the relative movement between the 80

81 subcutaneous tumor and the coil was minimized. Further, motion compensation by averaging was performed for scans in which multiple acquisitions (=N acq ) were made. This means that instead of acquiring a single k-space line N acq times and averaging them before moving to the next line of k-space, all the lines of k-space were acquired before acquiring the next set of k- space data. Averaging was also used in order to improve the SNR by a factor of N acq. This is particularly important for gradient echo techniques because the signal amplitude is smaller than in spin-echo techniques. Using a small flip angle in the FLASH imaging sequence results in a T2* weighted sequence but also results in smaller signal (Ref. Equation 2-25). The problem is exacerbated by the use of a negative contrast agent (USPIO), which results in loss of signal due to dephasing of the transverse magnetization from those voxels that contain the USPIO particles. Therefore, averaging helps to improve SNR of acquired data Injection of ScDOPS-IO An injection volume of 250 µl of the ScDOPS-IO solution was used for all experiments except the pilot experiments. In the pilot experiment, a total of 400 µl ScDOPS- IO solution was used. Fresh preparations of the ScDOPS-IO solution as described in [Chapter 3] were made for each injection. Injections were done within a few hours of preparation of the solution after anesthetizing the animal and using a heat-lamp to warm the tail. After injection, typically, imaging was done several hours later, to allow for the clearance of the liposomes from blood, except in the case of the pilot experiment where the mouse was imaged immediately after injection and the followed up with imaging after 3 hours. 81

82 4.2 Pilot Experiments Animal Model For in vivo MR imaging, tumors were induced in the mice by injecting approximately 5x10 6 human neuroblastoma tumor cells (CHNLA-20) subcutaneously and tumors were established in approximately 5-6 weeks. All mice were housed at the Cincinnati Children s Hospital Research Center Animal facility. Animal experiment protocols were approved by the Institutional Animal Care and Use Committee Imaging High resolution MR imaging of the tumor was performed in vivo on mice at 7T using a 78 mm birdcage quadrature volume coil for transmit and surface coil placed on the tumor for receiving the MR signal. The surface coil was actively decoupled from the received and detuned using a PIN diode circuit during transmit pulses to the volume coil. T2* weighted 2D and 3D FLASH sequences (TE/TR=5 ms/20 ms, flip angle=10, 16 averages) were used with a 100µm isotropic resolution. The imaging time for the 2D and 3D scans was approximately 5 minutes and 50 minutes respectively Injection of ScDOPS-IO ScDOPS-IO solution was administered intravenously via a tail vein injection after acquiring reference images prior to injection. Image acquisition was repeated 3 hours and 24 hours post injection of 400 µl of the ScDOPS-IO solution. Follow up imaging was done 24 hours after the injection. The mouse was sacrificed after the follow up imaging and the tumor was collected for ex vivo analysis using inductively coupled plasma atomic emission 82

83 spectroscopy (ICP-AES). An identical experiment was performed on another mouse, with a reference contrast tube of deionized water taped to the mouse Image post processing and analysis Raw data acquired from the MR scanner was processed using Bruker software Paravison 4.0. Further analyses of MR was performed using an in-house image processing software, CCHIPS (Cincinnati Children s Hospital Image Processing Software) written in Interactive Data Language ( IDL) ICP-AES Analysis Inductively coupled plasma atomic emission spectroscopy (ICP-AES) was used to quantify the concentration of iron in the tumor. An Agilent 7500ce (Agilent Technologies, Tokyo, Japan) ICPMS equipped with shielded torch and collision/reaction cell technology was used after calibration of the instrument with a stock solution. Pulverized tumor tissue (0.5 ml) were placed in septum sealed glass tubes and treated with 0.5 ml of concentrated HNO 3 resulting in a 50% v/v acid concentration. The resulting solutions were digested in a water bath at 80 C for 2 hours. Digested samples were then diluted to 5.0 ml with DII water containing 250 µg/l Yttrium internal standard. Samples were then analyzed for total Fe with the Agilent 7500ce ICP-MS. Detailed procedures are provided in Appendix Imaging experiments in mice bearing human neuroblastoma xenografts Pilot experiments were performed to determine the optimum coil configuration and imaging protocols and also to determine how to assess change in tumor contrast. Since catheterizing a mouse tail vein is challenging and requires highly trained veterinary 83

84 professionals, an attempt was made to image a mouse treated with Sc-DOPS-IO immediately after injection. The injection was performed without removing the mouse from the cradle. After injection, the mouse was repositioned in the scanner and a few minutes were required to acquire imagines to localize the mouse. Images taken after injection of ScDOPS-IO did not show any signal change in the tumor. However, a darkening of the blood vessels or interstitial space in the tumor surrounding the tumor is observed as seen in Figure 4-2, and indicated by the white arrows. a b Figure 4-2: (a) High resolution in vivo imaging of mouse tumor. Image strip showing 2D T2* weighted FLASH image of tumor before injection of ScDOPS-IO. (TE=5ms/TR=20ms/FA=10 ), 16 averages; Resolution: 100µm in-plane resolution, 250µm slice thickness (b)image a few minutes after injection of contrast agent. The white arrows indicate changes in the image after injection where ScDOPS-IO is present in the tumor blood vessel or interstitial space. For the next pilot experiment, a baseline image prior was acquired prior to injection of Sc-DOPS-IO. The mouse was then injected with Sc-DOPS-IO and placed in the scanner immediately. After acquiring post-injection scans, the mouse was kept in the scanner for 3 hours, in order to acquire images of the tumor after washout of SC-DOPS-IO from the blood. Measurements of uptake of ScDOPS-IO solution in vivo in nude mice bearing a neuroblastoma tumor on the flank were made using MRI and ICP-AES methods. Results are 84

85 shown in Figure 4-3 Panel (a) shows the tumor imaged before, immediately after, 3 hours post and 24 hours post injection of ScDOPS-IO. A visual comparison of the signal intensity in the tumor in the T2* weighted scans at these time points shows a drop in signal intensity over the entire tumor volume. Quantitatively, the signal intensity was measured over the entire tumor volume from the 3D, T2* weighted scan in a manually drawn ROI containing the tumor. In Figure 4-3(b), the mean, normalized signal intensity over the 3D tumor volume and the standard deviation are shown. For normalization, an ROI of tissue directly below the tumor was selected. It can be seen that the MR signal intensity drops within the tumor ROI immediately after injection and gradually decreases further before rebounding slightly 24 hours later. The signal intensity of the tissue below the tumor remains relatively constant. a) (1) Pre -injection ( 2) Post injection (3) 3 hours post (4) 24 hrs post Normalized Tumor Signal Intensity Iron content in mouse tumor Iron content in treated mouse b) Normalized SI Time Point c) [Fe] in µg/mg tissue Iron content control mouse category Figure 4-3: a) In vivo imaging of mouse tumor. Image strip shows T2* weighted FLASH image of tumor before and after injection of contrast agent. (TE=5ms/TR=20ms/FA=10 ), 16 averages; Resolution: 100µm isotropic b) Tumor signal intensity plotted at the time points of image acquisition shown in figure (a), normalized against the signal intensity of an ROI selected below the tumor. c) ICP-AES data representing iron concentration in excised tumor of mouse treated with ScDOPS-IO vs. control mouse. 85

86 ICP analysis of the same tumor (c) indicates that, approximately five-fold iron concentration was found in compared to the same volume of tumor in a control mouse that was not treated( 244 µg Fe/mg of tumor vs 49.3 µg Fe/mg of tumor tissue). The tumor was sectioned and ICP was performed on the two halves (Figure 4-4). The iron concentration was found to be proportional to the percentage signal change. The top half of the tumor showed an average signal drop of 13% with respect to surrounding tissue and the corresponding iron concentration measured in the tumor was 23.5 µg. The bottom half of the tumor showed an average signal change of about 7% and the corresponding iron concentration was 13.2 µg of iron. Figure 4-4: A single slice of the 3D FLASH image of tumor immediately after injection of ScDOPS- IO (left) and 3 hours post injection (right). The horizontal line shows the approximate location where the tumor was cut into two halves for ICP analysis. The amount of iron in each half of the tumor was found to be proportional to the signal loss observed in the MR images shown above. In another mouse with a larger neuroblastoma tumor, the same pattern of signal loss was observed when the signal intensity of the tumor was normalized with the contrast of a tube of de-ionized water. ( Figure 4-5). Normalization with a reference contrast was done to 86

87 account for the variations from one scan to the next and also because change in contrast in the tumor was not always visually discernable. a) (1) Pre-injection ( 2) Post injection (3) 3 hours post (4) 24 hrs post Tumor SI (normalized) Normalized SI b) Time point Figure 4-5: a) In vivo imaging of mouse tumor with reference tube containing deionized water. Image strip showing T2* weighted FLASH image of tumor before and after injection of contrast agent. (TE=5ms/TR=20ms/FA=10 ), 16 averages; Resolution: 100µm isotropic. b) Tumor signal intensity plotted normalized relative to the intensity of reference vial at the time points of image acquisition shown in (a). 4.4 Experiments in mice bearing human pancreatic tumor xenografts Study design In order to study the uptake of ScDOPS-IO solution on an alternative tumor cell line, 5 mice bearing human pancreatic tumor (MiPaCa-2) xenografts were imaged. Due to different rates of tumor progression and availability of mice, the mice were imaged on 87

88 different days. Tumors were induced in the mice by injecting approximately 5x10 6 MiPaCa- 2 tumor cells subcutaneously and tumors were established in approximately 5-6 weeks. Two separate animals were imaged in control experiments. The control experiments included injection of a mixture of free USPIO solution and a DOPS-USPIO solution with no Saposin- C. Each animal also served as its own control, by imaging the animal prior to any injection. The animals were imaged a few hours after injection using the imaging protocols described in the following section. Imaging was performed about 3 to 4 hours after injection to allow for clearance of liposomes from the blood. Five animals were used for the initial part of the experiment comparing tumor T2 values at the above mentioned time points. One animal was followed longitudinally for 3 days, with additional treatment of ScDOPS-IO injections Imaging As in the pilot experiment, high resolution MR imaging of the tumor was performed in vivo on mice at 7 T using a 78 mm birdcage quadrature volume coil for transmit and surface coil placed on the tumor for receiving the MR signal. The surface coil was actively decoupled from the receiver and detuned using a PIN diode circuit during transmit pulses to the volume coil during transmit periods. T2* weighted 2D and 3D FLASH sequences (TE/TR=5 ms/20 ms, flip angle=10, 16 averages) were used with a 100 µm isotropic resolution. The imaging time for the 2D and 3D scans was approximately 5 minutes and 50 minutes respectively. In addition, a T2 mapping protocol was used to measure the change in tumor T2 values before and after injection of the ScDOPS-IO solution. A multiple echo-spinecho sequence available in Bruker Paravision 4.0 was modified with the following parameters: TE=10ms to 160 ms (16 echoes), TR=2000 ms, resolution 256x192 pixels (approximately 200 µmx200 µmx1 mm) 88

89 4.4.3 Image post processing and analysis Raw data acquired from the MR scanner was processed using Bruker software Paravison 4.0. Further analyses of MR was performed using an in-house image processing software, CCHIPS (Cincinnati Children s Hospital Image Processing Software) written in Interactive Data Language ( IDL) to estimate T2 values as a quantitative measure of iron uptake in the tumor. 4.5 Results In control mice, injected with free USPIO and DOPS-USPIO not containing Saposin- C, there was very little change in tumor SI before and 4 hours post injection (Figure 4-6). Optical imaging experiments conducted independently using fluorescence labeled liposomes indicated that there was no accumulation of non-saposin-c liposomes in the tumor. Representative images of these experiments are shown in Appendix 1, section IV. Before injection 4 hours post injection Figure 4-6: 3D FLASH images of mouse tumor before and 4 hours after injection of solution containing free USPIO and DOPS-USPIO. No change in mean signal intensity was seen. Both control mice and mice treated with Sc-DOPS-IO solution showed uptake of contrast in the visceral organs. Figure 4-7 shows sagittal FLASH (TE=3.6 ms/tr=100 ms/ FA=30 ) images of the mouse abdomen show darkening in the liver, indicating that the 89

90 ScDOPS-IO liposomes in the blood are taken up primarily in the liver. Approximately 20 to 40 % drop in the SI of the liver was seen. Before injection After injection Figure 4-7: Sagittal gradient echo reference scans before and after injection of ScDOPS-IO shows darkening of the visceral organs, indicating uptake primarily of the liver( indicated by arrows). A FLASH imaging sequence (TE=3.6ms/TR=100ms/ FA=30 ) was used. In this particular example, approximately 36% drop in liver signal intensity was seen. Before injection After injection Figure 4-8: A single slice of a 3D FLASH T2* weighted images approximately the same location of a tumor at an isotropic resolution of 100µm isotropic. The image taken a few hours after injection shows dark spots in some locations. 90

91 In Figure 4-9 a typical color coded T2 map of the tumor is shown. The scale shown alongside shows the range of T2 values mapped by the rainbow color scale (0 ms to 130 ms). The tumor images indicate change in T2 of the tumor about 4 hours after injection. 130 T2 Pre-injection 4 hours post injection 0 Figure 4-9: Figure on the left shows a color coded T2 map a mouse tumor prior to injection of ScDOPS-IO. The figure on the right shows the T2 map of approximately the same slice of the tumor 4 hours after injection of ScDOPS-IO. The scale used for transforming the grayscale T2 map into a color coded one is shown on the far right. 140 Tumor T2 before and after injectioon of Sc-DOPS-IO Tumor T2 in ms Before injection 4 hours post injection Figure 4-10: Graph showing the average of the T2 values of tumors before and after injection of ScDOPS-IO in n=5 animals. A one tailed paired t-test shows that the difference in the means is significant ( p=0.011). 91

92 On measuring the tumor T2 values before and after injection in 5 mice, approximately 4 hours after injection, an average drop of 29 ms was seen ( from 91.4 ms to ms), in a manually drawn ROI of approximately the same slice of tumor. This result verifies the hypothesis that the mean of the post-treatment tumor T2 values would be lower than the mean tumor T2 prior to injection reflecting uptake of the contrast in the tumor. A one-tail paired t-test was done on the mean tumor T2 values obtained (n=5). The difference was found to be significant (p=0.011). However, there was some variability in the T2 values between animals. In some cases, a significant drop in T2 ( > 30 ms) was observed and in other cases, the change was much smaller ( < 2 ms). This could be attributed to several factors including tumor size, the nature of the tumor (homogenous tissue vs. heterogeneous tissue), the degree of permeability of the neovasculature (more leaky vs. less leaky) and variation from one animal to the next. A larger study with more animals and different tumor models would be required to unravel the dependence of these factors on uptake of ScDOPS- IO. An outline for such a study is presented in the next section Statistical analysis of In vivo results For the in vivo pilot experiments, descriptive statistics were reported i.e. mean ±standard deviation of the signal intensity of the tumor at various time points, computed over a 3D ROI. The SI was normalized against either the SI tissue surrounding tissue or against an external reference vial containing deionized water. At the 3 hour time point, the peak change in normalized, mean SI of the tumor, a drop of nearly 40% was observed, in the pilot study described in section 4.3. ICP-AES was also used to validate the results of the imaging experiments. ICP results indicated about 5-fold concentration of iron in the tumor when compared to the same mass of tumor from an untreated mouse. In the second pilot experiment, in the mouse with a larger tumor, the peak signal change was approximately 17% at the 3 hour time point, when normalized against the external reference vial. In 92

93 reference scans of the visceral organs, between 20 to 40 % drop in the signal intensity of the liver was seen. (Note that the reference scan parameters are different from the parameters of the high resolution scan of tumors acquired using the surface coil.) In order to study the effect of ScDOPS-IO on the tumor T2 in a pancreatic tumor model, n=5 mice were imaged prior to injection and 4 hours post injection of Sc-DOPS-IO. This time point was chosen based on the preliminary data from the pilot experiments. It was hypothesized that ScDOPS-IO would significantly lower the average T2 of the tumor, computed over an ROI of approximately the same imaging slice before and after injection of ScDOPS-IO (p<0.05). Therefore, one tail t-test was performed on the mean tumor T2 values obtained in the five animals. An average drop of 29 ms was seen (from 91.4± 9.6 ms to 62.68±7.2 ms), in a manually drawn ROI of approximately the same slice of tumor supporting the hypothesis at a significance of p= The Cohen s d statistic was determined using Microsoft Excel to be with a corresponding effect size correlation (r) of Based on the preliminary results, it is clear that ScDOPS-IO lowers the T2 of the tumor, which is reflected in the changes in signal intensity in T2 and T2* weighted images. However, the variability in the T2 values between animals reflects the influence of several factors including tumor size, the nature of the tumor (homogenous tissue vs. heterogeneous tissue), the degree of permeability of the neovasculature (more leaky vs. less leaky) and variation from one animal to the next. Data from recent studies [1]suggest that a concentration of about 1 µg of iron per gram of tissue would be sufficient for detection using optimized imaging techniques. For a mouse model, a 250 µl injection containing about 20 µg of iron would be detectable in a 1gm tumor even if only 5% of the injected iron is retained in the tumor. Ultimately, the sensitivity of the technique will depend on 1) Amount of iron oxide contrast delivered and retained in the tumor, 2) MR imaging sequence parameters: 93

94 Resolution of the imaging sequence, SNR of the imaging sequence and the degree of T2/T2* weighting of the imaging sequence. 4.6 Discussion The in vitro experiments described in the previous chapter showed that ScDOPS-IO liposomes were taken up by tumor cells in quantities that allowed for MR detection using high resolution gradient echo imaging methods. The limited numbers of animal experiments described in this chapter were conducted to provide proof of concept with the following objective in mind: To determine if the USPIO labeled Saposin-C liposomes could be detected in vivo using optimized, highly sensitive and high resolution MR imaging techniques. From the pilot experiments it can be seen that high resolution gradient-echo MR images show a drop in tumor signal intensity after injection of ScDOPS-IO. This change in tumor SI was correlated with the presence of iron, which was quantitatively determined using ICP-AES. A T2 mapping protocol to map the change in tumor relaxation after the injection of ScDOPS-IO was used in imaging experiments on 4 tumor bearing mice. Using these techniques, the delivery of the ScDOPS-IO liposomes to the tumor was demonstrated in a quantitative manner. Extending these techniques beyond the scope of this dissertation, the imaging techniques could further be used to estimate the delivery and uptake of ScDOPS-IO in the mouse model. For example, a larger imaging study involving more animals and control groups could be designed as follows: Group 1: Control group contains un-treated animals (i.e. PBS injection). Group 2: Animals treated with non-targeted USPIO liposomes (DOPS- USPIO). Group3: Animals treated with targeted, contrast enhanced liposomes (ScDOPS-IO). Group 4: Animals treated with Saposin-C-DOPS liposomes not carrying any USPIO. The imaging protocols could then be applied to the animals in a longitudinal study. Imaging could be used to monitor tumor size, tumor SI (normalized against a reference contrast), tumor T2 94

95 and contrast change in the visceral organs. These measures could then be used to determine more accurately the delivery, retention and effect of ScDOPS-IO on the tumor. If post mortem ICP analysis is performed, the actual iron concentrations could be used to determine the percentage efficiency of delivery of ScDOPS-IO to the tumor. This could then be correlated with T2 change per tumor volume, to obtain calibration curves for tumor T2 vs. tissue concentration of ScDOPS-IO. An important advantage of using non-invasive imaging modalities with targeted contrast enhancement is that tumors located deep within the body can be observed and its properties (tumor burden, size, growth rate, cellularity etc) can be quantitatively measured. For tumors supplied by leaky, neoplastic blood vessels, ScDOPS-IO could aid in the early detection of the tumor due to selective contrast enhancement. Finally, in order to evaluate the effect of Saposin-C on the tumor, diffusion MRI could be used. In several studies, the apparent diffusion coefficient (ADC) has been shown to correlate with cellularity of a tumor [85]. An effective cancer therapy which causes cell death in the tumor would result in an increase in the diffusion of water protons compared to the diffusion in intact cells. The ADC would correspondingly increase when the therapy is more effective and this could be used to assess the effect of Saposin-C on the tumor. The above strategy describes a road map that could be used for pre-clinical pharmaceutical drug trials for a targeted chemotherapy agent. There are, however, many challenges that remain to be addressed in the area of targeted contrast imaging of molecular targets using MRI. These will be addressed in the concluding chapter, with some thoughts on future directions for continuing the research related to this dissertation. 95

96 Chapter 5 Conclusion This chapter concludes the dissertation by briefly reviewing the experiments performed, and discussing the major findings. A few avenues for continuing research are also discussed. Finally, a summary of some of the latest work in the field of molecular imaging using MRI is presented. 5.1 A Summary and Discussion of the Major Findings Background As restated from Chapter 1, the broad goal of this dissertation was to develop and test an MR imaging probe that could be used to visualize the delivery of targeted nanoparticles to tumors in vivo. MRI is an inherently a low-sensitivity imaging modality, albeit capable of high resolutions of the order of 100 µm or better. In order to detect molecular targets, contrast agents are necessary. In this particular case, phosphatidylserine (PS) expressed on the exoplasmic leaflet of the plasma membrane of tumor cells, was the target of the imaging probe. The imaging probe comprised of liposomes made of di-oleyl-phosphatidylserine (DOPS) and the targeting protein Saposin-C loaded with an iron oxide based MRI contrast agent (USPIO) commercially known as AMI-227 or Ferumoxtran. The USPIO contrast agent was selected because of its small size and high relaxivity resulting in high negative contrast enhancement in T2/T2* weighted MR images. Although contrast agents based on gadolinium chelates provide positive contrast enhancement in T1 weighted images, there are two 96

97 disadvantages that make them unsuitable for molecular imaging applications. The first is their low relaxivity, requiring very large concentrations of gadolinium to be encapsulated in the liposome for delivery at the imaging target. The second is that encapsulation of gadolinium in liposomes leads to reduced relaxivity due to the fact that water molecules, in order to interact with the paramagnetic centers, have to cross the liposome membrane. On the other hand, since superparamagnetic iron oxide based contrast agents work by creating a magnetic field inhomogenieties that lead to dephasing of the transverse MR signal, the encapsulation does not reduce the relaxivity. In fact, relaxivity is usually enhanced based on the static dephasing regime theory[78]. In order to detect small quantities (of the order of 1 pg) of iron oxide particles sequestered in cells using MRI, high resolution, high SNR techniques are required. Therefore, this dissertation began with the development and characterization of a highly sensitive reporter construct ScDOPS-IO, followed by in vitro MR experiments in tumor cells and proof-of-concept in vivo experiments Encapsulation of USPIO in DOPS liposome A chemical coupling method was developed and optimized for efficient encapsulation of the USPIO nanoparticles in the liposomes. This involved oxidizing the dextran coating of the USPIO to generate aldehyde groups, followed by linking to amine groups on the lipid via ph dependant Schiff bonds. MR relaxometry was performed on the liposome-contrast agent solution, termed ScDOPS-IO and ICP-AES analysis was used to determine true values of encapsulated USPIO In vitro cell experiments Tumor cell cultures (CHNLA-20 and MiCaPa-2) were incubated with the ScDOPS- IO solution to measure uptake and MR detectability. Cells were exposed to varying concentrations of the ScDOPS-IO solution or to one concentration for different time periods. 97

98 Cells were collected, washed and fixed in a standard agarose preparation for MR imaging. Cells were detectable using a high resolution (100 µm), 3D gradient echo MR imaging sequence. Further, quantification of the uptake of ScDOPS-IO in cells was possible using MR relaxometry. ScDOPS-IO uptake in cells was found to be dependant on both initial concentration and time of exposure. This was validated using ICP-AES. Excellent correlation was found between the actual concentration of iron in the cells and the R2 and R2* values determined by MRI In vivo experiments Pilot imaging experiments performed on a mouse tumor xenograft model indicated that ScDOPS-IO accumulates passively in the tumor due to the enhanced permeability and retention (EPR) effect. MR imaging was validated by ICP-AES and also by independent experiments using fluorescence labeled liposomes on a spectro-photometric imaging system. Further, experiments indicated that tumor T2 dropped significantly after treatment with ScDOPS-IO, as determined by MR relaxometry of the tumor. 5.2 Discussion and Future research The findings of the experiments outlined in this dissertation provide a starting point for further research on the use of MRI to study the delivery, uptake and effect of Saposin-C- DOPS. A mechanism for combining a liposomal vector for the protein Saposin-C with a highly sensitive MR contrast agent by means of a chemical coupling method was developed. By developing this chemical coupling between the contrast agent and the liposome, the issue of encapsulation of sufficient contrast agent to enable MR detection was addressed. This was demonstrated conclusively by the in vitro experiments. This also means that since the USPIO contrast agent is covalently bound to the lipid, the detection of the USPIO in images would 98

99 indicate the presence of Saposin-C-DOPS, which would be useful in the determination of the biodistribution of Saposin-C-DOPS in vivo. Further, this technique is highly customizable for other applications involving binding of contrast agents to other lipids used for making liposomes. Making Saposin-C-DOPS MR detectable by combining it with USPIO allows several new avenues of research could be pursued: 1) The pharmacokinetic properties ScDOPS-IO could further be improved by incorporating PEG (poly-ethylene-glycol) in the lipid bilayer to improve circulation time of the liposomes. However, Saposin-C would then have to be combined with PEG to retain the targeting properties. 2) Dynamic contrast imaging for determining temporal and spatial biodistribution of the liposomes in the blood and the tumor, to determine pharmacokinetic and pharmacodynamic parameters. This would require catheterization of the mouse tail vein and also monitoring the clearance of the liposomes form the liver/kidney/spleen. 3) Tumor morphology, signal intensity and T2/T2* changes would give a quantitative estimate of the retention of the liposomes in the tumor. Tumor size can also be measured using 3D imaging techniques. This is especially useful if the tumors are not subcutaneous but are orthotopic. If the cells are loaded with sufficient iron oxide, it is also possible to detect metastases of the tumor using MR imaging. 4) Efficacy to Saposin-C as an anticancer agent: Cell density has been shown to be a biomarker for efficacy of the therapeutic agent. Diffusion weight MRI could be used to determine the apparent diffusion coefficient (ADC) an indicator of cell density. However, this would have to be done in animals treated with Saposin-C-DOPS not containing USPIO to avoid signal loss due to the iron-oxide contrast. 5) Some studies [46, 86] have shown that pulsed high intensity focused ultrasound could be used to locally enhance the delivery of systemically injected macromolecules. This 99

100 technique could be applied to order to enhance permeability of the tumors to liposomes, increasing ScDOPS-IO concentrations in the tumor. 6) While spectro-photometric techniques such as bioluminescence and fluorescence imaging are extremely sensitive and provide semi-quantitative information in small animals, they cannot be translated to larger animals or humans because of limited penetration of optical wavelengths in opaque biological tissue. Therefore, one of the most important advantages of MR imaging is the relatively straightforward translation of techniques developed for preclinical research into the clinic. 5.3 Conclusion Several exciting research areas have opened up in the field of molecular imaging using MRI in the past few years[87]. For example, sparsely expressed biological targets such as cell surface integrins [88], surface phospholipids on apoptotic cells[89], vascular adhesion molecules [90] and macrophage receptors [91] have been imaged non-invasively in vivo using contrast enhanced MRI. A recent advancement in the use of iron oxide contrast is the development of aminated cross linked iron oxide (CLIO), a highly stabilized form of monodisperse iron oxide that allows a large variety of ligands, e.g. annexin for the detection of apoptosis[92], to be conjugated to the nanoparticle. With the development of superparamagnetic contrast agents with even higher relaxivities that presently available, the imaging of the molecular targets in small animals is expected to become a routine and established technique in biological and pharmaceutical research. Magnetic relaxation switches, another class of MR contrast agents, designed to image dynamic events have recently been described by Perez et al [93]. They consist of iron oxide particles that undergo a reversible association or disassociation in the presence of a specific enzyme or compound, thereby producing a change in the relaxivity. These novel contrast 100

101 agents have been shown to be capable of detecting proteases[93], oligonucleotides[94], and also to detect the concentration of analytes such as glucose[95] and calcium[96] in solution. MRI has also been used to image enzyme activity using activatable MR contrast agents. A gadolinium-seratonin chelate, was used to image the activity of the enzyme myeloperoxidase [97, 98]. In the presence of myeloperoxidase, the serotonin moiety tends to aggregate into dimmers and oligomers, increasing its longitudinal relaxivity significantly more than the parent compound. Finally, multi-spectral MRI, for example, combining proton and fluorine MRI [99], enable multiple simultaneous readouts of individual probes. Fluorescence techniques have been combined with MRI using Bangs SPIO particles and more recently, using quantumdots with a paramagnetic coating[100]. An alternative approach is chemical exchange saturation transfer (CEST) imaging[101], exploiting the magnetization transfer between two pools of protons with an offset in their Larmor frequencies. Molecular imaging using MRI is rapidly growing, with advancements in development of nanoparticles, discovery of novel molecular targets and higher field strength MRI magnets with more advanced image acquisition and processing schemes. According to Weissleder et al [22], the fruits of today s molecular imaging research will have a direct effect on patient care within the next 5-15 years. Our current assessment of disease is based on anatomic changes or, more recently, in specialized cases, physiologic changes that are a late manifestation of the molecular changes that truly underlie the disease. Direct imaging of these molecular changes will directly affect patient care by allowing much earlier detection of disease. The coming years will see more applications of MRI in the field of molecular imaging, particularly with the development of sophisticated probes that are sensitive, activatable and provide quantitative measures of molecular activity in vivo. 101

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107 Appendix 1 Additional data, Graphs, Methods 107

108 I. Relaxation times of USPIO in saline and gelatin Table A.1-1: Table of T1, T2, T2* of various concentrations of Combidex in saline. Fe Conc. (µg/ml) T1 (ms) R1(s -1 ) T2(ms) R2(s -1 ) T2*(ms) R2*(s -1 ) T1 vs Concentration of USPIO T2 and T2* vs Concentration of USPIO T1 in milliseconds [Fe] in!g/ml T2 and T2* in milliseconds T2 350 T2* [Fe] in!g/ml R1 vs [Fe] for USPIO in saline R2 and R2* vs [Fe] for USPIo in saline R1 in 1/s R2 in 1/s y = 1.718x R 2 = y = x R 2 = y = x R 2 = [Fe]in!g/ml [Fe] in!g/ml 108

109 Fe Conc. (µg/ml) T1(ms) R1(s -1 ) T2(ms) R2(s -1 ) T2*(ms) R2*(s -1 ) Table A-1-2: Table of T1, T2, T2* of various concentrations of Combidex in gelatin. T1 vs concentration of USPIO in gelatin T2 and T2* vs concentration of USPIO in gelatin T1 in ms T2 in ms [Fe] in!g/ml [Fe] in!g R1 vs Concentration of USPIO in gel R2 and R2* vs Concentration of USPIO in gel R1 in 1/s y = 0.012x R 2 = [Fe] in!g/ml R2 in 1/s y = x R 2 = y = x R 2 = [Fe] in!g/ml 109

110 Table A-1-3: Relaxivities estimated from slope the best fit line in plots of relaxation times R 1, R 2 and R 2 vs concentration of USPIO in saline and gelatin Relaxivities (1/(s x µg/ml) In saline(1/(s x µg/ml) In gelatin(1/(s x µg/ml) r r r2*

111 II. ICP-AES Procedure Reagents A stock calibration standard (SPEX CertiPrep Inc., Metuchen, NJ) containing 1000 µg/ml Fe was utilized for the preparation of all external calibration standards. Calibration standards from µg/ml-5 µg/ml were prepared through serial dilution with 5% v/v HNO 3 (Pharmaco, Hartford, Ct) from the stock elemental standard. All DII water was prepared by passing through a NanoPure (18 MΩ ) treatment system (Barnstead, Boston, MA) prior to use. Helium gas (Matheson Gas Products, Parisppany, NJ), with a purity of %, was used as the collision/reaction gas throughout all experiments. Gas flow rate was optimized during instrument tuning prior to each experiment and controlled by the mass flow controller provided with the instrument. Sample Preparation Cell or tissue samples (0.5 ml) were placed in septum sealed glass tubes and treated with 0.5 ml of concentrated HNO 3 resulting in a 50% v/v acid concentration. The resulting solutions were digested in a water bath at 80 C for 2 hours. Digested samples were then diluted to 5.0 ml with DII water containing 250 µg/l Yttrium internal standard. Samples were then analyzed for total Fe with the Agilent 7500ce ICP-MS. Instrumentation Inductively coupled plasma mass spectrometer (ICPMS) An Agilent 7500ce (Agilent Technologies, Tokyo, Japan) ICPMS equipped with shielded torch and collision/reaction cell technology was used for the element specific 111

112 detection of Fe throughout this experiment. The collision/ reaction cell consisted of an octopole ion guide operated in rf only mode and also served for the removal of polyatomic interferences. A detailed description of ICPMS operating conditions is provided in Table 1. Table A-1.4 ICP-MS Instrumental Parameters. ICP-MS parameters Forward power Plasma gas flow rate Auxiliary gas flow rate 1500 W (with shielded torch) 15.6 L min L min -1 Carrier gas flow rate 1.20 L min -1 Nebulizer Spray chamber Sampling Depth Sampling and Skimmer Cones Dwell time Isotopes monitored (m/z) Octopole Reaction System Glass Expansion Microconcentric 2 C (Scott Double Channel) 8 mm Nickel 0.1 s 56 Fe, 57 Fe, 89 Y (Internal Std.) He (Flow Optimized Prior to Experiment) Digestion parameters Instrument Neslab RTE-110 Water Bath Sample Volume 0.5 ml w/ 0.5 ml HNO 3 Temperature Time 80 C 2 hrs 112

113 III. Freeze-fracture and electron microscopy of Saposin-C DOPS liposomes Figure A-3.1: Above figure shows a freeze-fracture electron microscopy image of the SaposinC-DOPS liposomes. Figure A-3.2: Above figure shows electron microscopy of ScDOPS-IO. 113

114 IV. Fluorescence imaging in mouse tumor model MiaPaCa -2 Control MiaPaCa -2 Control Before injection Immediately after MiaPaCa -2 Control MiaPaCa -2 Control 3 hours after injection 24 Hrs after injection Figure A3.3: Fluorescence imaging of DOPS-Saposin-C liposomes labeled with a far red fluorescent probe. The two mice bearing pancreatic tumor xenografts show accumulation of the liposomes in the tumor. The control mouse shows no accumulation after 24 hours. 114

115 Immediately after injection 2 hours after injection hours after injection 48 hours after injection Figure A-3.4: Control experiments: Mouse bearing MiCaPa-2 pancreatic tumor xenografts were injected with control solutions. Mouse 1 was injected with a solution containing DOPS and Saposin C (unbound). Mouse 2 was injected with non-targeted DOPS (no Saposin-C) solution. Mouse 3 was injected with PBS. There was no accumulation of the liposomes in any of the control mice. 115

116 Methods Preparation of Saposin-C-DOPS nanovesicles for electron microscopy and fluorescence imaging: Bath sonication was used to form Saposin-C-DOPS nanovesicles as described in Chu et al [102]. A small aliquot of ph 4.7 citric-phosphate buffer was added to the dry Saposin-C and DOPS (Avanti Polar Lipids, Alabaster AL) or DOPS alone followed by addition of large volume of PBS. The mixture was sonicated for 20 min at ~4ºC to form Saposin-C-DOPS nanovesicles. The sonicated samples were stored at 4 C for two days, and were characterized by freeze-fracture electron microscopy. To label the Saposin-C-DOPS, CellVue Maroon (CVM, a far-red fluorescent probe, PTI Research Inc., Exton, PA) in ethanol was dried along with DOPS. CVM labeled Saposin-C-DOPS nanovesicles for in vivo fluorescence imaging were formed by sonication. Freeze-fracture electron microscopy For freeze-fracture electron microscopy all samples were quenched using the sandwich technique and liquid nitrogen-cooled propane, described in Gregoriardis et al [103]. Using this technique a cooling rate of 10,000 Kelvin per second is reached avoiding ice crystal formation and artifacts possibly caused by the cryofixation process. The cryofixed samples were stored in liquid nitrogen for less than 2 hours before processing. The fracturing process was carried out in JEOL JED-9000 freeze-etching equipment and the exposed fracture planes were shadowed with platinum for 30 sec in an angle of degree and with carbon for 35 sec (2kV/ 60-70mA, 1x10-5 Torr). The replicas produced this way were cleaned with concentrated, fuming HNO 3 for 24 hours followed by repeating agitation with fresh chloroform/methanol (1:1 by vol.) at least 5 times. The replicas cleaned this way were examined at a JEOL 100 CX electron microscope. Freeze-fracture electron micrographs of the Saposin-C-DOPS nanovesicles show overall spherical vesicles. They display their shadows in front as well as behind their structures representing bilayer liposome vesicles. 116

117 In vivo fluorescence imaging In vivo fluorescent imaging was done using IVIS 200X imaging system (Xenogen Co., Alameda, CA). The filter set (passbands λ EX = nm and λ EM = nm) was used for imaging. Scanning time was 0.1 second. 117

118 Appendix 2 CCHMC Animal Imaging Protocol Approved by the Institutional Animal Care and Use Committee 118

119 All animals brought to the IRC for imaging procedures are covered by an existing IACUC approved Animal Use Protocol assigned to CCHMC investigators. All personnel listed on this Animal Use Protocol have received training by CCHMC Veterinary Services in the proper and ethical use of animals in research, or will be trained before being involved in animal-based imaging procedures. In addition to the IRC staff involved in animal imaging procedures, all animals brought to the IRC for imaging procedures will be accompanied and managed by trained animal technicians who have received appropriate training and certification by CCHRF Vet Services. The 7Telsa Bruker MR imaging instrument is dedicated to animal imaging procedures. A micro-ct scanner is also part of the new imaging facility and will be used strictly for imaging in small animal models. The goal of the this Animal Use Protocol section is to describe the facilities, species and procedures available in the new core facilities of the IRC and to insure that all animals brought to the center for imaging procedures are treated in accordance with appropriate guidelines for the humane treatment of animals as outlined in the Animal Welfare Act and the U.S. Government Principles for the Utilization and Care of Vertebrate Animals Used in Testing, Research and Training. Procedures The specific facilities and procedures to be used are as follows: Location All imaging procedures covered by this protocol will be performed in the IRC, located on the R-level of the CHRF building. Veterinary Services houses the species covered in this protocol on the 5 th 6 th and 7 th floors of this same building and an elevator between these floors is located within 30ft of the entrance to the IRC. 119

120 Imaging Modalities MRI and X-Ray CT have a long history of use in animal experimentation. MRI procedures will include the following methods: 1. Anatomical imaging using MRI relaxation time contrast. 2. Contrast enhanced anatomical imaging and perfusion imaging using USPIObased contrast injected intravenously (i.v.) or intraperitoneally(i.p.). Anesthesia --Animal technicians and investigators will be responsible for anesthesia of all species covered in the protocol and will adhere to the requirements approved by the specified Animal Use Protocol to which the animal belongs. The IRC maintains capabilities for isoflurane induction in air mixture and this system is available for use for the duration of the imaging procedure. This method and all other types of anesthesia used for imaging procedures must be explicitly approved in the investigator s primary Animal Use Protocol Imaging Methods For all imaging procedures included in this protocol, animals will be imaged either alive or dead. For in vivo imaging procedures animals will be anesthetized as described below. Animals will be placed on a special purpose cradle (see figure above) that fits into the MRI or CT scanner. The cradle has a built-in bite bar that is used to position the animal and keep it in a fixed position during the imaging procedure. The bite bar is put in position in between the animal s teeth after anesthesia has been administered and the animal is asleep. An RF coil for detection of the MRI signal is then placed over the animal in the location of 120

121 interest. The cradle contains a warm water circulation system to maintain the animal s body temperature during the imaging procedures. Once the animal is installed in the MR-compatible cradle as show in the picture to the left, the cradle and animal are positioned carefully within the MRI or CT scanner so as to align the appropriate anatomical region with the center of the image field of view. In this position, the animal is monitored continuously using a closed circuit video system with a camera aimed inside the scanner and zoomed on the animal and a monitor placed at the operator s console. Vital signs are monitored using remote monitoring equipment that provides periodic assessment of cardiac and respiratory rate, oxygen saturation and CO2. The animal technician continuously monitors the animal using these tools to insure that the animal remains in a physiological equilibrium appropriate to the experimental protocol. MRI procedures generally last from 1 hour to 6 hours depending upon the types of imaging and spectroscopy information needed and the physiological conditions intended. At the end of the imaging experiment, the animal is removed from the scanner and returned to its cage for recovery. Following recovery from anesthesia, the animal is transported to appropriate animal housing according to the procedures described below. If the imaging procedure is terminal, the animal technician will euthanize the animal before recovery from anesthesia. Contrast Agents Contrast agents mentioned above such as Gadolinium (Gd) and USPIO may be used to enhance contrast in MR imaging procedures. These contrast agents are approved for human use by FDA either under a full 510k approval or under an IDE approval. As such, these agents have been fully tested for adverse biological effects 121

122 in humans and animals and have been found to be safe for use in mammalian species. These contrast agents are administered i.v. in the anesthetized animal in a volume not to exceed 200 µl. Gadolinium can also be administered i.p. with a total volume not to exceed 1 ml (50% contrast agent/ 50% physiological saline). No pain is caused by the injection because the animal is fully anesthetized. This procedure is routinely used in humans without complications, therefore we do not anticipate any undo distress to the animals, nor do we anticipate lethality. Animals - At this time, this protocol will include imaging procedures for mice, rats and rabbits. Additional species may be requested later in an amended protocol. Animals may be brought to the IRC for imaging either as living specimens, freshly dead or fixed specimens. In the case of the deceased animals, the P.I. will be responsible for obtaining the needed approval for use of animal tissues. In the case of living animals, the P.I. must have an active IACUC protocol in place to cover the use of the animal as well as an amendment to permit the use of imaging procedures under this protocol. Transport of Animals to and from IRC Animals will be transported from the Veterinary Services animal housing facilities in CCHRF to the IRC in the same building in approved animal housing containers. Animals will be brought to the animal prep lab in the IRC where anesthesia will be administered and specific preparations such as i.v. line and breathing tubes may be inserted. Following survival imaging procedures, animals may be handled in one of the following ways. 1) For longitudinal imaging studies where the animals will be imaged repeatedly over a short span of time (up to 2 weeks) the animals may be housed within the IRC Bubble. This temporary housing facility is 122

123 infection controlled and provides appropriate enrichment for the animals during a short stay. Vet services personnel will monitor, feed, clean and treat animals in this location according to the same procedures used in Animal Care facilities. This facility will provide infection isolation for short-term studies in immunocompromised animals. 2) For longitudinal imaging studies where animals will be imaged repeatedly over a longer span of time (more than 2 weeks) the animals can be returned to animal care to a non sterile holding area where infection control restrictions do not prevent reentry of animals on a periodic basis. This arrangement will be utilized for long term imaging evaluation of animals that are not immunocompromised. 3) For studies involving the use of imaging in animals that will later be used in other procedures, reentry into the animal care facilities will be carried out according to the Vet Services guidelines. Animals may also come to the IRC for terminal imaging procedures. Such procedures may involve imaging of dead animals or termination of the animals following the imaging procedure. The status of the animal for imaging procedures will be determined on the basis of the scientific goals of the experiments and euthanasia, if required will be carried out according to the procedures outlined in each investigators Animal Use Protocol for the specific experiment. 123

124 Appendix 3 Abstracts and Conference Proceedings 124

125 MR detection of Tumor cells labeled with USPIO using DOPS Liposomes ISMRM 14th Scientific Meeting & Exhibition, Seattle, USA, May 2006 V. Kaimal 1,2, S. K. Holland 1,2,V.J. Schmithorst 2, and X. Qi 3 1 Biomedical Engineering, University of Cincinnati, Cincinnati, OH, United States, 2 Imaging Research Center, Cincinnati Children's Hospital Medical Center, Cincinnati, OH, United States, 3 Human Genetics, Cincinnati Children's Hospital Medical Center, Cincinnati, OH, United States, 4 Chemistry, University of Cincinnati, Cincinnati, OH, United States Introduction: Liposomes have been widely used to deliver drugs to tissue in vivo. They can also be used to label cells with MR contrast agents such as Ultrasmall SuperParamagnetic Iron Oxide (USPIO) nanoparticles. However, labeling of non-phagocytic cells with USPIO for MR detection requires that the liposomes encapsulate and deliver sufficient quantities of the contrast agent. Tumor-specific liposomes could potentially be used to deliver USPIO to the tissue, aiding in earlier detection and better visualization using MRI. Delivery and uptake of targeted drugs can also be estimated using contrast enhanced MR microimaging, by using liposomes as dual carriers for the drug and the contrast agent. In this preliminary study, we have adapted a method for encapsulating Combidex (Advanced Magnetics, MA, size ~20nm) in liposomes made of dioleylphosphatidylserine or DOPS (Avanti Polar Lipids, Alabaster AL) and demonstrate that these liposomes can be effectively delivered to human neuroblastoma cells in vitro for detection using MRI. Materials & Methods: a) Preparation of liposomes: Sonication of dextran coated USPIO particles in aqueous solution with DOPS does not yield sufficient encapsulation in the liposomes. In order to increase USPIO content in liposomes, a chemical coupling method as described by Bogdanov et al [1] was used with minor modifications. Briefly, the dextran coating on the USPIO particles is oxidized to generate aldehyde groups. Aldehydes form a covalent Schiff bond at high ph with amines of DOPS. Liposomes obtained have a mean size of 150nm as confirmed by N4 + Particle Sizer (Beckman Coulter, CA) analysis. The liposome solution is dialyzed against a low ph solution to detach USPIO bound to the external layer of the liposomes. Unencapsulated USPIO are removed by affinity R2(ms -1 ) (32,0.04) y = x R 2 = Figure 1: Standard R2 plot of USPIO and liposomes in PBS chromatography using a Con-A Sepharose 4B column (Amersham Biosciences Corp., NJ). The USPIO-DOPS liposome structure was confirmed by conventional electron microscopy. A standard R2 relaxivity curve generated using known quantities of free USPIO and DOPS liposomes mixture was used to estimate the iron concentration in the DOPS liposomes (Fig. 1). A maximum content of 32µg Fe/ml was achieved using 1 mm DOPS concentration. b) Uptake in neuroblastoma cells: Four samples of neuroblastoma cells were prepared with approximately 10,000 cells per group. The first and second samples were incubated with a 100 µm and 300 µm USPIO- DOPS liposome preparation in growth medium respectively. The third sample contained liposome-uspio solution prepared by sonication without chemical coupling and the fourth sample contained cells with no USPIO or liposomes. After incubation for 36 hours, the cells were washed 4 times, trypsinized and fixed in a mixture of 0.5% agarose solution and growth medium (1:1) in 4ml glass vials. c) Imaging: High resolution MR imaging of the cells was performed using a 7T Bruker Biospec scanner using gradient echo methods optimized for T2* weighting. A 3D FLASH imaging sequence with TR/TE/θ of 200ms/35ms/10 and a 320! 320! 64 matrix was used for a 3.2cm!3.2cm!0.64cm FOV resulting in an isotropic 100µm resolution. 125

126 Figure 2: MR image of vials containing cells fixed in agarose Results & Discussion: The MR images (Fig.2) indicate uptake of USPIO particles by cells in samples 1 and 2, with sample 2 showing an increased uptake corresponding to the higher concentration of USPIO- DOPS liposomes. A much lower number of cells was detected in sample 3, containing cells incubated with liposome-uspio solution prepared by sonication. An estimate of the number of cells detected in each vial was obtained using a post-processing algorithm written in IDL. Number of cells detected in vial 2 was approximately 1.4 times compared to vial 1. The average contrast-to-noise ratio (CNR) between the gel and hypo-intensity regions representing cells was (std. dev 11). Conclusion: At this stage, we have successfully demonstrated that USPIO can be coupled with DOPS in sufficient quantities to form liposomes that are internalized by neuroblastoma cells in vitro. The cells can be detected using high resolution gradient echo imaging methods. DOPS liposomes are also being used as carriers for a drug currently being investigated for its effect in inducing apoptosis in human neuroblastoma cells. In future studies, this drug will also be incorporated into the liposome and targeted to tumor cells. Ultimately, we aim to test this approach of using DOPS liposomes for the combined delivery of the drug and USPIO to tumors in vivo and use high resolution MR imaging to estimate the delivery efficacy and effect of the drug. References: 1. Bogdanov, A.A., Jr., et al., Trapping of dextran-coated colloids in liposomes by transient binding to aminophospholipid: preparation of ferrosomes. Biochim Biophys Acta, (1): p

127 MR imaging of targeted delivery of Saposin C and USPIO to tumor cells in vivo using liposomes ISMRM Cellular and Molecular Imaging Workshop, February 2007 V. Kaimal 1,2, S. K. Holland 1,2, Z. Chu 3, D. D. Richardson 4, and X. Qi 3 1 Biomedical Engineering, University of Cincinnati, Cincinnati, OH, United States, 2 Imaging Research Center, Cincinnati Children's Hospital Medical Center, Cincinnati, OH, United States, 3 Human Genetics, Cincinnati Children's Hospital Medical Center, Cincinnati, OH, United States, 4 Chemistry, University of Cincinnati, Cincinnati, OH, United States Introduction: The delivery of targeted drugs to tumor cells using liposomes can be estimated by via MR imaging by coupling it with contrast agents like USPIO (Ultrasmall Super-Paramagnetic Iron Oxide). Previously, we reported on a method for encapsulating Combidex (Advanced Magnetics, MA, size~20nm) in liposomes made of dioleylphosphatidylserine or DOPS (Avanti Polar Lipids, Alabaster AL) and demonstrated that these liposomes can be effectively delivered to human neuroblastoma cells in vitro and detected using MRI[1]. DOPS liposomes are being used as carriers for the protein Saposin C, currently being investigated for its effect in inducing apoptotic death of human neuroblastoma cells. In this study, we have incorporated Saposin C and USPIO in the liposome and demonstrate in vivo the combined delivery of the drug and USPIO to tumors using MR imaging. Passive targeting to tumors in vivo is enabled by some inherent properties of tumors leaky tumor vasculature resulting in enhanced permeability and retention of the liposomes [2], presence of phosphatidylserine (PS) in high concentrations on the surface of tumor cells and blood vessels [3], the low extracellular ph (~5-6.15) in tumors [4] and the fusogenic activity of Saposin C [5] in aiding in the fusion of PS lipid membranes, at low ph. Materials & Methods: a) Liposome preparation: USPIO and Saposin C laden liposomes were prepared using a chemical coupling method as described by Bogdanov et al [6] with minor modifications. Saposin C to DOPS molar ratio of 1:30 was used and a maximum iron content of 90µg Fe/ml was achieved using 1 mm DOPS concentration, as determined by inductivelycoupled plasma spectroscopy (ICP). b) In vivo imaging: Animal experiment protocols were approved by the Institutional Animal Care and Use Committee. Female athymic mice were implanted subcutaneously with 5 X 10 6 human neuroblastoma tumor cells and tumors established in about 6 weeks. High resolution MR imaging of the tumors was performed on a 7 Tesla Bruker Biospec scanner using a volume coil for transmitting and surface coil placed on the tumor for receiving the MR signal. T2* weighted 2D and 3D FLASH sequences (TE/TR=5ms/20ms, FA=10, 16 averages) were used with a 100µm isotropic resolution. Images were taken before and after two injections of 200µL of the DOPS-Saposin-C-USPIO solution over two days. c) ICP: The mice were sacrificed and the tumor was extracted for ICP analysis. Results: MR images of the tumor after injection (right) show marked hypo-intensity compared with the images before injection (left). ICP of tumor tissue indicated 6-fold iron content when compared tumor of a control mouse (18 µg vs. 3 µg). Conclusion: At this stage, we have demonstrated that the DOPS-Saposin-C-USPIO complex can be successfully targeted to tumor cells in vivo T2* weighted MR image of tumor before (left) and after (right) injection of the DOPS-Saposin C- USPIO liposome solution. 127

128 and can be visualized using high resolution MR imaging. In future experiments, we will attempt to use MR imaging to quantify the delivery efficiency of the DOPS-Saposin-C-USPIO complex in vivo and effect of Saposin C in inducing apoptotic cell death in various strains of tumors. References: 1. Kaimal et al, Proc. ISMRM Allen et al, Science, (5665): p Utsugi, T., et al. Cancer Res, (11): p Vaupel, P. et al,cancer Res, (23): p Wang, Yet al Arch Biochem Biophys, (1): p Bogdanov, A.A., Jr., et al., Biochim Biophys Acta, (1): p

129 Imaging the targeted delivery of the protein Saposin-C to tumors in mice Joint Molecular Imaging Conference, Providence Rhode Island, September 2007 V. Kaimal 1,2, S. K. Holland 1,2, Z. Chu 3, and X. Qi 3 1 Biomedical Engineering, University of Cincinnati, Cincinnati, OH, United States, 2 Imaging Research Center, Cincinnati Children's Hospital Medical Center, Cincinnati, OH, United States, 3 Human Genetics, Cincinnati Children's Hospital Medical Center, Cincinnati, OH, United States. In this study, we report on imaging the targeted delivery of a potential anti-cancer and targeting agent, the protein Saposin C, to tumors in mice using liposomes. Both MRI and fluorescence techniques were used to image the delivery of the Saposin-C- DOPS complex to tumors in nude mice bearing human neuroblastoma tumors. For high resolution MR imaging, liposomes fabricated from di-oleyl-phosphatidyl-serine (DOPS) with Saposin C molecules embedded in lipid bilayer, were tagged with Ultra-Small Super- Paramagnetic Iron Oxide particles (USPIO). Gradient Echo images tuned for sensitivity to the USPIO labels (FLASH, TR=5ms/20ms, flip angle=10, 16 averages, 100µm isotropic resolution) were obtained after tail vein injection of the Saposin-C-DOPS-USPIO particles in tumor-bearing mice. Images were acquired using a 7 Telsa small animal MRI scanner. Figures 1(a) and 1(b) show the tumor before and 3 hours after injection respectively. MR images acquired at different time points show evidence of uptake of Saposin-C-DOPS-USPIO in the mouse tumor, For fluorescence imaging, the Saposin-C-DOPS liposomes were labeled with a far-red fluorescent probe, (CellVue Maroon, PTI Research Inc.) and injected in tumor bearing mice. Imaging was performed on IVIS 200X (Xenogen) imaging system. Figure 2(a) shows that Saposin-C-DOPS liposomes are targeted to tumors while the tumors in control mice in Figure 2(b) do not show accumulation of the complex. Quantitative uptake Saposin-C-DOPS-USPIO nanoparticles in tumors is validated using an inductively coupled plasma (ICP) spectroscopy method. Results from MRI, bioluminescence and ICP will be presented to prove the feasibility of the in vivo methodology. Figure 1: 129

130 Figure 2: 130