DEVELOPMENT OF PARACEST MRI TO DETECT CANCER BIOMARKERS GUANSHU LIU. Submitted in partial fulfillment of the requirements

Transcription

1 DEVELOPMENT OF PARACEST MRI TO DETECT CANCER BIOMARKERS by GUANSHU LIU Submitted in partial fulfillment of the requirements For the degrees of Doctor of Philosophy Dissertation Adviser: Dr. Mark D. Pagel Department of Biomedical Engineering CASE WESTERN RESERVE UNIVERISITY May 2008

2 CASE WESTERN RESERVE UNIVERSITY SCHOOL OF GRADUATE STUDIES We hereby approve the dissertation of candidate for the Ph.D. degree *. (signed) (chair of the committee) (date) *We also certify that written approval has been obtained for any proprietary material contained therein.

3 Table of Contents Table of Contents...1 List of Figures...4 List of Tables...7 Acknowledgements...8 List of Abbreviations...10 Abstract...13 Chapter 1. Introduction Molecular imaging of cancer biomarkers Cancer biomarkers in cancer studies Molecular imaging for cancer biomarker detection Molecular imaging in theragnostics Basics of MRI contrast T 1 contrast agents for molecular imaging T 2 /T 2 * contrast agents for molecular imaging CEST contrast agents for molecular imaging Theory Advantage of CEST contrast enhancement Classification and applications of CEST contrast agent PARACEST contrast agents for molecular imaging Lanthanide induced shifts Lanthanide based PARACEST contrast agents Applications of PARACEST contrast agents Conclusion...41 References...45 Chapter 2. Design and characterization of a new irreversible responsive PARACEST MRI contrast agent that detects nitric oxide Introduction Materials and methods Sample preparation

4 2.2.2 PARACEST and T 1 NMR measurements MRI measurements Molecular modeling Results Discussion Conclusion...78 References...80 Chapter 3. Measurement of extracelluar ph using a single MRI contrast agent based on PARACEST effects Introduction Theory Materials and methods Chemicals General methods Synthesis MR acquisition procedures Animals Results and discussion Synthesis of DO3AoAA ligand CEST characteristics of different Ln (III) complexes ph dependencies of PARACEST signals ph dependencies of water exchange rates Accuracy of the ph measurement In vivo demonstration Conclusion References Chapter 4. In vivo PARACEST MRI with improved temporal resolution Introduction Theory Theoretical model for PARACEST contrast The dynamics of PARACEST contrast

5 The temporal resolution and efficiency of PARACEST MRI Materials and methods Chemicals Simulations Animals MRI acquisition procedures Results Simulations Phantom studies In vivo studies Discussion Conclusion References Chapter 5. Future studies Improvement in sensitivity Nanoparticle delivery systems Enzyme based amplification Improvement of cellular internalization Improvement of in vivo implementation by using a control agent Improvement of selectivity Reduction of direct water saturation Reduction of the interferences of multiple PARACEST sites Reduction of RF power deposition Improvement of quantification Conclusion References Appendix A1. The concentration of water A2. Resolutions of NMR spectrum and CEST spectrum A3. Scan average A4. Standard deviations of PARACEST images

6 List of Figures Figure 1.1. Schematic representation of molecular imaging for cancer biomarker detection and theragnostics 16 Figure 1.2. The chemical structures and commercial names of the ligands that are used as clinically approved Gd(III) based T 1 MRI contrast agents Figure 1.3. Schematic representation of structures of SPIO (a) and supermagnetism (b). 27 Figure 1.4. A schematic of CEST contrast enhancement in MRI Figure 1.5. The two pool model for CEST contrast Figure 1.6. Useful ranges of pseudo-contact shifts (PCS) for Ce(III), Yb(III), and Dy(III)- containing proteins Figure 1.7. Chemical structure of Ln(III)-DOTAMGly...38 Figure 2.1. The illustration of the reaction of Yb(1) with NO in the presence of oxygen converts aromatic amines to a triazene on Yb(2) Figure 2.2. The CEST spectrum of Yb(1)...65 Figure 2.3. The effect of saturation power and saturation delay on the PARACEST effect of Yb(1) Figure 2.4. The effect of a) temperature and b) ph on the PARACEST effect of Yb(1).. 66 Figure 2.5. The effect of concentration on the PARACEST effect of Yb(1) Figure 2.6. The MR images of 30 mm Yb(1) before reaction and after reaction Figure 2.7. The T 1 relaxivities of Gd(1) and Gd(2) Figure 2.8. The molecular model of Yb(1) Figure 3.1. The synthetic route of DO3AoAA Figure 3.2. The CEST spectra of 20 mm Yb-DO3AoAA in PBS with ph varied from 6.12 to 8.0 at 37 o C...98 Figure 3.3. The extraction of PARACEST effects of amine and amide from experimental data by a three-lorentzian model...99 Figure 3.4. The different ph dependencies of amide CEST and amine CEST (a), and the ph measurement calibration using the proposed PARACEST agent (b)

7 Figure 3.5. The maximal proton water exchange rates for amide (a) and amine (b) which are estimated by fitting the PARACEST signals with a time dependent PARACEST model Figure 3.6. The ph-exchange rate dependencies of amine and amide protons Figure 3.7. The demonstration of the proposed ph measurement MRI method in a mouse tumor model Figure 4.1. Schematic of the solutions to improve the temporal resolution of a PARACEST MRI study by using A: a multiple-echo strategy, and B: a short repetitive saturation strategy Figure 4.2. Structures of Eu-DOTAMGly (Eu(1)), Tm-DOTAMGly (Tm(1)) and Eu- DOTA-OBS2Gly2COOH (Eu(2)) Figure 4.3. A: Pulse sequence diagram of a PARACEST sequence with a multiple-echo imaging scheme. B: Pulse sequence diagram of a PARACEST sequence with a short repetitive saturation scheme Figure 4.4. Simulations of the water signal during the saturation scheme and after the removal of saturation pulses Figure 4.5. Simulation of a build-up and maintenance of PARACEST by using short repetitive saturation pulses Figure 4.6. The in vivo DCE MRI study of tumor tissue with a presat-rare MRI method with a RARE factor of 16 (80 sec/image) Figure 4.7. The in vivo DCE MRI study of liver with a presat-flash MRI method with a 300 ms TR (76 min/image) Figure 5.1. The schematic representation of a) liposome, b) polymer micelle nanoparticle and c) dendrimer as the delivery systems for imaging probes Figure The schematic representation of a Yb-DO3AoAA HPMA co-polymer system as a highly sensitive NO detecting agent Figure 5.3. The three major isomers of nitric oxide synthase and their functions

8 Figure 5.4. The schematic representations of a) pinocytosis b) phagocytosis c) receptor mediated endocytosis and d) transporter delivery mechanism for cell to uptake extracellular molecules Figure 5.5. The schematic representation of a Yb-DO3AoAA and Tm-DOTAMGly bifunctional HPMA co-polymer system Figure 5.6. CEST spectra of a mouse tumor ROI with 10 μt and 5 μt RF saturation power Figure 5.7. The CEST spectra of Tm-DO3AoAA at ph 6.93 and ph Figure 5.8. The CEST spectra of 40 mm Eu-DOTAMGly and PBS by a FISP-CEST method Figure 5.9. The demonstration of a keyhole approach for the reduction of RF power in a PARACEST study Figure The demonstration of the approach to quantify PARACEST based on the partial progress model Figure A1. The relative MRI signal for phantom, bladder and tumour measured in a DCE PARACEST MRI study Figure A2. The CEST spectra simulated by using RF sturation power of 128 Hz, 256 Hz and 512 Hz Figure A3. The effects of scan average number on CEST spectra of a 20 mm Yb- Do3AoAA NMR sample (ph 7.5, 37 o C)

9 List of Tables Table 1.1. Current molecular imaging modalities and their characterization...18 Table 1.2. Some SPIO contrast agents that are clinically approved and in clinical trials 26 Table 1.3. Magnetic moment, chemical shift and exchange rate of Ln(III)-DOTAMGly 39 Table Comparison of T 1, T 2, CEST and PARACEST contrast agents...42 Table 3.1. The PARACEST offsets and effects of the DO3AoAA complex with different lanthanide ions Table 3.2. The accuracies of the proposed ph measurement method Table 4.1. The simulated PARACEST effects of 20 mm Eu(1) with different PARACEST MRI methods that employ a long T s Table 4.2. The quantitative comparison of contrasts, temporal resolutions and efficiencies of different PARACEST MRI methods by using a 20 mm Eu(1) phantom Table 5.1. Nanoparticle systems for MRI molecular imaging and their typical characteristics Table 5.2. MRI pulse sequence candidates for high temporal resolution PARACEST MRI methods and their preferable main magnetic field strength (B 0 )

10 ACKNOWLEDGEMENTS My graduate study in the Department of Biomedical Engineering at Case Western Reserve is a four-year long journey with immense supports and encouragements of many kind people around me. I would like to take this opportunity to thank all these people: my advisor, my guidance committee, my colleagues, my friends, and my family. Without the support of these people, this thesis would not appear in its present form. First and foremost, I would like to express my sincerest gratitude to Dr. Marty Pagel, my advisor, for his patience and kindness, for his excellent academic expertise and for his sharing with enthusiasm and inspiration. I owe thanks for his guidance along my path from a student to a researcher. It has been a wonderful experience and real fortune to be his student, and I will enjoy his friendship for the rest of my life. I am also greatly indebted to my guidance committee, Dr. Mark Griswold, Dr. Jean Tkach, Dr. Xin Yu and Dr. Michael Zagorski. I also deeply appreciate their efforts on improving the quality of this work. Dr. Mark Griswold and Dr. Jean Tkach deserve many thanks for sharing their incredible depth of expertise in many aspects of magnetic resonance imaging. Dr. Xin Yu deserves many thanks for her kind help and enlightening courses. Dr. Michael Zagorski deserves many thanks for the time and energy he spent on my behalf. They are all inspirational academicians and brought in insightful critiques and advices throughout my thesis study. It really was a great honor to be a student under their mentorship. I am very happy about and grateful for the support from many colleagues at the Department of Biomedical Engineering and the Department of Radiology. I will give my special thanks to all of my colleagues in the Contrast Agent Molecular Engineering - 8-

11 Laboratory (CAMEL): Dr.Yuguo Li, Dr. Meser Ali, Dr. Byunghee Yoo, Rachel Rosenblum and Vipul Sheth, for their help, support and friendship. I would like to thank all the colleagues at the Case Center for Imaging Research (CCIR): Dr. Chris Flask, Deborah Sim, Kelly Covey and Jack Jesberger for their assistances, which guaranteed the progress of all my MR imaging research. Dr. Chris Flask deserves many thanks for his supports and expertise in MR pulse programming. I also would like to extend my sincerest gratitude to all my colleagues and friends from Dr. Xin Yu group, Dr. Wilson s group, Dr. Duerk group, Dr. Saidel s group, Dr. Flask s group and other research groups. They all are talented students and outstanding researchers. I also would like to express my deep appreciation to all my Chinese friends, wherever they are. Many thanks go to my friends in Cleveland for the innumerable altruistic care and help to me and our family especially during the most stressful phases after we just had our first baby. Last, but not least, I want to thank my dear family for their unfailing encouragement and support. They were always there for me to share my sorrows and joys. Seeing the pride in their eyes is the most meaningful thing to me in this world. No words can express my gratitude to my wife Junnan. She is the source of my happiness, and the source of my strength to survive even during the most frustrating moments. The truly sincere thanks also go to my dear parents for their absolute confidence in me, go to my grandparents for their forever love, and go to our young daughter Claire for her most beautiful smile in this world. - 9-

12 List of Abbreviations BLI: Bioluminescence Imaging BW: bandwidth CA: Contrast Agent CEST: Chemical Exchange Saturation Transfer CF: Cystic Fibrosis CNR: Contrast to Noise Ratio CPP: Cell Penetrating Peptide CT: Computed Tomography CW: Continuous Wave DANTE: Delay Alternating with Nutations for Tailored Excitation DCE MRI: Dynamic Contrast Enhanced MRI DO3A: 1,4,7,10-Tetraazacyclododecane-1,4,7-triacetatic acid DO3A-oAA: 1, 4,7,10-tetraazacyclododecane-1,4,7,10-tetraacetic acid o-aminoanilide DOTA: 1,4,7,10-Tetraazacyclododecane-1,4,7,10-tetraacetic acid enos: epithelial Nitric Oxide Synthase EPI: Echo-Planar Imaging FDA: Food and Drug Administration FI: Fluorescence Imaging FISP: Fast Imaging with Steady-state Precession GRASS: Gradient-Recalled Acquisition in the Steady State GRE, GE: Gradient Echo - 10-

13 HGP: Human Genome Project HPMA: N-(2-hydroxypropyl) methacrylamide htfr: human Transferrin Receptor inos: inducible Nitric Oxide Synthase MION : Monocrystalline Iron Oxide Nanoparticles MRI: Magnetic Resonance Imaging MRS: Magnetic Resonance Spectroscopy MSME: Multiple Slice Multiple Echo MT: Magnetization Transfer NMR: Nuclear Magnetic Resonance nnos: neuronal Nitric Oxide Synthase NO: Nitric Oxide OATP: Organic Anion Transport Protein ORI: Off-Resonance Imaging PARACEST: Paramagnetic CEST PBS: Phosphate Buffered Saline PET: Positron Emission Tomography ph e : extracellular ph ph i : intracellular ph RARE: Rapid Acquisition with Relaxation Enhancement RF: radiofrequency ROI: Region Of Interest SAR: Specific Absorption Rate - 11-

14 SE: Spin Echo SNR: Signal to Noise Ratio SPECT: Single Photon Emission Computed Tomography SPIO: Super Paramagnetic Iron Oxide TE: echo time TR: repetition time US: Ultrasound Imaging - 12-

15 Development of PARACEST MRI to Detect Cancer Biomarkers Abstract by GUANSHU LIU Molecular imaging has become one of the most significant developments in MRI in the last decade with dramatic impacts for disease diagnosis and therapy. A novel MRI contrast strategy named PARAmagnetic Chemical Exchange Saturation Transfer (PARACEST) can exploit the NMR chemical shift to selectively detect molecular biomarkers, while retaining the sensitivity and spatial resolution of MRI. The development of this novel approach for cancer biomarker detection is described in the following chapters. Chapter 1: Background information is presented about the relevance of biomarker detection by molecular imaging and MRI molecular imaging. The mechanisms of MRI contrast agent including T 1 contrast agents, T 2 /T 2 * contrast agents and CEST/PARACEST contrast agents are introduced. The PARACEST contrast mechanism is then emphasized for the potential biomedical applications. Chapter 2: A new irreversible nitric oxide (NO) responsive PARACEST MRI contrast agent, Yb- DO3oAA, has been designed, synthesized and characterized. The PARACEST effects - 13-

16 have been investigated with respect to ph, temperature, and concentration for the in vivo applicability. The ability to detect NO has been demonstrated in vitro. Chapter 3: A new MRI method has been developed for assessing in vivo ph by using a PARACEST MRI contrast agent Yb-DO3AoAA. A ratiometric approach has been employed based on the two intra-molecular PARACEST signals. Our study has demonstrated that this method can be used for measuring extracellular ph within in vivo animal models without the need for a second control agent. Chapter 4: New MRI pulse sequences have been developed to address the poor temporal resolution challenges for the in vivo applications of PARACEST agents. Different strategies have been developed for the applications in different in vivo environments. These new MRI methods have been developed for high field small animal studies and tested both within in vitro and in vivo models. Chapter 5: The major drawbacks and technical hurdles in PARACEST studies are presented and followed by the discussion of future developments. A number of potential projects thereby are proposed as future studies

17 CHAPTER 1 INTRODUCTION 1.1. MOLECULAR IMAGING OF CANCER BIOMARKERS Cancer Biomarkers in Cancer Studies Cancer is a major cause of mortality and the worldwide incidence of cancer still continues to increase. Cancer has been categorized as a collection of diseases induced by genetic mutations and molecular alterations. Because of the diversity of the underlying mechanisms and phenotypes, the accurate diagnoses in the early stages and efficient therapies in the late stages are still formidably challenging. For both purposes, new techniques that can effectively target the specific biomarkers of cancer cells are believed to be the keys to a revolution in cancer diagnosis. With the complete of Human Genome Project (HGP) in 2003 and tremendous ever-growing developments in proteomics 1 and metabonomics, 2,3 there are renewed interests in discovering cancer biomarkers for early cancer detection. Recently tremendous discoveries of either individual or panel cancer biomarkers have been reported based on the modern analytical technologies, such as DNA microarrays, 4,5 mass spectrometry, 6,7 nuclear magnetic resonance (NMR) spectroscopy 8 and biosensors. 9, 10 General speaking, a biomarker is a laboratory measurement or physical sign that can be used as an indication for a clinically meaningful endpoint. 11 If a biomarker can be used as a substitute for a clinically meaningful endpoint, the biomarker is called a surrogate endpoint. 11 For example, low extracellular ph has been discovered to be highly correlated with carcinogenesis and metastasis. Therefore ph is a biomarker for tumor and its - 15-

18 metastasis. ph has been validated as a surrogate endpoint by preclinical and clinical studies. Ideally, a biomarker is strongly related to a clinically meaningful endpoint, and therefore reflects changes in clinically meaningful endpoint. However, it is not always the case in practice. The validation step therefore is critical to link a biomarker to a clinical endpoint with respect to the sensitivity, specificity and reproducibility. 12 Figure 1.1. Schematic representation of molecular imaging for cancer biomarker detection and theragnostics. The molecular imaging targets the molecular biomarkers of carcinogenesis, whereas the traditional anatomic imaging and functional imaging are related to clinical endpoints. The molecular imaging can be designed for theragnostics by monitoring changes in cancer molecular biomarkers at the earliest time after therapy Molecular Imaging for Cancer Biomarker Detection Molecular imaging is emerging as a novel multidisciplinary approach that conbines biomedicine (basic cellular and molecular biology, medicine, pharmacology, - 16-

19 biomathematics, and bioinformatics), chemistry (including biochemistry) and radiology (medical physics and image processing) into a new imaging paradigm. Molecular imaging has been developed for the visualization, characterization, and quantification of pathophysiological processes at the cellular and subcellular levels within intact living organisms. Therefore, it endows biomedical images with the ability to render abnormalities based on underlying molecular mechanisms in addition to the anatomic or functional delineation of abnormal tissues. Moreover, the tracking of the changes in biomarkers by an efficient molecular imaging technique can be both temporal and the spatial. So far, a number of imaging modalities have been developed for this novel approach. The earliest molecular imaging studies were based on the nuclear imaging, i.e. Positron Emission Tomography (PET) and Single Photon Emission Computed Tomography (SPECT), owing to the ultra-high sensitivity. The other modalities, such as optical imaging (i.e. Bioluminescence Imaging (BLI) and Fluorescence Imaging (FI)) and Magnetic Resonance Imaging (MRI), Ultrasound imaging (US) and Computed Tomography (CT), also have been brought to the field of molecular imaging with great interests. The prevailing modalities for molecular imaging are summarized in Table 1.1. The molecules or molecular events that are targeted by molecular imaging include gene expression, artificial/reporter genes, 16, 17 messanger RNA, 18 receptors on a cell membrane surface, 19, 20 other components on a cell membrane surface, 21,22 intercellular and extracellular enzymes, 14,23-25 apoptosis, 26,27 antibodies/antigens, 28 metabolites, 29 ions, and other key molecules 33 or changes in micro-environments 34,35 for pathophysiological processes

20 Table 1.1. Current molecular imaging modalities and their characterizations (modified from Ref. 36 and Ref. 37 ) Nuclear imaging (PET/SPECT) Optical imaging (BLI/FI) Sensitivity Spatial Temporal Penetrating Specific resolution resolution depth applications Up to10-12 M 1-2mm seconds minutes No limit Gene expression Receptor targeting Enzyme detection Antibody/antigen metabolism Up to mm seconds Up to 2 cm Gene expression M minutes Receptor targeting Enzyme detection Antibody/antigen MRI Up to 10-5 M μm minuteshours No limit Gene expression Receptor targeting Enzyme detection Antibody/antigen Metabolites CT minutes No limit 3D microarchitecture μm Multimodality US μm seconds minutes mm-cm Receptor targeting - 18-

21 Most of the targets that are used in molecular imaging are biomarkers or surrogate endpoints for a variety of diseases including cancers. The ultimate goal of molecular imaging is to use meaningful molecular biomarkers in clinical diagnosis and therapy. Therefore, the development of molecular imaging depends on the studies of biomarkers. Conversely, newly discovered biomarkers can also be quickly validated through visualization and quantification within in vivo models by appropriate molecular imaging techniques. Therefore, molecular imaging also plays a critical role in linking the in vitro discoveries of new biomarkers with in vivo characterization and validation of oncological pathology for new surrogate endpoints or clinical endpoints Molecular Imaging in Theragnostics One of the most promising developments in medicine is to combine in vivo diagnostic imaging with therapy to predict early therapeutic output and to determine the appropriate therapeutics. This approach, theragnostics, is believed to open a new avenue to the era of personalized medicine in the future. However, such an approach requires significantly high sensitivity and quantitative robustness. And biomarkers that are either the targets of therapeutic intervention or responsive to the therapy need to be precisely determined in vivo. Molecular imaging is considered to be a powerful tool in the development of theragnostics owing to the ability to correlate the in vivo imaging data with the underlying genotype-phenotype of the imaged tissues. For instance, recent advances in MRI molecular imaging have made it possible to detect cell apoptosis in presence of therapy. 26,27 Because apoptosis is a key mechanism for many anti-cancer drugs and therapeutics, the in vivo apoptosis imaging techniques can not only greatly facilitate the discovery and screening of new anticancer drugs, but also can aid the improvement of - 19-

22 efficacy of therapy by predicting the early therapeutic outcome. However, the responsive and targeting contrast agents are typically required in theragnostic studies. Since most of the responsive/targeting agents are still in the pre-clinical stage, clinical studies and applications are still challenging and expected to be achieved only in the distant future MRI Molecular Imaging for Cancer Biomarker Detection Among all the modalities available for molecular imaging, Magnetic Resonance Imaging (MRI, including Magnetic Resonance Spectroscopy (MRS) and MRS Imaging (MRSI)), is one of the most versatile non-invasive techniques with superb spatial resolution (<100 μm) and excellent soft tissue contrast. Regardless of inherently low sensitivity and high cost, MRI still is a preferred imaging modality for many clinical studies. The clinical translation of new MRI techniques is very feasible and has great potential for immediate impact on health care. MRS/MRSI is also a traditional technique for metabolic studies. MRI molecular imaging techniques so far have been developed to cover almost all the aspects of molecular imaging applications, such as visualizing gene expression, 14,15,38 labeling and tracking cell migration, 39,40 targeting cell surface receptors 41,19 and determining cell apoptosis. 26,27,42 With the recent advances in instruments, methodologies and new contrast agents, MRI molecular imaging has become one of the most important techniques for the development of cancer diagnostics and therapeutics BASICS OF MRI CONTRAST The signal of MRI is determined by both the MR measurement parameters and sample characteristics. The sample characteristics includes spin density, T 1 or longitudinal relaxation time, T 2 /T 2 * or transverse relaxation time, and chemical shifts of components - 20-

23 in the sample under study. For instance, by using a π/2 radiofrequency (RF) pulse, the MRI signals produced by a typical 1 H spin echo (SE) and gradient echo (GE) MRI sequence are shown in Eq. [1.1a] and [1.1b] respectively. S = k (1- e ) e SE -T R/T1 -T E/T2 ρ0 [1.1a] S = k (1- e ) e -T R/T1 -T E/T2* GE ρ0 [1.1b] where S is the detectable MRI signal, k is constant, ρ 0 is the spin density, T R is the repetition time and T E is the echo time of the MR measurement. The Signal to Noise Ratio (SNR) is the criterion to evaluate the image quality of a MRI method. The SNR is defined as the ratio of measured signal (S) and unavoidable noises (N) in the MR measurement (Eq. [1.2]). SNR=S/N [1.2] In a MRI study, contrast (C) is defined as the difference in MRI signals between different tissues or between normal and damaged tissues (Eq. [1.3a]). In order to better visualize the pathological and/or physiological variation, a significant contrast has to be obtained. Analogous to SNR, the term Contrast to Noise Ratio or CNR is often used for the ability of a MRI method to distinguish the targeted tissue (with MRI signal of S 1 ) from others (with MRI signal of S 2 ) (Eq. [1.3b]). C=S1-S 2 CNR=SNR1-SNR 2 [1.3a] [1.3b] With the appropriate manipulation of pulse sequences, the contrast can be made to reveal much desirable information such as anatomical characterization, tissue function, and molecular events. For instance, diffusion MRI that detects water molecules diffusional movement rate has been developed to distinguish the tissues with different micro

24 architectures and applied to evaluate therapeutic efficacy, where significant changes in cellular compartments occur. 43 Another example is that the abnormalities can be characterized by investigating the change in macromolecular content by using a Magnetization Transfer (MT) method. 44 However, insufficient contrast is nevertheless a major drawback in many applications of MRI due to the intrinsic low SNR and low CNR between normal and pathological tissues. In such cases, contrast agents can be used to achieve dramatic enhancements in MRI contrast. For instance, Dynamic Contrast Enhanced MRI (DCE MRI) has been widely used to investigate the angiogenesis based on the distinct pharmacokinetics of contrast agents between tumor and normal tissues. 45 More efficient MRI contrast agent systems have been developed with higher specificity (targeting agents and responsive agents) and higher sensitivity (nanoparticle delivery systems). Although numerous MRI contrast agents have been reported, the prevalent mechanisms for contrast enhancement fall in three major categories: T 1 contrast, T 2 contrast and Chemical Exchange Saturation Transfer (CEST) contrast. The mechanisms and the design guideline of each category will be introduced in the following sections and a quick comparison will be presented in the Section T 1 CONTRAST AGENTS FOR MOLECULAR IMAGING One of most successful paradigms of MRI contrast agents so far is the Gadolinium (Gd) based T 1 contrast agent. Gd(III) is a paramagnetic lanthanide ion with a high magnetic moment (μ eff = 7.94) and long electronic relaxation time (T 1e, T 2e = sec). 46 Because the free Gd(III) is toxic, its complexes that are kinetically inert and thermodynamically stable during the circulation are used instead for in vivo contrast - 22-

25 enhancements. Lanthanide complexes typically possess the most exchange-labile characteristics with reported first order water-exchange rate constants of 4.7 x10 7 and 8.3 x10 8 s Therefore the relaxation enhancement is often maintained despite the complexation. Clinically approved Gd(III) complexes are based on derivatives of either 1,4,7,10-tetracarboxymethyl-1,4,7,10-tetraazacyclododecane (DOTA) or diethylenetriaminepentaacetic acid (DTPA) as listed in Figure 1.2. Figure 1.2. The chemical structures and commercial names of the ligands that are used as clinically approved Gd(III) based T 1 MRI contrast agents. 48 Observed T 1 is the combination of the intrinsic T 1 (T 1,0 ) and the contribution from contrast agent (r 1, relaxivity and contrast concentration molar concentration [CA]) and - 23-

26 can be described by Eq. [1.4]. This equation also implies that the relaxivity (mm -1 sec -1 ) is the determinant for the efficiency of the T 1 contrast agent system. 1 1 = + r1 [ CA] [1.4] T T 1, obs 1,0 The two major contributions to relaxation enhancement of hydrated Gd(III) are the through-bond effect (contact, scalar, or hyperfine) and the through-space effect (dipolar). 46 In practice the scalar interaction is negligible for the most popular magnetic field strengths (>10 MHz) used in clinical and pre-clinical MRI scanners. Both the innersphere and outer-sphere coordinated water can be induced to a relaxation enhancement. However, the former mechanism has greater contribution to the relaxation enhancement 49, 50 except for specific agents. The rationale to design new contrast agents requires the elucidation of the relationship between the structure of the ligand and the relaxation enhancement. The quantum chemistry mathematical model, i.e. Solomon-Bloemgergen equations, 51 has been used to illustrate the mechanism of relaxation enhancements. The number of Gd(III) coordinated water molecules, or q, is expected to be high for improving the relaxivity. However, this strategy is often impractical due to the increased toxicity associated with higher q. The decrease of the distance between the Gd(III) and protons of coordinated water, r GdH, will increase relaxivity. τ m is the proton residence time in bound water molecules. A short τ m that allows a rapid exchange process for relaxation enhancement is desirable in practice, where the electronic relaxation rate is in the 1-10 GHz range. τ r is the re-orientational correlation time or tumbling time. In general, the slower tumbling (rotating) species will have higher relaxivity. At the magnetic field strengths of T, τ r is the most important parameter for determining relaxivity. 52 In brief, q, r GdH, τ m, and τ r are the - 24-

27 parameters that can strongly affect the relaxivity of a contrast agent, therefore should be taken under consideration in the design of new T 1 contrast agents. The Gd(III) based T 1 contrast agents are the most widespread MRI contrast agents for both pre-clinical studies and clinical diagnoses. Despite the great success of these agents as non-specific blood pool agents, the development of Gd(III) based targeted/responsive contrast agents are still challenging due to the inherent low sensitivity of MRI techniques. For example, the lowest calculated detection threshold of GdHPDO3A is 0.5 mm, whereas the typical concentration of tumor surface receptors is about mm. 41 New developments have to improve sensitivity as well specificity and pharmacokinetic features for realistic in vivo applications. With the recent developments of the new generation of agents (high load, targeting and responsive), Gd(III) based contrast agents have been used for numerous molecular imaging applications, such as measuring gene expression through a reporter, 14 enzyme activity, 53, 54 cell labeling and trafficking, 55 ph, 56 and ions T 2 /T 2 * CONTRAST AGENTS FOR MOLECULAR IMAGING T 2 or T 2 * contrast is another dominant contrast mechanism in MRI. A T 2 contrast agent can improve MRI contrast by darkening the region of interest (ROI) by significantly shortening the T 2 /T 2 * relaxation times (Eqs. [1.1a] and [1.1b]). So these agents are also traditionally called negative contrast agents. Most T 2 contrast agents are super paramagnetic iron oxide (SPIO) species, such as oral (large) SPIO ( nm), standard SPIO (SSPIO) ( nm), ultrasmall SPIO (USPIO) (10-40 nm) and monocrystalline iron oxide nanoparticles (MION) (3-10 nm). Agents of all types but not - 25-

28 the MION have either completed or are currently undergoing clinical testing (Table 1.2). Generally, small size SPIO contrast agents (<50 nm) are more favorable for molecular imaging due to the favorable pharmacokinetics and the feasibility of cellular internalization. 57 Table 1.2. Some SPIO contrast agents that are clinically approved or in clinical trials 57,58 Agent Class Trade and common names Status Mean particle size AMI-121 Oral Lumirem, Gastromark, Approved >300 nm SPIO Ferumoxsil OMP Oral SPIO Abdoscan Approved 3.5 μm AMI-25 SSPIO Feridex, Endorem, Ferumoxide Approved nm SHU555A SSPIO Resovist Phase III 62 nm AMI-227 USPIO Sinerem, Combidex, Ferumoxtran Phase III nm NC USPIO Clariscan Completed phase II (discontinued) 20 nm CODE USPIO Advanced magnetics Phase II nm

29 SPIO contrast agents are typically composed of crystalline core (usually magnetite (Fe 3 O 4 ) or maghemite (γ-fe 2 O 3 ) and surface modifying moieties (dextran or starch derivatives). The SPIO crystalline core consists of a number of magnetic domains (Figure 1.3), which are sufficiently large (on the order of a nanometer) crystal-containing regions of unpaired spins and thus can be regarded as thermodynamically independent. A magnetic domain has a net magnetic moment larger than the sum of individual magnetic dipoles that it contains, which is so called supermagnetism. Figure 1.3. Schematic representation of structures of SPIO (a) and supermagnetism (b). A, spinal crystal structure of SPIO magnetic domain; B, SPIO core with multiple magnetic domains; C, a SPIO particle with multiple crystal cores; D, the SPIO crystalline core is ferromagnetic when no external magnetic field; E, the SPIO crystalline core is supermagnetic when a external magnetic field B 0 is applied, their magnetic moments align in the direction of B

30 Similar to observed T 1 relaxation, observed T 2 is the combination of the intrinsic T 2 (T 2, 0 ) and the contribution from the contrast agent (r 2, relaxivity, and concentration [CA] (Eq. [1.5]). T 2 can be measured with a Carr-Purcell-Meiboom-Gill (CPMG) pulse sequence or from the line width of the peak in the MR spectrum (line-width = 1/(πT * 2 )). 1 1 = + r2 [ CA] [1.4] T T 2, obs 2,0 Compared to paramagnetic contrast agents, SPIO has much higher efficiencies for relaxation enhancements for both T 2 and T 1. It has been demonstrated that the r 2 (100 mm -1 sec -1 ) and r 1 (100 mm -1 sec -1 ) of SPIO were markedly higher than the r 2 (6 mm -1 sec - 1 ) and r 1 (2 mm -1 sec -1 ) of Gd-DTPA at 37 o C and 0.47 Tesla. 59 Concurrently, the SPIO contrast agents are the most sensitive MRI contrast agent so far (on the order of 10-6 M, 60 compared to Gd(III) agent on the order of 10-5 M 41 ) SPIO even can be used for single cell detection with a sufficient contrast to noise ratio (CNR). 61 Additionally, SPIO contrast agents are biocompatible since they are composed of biodegradable iron, which can be recycled by cells and thus rarely cause significant toxicity. Nevertheless, SPIO particles are visible by using light and electron microscopy, and therefore are easy to validate in vivo. SPIO agents have been extensively used for studies of abnormalities in liver, spleen and lymph nodes, imaging macrophages and detecting tumor angiogenesis. 57 Non-specific SPIO agents also have been used for cell labeling and trafficking. 62 Recently, specific targeting/responsive SPIO agents also have developed for imaging gene expression, targeting cell surface receptors, 20 detecting apoptosis, 26,27 sensing ions 32 and detecting 66, 67 enzyme activity

31 In spite of the ultra high sensitivity, the T 2 contrast still suffers from the poor quantification due to extremely high local concentration. Approaches that create positive contrast by using SPIO agents have been developed in the last few years. 68,69 With such so called off-resonance imaging (ORI) techniques, the inherent low SNR limitation due to negative contrast can be overcome and quantification can be greatly improved CEST CONTRAST AGENTS FOR MOLECULAR IMAGING The first proof-of-principle study of Chemical Exchange Saturation Transfer (CEST) was demonstrated by Balaban and his colleagues in This novel contrast mechanism soon become an attractive alternative to T 1 and T 2 contrast mechanisms, particularly at high magnetic field strengths. 71 CEST MRI agents possess labile protons that exchange with water molecules at moderate to slow exchange rates. Continuously applying RF saturation at the MR frequency of labile protons and followed by exchange of the protons with solvent water can significantly attenuate the MR signal of the water owing to the saturation transfer process (Figure 1.4). Because of the high molar concentration of water protons (~110 M), the CEST signal, even only at percentages of water signal, still can be sufficiently detectable by using milimolar and even sub-milimolar concentrations of contrast agents. In other words, by such an amplification strategy, significant contrast enhancement can be achieved by using a low concentration CEST contrast agent that allows the least perturbation to the biological system Theory The dynamic process of CEST can be described by a two-site exchange model as illustrated by Figure 1.5 for most cases or can be simplified to a two-site exchange - 29-

32 model in more complicated situations. 74 In brief, a two-site exchange PARACEST system consists of a bulk free water proton pool (W) and a small labile proton pool on a contrast agent (CA), which is water-exchangable and has a signal can be turned on by applying RF saturation pulses. O O O O H 2 O H-OH O H 2 O H 2 O H 2 O O O O H H N N N N N N R N N R Eu 3+ O Eu 3+ O O Yb O Eu Yb 3+ O Yb 3+ N N N N N N O O O O O O O O HOH H N R Figure 1.4. A schematic of CEST contrast enhancement in MRI. The water signal is attenuated due to the chemical exchange of saturated protons (in red) from contrast agents to water molecules. Figure 1.5. The two pool model for CEST contrast. The small contrast agent pool possesses labile protons that can exchange with the bulk water pool with the forward exchange rate of k CA-W and backward exchange rate of k W-C

33 A series of Bloch-McConnell equations has been commonly used to quantitatively describe the CEST process (Eqs. [1.6]). 75,76 In such a framework, the RF pulse is assumed to apply at an arbitrary offset of ω along the X direction in the rotating frame with irradiation strength of B 1 (Hz). dm dt dm dt dm dt dm dt dm dt dm dt W Z W X W Y CA Z CA X CA Y = M / T - M / T + k M k M + 2π BM W W CA W W 0 1Wsat Z 1Wsat CA W Z W CA Z 1 Y = Δ ω M - M / T + k M k M W W CA W W Y X 2W CA W X W CA X =Δ ω M + 2πBM - M / T + k M k M W W W CA W W X 1 Z Y 2W CA W Y W CA Y = M / T - M / T k M + k M + 2π BM CA CA CA W CA 0 1CA Z 1CA CA W Z W CA Z 1 Y = Δω M - M / T k M + k M CA CA CA W CA Y X 2CA CA W X W CA X =Δ ω M + 2πBM - M / T k M + k M CA CA CA CA W CA X 1 Z Y CA CA W Y W CA Y [1.6] M : Z, Y and X components of the magnetizations of pool CA and pool W WCA, Z, Y, X respectively; M : equilibrium magnetization of pool W and pool CA respectively; WCA, 0 T 1Wsat : longitudinal relaxation time of pool W in the presence of saturation; T 1CA : longitudinal relaxation time of pool CA; T 2W : transverse relaxation time of pool W; T 2CA : transverse relaxation time of pool CA; kca W : the exchange of protons from pool CA to pool W; kw CA: the exchange of protons from pool W to pool CA; ω W : the difference in frequency between on resonance frequency of water (ω W ) and applied irradiation offset (ω); - 31-

34 ω CA : the difference in frequency between chemical shift of labile protons of pool CA (ω CA ) and applied irradiation offset (ω); T 1 of the pool W (T 1W ) can be altered in the presence of saturation RF pulses, thereby T 1Wsat should be used for a restricted description. 77 However, the effect of saturation pulses on T 1W is typically negligible and thereby the T 1W is used instead as a good approximation of T 1Wsat. Only longitudinal magnetization of pool W contributes to the MRI signal, hence only the first equation in the Bloch-McConnell equations is most relevant. In practice, the Bloch- McConnell equations can be simplified by assuming that there is a complete saturation of W pool CA ( M = 0) and no direct saturation of pool W ( M = 0) (Eq. [1.7]). CA Z Y dm dt W = M / T - M (1/ T + k ) [1.7] W W 0 1W 1W W CA Because these assumptions cause an omission of M Y in Eq. [1.7], the subscript Z was removed from this equation and for all following equations for simplification. At thermal equilibrium, the exchange rates between the two pools have a mass balance (Eqs. [1.8]). M / M = k / k [1.8 a] W CA 0 0 CA W W CA or [ proton] CA nca [ ] kw CA= kca W = k [1.8 b] CA W [ proton] 2[ H O] W 2 In Eqs. [1.8], n is the number of exchangeable protons on the PARACEST agent and [H 2 O] and [CA] are the concentrations of water and the PARACEST agent, respectively. The Eq. [1.7] can be solved using numerical 78,74 or analytical 79 methods. A steady state analytical solution (Eq. [1.9]) is most often used to provide contrast quantification while a transient analytical solution can provide more useful information for the exchange rate - 32-

35 estimation 72,78 as well as for method optimization (Eq. [1.10]). Where τ CA is the life time of protons in pool CA (τ CA = 1/ kca W ) M M S τw = T + τ 0 1W W [1.9a] or M S nt [ CA] M 2 τ [ H2 O] 0 = 1W 1/(1 + ) [1.9b] CA M M τ T 1 1 (- t / τ1w ) = + [1.10] S W W 0 1W τ τ W e where = + τ T τ 1w 1w w [1.11] In Eq. [1.9] and Eq. [1.10], M S and M 0 represent the coherent magnetization of the pool W with or without saturation, respectively. Also, τ1w lifetime of pool W (Eq. [1.11]). represents the effective residence A quantitative CEST study is typically conducted by an asymmetric analysis approach, i.e. two image acquisitions with irradiation at the CEST offset ( Δ ω ) and the opposite offset(- ω) related to water offset (set ω W =0) are collected and computed to determine the %CEST ratio (Eq. [1.12]), % CEST( Δ ω ) = 1- M S ( Δ ω )/M S ( Δ ω ) [1.12] where M S ( Δ ω ) and M S ( Δ ω ) are the MR signals with the RF saturation pulses applied at Δ ω and Δ ω respectively Advantage of CEST Contrast Enhancement The CEST MRI contrast has been demonstrated with a number of potential advantages compared to T 1 and T 2 contrast agents. Foremost, the CEST signal can be turned on or off by modulating the saturation pulses. Therefore, endogenous MR contrast may be - 33-

36 continually monitored in the presence of PARACEST contrast agents by neglecting to saturate the MR frequency of the exchangeable proton (and assuming that the T 1 relaxation of the PARACEST agent is negligible). Additionally, the CEST signal is only turned on by applying a RF saturation pulse at the resonance frequency (CEST offset). By developing contrast agents with distinct CEST offsets, it is possible to realize the simultaneous detection of multiple agents during a single scan session by using an interleaved selective detection scheme. 80 For example, the CEST offsets of lanthanide based Paramagnetic CEST agents can be easily modulated by chelating different Ln(III). 72 Nevertheless, CEST sites, such as -NH or OH, are often involved in many biological reactions. This endows the CEST contrast agents with the ability to target important biomarkers such as proteases. 81 Moreover, not only can the specificity be greatly improved with the design of enzyme responsive CEST agent, the sensitivity can also be dramatically increased as the result of the accumulation of CEST signals upon the 42, 81 continuous catalysis of proteases Classification and Applications of CEST Contrast Agent According to the excellent review article by Zhou, 71 CEST contrast agents can be classified into three subsets, i.e. diamagnetic CEST (DIACEST), paramagnetic CEST (PARACEST) and amide proton transfer (APT). DIACEST was first reported by using small molecule ammonium chloride. 70 A more systematic study of DIACEST then was carried out and reported in based on numerous exchangeable protons (-OH, -NH and NH 2 ) carried by sugar, amino acids and other endogenous small molecules. Recently, macromolecular DIACEST agents have been reported with greatly improved sensitivity based on poly(l-lysine) 83 or poly(ru). 84 DIACEST agents have been explored - 34-

37 for ph measurements, 85 reporting in vivo gene expressions, 16 imaging the in vivo distribution of glycogen 86 and studying cartilage degradation. 87 APT imaging is based on the CEST properties of amide protons in endogenous proteins and peptides. Therefore, APT imaging has potential advantages over other exogenous CEST contrast agents owing to the minimal perturbation to biological systems. APT imaging has been successfully employed to in vivo ph measurements, 88 brain tumor imaging 89,90 and ischemia detection PARACEST CONTRAST AGENTS FOR MOLECULAR IMAGING The PARACEST (PARAmagnetic CEST) contrast is a subset of the CEST contrast mechanism. The sensitivity of CEST agents can be improved by increasing the exchange rate as evident in Eq. [1.9b]. However, the exchange rate of labile protons cannot be further improved when the limit of slow to moderate exchange rate (Δω k ex ) is reached. One practical way to overcome this limitation is to improve Δω by using a paramagnetic NMR shift agent. The paramagnetic lanthanide ion can shift the MR frequencies of the exchangeable proton to unique values and hence greatly facilitate selective detection. 71,72 Lanthanide based PARACEST contrast agents have been recently developed for this approach Lanthanide Induced Shifts Lanthanide ions have a long history as shift agents in NMR studies since 1960s 92 and are still widely used as shift reagents in chiral NMR studies. 93 Ln(III) ions have electronic configurations of [Xe] 4 f n (n=0-14). Except for diamagnetic La(III) and Lu(III), all other Ln(III) ions are paramagnetic due to unpaired 4f electrons. Ln(III) ions therefore can be - 35-

38 added to diamagnetic samples and induce huge chemical shift changes via the interactions between the high density 4f electrons and nearby nuclei. The apparent chemical shift of a nucleus of a ligand that is coordinated Ln(III), δ app, can be expressed as the sum of the diamagnetic chemical shift (δ dia ) and the paramagnetic chemical shift ( para ) (Eq. [1.13]). The latter contribution can be further separated into through-bond effect ( c, contact or scalar), through-space effect ( p, pseudo-contact or dipolar), and bulk magnetic susceptibility effect ( χ ) (Eq. [1.14]). δ = δ +Δ [1.13] apprent dia para Δ para =Δ c +Δ d +Δ χ [1.14] As far as the CEST effect ( ω= ω CA - ω W ) is concerned, the χ contribution can be eliminated because BMS effects are equivalent for all species in the same compartment, including the protons on water and CEST sites of contrast agents. The mechanism of contact shift ( c ) works by through-bond transmission of unpaired spin density of Ln(III) to the ligand nuclei. The contact shift is the dominant mechanism for those nuclei that are directly bonding to the Ln(III). The pseudo-contact shift ( p ) is caused by the throughspace dipolar interaction between the magnetic moments of unpaired electrons in Ln(III) and the nuclei under study. (Eq. [1.15]). 46 p = C (3cos 2 θ-1) /r 3 [ 1.15] where C is a constant, r and θ are the spherical coordinates of the observed nucleus with respect to Ln(III) at the origin and the principal magnetic axis of the system as the z- axis. 94 As is evident in Eq. [1.15], the pseudo-contact shift will be maximized if a nucleus has an orientation along the symmetry axis and/or a high proximity to Ln(III). However, the - 36-

39 short distance between the Ln(III) and the nucleus also results substantial line broadening. Figure 1.6 demonstrates the relationship between the pseudo-contact shifts (PCS) and line widths with the metal-to-proton distance. As evident in this demonstration, the useful distance for Yb(III) as a shift reagent in an NMR study is between 9-25 Å. 95 However, with a tradeoff of line broadening, a high shift ( >10 ppm) can be achieved within 2-3 bond lengths (<4 Å). Figure 1.6. Useful ranges of pseudo-contact shifts (PCS) for Ce(III), Yb(III), and Dy(III)-containing proteins. The dashed lines indicate the observed line widths Lanthanide Based PARACEST Contrast Agents So far most reported PARACEST MRI contrast agents are based on the chelator macrocyclic DOTA and its derivatives. The prototype PARACEST agent that has been well studied is Ln(III)-DOTAMGly (Fig. 1.7). 72, 73, 96,

40 HOOC HOOC HN HN O N N O OH 2 N Ln 3+ NH COOH O N O COOH NH Figure 1.7. Chemical structure of Ln(III)-DOTAMGly Two types of labile protons, bound water protons and amide protons, have been characterized to provide significant PARACEST signals (Table 1.3). Upon the coordination with Ln(III), amide protons have demonstrated very high paramagnetic shifts up to 77 ppm (Dy(III)) while much higher shifts have been observed for water protons. 96 Owing to such huge shifts, the limitation of slow to moderate exchange rate (Δω >> k CA ) still can be satisfied even with very high exchange rates (e.g. k CA >10 +5 Hz). As a result, the PARACEST effects of amide protons with all Ln(III) and bound water protons with four Ln(III) (Eu, Pr, Nd and Tb) are observable even though the exchange rates are quite fast (on the order of Hz). The paramagnetic shifts of bound water resulting from contact shifting by Ln(III) are typically greater than the shifts of amide protons, where the pseudo-contact shift mechanism takes place. As a result, PARACEST effects from bound water protons and amide protons have demonstrated distinct temperature and ph dependencies

41 Table 1.3. Magnetic moments, chemical shifts and exchange rates of Ln(III)- DOTAMGly 96 Amide protons Ln-water protons Ln μ eff ω (ppm) k CA (Hz) ω (ppm) k CA (Hz) Pr x10 5 Nd x10 4 Eu x10 4 Tb x10 5 Dy At coalescence Ho At coalescence Er At coalescence Tm At coalescence Yb At coalescence * Values determined at 7.05T, 312K and ph 7.4 The nature of Ln(III) can dramatically affect the efficiency of PARACEST systems and thus has to be taken into account for the design of new PARACEST contrast agents. In the same study as above, 96 Eu(III) was demonstrated as the most efficient system to provide a bound water proton PARACEST effect while Yb(III) is the most efficient system based on amide protons. The high efficiency of Eu(III) is due to the best trade-off between Δω and k CA. 96 Because the exchange rate of amide protons is mainly determined by the nature of ligand rather than Ln(III), the different apparent efficiencies are considered due to the different T 1 relaxivities of lanthanides. As far as inner-sphere - 39-

42 relaxation enhancement mechanism is concerned, the relaxivity is approximately proportional to the square of μ eff. 46 A short T 1 relaxation time due to high relaxivity will sacrifice the PARACEST effect as evident in Eq. [1.9b]. Indeed, Yb(III) has the smallest μ eff among heavier lanthanide ions and has been demonstrated to be the most efficient lanthanide for amide proton PARACEST effect. In addition to achieving higher sensitivity by permitting a higher exchange rate, Ln(III) coordination also avoids the inferences from conventional magnetization transfer effects and direct water saturation owing to a CEST offset at a much higher frequency range. Nevertheless, at a higher exchange rate, the optimal saturation RF power (B 1,opt ) can be even higher, which typically leads to a higher contrast. 74 B 1,opt =k CA /2π [1.16] Applications of PARACEST Contrast Agents Owing to the highly shifted CEST offsets, PARACEST contrast agents can be designed with good detection sensitivities. Since the first demonstration of PARACEST contrast agent by Zhang and his co-workers in 2001, 98 PARACEST contrast agents have been explored for a number of biomedical applications for different targets, such as enzymes, 42 metabolites, 99, 100 metal ions, 101 ph 34,72 and temperature. 102 However, the in vivo application of PARACEST contrast agents is plagued by the relative low sensitivity and poor temporal resolution for detection. A variety of strategies have been explored to 100, 103 improve the sensitivity by using payload systems, such as liposomes, polymers and dendrimers

43 1.7. CONCLUSION The development of MRI contrast agents can greatly facilitate the MRI molecular imaging for cancer biomarker detection and can ultimately be used for clinical diagnosis and theragnostics. Among the families of MRI contrast agents, non-specific T 1 and T 2 agents have been studied for decades and many of them have been approved by the FDA. Targeted and responsive contrast agents, however, have been only developed very recently for MRI molecular imaging. By enormous recent endeavors, all the three contrast mechanisms, including T 1, T 2 and CEST/PARACEST, have been used to develop the responsive/targeting molecular imaging probes and explored for appropriate biomedical applications. Each contrast mechanism has specific advantages and potential pitfalls, therefore may be favorable for some particular applications but not for others. General speaking, a molecular imaging technique can be evaluated by a number of metrics such as temporal resolution, spatial resolution, sensitivity, specificity and applicability. Additionally, the pharmacokinetics has to be considered for preclinical or clinical applications. A comprehensive comparison of T 1, T 2, CEST and PARACEST contrast agents is summarized in Table

44 Table Comparison of T 1, T 2, CEST and PARACEST contrast agents Contrast mechanism T 1 T 2 /T 2 * CEST PARACEST Prototype Gd-DTPA Gd-DOTA SPIO Ploy-lysine Ploy-(rU) Ln- DOTAMGly Major Modulation mechanism Temporal resolution (1/TR) for responsive agent 105 Tumbling time Hydration number Fastest (~100ms) size Exchange rate Exchange rate Paramagnetic shifts Slow (1-100s) Slow (1-10s) Slow (1-10s) Spatial resolution Highest Low High High Sensitivity >100 μm >10 fm μm per polymer 84 >500 μm 72 ~2.8 nm per liposome 103 Specificity High Low Low High Endogenous/exogen ous Exogenous Exogenous Endogenous Exogenous Exogenous Multiple agents detection No No Yes Yes Limitation in applied tissue type No ShortT 2 Short T 1, low water content Short T 1, low water content Applications in molecular imaging Gene expression Enzyme detection Receptor targeting Cell labeling Apoptosis Cell labeling Gene expression Receptor targeting Enzyme detection Apoptosis Gene expression ph gene therapy delivery ph enzyme cell labeling metabolite ion receptor targeting Ion Type of agents Targeting Targeting Responsive Targeting Responsive Responsive Responsive - 42-

45 All the characterizations that have been summarized here are according to many published reports. Considering the ever-growing developments in all aspects for each contrast mechanism, some of the limitations of these agents may be overcome in the near future. For instance, with the use of a nanoparticle carrier, the sensitivity limitation may be overcome to reach the μm - nm range for detecting a PARACEST contrast agent. However, there are still some inherent limitations that cannot be overcome at least in the near future. For example, multiple T 1 contrast agents may still not be simultaneously detected. The applicability in T 1 tissues with short T 1 relaxation times is still poor for PARACEST and CEST contrast agents. Among these contrast mechanisms, PARACEST is very promising for molecular imaging regardless of inherent low temporal resolution and low sensitivity. One of advantages of PARACEST is the high specificity. The PARACEST offsets are often beyond the chemical shift range for endogenous components and ensure the least interferences compared to other contrast mechanisms. The signal can be turned on and off with the modulation of the saturation pulse therefore allowing the detection of inherent contrasts and even the T 2 contrast agent (if there is no significant shortening of T 1 relaxation time) in the presence of PARACEST contrast agents. Furthermore, the quantification of PARACEST agents can be easy and precise with the use of inter-molecular or intramolecular multiple PARACEST sites owing to the ability of multiple agent detection. Due to these merits, the exploration of PARACEST for cancer biomarker detection is very promising. The future developments for PARACEST contrast will focus on the improvement of SPECIFICTY, TEMPORAL RESOLUTION, and SENSITIVITY

46 In this thesis study, a new PARACEST agent has been developed for the detection of nitric oxide as an example of improvement of SPECIFICTY of PARACEST agent, which will discussed in Chapter 2. A novel ph measurement method then has been developed based on the same agent to demonstrate the improvement of SPECIFICTY and quantification, which will be introduced in Chapter 3. The challenges in TEMPORAL RESOLUTION have been also studied and new MRI methods have been developed to fit the requirements of in vivo application s, which will be described in Chapter 4. Issues of SENSITIVITY will be addressed in Chapter 5 as part of future studies

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60 P.; Meijer, E. W.; Grull, H., Dendritic PARACEST contrast agents for magnetic resonance imaging. Contrast Media Mol Imaging Shapiro, M. G.; Atanasijevic, T.; Faas, H.; Westmeyer, G. G.; Jasanoff, A., Dynamic imaging with MRI contrast agents: quantitative considerations. Magnetic Resonance Imaging 2006, 24, (4),

61 CHAPTER 2 DESIGN AND CHARACTERIZATION OF A NEW IRREVERSIBLE RESPONSIVE PARACEST MRI CONTRAST AGENT THAT DETECTS NITRIC OXIDE 2.1. INTRODUCTION Responsive MRI contrast agents, which change their efficacies for generating image contrast in response to changes in their chemical or biochemical environments, have shown promise as a powerful strategy for biomedical diagnostics. 1,2 Responsive relaxivity-based MRI contrast agents can undergo reversible conformational changes that modulate water accessibilities to the agents metal ions, 3,4 or that undergo reversible intermolecular associations that change the agents rotational correlation times. 5,6 However, these reversible responsive MRI contrast agents require relatively high concentrations of molecular targets to produce a detectable response. Rapid conformational changes or intermolecular dissociations that are faster than the relatively slow MRI acquisition timeframe can produce an averaging effect that further limits the sensitivity of the agent s response. Responsive relaxivity-based MRI contrast agents can undergo irreversible covalent changes through enzymatic cleavage of a ligand, 7,8 or through enzyme-catalyzed polymerization or depolymerization Non-covalent changes may also be considered irreversible if the non-covalent change is absolutely stable during the timeframe of the study. Rapid catalysis can generate a detectable response from a relatively low - 59-

62 concentration of enzymes, and the irreversible nature of covalent changes can reduce averaging effects during the MRI acquisition timeframe. Therefore, irreversible responsive MRI contrast agents have significant advantages for molecular imaging studies. Chemical Exchange Saturation Transfer (CEST) MRI has been shown to be a promising alternative to relaxivity-based contrast mechanisms. 12 CEST MRI contrast agents possess one or more hydrogen protons that exchange with water. Saturation of the MR frequency of one or more of these protons, followed by chemical exchange, can reduce the MR signal of the water. PARACEST (PARAmagnetic CEST) agents include a paramagnetic lanthanide ion that shifts the MR frequency of the exchangeable proton to unique values to facilitate selective detection. 13,14 This selective saturation method provides a tremendous advantage relative to relaxivity-based MRI contrast agents, because multiple agents with unique MR frequencies may be selectively detected as needed, or standard MR images that are not influenced by the presence of the contrast agent may be acquired by neglecting selective saturation. 15 Reversible responsive PARACEST MRI contrast agents have been developed that change their chemical exchange rates after binding to glucose, 16 polyarginine, 17 or zinc, 18 or that changes the chemical shift of the amide protons after binding to lactate. 19 Because amide proton chemical exchange is catalyzed by hydroxide ions, modulations in chemical exchange rates have also been exploited to detect ph using the PARACEST effect. 15,20 These reversible responsive PARACEST MRI contrast agents share similar advantages and disadvantages with reversible responsive relaxivity-based MRI contrast agents

63 Very recently, we have reported an irreversible responsive PARACEST MRI contrast agent that converts an amide to an amine during enzyme catalysis. 21 This irreversible change was monitored with MRI by observing the disappearance of the PARACEST effect from the amide and the appearance of the PARACEST effect from the amine. To further investigate the properties of this new class of contrast media, we have developed a new PARACEST MRI contrast agent, Yb-1,4,7,10-tetraazacyclododecane-(N,N,N - triacetic acid)-(n -orthoaminoanilide) (Yb-DO3A-oAA; Yb(1) in Figure 2.1). Figure 2.1. The illustration of the reaction of Yb(1) with NO in the presence of oxygen converts aromatic amines to a triazene on Yb(2). This contrast agent undergoes an irreversible covalent change through a chemical reaction with an N-nitroso intermediate produced by the endogenous autooxidation of Nitric Oxide (NO). Aromatic amines are known to react in the presence of oxygen to produce triazenes, which causes a loss of the exchangeable protons. 22 The loss of amine protons and the loss of the proximity of the amide proton to the lanthanide ion in Yb(2) cause a loss or deactivation of the PARACEST effect exhibited by Yb(1). This chemical reaction mechanism is similar to the mechanism used by DAF-FM, which changes fluorescence properties in the presence of NO and O The following report - 61-

64 describes the PARACEST MR properties of this contrast agent, the potential utility and remaining challenges of this contrast agent for biomedical imaging applications, and the advantages and disadvantages of irreversible responsive PARACEST MRI contrast agents for molecular imaging MATERIALS AND METHODS Sample Preparation 1,4,7,10-tetraazacyclododecane-1,4,7,10-tetraacetic acid o-aminoanilide (DO3A-oAA) was purchased from Macrocyclics (Dallas, TX). The Yb(III) DO3AA-oAA complex Yb(1) was prepared by mixing a 1:1.01 molar ratio of DO3A-oAA and YbCl 3 in ph 6 aqueous solution at 60 o C for 2 hours. The ph was increased to 8 and then allowed to react for an additional 30 minutes. After precipitating the excess free Yb 3+ ions by adjusting the ph to 12, the solution was centrifuged, filtered and lyophilized to collect a pure complex. The structure of Yb(1) was confirmed by MALDI-TOF mass spectroscopy and NMR spectroscopy. The same procedure was used to chelate Gd(III)to form Gd(1). NONOate (Cayman Chemical Company, Ann Arbor, MI) was used to produce NO. 24 The reaction of the PARACEST MRI contrast agent with NO was conducted at 37 C by combining 1 ml of 40 mm Yb(1) in ph 7.2 PBS buffer with 40 mg NONOate for 1 hour. This reaction time ensured that all NONOate had reacted to form NO, and that no significant concentrations of NO gas remained in the solution. The same procedure was used to convert Gd(1) to Gd(2) PARACEST and T 1 NMR Measurements The 1 H NMR spectrum of 1 in D 2 O was acquired at room temperature with a 600 MHz Varian Inova NMR spectrometer (Varian, Inc., Palo Alto, CA). The 1 H NMR spectrum - 62-

65 of Yb(1) in H 2 O was obtained by a modified water suppression WET-1D method 25 at room temperature using the same NMR spectrometer. PARACEST spectra were acquired with a modified presaturation pulse sequence that included a continuous-wave selective saturation pulse applied for seconds at μt saturation RF power, at frequencies spanning from 35 ppm to -35 ppm in 1 ppm increments. Unless otherwise noted, all PARACEST spectra were acquired at ph 7.0 and 37 C. NaOH and HCl were used to adjust ph in all studies. When the effect of ph on PARACEST was studied, 40 mm samples of Yb(1) were adjusted to ph values between 3 and 10 with increments of 1 ph unit. When the effect of temperature on PARACEST was studied, 40 mm samples of Yb(1) were adjusted to 8 C, 15 C, 25 C and 37 C. To study the sensitivity of the PARACEST effect, the concentrations of samples of Yb(1) with 20 mm and 80 mm were validated with ICP analysis of Yb(III) content (Chemical Solutions Ltd., Mechanicsburg PA). To account for direct saturation of water, each spectrum was fit to a single lorentzian lineshape using Origin v7.5 (OriginLab Corp. Northampton, MA), excluding a 5ppm-wide region centered at the PARACEST effects of the amide and amine. The difference between the lorentzian fit and the experimental data was used to calculate the magnitude of each PARACEST effect. The NMR data were processed and plotted using VNMR (Varian, Inc., Palo Alto, CA) and Microsoft Excel (Microsoft Co., Seattle. WA) MRI Measurements A Bruker Biopsin 9.4T small animal MRI scanner (Bruker Biopsin Co. Billerica, MA) was used to acquire MR images of contrast agents in phosphate buffered saline (PBS). A MSME pulse sequence with TR/TE = 10,005/10.9 ms was used with selective saturation - 63-

66 applied at +11 ppm, +8 ppm, -8 ppm and -11 ppm through a 4.5 sec train of Gaussianshaped saturation pulses applied at a saturation power of 30 μt and a bandwidth of 0.5 ppm. Only one image slice (0.5 mm thickness) was acquired for each scan with a FOV of 30mm x 30mm. T 1 relaxation time constants were measured using a MSME pulse sequence without selective saturation and with TR spanning 20 to 9000 msec Molecular Modeling A model of 1 was constructed with CS MOPAC Pro v8.0 within Chem3D Ultra v8.0 (CambridgeSoft Corp., Cambridge MA). A metal atom was inserted in the binding pocket of the model, distances between the chelating amines, carboxylates, and carbonyl were constrained to 2.51 angstroms, and the resulting model of Yb(1) was energyminimized using the MM2 force field RESULTS To form the PARACEST MRI contrast agent Yb(1), Yb(III) was chelated with a derivative of DOTA (1,4,7,10-tetraazacyclododecane-1,4,7,10-tetraacetic acid) that included an ortho-aminoanilide motif (Figure 2.1). The structure of 1 was confirmed by NMR spectroscopy and mass spectroscopy (MALDI-TOF m/z =666.11). The NMR spectra in H 2 O and D 2 O also indicated that the chemical shifts of the aromatic amide and amine were shifted to -11 and +8 ppm relative to the MR frequency of water. The CEST spectrum of Yb(1) showed two PARACEST peaks at +11 ppm and -8 ppm, which matched the chemical shifts of exchangeable protons obtained from 1 H NMR spectra (Figure 2.2). CEST spectra were acquired with different saturation powers and times to optimize the PARACEST effects, which demonstrated that 4.2 μt and 4 sec - 64-

67 saturation time produced the greatest combined CEST effects from the amide and amine (Figure 2.3). Figure 2.2. The CEST spectrum of Yb(1). A continuous wave saturation was applied to a sample containing 40 mm Yb(1) in PBS at 4.6 μt for 4 seconds in 1 ppm increments. The height of each water signal was normalized relative to the maximum water signal acquired during this experiment. By using the experimental determined optimal saturation power and time, the similar acquisition & analysis procedure was used to measure the magnitude of the CEST effect over a range of ph values (Figure 2.4a) and temperatures (Figure 2.4b). To determine the sensitivity of detecting Yb(1), the PARACEST effect of the agent was correlated with concentration using modified Bloch equations for two proton pools that 14, 26 undergo exchange (Eq.[2.1])

68 a b Figure 2.3. The effects of saturation power and saturation delay on the PARACEST effect of Yb(1). Each curve is labeled with the saturation power. The curves were fit to the data according to: % CEST = 1-τ/T 1sat - τ/τ water e (-t/τ), where τ = τ water T 1sat / (τ water + T 1sat ), τ water = water residence lifetime and T 1sat = T 1 relaxation time constant during saturation. a b Figure 2.4. The effects of a) temperature and b) ph on the PARACEST effect of Yb(1). A continuous wave saturation was applied at 4.6 μt for 4 seconds at +8 or -11 ppm. Each PARACEST spectrum was fit to a single lorentzian lineshape, and the difference between the lorentzian fit and the experimental data was used to calculate the magnitude of each PARACEST effect

69 Ms 1 = M0 n [CA]T CA 1+ n [H2O]τ H2O 1sat CA [2.1] M s : MR signal of water during selective saturation of the contrast agent M 0 : MR signal of water without selective saturation n CA : number of exchangeable protons of the contrast agent n H2O : number of exchangeable protons of water [CA]: concentration of contrast agent [H 2 O]: concentration of water (~55 M) T 1sat : T 1 relaxation time constant of water during selective saturation of the contrast agent τ CA : average lifetime of the protons on the contrast agent a b Figure 2.5. The effect of concentration on the PARACEST effect of Yb(1) a) calibration of concentration and CEST effect according to Eq. [2.1], b) calibration of CEST effect vs concentration based on Eq. [2.2] - 67-

70 1/T 1sat of Yb(1) was found to be linearly related to contrast agent concentration by using a T 1 inversion recovery method with selective saturation at the amide or amine chemical shifts. By substituting T 1sat with a linear relationship based on [CA], the modified Bloch equations can be further simplified (Eq. [2. 2]), where m and b represent the slope and intercept of the linear relationship between 1/T 1 and [CA]. This equation was exploited to determine the sensitivity of detecting contrast agent Yb(1) (Figure 2.5). 1 1 nca m = [ ]- [CA] M0 ( -1) b n H2O [H2O] τm b M S [2. 2] To determine whether NO can modulate the PARACEST effect of Yb(1), 40 mm of this agent was combined with an excess amount of NONOate at ph 7.0 and 37 C for 1 hour. Complete conversion of the aromatic amine to a triazene was confirmed by mass spectrometry (MALDI m/z 1339). MR images of reactant Yb(1), reaction product Yb(2), Tm-(DOTAM-Gly), and a phantom of PBS were acquired with selective saturation at +8 or -11 ppm. PARACEST parametric maps were generated by subtracting the image of a contrast agent from the image of PBS acquired with the same saturation frequency (Figure 2. 6). A PARACEST effect is evident from Yb(1), and a control agent Tm- (DOTAM-Gly), but a PARACEST effect is not present throughout the PARACEST map of Yb(2), as only susceptibility artifacts are present in this map. To determine whether NO can also modulate the T 1 relaxivity of Gd(1), T 1 relaxation time constants of Gd(1) and Gd(2) were measured using a T 1 inversion recovery method without selective saturation (Figure 2.7). These results showed a change in relaxivity from 2.9 mm -1 sec -1 to 4.2 mm -1 sec -1, which amounts to a 44.8% change in relaxivity caused by the chemical reaction

71 Figure 2.6. a) The MR images of 30 mm Yb(1) (before reaction) and 30 mm Yb(2) (after reaction) in phosphate buffered saline (PBS) with selective saturation at -11ppm. The PARACEST map is generated by subtracting these images from an image of PBS without the contrast agent. Each PARACEST map was independently scaled to demonstrate that only susceptibility artifacts are present in the PARACEST maps of Yb(2). b) The PARACEST map of 30 mm Yb(1) (before reaction) and 30 mm Yb(2) (after reaction) with selective saturation at +8 ppm. c) The PARACEST map of 10 mm Tm-(DOTAM-Gly) before and after applying the same reaction conditions with selective saturation at -51 ppm. Molecular modeling of Yb(1) was conducted to estimate the proximity of the amine, amide, and water to the lanthanide ion (Figure 2.8). Based on this model, the exchangeable amide proton and the amine protons are 5.6 Å and 4.6 Å from the lanthanide ion, respectively, indicating that each of these protons is sufficiently close to the ion to experience a hyperfine contact shift. Furthermore, an amine proton lays 2.4 Å - 69-

72 Figure 2.7. a) The T 1 relaxivities of Gd(1) and Gd(2). The lines represent a leastsquares fit to the data, with r2 > 0.99 for each fit. Error bars represent the standard deviation of T 1 values determined from parametric maps of each sample. b) T 1 weighted images of 2.5 mm Gd(1) and Gd(2). TR/TE= 6000 ms/ 14.1 ms. a b Figure 2.8. The molecular model of Yb(1). The two orientations of the same model differ by a 70 o rotation

73 from the carboxylate oxygen of an adjacent acetate ligand, and the other amine proton lays 2.6 Å from the carbonyl oxygen of the oaa ligand, which indicates that the amine may form hydrogen bonds with these oxygens. The oaa ligand lies on the side of the basket-shaped macrocyclic ring (Figure 2. 8b), and water accessibility to the lanthanide ion is not physically inhibited by the presence of this ligand. An alternative model was constructed by rotating the oaa ring about the amide nitrogen-ring carbon bond, which is the only major degree of freedom in this model. Although the interatomic energies of this model were not as favorable as the initial model due to relative orientations of the acetate ligands, the amine hydrogens were still proximal to the carboxylate oxygen of an adjacent acetate ligand and the carbonyl of the oaa ligand, and the ring was still oriented on the side of the macrocyclic ring. Although other conformational adjustments are possible, this model adequately represents the gross morphological features of Yb(1) DISCUSSION The PARACEST effect of Yb(1) at -11 ppm was assigned to the amide, based on similar PARACEST results reported for amides within Yb-chelates. 15 To confirm this assignment, the two PARACEST effects were measured over a range of ph values (Figure 2.4a). The amide showed increasingly greater PARACEST with increasing ph, reaching the greatest effect at ph 7. The base-catalyzed proton chemical exchange rate between an amide and water is relatively slow on the MR time scale (characterized by the chemical shift difference between the amide and water), and increasing hydroxide ion concentration accelerates this rate to improve the PARACEST effect. 27 The PARACEST effect at +8 ppm showed increasingly greater PARACEST with decreasing ph, which is - 71-

74 consistent with acid-catalyzed proton chemical exchange between an amine and water. 28 The ph dependencies also proved that these PARACEST signals originated from amines and amides instead of bound water molecules, which typically exhibit ph-invariant PARACEST effects. The PARACEST effect from the amine group of Yb(1) is unusually strong and selective (i.e., spectroscopically narrow in Figure 2.2). The observation of a PARACEST effect requires MR frequency differences between water and the PARACEST functional group that are faster than chemical exchange rates between water and the functional group. 14 Typical amines exchange protons with water at approximately 3000 sec The chemical shift of an amine must be at least 5 ppm from water at 14.1T magnetic field strength (3000 Hz at 600 MHz proton MR frequency), or 7.5 ppm at 9.4T (3000 Hz at 400 MHz proton MR frequency). Therefore, the unusually large chemical shift difference of the amine of Yb(1) is essential for observing this PARACEST effect. The molecular model supports the observation of an unusually large hyperfine contact shift of the amine, because the amine may be positioned near the lanthanide ion. In addition, the molecular model indicates that the amine protons may participate in intramolecular hydrogen bonding that slow the exchange rate to generate a stronger and more selective PARACEST effect. A strong and selective PARACEST effect from an amine expands the paradigm of irreversible responsive PARACEST MRI contrast agents, because covalent reactions that consume or produce amines are prevalent in many biological processes. The sensitivity threshold for detecting a PARACEST MRI contrast agent greatly depends on the saturation conditions of the MRI protocol. Our in vivo studies with other - 72-

75 PARACEST agents have demonstrated negligible changes in core body temperatures of mouse models with saturation powers of 30 μt. Saturation studies over a range of power levels indicate that sufficient detection sensitivity can be achieved at 4.6 μt (Figure 2.4), which indicates that the detection of Yb(1) is not limited by problems with saturation powers during in vivo PARACEST MRI methods. Although the sensitivity threshold for detecting a MRI contrast agent also depends on the performance of the MRI scanner, 29 a 5% change in image contrast is a generally accepted detection threshold. The dependence of the PARACEST effect on agent concentration was evaluated by using modified Bloch equations, and accounting for T 1 relaxation and direct saturation of water. These results indicate that the detection threshold of Yb(1) is 5.9 mm when the PARACEST effect of the amide is detected, and 2.9 mm when the effect of the amine is detected. These studies were conducted at ph 7.0 and 37 C to simulate physiological conditions. Variances in temperature change the absolute PARACEST contrast change by 0.44% per degree. These temperature-dependent modulations are minor under standard physiological conditions, but must be considered when approaching the detection threshold. An evaluation of Equation 1 under practical experimental conditions indicates that the PARACEST effect has a linear relationship with the endogenous T 1 relaxation time of the sample. These reported concentrations that generate a 5% PARACEST effect were measured with aqueous solutions that have relatively long T 1 relaxation times. Samples or tissues with shorter T 1 relaxation times will require higher concentrations to generate a 5% PARACEST effect. MRI of Yb(1) detected PARACEST effects from the amine and amide before reaction with NO, but no PARACEST effect from Yb(2) was observed after the reaction. The - 73-

76 formation of the triazene caused a disappearance of the amines and corresponding absence of an amine-derived PARACEST effect from Yb(2). However, the absence of a PARACEST effect from the amide of Yb(2) was unexpected. Conformational changes that relocate the amide farther away from the lanthanide ion are presumed to account for the disappearance of this amide-derived PARACEST effect. PARACEST maps were generated by subtracting the selectively saturated MR image of a contrast agent in PBS from a similar image of PBS without contrast agent. This procedure accounted for direct saturation of water, and simulated the difference of MR images acquired before and after accumulation of Yb(1) within a tissue of interest during an in vivo application. Measurements of other CEST effects routinely compare selective saturation at frequencies on opposite sides of the water frequency to account for direct saturation of water, but this comparison is not possible for Yb(1). For example, saturation at -8 ppm for comparison with saturation of the amine resonance at +8ppm may potentially saturate the amide resonance at -11ppm. Although the comparison of Yb(1) with PBS demonstrated one procedure for detection with CEST MRI, improved methods are needed that selectively saturate frequencies close to 0 ppm without directly saturating water or other CEST frequencies, in order to develop other CEST MRI procedures for in vivo studies. Differential pharmacokinetics and biodistributions can complicate the evaluation of responsive contrast agents during in vivo molecular imaging applications. 4 For example, the absence of a detectable PARACEST effect after administering Yb(1) may be due to insufficient concentrations of this imaging agent in the tissue of interest, rather than the response to NO. Fortunately, the selective detection of PARACEST agents can be - 74-

77 exploited to solve this problem. An unresponsive contrast agent with a unique PARACEST frequency can be combined with the responsive PARACEST agent, so that the unresponsive agent can be used to track pharmacokinetics and biodistribution. 20 To demonstrate this approach, the NO-unresponsive Tm-(DOTAM-Gly) was subjected to the same reaction conditions as Yb(1), and MR images were acquired with selective saturation at -51 ppm to detect the PARACEST effect of Tm-(DOTAM-Gly). No differences were detected between reactant and product, which demonstrated that this unresponsive PARACEST MRI contrast agent can be used to monitor total concentration of the imaging agent during in vivo applications. The addition of a second, unresponsive PARACEST agent will decrease the signal-to-noise ratio of the analysis, which further warrants improvements in detection sensitivities for in vivo biomedical applications. To ensure that the in vivo pharmacokinetics of responsive and unresponsive PARACEST agents are identical, this strategy must be extended by covalently linking both agents or by packaging both agents in the same delivery vehicle. Although many previous studies have limited the conjugation of only a single type of DOTA-based MRI contrast agent to polymers and dendrimers, these methods appear to be easily extensible to packaging multiple DOTA-based agents on the same delivery vehicle. 30 The same chemical reaction was used to convert Gd(1) to Gd(2), which caused a modest 44.8% increase in the T 1 relaxivity of this responsive relaxivity-based MRI contrast agent. The molecular model supports this observation, as the presence of the oaa ligand in Gd(1) doesn t physically inhibit the association of water with the lanthanide ion, and conformational changes of this ligand in Gd(2) would cause only minor improvements in water exchange. This modest increase in relaxivity may also be attributed to a modest - 75-

78 increase in rotational correlation time of the larger Gd(2) relative to Gd(1). The MR images of Gd(1) and Gd(2) (Figure 2.6b) reflect this modest difference in relaxivities. In addition, the problems of differential pharmacokinetics and biodistributions discussed above will also complicate the interpretation of the responsiveness of Gd(1), as the absence of a change in MR relaxation may be due to insufficient concentrations of the agent instead of an absence of the response. Unfortunately, relaxivity-based MRI contrast agents can t be selectively detected, so that the inclusion of an unresponsive relaxivity-based contrast agent can t be exploited to solve this problem. NO is produced by inducible Nitric Oxide Synthase (inos) at 1 10 um concentrations, and has an approximate lifetime of sec before undergoing oxidation in physiological conditions. 31 This relatively dilute and fleeting messenger molecule is a poor objective for a reversible responsive MRI contrast agent. A 1:1 interaction of NO and a reversible contrast agent will require signal amplification strategies or improved signal detection strategies to detect 1 10 um of the responding agent, or strategies must be devised that increase the ratio of responding agents relative to each NO molecule. For example, EPR can easily detect < 1 um of a spin-trapped adduct of NO, but up to two hours are required to generate sufficient quantities to produce a marginal change in MRI contrast using the same approach. 32,33 Also, the short lifetime of NO may drive dissociation of the NO-agent complex so that the response of the contrast agent may also be fleeting. For example, the same spin-trapped NO adduct has a half-life of 40 min under simulated physiological conditions, while MR images were acquired for 120 minutes to adequately assess image contrast changes

79 Irreversible responsive MRI contrast agents may be an advantageous strategy for detecting NO. As shown in Figure 2.8, sufficient amounts of irreversible responsive agents can be accumulated over time for adequate detection with MRI. Also, this response remains after NO has disappeared from the sample or site of interest, which may facilitate detection of this fleeting messenger molecule. These advantages are manifested by irreversible responsive fluorescence agents that undergo covalent modifications through a reaction with NO. 24 The in vivo production of 2.9 mm of NO (the threshold to generate a 5% PARACEST effect from the amine) is estimated to require as long as 2.5 minutes assuming that physiological concentrations of L-arginine and O 2 substrates are present, 34 so that sufficient quantities of NO can be produced just prior to MRI studies. Yb(1) is inert during the MRI scan session until it encounters NO in the presence of oxygen. Yb(1) presumably has the same raction mechanism as the fluorescence dye DAF-FM, which reacts with an N-nitroso intermediate of NO autooxidation. 24 Therefore, the in vivo administration of large amounts of Yb(1) won t perturb a biological system by directly consuming NO or O 2, and without interference from other in vivo free radical species, such as OH -, H 2 O 2 and O - 2, NO - 2, or NO - 22, This highly specific reaction mechanism ensures that this irreversible responsive PARACEST MRI contrast agent can detect NO without ambiguity. Due to the fleeting lifetime of NO, this messenger molecule is estimated to diffuse only μm from its site of production during one half-life. 35 Therefore, detection of NO with Yb(1) may also localize Nitric Oxide Synthase activity. Although inos exhibits the greatest production rate of NO, other NOS enzymes also express NO, including enos and nnos, and the exact form of NOS can t be identified through NO detection

80 Diffusion or perfusion of the responsive and unresponsive agents must also be considered in order to localize NO and NOS enzymes using this PARACEST strategy. Agents in the vasculature may be rapidly perfused from the tissue of interest, which may obviate studies of NO in highly vascularized tissues. Based on the diffusion rate of free water, agents can diffuse no more than approximately 500 μm during the 2.5 minutes that are required to generate sufficient production of NO for detection. This diffusion distance is on the order of the spatial resolution of pre-clinical and clinical MRI, and therefore only has a minor impact on localizing NO production. This report indicates some remaining challenges that must be overcome before irreversible responsive MRI contrast agents can be applied to in vivo biomedical applications. The need for retention at the site of NO production must be addressed by exploiting active or passive retention strategies, such as packaging agents within a delivery vehicle or conjugating agents to a molecular targeting probe. These nanocarrier and active targeting strategies may also improve sensitivity and reduce potential acute toxicity by delivering higher payloads of the PARACEST agents to the tissue of interest, rather than delivering PARACEST agents to other tissues for sequestration or excretion. These strategies for delivering and retaining responsive contrast agents within specific tissues will provide important improvements for developing in vivo applications with irreversible responsive PARACEST MRI contrast agents CONCLUSION Yb(1) represents a fundamentally new type of PARACEST MRI contrast agent that undergoes an irreversible covalent change that causes an irreversible change in the - 78-

81 PARACEST effect. This irreversible responsive PARACEST MRI contrast agent was selectively detected at two distinct PARACEST frequencies. The PARACEST effect from an amine expands the paradigm for designing PARACEST MRI contrast agents. Yb(1) specifically reacted with NO to cause a 100% disappearance of both PARACEST effects. A responsive relaxivity-based MRI contrast agent, Gd(1), showed only a 44.8% change in relaxivity in response to NO. The response of Yb(1) may be distinguished from pharmacokinetic effects by selectively detecting an unresponsive agent that has a different PARACEST frequency. The irreversible covalent change caused by the reaction facilitated detection, as an accumulation of the reaction product before the MRI acquisition resulted in high concentrations of Yb(2). Although this report establishes thresholds for detection sensitivity within in vitro samples, the delivery and retention of these PARACEST agents within in vivo tissues must be investigated before this approach can be routinely applied to in vivo studies. Because a variety of amides and amines are modified by many biochemical events in many physiological processes, this initial demonstration may represent a powerful platform technology for designing new irreversible responsive molecular imaging agents to address many biomedical applications

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87 CHAPTER 3 MEASUREMENT OF EXTRACELLUAR PH USING A SINGLE MRI CONTRAST AGENT BASED ON PARACEST EFFECTS 3.1. INTRODUCTION The homeostasis of extracellular ph (ph e ) is critical for the maintenance of normal cell metabolism and function. Changes in ph e have been observed for many pathological processes such as cancer, 1-3 renal failure, 4,5 ischemia 6-8 and hyperglycemia. 9,10 The measurement of ph e has significant clinical value for both pathological diagnoses and therapy. Non-invasive methods for in vivo ph e measurements have been developed based on Magnetic Resonance Spectroscopy (MRS), such as 31 P MRS, F MRS, 12 and 1 H MRS. 13,14 However, MRS has limited spatial resolution, which can limit the utility of these methods for in vivo diagnoses. For example, the low ph e within tumor tissues is heterogeneously distributed and may vary by 0.5 ph units across an 8 mm distance, which requires high spatial resolution for proper evaluations. Therefore, new MRI methods with high spatial resolution, such as contrast enhanced MRI, 18 are desirable for in vivo applications. Sensitive ph e measurements have been performed with contrast enhanced MRI by using an exogenous T 1 relaxation contrast agent However, the quantification of ph e using an exogenous contrast agent requires a second ph-unresponsive agent as a control to account for other factors that affect MRI contrast, such as the concentration of the agent. Although a T 2 * relaxation contrast agent may serve this purpose, the correlations between - 85-

88 T 1 relaxation and T 2 * relaxation create difficulties in quantifying each agent in a mixture, which makes this approach nearly impractical. MRI contrast agents that are detected via Chemical Exchange Saturation Transfer (CEST) have also been employed to measure ph e. CEST contrast agents possess protons that exchange with water at slow to moderate rates (relative to the MR frequency differences between the water and the agent). The selective saturation of the MR frequency of these exchangeable protons, and subsequent exchange with water, can transfer saturation to the water MRI signal, which leads to negative contrast in the MR image. The CEST effect can be measured from the amide protons of endogenous proteins and peptides to measure ph e to assess ischemia 8 and gliomas, 22 though the CEST effect must be separated from magnetization transfer effects. More specific and sensitive ph measurements have been reported by using exogenous CEST contrast agents 23 24, 25 and paramagnetic CEST (PARACEST) contrast agents. PARACEST contrast agents contain a lanthanide ion that greatly shifts the MR frequency of the exchangeable amide protons from the MR frequency of water, which avoids MT effects and expands the range of MR frequencies exhibited by a single agent. PARACEST agents have been designed that contain two proton exchange sites with different MR frequencies and that generate a ph-responsive and ph-unresponsive CEST effect. 23 A ratio of the two PARACEST effects can then be used to measure ph e without complications from varying concentrations of the agent. Unfortunately, these PARACEST agents only measure ph e above ph 7.0, while many pathological conditions must be evaluated in an acidic range. In addition, extremely high saturation powers are - 86-

89 required to detect both PARACEST effects, so that these agents cannot be applied to in vivo measurements. In Chapter 2, we have proposed a new PARACEST contrast agent, Yb-DO3oAA for detecting NO. This PARACEST agent possesses two PARACEST sites with distinct PARACEST offets. Our study also indicated the ph-responsive properties of these PARACEST sites were also markedly different. Herein we propose that the extracellular ph measurement could be realized by selective detecting these two PARACEST effects at their distinguishable PARACEST offsets. The ratio of the two PARACEST effects can measure the entire physiological range of ph e from , using saturation powers that are safe for in vivo studies. To the best of our knowledge, these studies represent the first demonstration of in vivo ph measurements with a single MRI contrast agent THEORY The PARACEST effect (MRI signal attenuation, M S /M 0 ) is determined by the chemical exchange rate with water (k ex or k CA ), the local T 1 relaxation time during the saturation (T 1sat, ), the concentration of the PARACEST agent [CA] and the number of labile protons at the exchange site in the PARACEST agent (n CA ) as shown in Eq. [3.1]. M M S 0 1 = n k [ CA] T 1 + n [ H2 O] CA CA 1sat H2O [3.1] While the other parameters (n CA, n H2O, [H2O] and T 1sat ) are kept constant during the study, the [CA] and k CA are the determinants for PARACEST. In other words, if [CA] is known, then k CA can be determined by measuring the PARACEST according to Eq. [3.1]

90 The exchange rate k CA is defined as the first order reaction constant for a proton to transfer from the contrast agent pool to the bulky water pool. The k CA is determined by physical parameters such as ph and temperature. The chemical exchange rate between an amide proton and water or aryl amine protons and water, is typically catalyzed by hydroxide ions or protons, respectively, according to acidic-basic catalysis theory (Eq. [3.2]) k = k0 + k [ H ] + k [ OH ] [3.2a] obs a b or ph ( W ) pk ph obs = + a + b [3.2b] k k k k k obs : the observed apparent exchange rate, k obs = k CA in a PARACEST study k 0 : spontaneous exchange constant k a :acid catalyzed exchange constant k b : base catalyzed exchange constant pk W : negative logarithm of ionization constant of water, pk W =15.4 at 37 o C 27 The temperature is another environmental variable that can affect the exchange rate according to the Eyring relationship (Eq. [3.3]). 28 ΔHex 1 1 ( kex ) T = ( kex ) 312 ( T / 312) exp( ( )) [3.3] R 312 T T: temperature in Kelvin; (k ex ) T : exchange rate at temperature T; H ex : the enthalpy difference involved in the exchange process R: molar gas constant, 8.314J K -1 mol

91 Under physiological circumstances, the temperature effect is negligible because the in vivo temperature is generally constant or only subjective to a small change. Therefore the ph is the predominant determinant of k CA. The ph can be indirectly determined by measuring k CA through the detection of PARACEST effects according to Eq. [3.1]. However, such an approach requires additional measurements of [CA] and T 1sat, and is impractical for in vivo applications because the accurate estimation of [CA] becomes very difficult. Alternatively, a secondary control agent that is unresponsive to ph can be used to account for the concentration and T 1 in the PARACEST study (Eq. [3.4]). But the accurate determination of the concentration and exchange rate of unresponsive agent is still time consuming and difficult to apply. M =. M ( 1) [ CA] n T ( k ) M 0 ( 1) A M S [ CA] A na T1 sat ( kca) A 0 B B B 1sat CA B S [3.4] The solution can be further simplified to Eq. [3.5] when PARACEST sites A and B are intra-molecularly located, 23 which allows the cancellation of [CA] and T 1sat by using a ratiometric approach. M M na ( kca) = M ( 1) n ( k ) M 0 ( 1) A S A 0 B CA B B S [3.5] Although the n CA values of an amine and an amide are also ph-dependent, the pka value of an amide is greater than 8, so that n CA is essentially invariant for an amide within the physiological ph range. Yet the n CA of amine may be variable in the ph range of 6-8, the - 89-

92 n CA is still deterministic in a certain ph range. Therefore ph can be simply measured by a single agent because the ratio of PARACEST effects from intra-molecular amide protons and aryl amine protons is only related to the ph (Eq. [3.6a]). M 0 ( 1) A M S M 0 ( 1) B M S = N f( ph) [3.6a] where N=n A /n B and f(ph) stands for a function of ph. In order to accurately determine f(ph), Eq. [3.5] and Eq. [3.2a] were combined into Eq. [3.6b] M K [ ] K [ ] 0 A A + A W ( 1) A k0 + ka[ H ] + kb + M S H M = N 0 B B + B W ( 1) B k0 + ka[ H ] + kb + M S H [3.6b] where superscripts A and B stand for the PARACEST sites A and B respectively, K W is -pkw the ionization constant of water with K W = 10. Eq. [3.6b] can be used to precisely describe the relationship between ph and measurable PARACEST effects abased on the accurate determination of the catalyzed exchange constants in Eq. [3.2a]. In practice, the conditions for a steady state PARACEST effect (Eq. [3.1]) may not be satisfied due to incomplete RF saturation duration or strong saturation power. As far the saturation duration (or saturation time, saturation delay) is concerned, the PARACEST effect can be described by a transient solution of Bloch-McConnell Equations (Eq. [3.7]). By the means of varying saturation times, parameters such as T 1 relaxation under saturation (R 1sat, R 1sat =1/T 1sat ) and water exchange rates (k water, k water = 1/τ water, where τ water is the water proton lifetime) can be extracted according to Eq. [3.7], which provides a practical way to measure the exchange rates of CEST contrast agents , 29

93 M ( Ts) R k W M R R S 1Sat water (- R1 Ts) = + e [3.7] 0 1W 1W where R1 W = R1 Sat + kwater [3.8] At equilibrium, k water is proportional to the contrast agent proton life time (Eq. [3.9]), which is independent of the concentration of CEST contrast agents. k water n = kca n CA [ CA] [ H O] HO 2 2 [3.9] 3.3. MATERIALS AND METHODS Chemicals DO3A-tBu was purchased from Macrocyclics Co. (Dallas, TX), and all other reagents were purchased from Aldrich (St. Louis, MO) and Fisher Scientific (Pittsburgh, PA). All the reactions were carried out under argon atmosphere. All the organic solvents were freshly distilled before reactions. NaOH solution and HCl solution at a variety of concentrations ( M) were used to adjust ph in all studies. During the adjustment, the volume of titrating solution added were limited to 1% (<5 μl) of the original volumes (~500 μl) and recorded for the accurate calculation of contrast agent concentrations General Methods The 1 H NMR spectrum of a contrast agent in D 2 O solvent was acquired at room temperature with a 600 MHz Varian Inova NMR spectrometer (Varian, Inc., Palo Alto, CA). The 1 H NMR spectrum of same contrast agent in H 2 O solvent was obtained by a modified water suppression WET-1D method 30 at room temperature using the same NMR spectrometer. The chemical structures of synthesized contrast agents were confirmed by using ESI MS (Bruker Daltonics Esquire HCT) or MALDI-TOF MS - 91-

94 (Bruker BIFLEX III). The matrix used in MALDI MS was a saturated solution of 2,4,6- trihydroxyacetophenone (THAP) in 50% acetonitrile. 31 The concentrations of lanthanides were determined by ICP-MS analysis (Chemical Solutions Ltd., Mechanicsburg, PA) Synthesis Synthesis of DO3oAA Ligand The DO3AoAA was synthesized from 2-nitrobenzenamine in 4 steps (Figure 3.1). In brief, 2-nitrobenzenamine (in 50 ml CHCl 3 ), 1.5 equivalent 2-bromoacetyl bromide (10 ml CHCl 3, add dropwise) and1.5 equivalent (CH 2 CH 3 ) 3 N (as base) were mixed and stirred for 1 hour at 0 o C. The reaction was continued by stirring for another 30 minutes at room temperature. The product was washed with 10 ml 5% aqueous NaHCO 3 solution for 3 times and filtered after the removal of residual water (by 10g Na 2 SO 4 ) and CHCl 3 solvent (by rotary evapoation). A chromatographic column was used to purify the product by using mixture of ether (or ethyl acetate) and petroten (1:4). The yield of the first step was 50%. Then 1 equivalent 1 was mixed with 0.7 equivalent DO3A-t-Bu-ester and 6 equivalent K 2 CO 3 in acetonitrile and heated at 70 o C for 24 hours. The product 2 was purified with a chromatographic column by using ether and methanol (95:5). The yield was 30%. The protecting t-bu groups in 2 were then removed in a TFA solution for 6 hours with a yield of 98%. After methanol/ether extraction, the nitro group in 3 was reduced to an amine group by using a hydrogenator (Pd/C (10%), water and 3 atm H 2 ) and stirring for 24 hours (yield = 90%). The final product 4 was lyophilized. The structure of synthesized DO3AoAA was confirmed by 1 H NMR (δ 2.54 (s, 2H), 2.78 (s, 6H), 3.34 (t, 16H), 5.98 (s, 1H), 7.37(t, 1H), 7.76 (t, 1H), 8.07 (d, 2H), 9.60 (s, 1H) ) and mass spectroscopy (MALDI-TOF m/z (M+H) + =495.83)

95 O O O O HO O O OH NO 2 NH 2 NO 2 a H b N N c N Br N N O N N N N 1 N NO H O O O N 2 NO H O O OH HO O O OH d N N N N NH 2 N H O O OH 4 Figure 3.1. The synthetic route of DO3AoAA. a) 2-bromoacetyl bromide (1.5 eq.), (CH 2 CH 3 ) 3 N (1.5 eq.), CHCl 3, 0 o C for 1 hour, RT. for 30 min, 50%. b) DO3A-t-Buester(0.7 eq.), K 2 CO 3 (6eq.), acetonitrile, 70 o C, 24 hr, 30%. c) 100% TFA, 6 hr, 98%. d) Pd/C (10%), water, H 2 (3 atm), 24 hr, 90% Synthesis of Ln(III) Complexes The Yb-DO3AA-oAA complex was prepared by mixing DO3A-oAA and YbCl 3 at a molar ratio of 1:1.01, in ph 6 aqueous solution at 60 o C for 2 hours. The ph was increased to 8.0 for an additional 30 minutes. After precipitating the excess free Yb 3+ ions by adjusting the ph to 12, the solution was centrifuged, filtered and lyophilized to collect a pure complex. The structure of Yb-DO3AoAA was confirmed by MALDI-TOF mass spectrometer and NMR spectroscopy. The concentrations of the chelated Yb(III) complex were determined by ICP MS analysis. The same procedure was also used to chelate Eu(III), Tm(III), Dy (III), Ho(III) and Er(III)

96 MR Acquisition Procedures PARACEST NMR measurements PARACEST NMR studies were conducted with a 600 MHz Varian Inova NMR spectrometer (Varian, Inc., Palo Alto, CA). PARACEST spectra were acquired with a modified presaturation pulse sequence 24 that included a selective continuous-wave (CW) saturation pulse applied for 3-6 seconds at 12.2 μt RF saturation power. Initial spectra were acquired with saturation frequencies spanning from 150 ppm to -150 ppm in 1 ppm increments, and subsequent spectra acquired with saturation frequencies spanning a range that was adjusted according to the PARACEST offsets of each individual Ln(III) under study. Unless otherwise noted, all PARACEST spectra were acquired at ph 7.0 and 37 C. To account for direct saturation of water, each spectrum was fit to three Lorentzian curves (centered at 0ppm, -11ppm and +8 ppm) by using a customized Matlab code (Matlab R2007a, Mathworks Corp, Natick, MA) to extract the PARACEST effects of the amide and amine. The difference between the Lorentzian fit centered at 0 ppm and the experimental data was used to calculate the magnitude of each PARACEST effect at 8 ppm and -11 ppm. The NMR data were processed and plotted using VNMR (Varian, Inc., Palo Alto, CA) and Microsoft Excel (Microsoft Co, Seattle. WA). The standard ph responsive curve was determined by using a series of samples containing 20 mm Yb-DO3A-oAA in PBS at 37 C, with ph values ranging from 6.0 to 8.0 in increments of 0.25 ph units. To validate these results, all samples were measured at least three times at different days with a ph meter (ThermoElectron, Waltham, MA) that were carefully calibrated with standard ph buffers (ph 4.0, ph 7.0 and ph 10.0). Samples with different concentrations (5 mm to 80 mm) and T 1 relaxation times (0.5 s to - 94-

97 3 s, adjusted by adding Gd-DTPA solution (Magnevist TM, Berlex Inc) were used to validate the standard curve and the potential influence of concentration and T 1 relaxation times Water Exchange Rate Measurements The same samples were also used to determine chemical exchange rates by using a 12.2 μt CW saturation pulse with durations of 0.1, 0.2, 0.5, 1.0, 2.0, 4.0 and 6.0 sec. The data were fitted to Eq. [3.7] to quantify the exchange rates of labile protons and T 1sat by using 24, 32 customized Matlab codes PARACEST MRI Procedures MRI scans were performed with a 9.4T Bruker Biospec animal MRI scanner equipped with a 35 mm birdcage RF coil. A RARE MRI pulse sequence with a RARE factor of 16 (TR/TE= 6.0 sec /8.6 msec) was used for all the MRI studies, which was prepended with a train of 300 Gaussian RF pulse applied at 21 μt for 5.49 sec to create steady-state selective saturation. The total acquisition time was 48 sec for a single MR image (128x128, single slice with a thickness of 1 mm) with single average. All the MRI phantoms were prepared in 200 μl centrifuge tubes Animals All in vivo studies were conducted according to approved procedures of the Institutional Animal Care and Use Committee of Case Western Reserve University. Mouse models with subcutaneous tumors of MCF-7 human mammary carcinoma were prepared by injecting 1.5M MCF-7 tumor cells in 0.5 ml of 50% Matrigel into the right lower flank of a 6-week-old female athymic NCR nu/nu mouse. To prepare for the MRI exam, each - 95-

98 mouse was anesthetized with % isoflurane delivered in 2 L/min oxygen gas ventilation. 60 mm of Yb-DO3AoAA in 50 μl volume was subcutaneously injected to the center of tumor. MRI scans were performed at 2 minutes after the injection with selective saturations sequentially applied at +8 ppm, -11ppm and 50 ppm respectively by using the same RARE sequence as described above RESULTS AND DISCUSSION Synthesis of DO3oAA Ligand The DO3AoAA ligand used in the Chapter 2 was purchased from Macrocyclics. After the successful application of the agent to the detection of NO, we endeavored to develop a synthetic method to synthesize DO3AoAA in our laboratory. The synthesized DO3AoAA was characterized as same as the purchased DO3AoAA CEST Characteristics of Different Ln (III) Complexes Paramagnetic Lanthanide ions markedly influence PARACEST offsets of the labile amide protons and PARACEST signal amplitudes. 24 We have investigated six Ln(III) DO3oAA complexes that chelate Yb(III), Tm(III), Dy(III), Eu(III), Ho(III) or Er(III), to determine the most efficient PARACEST system (Table 3.1). There was no observable PARACEST signal for DO3oAA complexes with Eu(III), Ho(III) and Er(III). The Yb(III) complex has been demonstrated to be the most efficient systems among the Ln(III) complexes under study. Therefore Yb-DO3AoAA was chosen as the PARACEST contrast agent for the ph measurement for the remainder of this study. Yb-DO3oAA demonstrated two distinct PARACEST effects at +8 ppm and -11 ppm, corresponding to the labile amine protons and amide protons respectively (Table 3.1, - 96-

99 Figure 3.2). The PARACEST offset of the amide proton (-11 ppm) agreed closely with similar studies with amide-containing derivatives of Yb-DOTA (-16ppm). 24 The relatively large MR frequency shift of the amine (+8ppm vs ~ ppm for diamagnetic chemical shift of an amine with respect to the water Larmor frequency) was unexpected, and was presumably due to conformational constraints of the oaa ligand that positions the amine near the Yb(III) ion. This large MR frequency shift was critical for this first practical demonstration of a PARACEST effect from an amine group, which expands the functionalities that may be exploited for PARACEST MRI. Table The PARACEST offsets and effects of the DO3AoAA complex with different lanthanide ions. Lanthanide PARACEST offset PARACEST effect * Yb -11 ppm 21.8% +8 ppm 14.5% Tm -52 ppm 8.5% -36 ppm 5.1% +16 ppm 4.0% Dy -49 ppm 5.9% -36 ppm 3.8% Er - - Eu - - Ho - - *40 mm Ln-DO3oAA complex was used, ph =7.0, temperature=37 o C - 97-

100 When an asymmetric analysis is used, significant interferences can be caused due to the spillover of saturation of one PARACEST offset to its neighbors, i.e. -11ppm to -8 ppm and +8 ppm to +11 ppm. In order to account for the direct saturation of water, the experimental NMR data were corrected by fitting a three-lorentz model (centered at 0 ppm, -11ppm and +8 ppm) to account for the interference (Figure 3.3). The PARACEST effects of amide and amine that would use in the remainder of this study were extracted from the experimental data by the same procedure. Figure 3.2. The CEST spectra of 20 mm Yb-DO3AoAA in PBS with ph varied from 6.12 to 8.0 at 37 o C. A CW pulse was used to saturate a frequency from +20 ppm to -20 ppm in 1 ppm increments. The selective RF saturation was applied for 6 sec with a power of 12.2 μt. The arrows show PARACEST signals at -11ppm (blue) and 8 ppm (red) for amide proton and amine protons respectively

101 1 0.8 Ms/M saturation offset(ppm) water amide amine experimental data sum of fittings Figure 3.3. The extraction of PARACEST effects of amine and amide from experimental data by a three-lorentzian model (centered at 0 ppm, -11ppm and +8 ppm). The experimental data shown here (black solid line) is the CEST spectrum for a 20 mm Yb-Do3AoAA NMR sample (ph 7.5, 37 o C). The PARACEST NMR study is conducted with a selective CW saturation pulse (at 12.2 μt for 6 sec) sweeping from +20 ppm to - 20 ppm in 1 ppm increments. The model output showed a symmetric water Lorentzian curve (pink dash line), and asymmetric Lorentzian curves of amine (blue circle line ) and amide (green star line) ph Dependencies of PARACEST Signals As is evident in Eq. [3.2], chemical exchange of the amine and amide with water is catalyzed by hydrogen or hydroxide ions, and therefore is dramatically affected by ph. Consequently, the PARACEST signals are significantly changed due the changes in exchange rates. Moreover, the ph dependencies may be significantly different due to the - 99-

102 differences in the nature of bonding and surrounding electron environments for a variety of protons with PARACEST signals. The different ph dependencies of the two PARACEST effects of Yb-DO3AoAA have been investigated in the ph range of (Figure 3.2 and Figure 3.4a). The distinct ph dependencies endow the PARACEST agent with the ability to measure ph according to Eq. [3.6]. The standard ph responsive curves of the PARACEST effects of Yb- DO3oAA has been measured and presented in Figure 3.4a. The ratiometric approach was conducted according Eq. [3.6]. Since a linear calibration curve is always desirable for a measurement, we have investigated several post-processing methods in order to obtain an approximately linear relationship between the ph and the ratio of PARACEST signals. The logarithm of the (M0/Ms-1) ratio was chosen as the best solution (Figure 3.4b) and provided a broad linear relationship range approximately from ph 6.5 to ph 7.5, which fits the requirements for the most measurements in physiological ph and pathological ph ph Dependencies of Water Exchange Rates The exchange rates were estimated by fitting the theoretical model of the experimentally obtained PARACEST effects at different ph values with varying RF saturation durations( Figure 3.5). 24, 32 The maximal exchange rates in the ph range under study were measured as Hz (ph 8.0) and Hz (ph 7.0) for amide and amine protons respectively. The unusually slow exchange rate of the amine is presumably due to hydrogen bonds that are formed between the amine and the carboxylate ligands of the agent

103 a b. Figure 3.4. The different ph dependencies of amide CEST and amine CEST (a), and the ph measurement calibration using the proposed PARACEST agent (b). The phantoms consisted of 20 mm Yb-DO3AoAAA in PBS with ph values ranging from 6.0 to 8.0 in 0.25 increments. The PARACEST NMR study is conductedd by a modified presaturation pulse sequence with a selective CW saturation pulse (at 12.2 μt for 6 sec) sweeping from +20 ppm to -20 ppm in 1 ppm increments. The estimated exchange rates of amide and amine were fitted to the acid-base catalyzed exchange model (Eq. [3. 2]) to reveal the underlying catalysis mechanisms. 33 The amide proton showed a typical base-catalyzed characteristics 34 with a base catalyzed exchange constant K b of 6.71 x10 9 s -1 (K 0 = s -1 ), which was very close to the previously reported value of 5.57 x10 9 s The ph dependency of the amine was found to be more complicated. In the ph range of 6.0 to 7.0, the agent showed a base catalyzed water exchange behavior with base catalyzed exchange constantt K b of 3.19 x10 10 s -1 (K 0 = s -1 ). In the ph range of 7.0 to 8.0, the experimental data demonstrated acid catalyzed exchange characteristics 33 with an acid catalyzed exchange constant K a of

104 Figure 3.5. The maximal proton water exchange rates for amide (a) and amine (b) estimated by fitting the PARACEST signals with a time dependent PARACEST theoretical model according to Eq. [3. 6]. Figure 3.6. The exchange rate dependencies on ph of amine and amide labile protons. The measured exchange rates of the amide (red ) and the amine (blue ) were fitted to Eq. [3.2] to extract the acid-base catalyzed exchange constants. The red solid line was the estimated behavior for amide exchange process. The green dash line and blue dash line were the estimated results for the exchange of amine protons in ph ranges of and respectively

105 x10 9 s -1 (K 0 = s -1 ) (Figure 3.6). The underlying mechanism for the aryl amine proton to swap from base-catalyzed to acid catalyzed exchange around ph 7 is still under investigation, but very likely is due to the changes in prontonated state of the aryl amine group at ph The spontaneous exchange constants also slightly changes mainly due to the same reason Accuracy of the ph Measurement The ph of a series of MRI phantoms have been measured with respect to the concentration of Yb-DO3A-oAA and T 1 relaxation times. The accuracy of this ph measurement method thereby was evaluated with respect to measurement by a calibrated ph meter, which can be used as the standard method. The comparison showed that most measurements fell in the error range of 0.2 ph units with a mean error of ph units (Table 3.2). Therefore the accuracy of this method was close to the MRS methods, which typically have the standard deviation of ph units. 18 However, when the T 1 relaxation time was less than 0.5 sec (data not shown), the error was too high to produce a reliable result due to the low PARACEST contrast to noise ratio. The ph measured with low concentration (<24 mm) were also associated with high level errors (>0.2 ppm) and indicated a higher concentration of PARACEST contrast agent that is able to produce higher PARACEST contrast should be used for the in vivo applications. Nevertheless, these results verified that ph measurements determined from the PARACEST effects were independent of T 1 and [CA]

106 Table The accuracies of the proposed ph measurement method concentration T 1 (sec) Ph Difference (mm) ph meter PARACEST In Vivo Demonstration An in vivo study was conducted by directly injecting a solution of Yb-DO3A-oAA directly into a solid tumor in a mouse model of mammary carcinoma. The xenograft flank tumor was very mature (~4 weeks old) and big in size (4.55 mm), and therefore was expected to exhibit low ph e. 36 A parametric map of ph was calculated by comparing the water signal at 2 minutes after injection (M s ) with the water signal before injection (M 0 ) at saturations applied at +8 ppm and -11ppm, using Eq. [3.5] (Figure 3.7). Most of the tumor region showed a ph e of 6.5 or lower, while the phantom containing the PARACEST agent showed a ph of 6.9 as expected. Although the magnitudes of the

107 PARACEST effects indicated that the PARACEST agent may be heterogeneously distributed throughout the tumor, this heterogenous distribution did not affect the ph map. The successful tumor ph measurement in a live subject demonstrated the feasibility to establish a simple ph measurement MRI method without the use of a second control agent to account for the pharmacokinetics and biodistribution. However, there are still a number of drawbacks that need to be overcome. The major limitation to apply this method to in vivo studies is still the relatively low sensitivity of the MRI method. A high concentration of contrast agent (>24 mm) may be needed to obtain sufficient contrast and high accuracy, which may be beyond the clinical limitation because of the potential toxicity. A polymerization approach can be used in the future to improve the sensitivity and to reduce the potential side effects caused by a high concentration of agents. Another concern when applying this method to in vivo studies is that there may be nonspecific or covalent bonding between the PARACEST agents and in vivo macromolecules such as serum albumin. As a result, the chemical exchange processes of PARACEST agents may behave differently relative to in vitro models. The standard ph response curve of PARACEST agents thus needs to be recalibrated according to circumstances encountered in vivo. Lastly, the B 0 inhomogeneity and B 1 inhomogeneity, especially in a higher magnetic field (>3Tesla), can seriously compromise the accuracy of the ph measurement. Indeed, the non-uniform distribution of the ph map in the contrast agent phantom and water phantom (Figure 3.7d) has indicated that the artifacts caused by the B 0 inhomogeneity and B 1 inhomogeneity were not negligible. Therefore further developments to reduce the artifact and improve the image quality will be essential and

108 can greatly facilitate the in vivo applications of the proposed ph measurement method with PARACEST MRI. Figure 3.7. The demonstration of the proposed ph measurement MRI method in a mouse tumor model. The contrast agent was subcutaneously injected into the approximate center of tumor. a) the proton density weighted MRI images with the phantoms (water and 20 mm contrast agent at ph 6.9) and tumor ROI were marked, b) the parametric map of the ratio of the PARACEST effects of amine and amide, d) the pseudo ph image by converting the ratio to ph based on the standard ph responsive curve, and c) the zoomed ph map of tumor ROI

109 3.5. CONCLUSION In summary, a new PARACEST MRI contrast agent Yb-DO3A-oAA showed two PARACEST effects that can be used to accurately measure ph with a single agent. The intra-molecular dual PARACEST effect enabled the measurement of ph without the need for a second control agent to account for agent concentrations or T 1 relaxation. A preliminary in vivo study demonstrated the accuracy and feasibility of applying this method to in vivo studies of pathological tissues that require high-resolution maps of the physiological ph range

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115 CHAPTER 4 IN VIVO PARACEST MRI WITH IMPROVED TEMPORAL RESOLUTION 4.1. INTRODUCTION Chemical Exchange Saturation Transfer (CEST) is a novel MRI contrast mechanism that is an attractive alternative to T 1 and T 2 contrast mechanisms, 1 particularly at high magnetic fields. 2 CEST agents possess a hydrogen proton with a moderate to slow exchange rate with water. Selective saturation of the MR frequency of this proton, followed by exchange with solvent water, reduces the MR signal of the water. PARACEST (PARAmagnetic CEST) agents include a paramagnetic lanthanide ion that shifts the MR frequencies of the exchangeable proton to unique values to facilitate selective detection. 3,4 Endogenous MR contrast may be continually monitored in the presence of PARACEST agents by neglecting to saturate the MR frequency of the exchangeable proton (and assuming that the T 1 relaxation of the PARACEST agent is negligible). Selectively-detectable PARACEST agents have been designed to respond to enzymes, 5 metabolites, 6,7 metal ions, 8 tissue ph, 3,9 and temperature. 10 Therefore, responsive PARACEST agents may have particular advantages for molecular imaging applications. In vivo applications of off-resonance PARACEST MRI have not been established due in part to the formidable challenges of low sensitivity and poor temporal resolution. In general, the PARACEST effect is measured through asymmetric analysis of coherent

116 magnetization M S following selective saturation at Δ ω (frequency with respect to water resonance frequency) and Δ ω (Eq. [4.1]). %PARACEST( Δ ω ) = 1- M S ( Δ ω )/M S ( Δ ω ) [4.1] Therefore two image acquisitions at saturation frequencies of Δ ω and - Δω are required for a quantitative PARACEST study. In practice, a long saturation scheme is also required in addition to the imaging scheme for full saturation transfer between two proton pools prior to each k-space acquisition. 3 As a example, with the use of a saturation pulse of approximately 2 s, the total scan time for a standard quantitative PARACEST study can be as long as 10 minutes, which includes the acquisition of two MR images (M S ( Δ ω ) and M S ( Δ ω ) ) with 128 phase encoding steps. It is impractical to apply PARACEST agents to an in vivo dynamic study with such a poor temporal resolution. Therefore the development of PARACEST MRI methods with improved temporal resolution is highly desirable for in vivo applications of PARACEST agents. Some biomedical applications require the monitoring of biological processes with a temporal resolution that is much faster than 10 minutes. For example, the accumulation of small molecule MRI contrast agents in tumors and normal tissues can often occur within the first 10 minutes after the agent is injected into the vasculature, followed by a wash-out of the agent from the tissue during the subsequent minutes. 11 Although temporal resolutions approaching 1 s can most accurately determine the rates of these pharmacokinetic processes, 12 a temporal resolution on the order of s can be sufficient to estimate these rates. 13,14 Therefore, PARACEST MRI methods with a temporal resolution of approximately s could be used for Dynamic Contrast Enhanced (DCE) MRI studies of pharmacokinetics processes

117 As we describe here, the temporal resolution of PARACEST studies may be greatly improved by employing faster MRI pulse sequences, such as a multiple-echo acquisition scheme (Figure 4.1A). The temporal resolution may also be improved by employing MRI pulse sequences with faster selective saturation periods, such as an interleaved repetitive short saturation-acquisition scheme (Figure 4.1B). In this report, both methods are compared with each other and with a standard PARACEST MRI method to assess the impact of improved temporal resolution on PARACEST contrast. Furthermore, both methods are applied to in vivo PARACEST MRI studies to demonstrate that the choice of the PARACEST MRI method is dependent on the properties of the PARACEST agent and the endogenous tissue. Figure Schematic of solutions to improve the temporal resolution of a PARACEST MRI study by using A, a multiple-echo strategy; or B, a short repetitive saturation strategy. 4.2.THEORY Theoretical Model for PARACEST Contrast

118 PARACEST can be described by a two-site exchange model for most cases, 3,15 or can be simplified to be a two-site exchange model in more complicated situations. 16 A two-site exchange PARACEST system consists of a bulk free water proton pool (W) and a small labile proton pool on a contrast agent (CA) that is water-exchangeable. The longitudinal magnetization of bulk water can be described by a simplified Bloch-McConnell equation 17 with the assumption of complete saturation of pool CA and no direct saturation of the pool W (Eq. [4.2]). dm dt W = M / T - M (1/ T + / τ ) [4.2] W W 0 1W 1W W W M : MRI signal or the magnetization of pool W W M : the equilibrium magnetization of pool W 0 T 1W: longitudinal relaxation time of pool W in the presence of saturation τ W : the proton residence life time of pool W At thermal equilibrium, the exchange rates between the two pools have a mass balance (Eqs. [4.3]). M / M = τ / τ [4.3.a] W CA 0 0 W CA [ proton] 2[ or W HO 2 ] τw = τca = τ [4.3.b] CA [ proton] n[ CA] CA where τ CA is the proton residence life time of pool CA In Eq. [4.3b], n is the number of exchangeable protons on the PARACEST agent and [H 2 O] and [CA] are the concentrations of water and the PARACEST agent, respectively. The Bloch-McConnell equation can be solved using numerical 4,16 or analytical 18 methods. A steady state analytical method (Eq. [4.4]) is most often used to provide contrast

119 quantification while a transient analytical solution can provide more useful information for the exchange rate estimation 3, 4 as well as for method optimization (Eq. [4.5]). M M S τw = T + τ 0 1W W [4.4.a] M S nt1 W[ CA] or 1/(1 ) M = + 2 τ [ H2 O] [4.4.b] 0 CA M ( Ts) τ τ W M T τ S 1W 1W (- Ts / τ1 ) = + e [4.5] 0 1W W In Eqs. [4.4] and [4.5], M S and M 0 represent the coherent magnetization of the pool W with or without saturation, respectively. Also, τ1w represents the effective residence lifetime of pool W (Eq. [4.6]) = + τ T τ 1w 1w w [4.6] The Dynamics of a PRACEST Contrast As is evident from Eq. [4.5], the attenuation of the MR signal of the water pool depends on the length of the RF saturation pulse (T S, time of saturation). To maximize the PARACEST contrast, one should maximize the time of saturation, as is typically done in PARACEST experiments which use a very long T S (> 2 s) to ensure the establishment of the steady state of M S (Eq. [4.4]), especially when T 1 is also long. 3,15 Upon the removal of the saturation RF pulse, the M S of pool W relaxes to the thermal equilibrium state (M 0 ) at a rate equivalent to its native longitudinal relaxation rate (R 1W = 1/T 1W ). 18 The recovery of magnetization following saturation causes the loss of PARACEST contrast. Therefore the ideal PARACEST MRI acquisition would record the signal immediately

120 after the steady state of CEST is established, which is the typical scheme used for a NMR PARACEST study. This formalism can be extended for MRI studies with an effective echo time (T E, the time between excitation pulse and the center of echo; Eq. [4.7]). This full temporal dependence of M S can be modeled with numerical methods. Two boundary conditions can provide an intuitive understanding of this temporal dependence. When T E << T 1w, the PARACEST contrast is maximized when T S >> τ 1w. When T E >> T 1W, the PARACEST contrast approaches zero. Therefore, a short echo time and a long saturation time provide ideal PARACEST contrast. M(T+T) M τ τ S S E 1w (-T E /T 1w ) (-Ts/τ 1w =1- e (1-e ) ) 0 w [4.7] The Temporal Resolution and Efficiency of PARACEST MRI The temporal resolution for a MRI sequence is typically described as 1/TR. 19 The total time to acquire a MR image sequence, T total, is a function of TR and the numbers of slices, the number of phase encoding step and the number of averages. Moreover, the current fashion for a quantitative PARACEST study is to use asymmetric analysis (Eq.[4.1]), which requires 2 acquisitions at Δ ω (often referred as on-resonance offset for a PARACEST agent) and Δ ω (often referred as off-resonance offset for a PARACEST agent). Therefore, the T total of a PARACEST image is twice of the T total for a single MR image. The Contrast-to-Noise (CNR) Efficiency was used to quantitatively compare different PARACEST MRI methods in addition to the PARACEST contrast and temporal resolutions. The image contrast was obtained by subtracting the pre-contrast image from

121 the post-contrast image. The contrast was quantified as the mean of the ROI in the contrast image and the noise was estimated by the standard deviation (δ) of the air noise area and multiplied by The CNR efficiency was thereby calculated according to Eq. [4.8]. CNR Efficiency = contrast /( 2 δ t ) [4.8] total 4.3. MATERIALS AND METHODS Chemicals Eu-DOTAMGly (Eu(1)) and Tm-DOTAMGly (Tm(1)) (Figure 4.2) were synthesized according to published procedures 21 and used as model PARACEST agents for phantom studies. The labile protons that generate PARACEST signals are bound-water protons for Eu(1) and Eu(2), and amide proton for Tm(1) respectively. Tm(1) was also used for tumor PARACEST DCE MRI studies. Eu-DOTAOBS2Gly2COOH (Eu(III)1,7-Bis (2- (methylene benzyloxy ether)-acetic acid) acetamide-4,10-bis (acetamidoacetic acid) - 1,4,7,10-tetraazacyclododecane; Eu(2); Figure 4.2) was synthesized and characterized according to published procedures 22 and used for in vivo liver studies. These published procedures have also demonstrated that Eu(1) and Eu(2) possess very similar T 1 relaxation and PARACEST properties

122 Figure 4.2. Structures of Eu-DOTAMGly (Eu(1)), Tm-DOTAMGly (Tm(1)) and Eu- DOTA-OBS2Gly2COOH (Eu(2)). The protons (i.e. bound water or amide) marked in red color are the exchangeable protons that contribute the PARACEST signals. The PARACEST saturation frequencies of Eu(1), Eu(2) and Tm(1) were confirmed to be +50 ppm, +54 ppm and -51 ppm, respectively, via NMR spectroscopy of samples dissolved in H 2 O with 10% D 2 O. The concentrations of PARACEST MRI phantoms were 20 mm for Eu(1) and 5 mm for Tm(1). The concentrations of phantoms used to measure T 1 relaxivities were 0, 12.5, 25, and 50 mm for Eu(1) and 5, 10, 20 and 40 mm for Tm(1). All phantoms used phosphate-buffered saline to maintain a ph of All phantom studies were conducted at 37±0.2 C using a temperature probe and an air heater (SA Instruments, Inc.)

123 Simulations All the modeling and simulations were conducted by using customized Matlab codes. Simulation 1 was conducted according to Eq. [4.7] for a standard PARACEST sequence with a model contrast agent representing 20 mm Eu(1) with a reported chemical exchange rate of 33 KHz at 298K. 23 The model parameter T 1W was set to 1, 2 and 3 s to represent relatively short, moderate and long T 1 relaxation times at high magnetic field strengths. The saturation consisted of 1000 Gaussian pulses with a pulse length of 2.25 ms and inter-pulse delay of 10 μs (Ts =2.27 s). A TE up to 2.5 s was used to study the loss of PARACEST contrast caused by the post-saturation delay. Then PARACEST contrasts were calculated for different T 1W by using 320 ms TE time, which is effective TE time for a presat-rare sequence with RARE factor of 64. Simulation 2 was conducted by iterating Eq. [4.5] for 128 phase encoding steps of presat- FLASH sequences with a TR time of 100, 200 or 300 ms. The model parameter T 1W was set to 1 sec to simulate a relatively short T 1 relaxation time at high magnetic field strengths. The model contrast agent and saturation conditions were the same as described above except that the numbers of Gaussian pulses were adjusted so that the saturation filled TR Animals All in vivo studies were conducted according to approved procedures of the Institutional Animal Care and Use Committee of Case Western Reserve University. Mouse models with subcutaneous tumors of MCF-7 human mammary carcinoma were prepared by injecting 1.5M MCF-7 tumor cells in 0.5 ml of 50% Matrigel into the right lower flank of a 6-week-old female athymic NCR nu/nu mouse. An 18-week-old Balb/C mouse was

124 used to study PARACEST MRI in liver tissues. To prepare for the MRI exam, each mouse was anesthetized with % isoflurane delivered in 2 L/min oxygen gas ventilation. A 26g dental catheter was inserted in the tail vein to facilitate the administration of 3.2 mmol/kg mm Eu(2) or 4.0 mmol/kg Tm (1) in 100 μl injection volume. The mouse was then secured to a customized MRI-compatible cradle, probes for monitoring rectal temperature and respiration were connected to the mouse, and core body temperature was regulated using an automated feedback loop between the temperature probe and an air heater (SA Instruments, Inc). At the conclusion of the MRI scan, the mouse was removed from the MRI magnet and cradle, and euthanized with CO 2 asphyxiation prior to recovery from anesthesia MRI Acquisition Procedures In order to optimize various acquisition procedures for PARACEST, several different MRI pulse sequences were tested. To simplify the analysis, single slice acquisitions were used in all studies. A Gradient Echo MRI pulse sequence with a single echo was prepended with a selective saturation period to create a presat-gre pulse sequence (TR = 10 sec, TE = 2.3 msec). This selective saturation period consisted of a train of 1000 Gaussian RF pulses (power = 20 μt, duration = 2.25 msec (bandwidth = 1218 Hz), interpulse delay=10 μs) followed by a 258 μs spoiling gradient to remove residual transverse magnetization at the end of each T S period. To quantify the PARACEST effect according to Eq. [4.1], images were acquired with selective saturation applied at the on-resonance offset frequency ( Δ ω ), and the opposite saturation frequency (- Δ ω ) with respect to the water resonance. The presat-gre method employed a NMR-like saturation and

125 acquisition scheme, and therefore could be used as standard CEST method for quantitative comparison. To investigate the strategy of multiple-echo acquisition, a presat-rare MRI pulse sequence was developed that consisted of a conventional RARE (Rapid Acquisition with Refocused Echoes) sequence 24,25 (RARE factor = 4, 8, 16, 32 or 64, TR = 10 s, TE=10.9 msec) that was prepended with the same selective saturation period as the presat-gre (Figure 4.3A). To investigate the strategy of short saturation, a presat-flash MRI pulse sequence was developed from a conventional FLASH (Fast Low Angle Shot) sequence 26 (TR = 100 msec, 200 msec, or 300 msec, flip angle = 39 o, 41 o, 43 o, TE = 2.3 msec) that was prepended with selective saturation. This saturation consisted of a number of Gaussian RF pulses (43 to 132 pulses, power = 20 μt, pulse duration = 2.25 msec) that were varied to fill the given TR (Figure 4.3B). All in vitro PARACEST MRI measurements used a 128x128 matrix size and two averages (to reduce RF inhomogeneities through phase cycling). MRI studies were performed with a 9.4T Bruker Biospec animal MRI scanner (Bruker Biopsin Co. Billerica, MA) equipped with a 35 mm birdcage RF coil. In vivo DCE MRI studies of tumor tissue were conducted by using a presat-rare method with a RARE factor of 16 (TR = 5000 msec, 258x258 matrix size, single average). A continuous series of PARACEST MR images were acquired with selective saturation applied at -51 ppm for Tm(1). The saturation pulses and other MR parameters were the same as the in vitro presat-rare method as described above. The total acquisition time for one presat-rare image was 80 sec

126 Figure 4.3. A: Pulse sequence diagram of a PARACEST sequence with a multiple-echo imaging scheme. B: Pulse sequence diagram of a PARACEST sequence with a short repetitive saturation scheme. For presat-rare, N represents the product of the numbers of phase encoding steps/echo train length, slices and averages. For presat-flash, N represents the product of the numbers of phase encoding steps, slices and averages

127 In vivo DCE MRI studies of liver tissue were conducted by using a presat-flash method with a TR of 300 ms (128x128 matrix size, two averages). A continuous series of PARACEST MR images were acquired with selective saturation applied at +54 ppm for Eu(2). The saturation pulses and other MR parameters were same as the in vitro presat- FLASH as described above. The total acquisition time for one presat-flash image was 76.8 sec. For all in vivo studies, the PARACEST agent was injected immediately after the third PARACEST MR image was acquired. Unlike previously published PARACEST MRI studies, selective saturation was not applied at -54 ppm or +50 ppm to measure M S (-Δω) for quantifying each PARACEST effect. Instead, the PARACEST contrast was quantified within each image by substituting the average M S during the first 3 images for M S (-Δω) in Eq. [4.1]. This procedure avoided an interruption in continuous selective saturation at a single MR frequency. A MSME single-spin-echo MRI experiment was used to determine T 1 relaxation times of phantoms of Eu(1) and Tm(1) (TR = 0.2, 0.5, 1, 1.5, 2, 3, 4.5, 6, 8, 12, 15, and 20 sec, TE= 11.2 ms, flip angle 90, one average). A similar MRI experiment was used to determine T 1 relaxation times of an in vivo flank tumor and liver tissue (TR = 0.1, 0.3, 0.6, 1, 1.5, 2.25, 3.5, 5, 7.5, and 10 sec, TE = 11.2 ms, flip angle = 90, one average). To calculate T 1 relaxation times, T 1 -weighted MR images were processed with PARAVISION (Bruker Biospin Inc) or the MRI analysis calculator package of ImageJ (NIH). 27,28 The PARACEST MR images were analyzed with ImageJ and MS Excel (Microsoft, Seattle, WA)

128 4.4. RESULTS Simulations The dynamics of PARACEST were depicted by Simulation 1 (Figure 4.4). These dynamics heavily depend on a long saturation scheme (pulse number >> 1 ) to reach a steady state with respect to conventional Magnetization Transfer (MT) methods. 29,30 This long saturation scheme not only increases the acquisition time but may also cause a high specific absorption rate (SAR). Therefore, a shorter and more practical saturation scheme should be used for in vivo PARACEST MRI applications. Figure 4.4. Simulations of the water signal during the saturation scheme and after the removal of saturation pulses. The parameters are τ w = 300 μs, T 1w = 1 sec, 2 sec and 3 sec. For simulation 1, a train of 1000 Gaussian pulses (Ts = 2.27 sec) was used for a simulation, which generated PARACEST effects that were very close to steady state CEST for all values of T 1w (Table 4.1). For comparison, a train of 1500 Gaussian pulses

129 (Ts=3.4 s) was used for a simulation, and the further improvement in CEST was only 0.6%, 1.9% and 2.7% (Table 4.1) relative to steady state values for T 1W of 1, 2 and 3 sec, respectively. Table 4.1. The simulated PARACEST effects of 20 mm Eu(1) with different PARACEST MRI methods that employ a long T s. Steadystate % PARACEST presat-rare %PARACEST Ratio Presat-RARE/Steady-state τ S = τ S =3.4 τ S =2.25 τ S =2.25 τ S =3.4 τ S =2.25 τ S =2.25 sec sec sec sec sec sec sec TE eff =0 TE eff TE eff =0 TE eff TE eff =0 TE eff =0 TE eff =0.32 sec =0 sec sec =0.32 sec sec sec sec T 1W =1 sec T 1W =2 sec T 1W =3 sec 54.8% 54.8% 54.4% 39.6% 99.9% 99.3% 72.2% 70.8% 70.6% 69.3% 59.1% 99.7% 97.9% 83.5% 78.4% 78.0% 76.1% 68.4% 99.5% 96.7% 87.9%

130 The recovery of saturated magnetization after the removal of irradiation pulses was also dependent on T 1W (Figure 4.4). For PARACEST contrast agents with low T 1 relaxivity and tissue with long endogenous T 1 relaxation time, the longitudinal relaxation was sufficiently slow that the loss of PARACEST was negligible during the time scale of hundreds of milliseconds, which would allow for a multiple echo imaging acquisition scheme to be fully accomplished. As an example, the simulation showed that the PARACEST effect was only slightly reduced by 8.8% in contrast after a 320 msec delay (which is an effective TE for RARE factor of 64) for environmental T 1w of 3sec. However, under the circumstances of relative short environmental T 1 relaxation time, the relaxation is too fast to avoid severe loss of PARACEST contrast. Therefore, an alternative saturation scheme must be used to compensate for rapid relaxation. We proposed that a presat-flash sequence could be used to improve the temporal resolution for circumstances with relatively short environmental T 1 s. Simulations of a short repetitive saturation scheme prior to the acquisition of each k-space line were investigated as a mechanism for this T 1 compensation (Figure 4.5). The simulation showed that this strategy quickly achieved a steady state of saturation of pool W within 30 repetitions with a TR as short as 100 ms, and effectively maintained the steady state PARACEST by compensating for relaxation loss of PARACEST contrast prior to each acquisition (Figure 4.5). The simulated PARACEST effects were 53.1%, 54.0% and 54.3% with a TR of 100 ms, 200 ms, and 300 ms respectively, which corresponds to 99.1%, 98.5% and 96.7% of the theoretical steady state value of 54.8%. Most importantly, these simulated results demonstrated an overall decrease in the PARACEST effect by using short TR presat-flash relative to presat-rare for environments with long T

131 relaxation times of 3 sec, similar PARACEST effects from each MR sequence for moderate T 1 relaxation times of 2 sec, and an improved PARACEST effect by using short TR presat-flash relative to presat-rare for short T 1 relaxation times of 1 sec %PARACEST TR=100ms TR=200ms TR=300ms PARACEST ss number of repetition Figure 4.5. Simulation of a build-up and maintenance of PARACEST by using short repetitive saturation pulses. The parameters are τ w = 300 μs, T 1w = 1 sec, TR = 100 ms, 200 ms, and 300 ms respectively. PARACEST SS is the theoretically calculated steady state PARACEST effect. Note that only the first 1-32 acquisitions of 128 phase encoding steps are displayed to show the steady state build-up Phantom Studies Quantitative studies were conducted with a series of phantoms containing different concentrations of the contrast agents Eu(1) and Tm(1). Eu(1) had a very low T

132 relaxivity (r1 = mm -1 s -1 ) and the addition of 5 mm of this agent to PBS caused a nearly negligible change in T 1 relaxation time from 5.0 s to 4.8 s. Therefore, Eu(1) represented a model PARACEST agent with low T 1 relaxation for the study of presat- RARE and presat-flash sequences. Tm(1) had a relative high relaxivity (r1= mm -1 s -1 ), and the addition of 5 mm of this agent to PBS caused the T 1 relaxation time to be drastically shortened from 5.0 s to 1.1 s. This result indicated that 1.9% of the 5mM Tm(III) dissociated from Tm(1) to form the aquo Tm(III) ion, based on a reported relaxivity of Tm(III) of 3.69 mm -1 sec -1. Therefore, this sample of 4.91 mm Tm(1) and mm Tm(III) ion. Therefore, Tm(1) was selected as an example of a PARACEST agent with relatively high T 1 relaxation for the study of a presat-flash sequence. A presat-rare method with a RARE factor of 64 delivered a 64-fold improved temporal resolution compared to the presat-gre method with one echo. Using the RARE sequence with a factor of 64, PARACEST MRI was achieved with a temporal resolution of sec per PARACEST image for, 128 phase encoding steps, and 2 averages, which is sufficient for monitoring changes in PARACEST contrast during dynamic MRI studies. This method retained 90.0% of the PARACEST contrast of the presat-gre method for 20 mm Eu(1), and the overall CNR efficiency was improved from to (Table 4.2), which indicates that the CNR of the presat-rare method would exceed the CNR of the gradient-echo method if both methods were acquired for the same total time. A presat-rare method with a RARE factor of 32 achieved the best PARACEST contrast, as it retained 93.6% of the PARACEST contrast and achieved a CNR efficiency of at the expense of a longer temporal resolution of sec per PARACEST image. The PARACEST contrast from the Tm(1) phantom was measured to be less than 1% with a

133 presat-rare method with a RARE factor of 4 to 64, which demonstrated that this presat- RARE sequence was unable to detect reliable PARACEST contrast from this agent. The presat-flash method also showed improved temporal resolution compared to the presat-gre method with one echo (Table 4.2). With TRs ranging from 100 to 300 msec, the steady-state saturation of the water pool was quickly reached during one image acquisition with 64 phase encoding steps. Therefore, the second image and all subsequent images of a continuous series of image acquisitions reached steady-state saturation and could be used to quantitatively assess PARACEST contrast. The presat-flash sequence retained 59.1%, 68.7% and 79.7% of the PARACEST contrast relative to the presat-gre sequence for 20 mm Eu(1) with TRs of 100, 200, and 300 ms, respectively. Yet the CNR efficiency was 0.261, and for a TR of 100, 200 and 300 ms respectively, which preserved or was greater than the CNR efficiency of the presat-gre sequence. The temporal resolution of presat-flash with a TR of 100 msec was 51.2 sec, which is sufficient for monitoring changes in PARACEST contrast during dynamic MRI studies. The PARACEST effect from a 5 mm Tm(1) phantom was measured to be 1.85%, 2.02% and 2.35% with a TR of 100, 200, and 300 msec, respectively. However, the CNR efficiency was only 0.015, and due to very low CNR (0.075, and 0.113) for a TR of 100, 200 and 300 ms respectively. For comparison, the presat-gre sequence detected a 2.77% PARACEST effect with a CNR of 0.79 and a CNR efficiency of In vivo Studies Both presat-rare and presat-flash methods were used to perform in vivo PARACEST MRI studies, particularly to investigate the applicability of PARACEST agents for DCE MRI studies. The presat-rare method was successfully applied to a

134 Table 4.2. The quantitative comparison of contrasts, temporal resolutions and efficiencies of different PARACEST MRI methods by using a 20 mm Eu(1) phantom. Methods Presat- GRE Presat- RARE Rare factor=4 Presat- RARE Rare factor=8 Presat- RARE Rare factor=16 Presat- RARE Rare factor=32 Presat- RARE Rare factor=64 Presat- FLASH TR=300ms presat- FLASH TR=200ms Presat- FLASH TR=100ms % PARACEST PARACEST contrast relative to presat-gre Time for a single CEST image (sec) Temporal resolution improvemen t CNR CNR Efficiency 37.4% 100% % 93.5% % 91.6% % 92.2% % 93.6% % 90.0% % 79.7% % 68.7% % 59.1%

135 PARACEST DCE MRI study of a subcutaneous flank tumor using Tm(1). The endogenous T 1 relaxation time of the flank tumor was measured to be 3.08 s, which was sufficiently long for a presat-rare sequence with a RARE factor of 16 while still retaining good PARACEST detection (Figure 4.6). To achieve a 95% and 99% probability that the contrast before and after injection of the agent was different, the CNR must reach 2 2 and 3 2, respectively. 20 The 99% CNR probability threshold was achieved 21 minutes after injection of the agent, when a 1.97% PARACEST signal was observed. The presat-flash method was applied to a DCE MRI study of liver tissue (Figure 4.7), which had a relatively shorter T 1 relaxation time of 1.27 s. This in vivo study used Eu(2), a derivative of Eu(1) that included two hydrophobic o-benzyl functionalities (obzl ligands) to improve the retention time in liver tissue. Furthermore, amounts of injected agents were times higher than used in most pre-clinical and clinical studies (typically mmol/kg), which also improved the tissue accumulation and lengthened the retention time of each agent. The 95% CNR probability threshold was achieved minutes after injection of the agent, when a 7.04% PARACEST signal was observed and the 99% CNR probability threshold was achieved minutes after injection of the agent, when a 22.09% PARACEST signal was observed. Each post-injection image was compared with the average of the pre-injection images to quantify PARACEST contrast, which required selective saturation at only one MR frequency. Using the average of the pre-injection images was particularly important for minimizing the effect of motion artifacts during in vivo liver MRI studies. The temporal resolutions were 80 sec per image and 76.8 sec per image, which greatly aided the visualization of contrast agent uptake

136 4.4. DISCUSSION The consequences of the theoretical analysis are best assessed by comparing Eq. [4.7] under practical conditions with the theoretical steady-state PARACEST effects calculated with Eq. [4.4b]. To improve the temporal resolution under the same saturation conditions, an acquisition scheme with a longer effective echo time may be employed. For example, A typical effective TE (the time for half of the echo train duration, which represents the TE time for acquiring the center k-space line) for a RARE spin-echo acquisition scheme may be as long as 320 msec. These conditions generate simulated PARACEST effects that are very close to the maximum values of simulated steady state CEST contrast for each value of T 1w (Figure 4.4; Table 4.1). For comparison, an analysis of a train of 1500 Gaussian pulses during a T S of 3.4 sec shows only very minor additional improvement in CEST contrast for each T 1W value, which demonstrates that a T s of 2.27 sec had effectively reached the steady state. The loss of PARACEST contrast after the end of T S is strongly dependent on T 1W (Figure 4.4; Table 4.1). For a T 1W of 3 sec, the PARACEST contrast suffers only a relatively minor loss of 10.0% at the end of an effective TE of 320 msec. However, for a T 1W of 1 sec, the longitudinal relaxation is too fast to avoid a more substantial 27.4% loss of PARACEST contrast. To improve the temporal resolution while retaining a short echo time, a shorter saturation time may be employed. For example, a FLASH sequence with Gaussian saturation pulses that fill TR times of 100, 200 or 300 msec can be evaluated with a T 1W of 1 sec and the same properties of 20 mm Eu(1) as model agent. These conditions quickly achieve a steady state of saturation of pool W within 30 repetitions with a TR as short as 100 msec, and effectively maintain the steady state PARACEST by compensating for relaxation loss

137 Figure 4.6. The in vivo DCE MRI study of tumor tissue with a presat-rare MRI method with a RARE factor of 16 (80 sec/image). The contrast agent was Tm(1). A: A schematic of the images shows the difference between post-injection images and the average of the pre-injection images at different time points. B: a representative image with marked tumor ROI area for quantitative analysis. C: the dynamic change in PARCEST contrast of tumor ROI. D: the corresponding CNR of the PARACEST contrast for the same ROI. Horizontal lines represent the 95% and 99% probability levels that the CNR was generated from the PARACEST agent

138 Figure 4.7. The in vivo DCE MRI study of liver with a presat-flash MRI method with a 300 ms TR (76 sec/image). The contrast agent was Eu(2). A: a schematic of the images shows the difference between post-injection images and the average of the pre-injection images at different time points. B: a representative image with marked ROI area for quantitative analysis. C: the dynamic change in PARCEST contrast of liver ROI. D: the corresponding CNR of the PARACEST contrast for the same ROI. Horizontal lines represent the 95% and 99% probability levels that the CNR was generated from the PARACEST agent

139 of PARACEST contrast prior to each acquisition (Figure 4.5). The simulated PARACEST effects were very similar to the theoretical maximum steady state value. Most importantly, these assessments of Eqs. [4b], [5] and [7] demonstrate an improved PARACEST effect by using a presat-rare sequence relative to a short TR presat- FLASH sequence for environments with long T 1 relaxation times of 3 sec. These assessments also demonstrate an improved PARACEST effect by using short TR presat- FLASH relative to presat-rare for short T 1 relaxation times of 1 sec. Therefore, the best PARACEST MRI method depends on the T 1W of the endogenous tissue and T 1 relaxation caused by the PARACEST agent. The T 1 relaxation times of the PARACEST agent and the sample dictate the selection of the fast method that will produce the greatest PARACEST effect. A presat-rare method provides better PARACEST contrast when the T 1 relaxivity of the agent is low and the endogenous T 1 relaxation time of the water pool is long. Although the CNRs of the presat-rare methods were lower than the CNR of the presat-gre method for the Eu(1) phantom study, a RARE factor up to 32 still provided a PARACEST CNR that was greater than the 99% probability threshold. In this study, the CNR efficiency was greatest when a RARE factor of 32 was used, and represented a 2.14-fold improvement relative to the presat-gre method. Under similar long T 1 relaxation conditions, the presat-flash method produces a lower PARACEST contrast relative to the presat-rare method, which results in a lower CNR and CNR efficiency. The presat-flash method provides better PARACEST contrast when the T 1 relaxivity of the agent is high and/or the endogenous T 1 relaxation time of the water pool is short. The phantom study of 5 mm of Tm(1) detected % PARACEST contrast when

140 presat-flash was used, while presat-rare detected almost no PARACEST contrast. The presat-flash sequences detected PARACEST contrast that was a comparable 67-84% of the PARACEST contrast detected by the presat-gre sequence, which indicated that the low PARACEST effect detected by presat-flash was predominantly due to the relatively low concentration of Tm(1). A similar comparison of results showed that the presat-flash acquisitions generated a CNR that was a very low (10-15% of the CNR of presat-gre), and yet the CNR efficiencies of presat-flash were a more comparable 61-83% of the CNR efficiency of presat-gre. Although the PARACEST contrast of % can be visualized in the image, the CNRs of all tested pulse sequences were much lower than the 2 2 or 3 2 thresholds for assigning the contrast to the presence of the PARACEST agent with 95% or 99% probability, respectively. 20 These analyses here show clearly that the exact PARACEST MRI pulse sequence used has a strong influence on the CNR and CNR efficiency, and that one must take into account the full CNR of the image, not simply the contrast level. For example, a presat- RARE with a RARE factor of 64 detected a 33.6% PARACEST effect from Eu(1), which was greater than the 29.8% PARACEST effect detected with presat-flash using a TR of 300 msec. These results may lead to the conclusion that this particular presat-rare experiment produced better results than the presat-flash experiment. Yet the statistical significance of MRI contrast is a function of CNR and not just the contrast level. This presat-rare experiment yielded a CNR of 2.46, which was less than the 3.26 CNR of the presat-flash experiment, leading to the proper conclusion that the presat-flash experiment outperformed the presat-rare experiment. Therefore, reports of PARACEST MRI should include the CNR and a comparison to the 95% or 99% CNR

141 probability levels (or other CNR levels of statistical significance as needed), along with the pulse sequence and parameters. These reports of CNRs will improve on the current practice of considering a relatively arbitrary 5% contrast threshold as sufficiently 7, 31 significant for PARACEST MRI studies. As with other RARE methods, 32 the spatial resolution in a presat-rare image is also compromised as a consequence of the T 2 relaxation over the course of the echo train, especially when the T 2 relaxation time is relatively short. A RARE factor of 16 was used for in vivo studies in this report to avoid potential T 2 relaxation effects during PARACEST MRI acquisitions. In these cases, presat-flash can be used as an alternative to presat-rare. The improved temporal resolutions of presat-rare and presat-flash provided the opportunity to perform practical in vivo PARACEST DCE MRI studies. Although both the endogenous T 1 relaxation time and the T 1 relaxivity of the agent can dictate the choice of the imaging sequence, the in vivo PARACEST MR imaging of the tumor with Tm(1) indicated that the endogenous T 1 relaxation time is more critical to this choice of sequence. The acquisition of multiple images before and after the injection of the PARACEST agent over the DCE MRI study provided several important advantages. Foremost, the pre- and post-injection images with selective saturation only at Δω were compared to quantify PARACEST contrast. This obviated the need to acquire images with selective saturation at Δω, which inherently doubled the temporal resolution and the CNR efficiency. For example, the temporal resolution of a single in vivo presat- RARE image with single average was 80 s/image, and the temporal resolution of a single in vivo presat-flash image with two averages was 76.8 s/image, which were half of the

142 times required for comparable in vitro studies that required saturations at Δω and -Δω. Selective saturation only at Δω maintained steady-state saturation during presat-flash, which was critical for accurate quantification of PARACEST contrast with this MRI pulse sequence. Finally, magnetic field inhomogeneities within in vivo animal models may skew the frequency profile of the water pool, so that direct saturation of water may not be cancelled by comparing M S (Δω) with MS(-Δω). This method of subtracting postcontrast images from pre-contrast images effectively cancels contrast mechanisms and other effects that are identical before and after the injection. The agent can cause bulk magnetic susceptibility effects, line broadening due to chemical exchange, and changes to the magnetization transfer effect. However, each of these effects is negligible for saturations applied at +50 ppm and -51ppm as conducted with these PARACEST agents. The first image of a series of presat-flash acquisitions is used to establish steady-state selective saturation, and should not be used for quantitative PARACEST MRI. With this in mind, the temporal resolution of a presat-flash sequence is maximized when applying continuous acquisitions for dynamic PARACEST MRI studies. In addition, this interleaved saturation-acquisition scheme has a short saturation period relative to the total saturation-acquisition time, and therefore has a relatively lower SAR. Very recently, an analogous method for indirectly detecting PARACEST agents, termed on-resonance PARACEST MRI or OPARACHEE, has been demonstrated in vivo with a temporal resolution of s. 33 To the best of our knowledge, this study represents the first demonstrations of in vivo off-resonance PARACEST MRI, which were greatly facilitated by using the PARACEST MRI methods with improved temporal resolutions. The long T 1 relaxation time of the tumor tissue allowed the use of the presat-rare pulse

143 sequence, which generated a maximum 6.7% PARACEST effect. The relatively short T 1 relaxation time of liver tissue required the use of a presat-flash pulse sequence, which generated a maximum 28.5% PARACEST effect (this high level was also facilitated by using liver-avid Eu(2)). More importantly, the CNR of the PARACEST contrast in the in vivo tumor and liver tissue studies reached the 2 2 threshold for assigning the contrast to the presence of the PARACEST agent with 95% probability after 21 minutes and minutes, respectively (Figures 4.6D and 4.7D). To reach this CNR threshold at earlier time points, methods to increase image SNR are just as critical as methods to improve PARACEST contrast. For example, acquisition gating may remove cardiopulmonary motion artifacts from MR images of liver tissue, which would improve the CNR level without requiring the generation of a greater PARACEST effect. These in vitro and in vivo studies have demonstrated that a PARACEST MRI acquisition method must be tailored to the properties of PARACEST agents, in addition to other properties of the MRI experiment such as MRI relaxation of the tissue of interest. Other MRI pulse sequences with faster acquisition schemes or shorter saturation periods may be more applicable for particular biomedical applications than the specific RARE and FLASH sequences used here. However, the underlying principles used to establish these sequences are still applicable to the design of more general PARACEST MRI pulse sequences for in vivo applications CONCLUSION Two strategies have been developed that improve the temporal resolution of PARACEST MRI. In one case, a relatively long saturation is followed by a rapid readout

144 Alternatively, one could use relatively short saturation periods with interleaved short acquisition periods. Under ideal environmental T 1 relaxation conditions, the temporal resolution can be reduced to 31.2 sec per image using these methods with high spatial resolution and improved CNR efficiency in PARACEST contrast. Under less favorable environmental T 1 relaxation conditions, the temporal resolution can be reduced to 51.2 sec per image with preserved CNR efficiency in PARACEST contrast. These temporal resolutions enabled the first demonstrations of in vivo PARACEST DCE MRI studies. Therefore, these PARACEST MRI methods with improved temporal resolutions may accelerate the development of other PARACEST MRI methods and the eventual translation of PARACEST MRI to clinical applications

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150 CHAPTER 5 FUTURE STUDIES As discussed in the conclusion of Chapter 1, the majority of the current and future developments of PARACEST based MRI molecular imaging must address SPECIFICITY, SENSITIVITY and TEMPORAL RESOLUTION. In Chapter 2, we have proposed an irreversible responsive PARACEST contrast agent, Yb-DO3AoAA, for the detection of NO. This contrast agent possesses a highly specific functional group that undergoes a chemical reaction only in the presence of NO and O 2, and therefore is able to detect NO with high SPECIFICITY. In Chapter 3, we have proposed a new MRI method to measure ph based on the same contrast agent by utilizing the two intramolecular PARACEST signals. This method depends on the exchange rates of the two PARACEST sites, which are determined only by local ph when temperature is constant. Therefore, the specificity and accuracy of the ph measurement MRI methods are greatly improved. In Chapter 4, the improvement of TEMPORAL RESOLUTION of PARACEST was investigated to fit the requirements of in vivo dynamic studies. These studies demonstrated the feasibility of PARACEST based MRI molecular imaging and built the foundation for a PARACEST technique platform. However, in spite of the success in improving specificity and acquisition speed, poor SENSITIVITY may still limit the application of PARACEST based MRI molecule imaging techniques for preclinical studies and translation to clinical studies. Additionally, other challenges may also need to be considered when developing PARACEST for in vivo applications. These difficulties exist in intracellular delivery of agents, the selectivity

151 of detection, in vivo pharmacokinetics, new pulse sequence implementation, safety/rf power deposition, and quantification of PARACEST. In this chapter, these challenges will be discussed in detail and potential solutions will be proposed to lead to future studies IMPROVEMENT IN SENSITIVITY The inherent low sensitivity is one of the major drawbacks of MRI molecular imaging. Recently the sensitivity of MRI signals has been dramatically improved by several orders of magnitude by using hyperpolarized noble gases. 1 However, the technical hurdles in producing hyperpolarized agents and applying them to in vivo animal models is still formidably challenging and cannot be immediately transferred to most in vivo studies. More practical and widely used sensitivity improvement strategies include the utility of drug delivery systems and enzyme based amplification. A brief introduction to each strategy will be followed by the discussion of the implications to our future studies in subsections and Other amplification strategies, such as a biotin-avidin system and a multiple-step targeting and prelabeling system, 2 nonetheless are also potentially important. However, we won t discuss them in this chapter due to the lack of immediate impact Nanoparticle Delivery Systems For all of the contrast agents that currently have been developed, one of most straightforward ways for generating high contrast is to increase the local concentration of the contrast agent by increasing the amount administered. However, a higher

152 concentration may significantly increase the osmotic pressure and potential risk of cytotoxicity. Therefore the ability of increasing concentration to enhance sensitivity is very limited. Alternatively, conventional drug delivery systems can be used to deliver high loads of imaging probes to markedly improve the sensitivity of MRI molecular imaging methods by several orders of magnitude without serious side effects. These delivery systems are also referred to as payload systems. Additionally, a delivery system can also be effectively modified to improve the targeting ability, pharmacokinetics, biocompatibility and safety, and therefore lead to further improvement of contrast agent effects. Drug delivery systems have been developed for decades with enormous impacts on the improvement of therapeutic efficacies of many existing drugs. Drug delivery systems that carry small molecules, nucleic acids, peptides or proteins to disease sites, can be classified into macroscale (>1 mm), microscale ( μm) or nanoscale (100 1 nm), based on the size of the implanted or injected devices. Nanoscale delivery systems appear to be the most interesting category with numerous advantages. The sizes of nanoparticles or carriers are typically ,000 times smaller than cells but similar to large biomolecules (e.g. proteins). The small size allows for quick circulation and absorbance by cells with a sufficient load capacity for delivering agents. 3 Indeed most widely used nanocarriers have a load capacity of tens to thousands for small molecule agents (Table 5.1). Recently, nanocarriers have been successfully used for diagnostic imaging with the ability to markedly boost contrast by delivering a sufficient amount of imaging agents on site. As a consequence, the low sensitivity of MRI contrast agents can become relatively sensitive, even for single cell detection without compromising cell proliferations

153 Currently, the most popular nanoparticle scaffolds for MRI molecular imaging are dendrimers, 5,6 polymers 7-9 and liposomes 10, 11 (Figure 5.1). A liposome (Figure 5.1a) is defined as a spherical vesicle composed of a phospholipid bilayer membrane and an aqueous solution core. Liposomes have been widely used as pharmaceutical carriers for decades due to the high capacity and the feasibility of targeting diseased tissues. 16 To the best of our knowledge, there are at least 11 clinically approved liposome based therapeutics and many others in clinical trials. 17 Recently, a responsive liposome based contrast system was also proposed and tested in vitro, 18 implying that more specific and sensitive liposome based contrast media can be developed for future studies. However, the fast clearance of liposomes from the blood stream by phagocytes often has to be taken under consideration for in vivo applications. 19 The big diameter of a liposome particle (200nm) also limits the delivery of imaging agents to targeted tissues due to the low vascular permeability, especially for those with small vasculatures (~ nm). The polymer/copolymer (Figure 5.1b) drug delivery system is another widely used and efficient technique platform. 9 A polymer system has the feasibility to manufacture targeting, biocompatible and biodegradable drug/imaging probe delivery system. 20 The size and shape of polymer systems may vary corresponding to the hydrophobic/hydrophilic characteristics, which can be adjusted by the modifications of side chains. Moreover, the feasibility of bifunctional polymer systems has recently been demonstrated, 21 which will greatly promote the therapeutic efficacy by spatially and temporally trafficking therapeutic drugs in the presence of imaging agents. The dendrimer system is another very promising nanoparticulate drug delivery vehicle. The drug/imaging probes delivered by using dendrimer systems has emerged as a hot

154 Table 5.1. Nanoparticle systems for MRI molecular imaging and their typical characteristics Polymer/copolymer Polymer micelles Liposome Dendrimers Protein carrier Low density lipoprotein (LDL) nanoparticles Prototype HPMA copolymer Ploy-(L-glutamic acide) paramagnetic liposome 12 PAMAM generation 5 Load of paramagnetic metals (i.e. Gd 3+ and Fe) Varied Molecular weight ~60KDalton(HPMA) ~28 Dalton (poly(lglutamic acid) Size nm ~50,000 ~MegaDalton 200nm -128 ~60,000 Dalton 5-8 nm 19 /65 ~80,000Dalton (albumin) ~52,000 Dalton (poly-l-lysine) 8 nm (albumin) LDL ,400-3,900 KDalton ~22 nm SPIO nanoparticle MION ~2,000 ~MegaDalton 20 nm Nanoparticle emulsions Carbon nanotube ACPL(antibodyconjugated Albumin/Poly-Llysine Gd-DTPA-SA- Gdperfluorocarbon ~74,000 ~MegaDalton 250 nm nanoparticles 14 Gd 3+ n@ustubes 15 <10 per cluster(2-5 nm) nm

155 area due to the well-defined architecture and unique characteristics of dendrimer particles. 22 Dendrimers are highly branched and multivalent macromolecules in spherical structures (Figure 5.1c). A number of preclinical studies have demonstrated dendrimer systems with high delivery efficiency (e. g. >100 functional surface amino groups that can covalently link to drugs or imaging probes 23 ) and excellent tumor targeting ability for MRI molecular imaging Figure 5.1. The schematic representation of a) liposome, b) polymer micelle nanoparticle and c) dendrimer as the delivery systems for imaging probes (shown as the red dots or balls in the structures).the architectures, relative load efficiencies and relative size of each nanoparticle are depicted in the figure. All these three scaffolds can be used to deliver PARACEST agents and circumvent the low sensitivity limitations of PARACEST contrast agents for in vivo applications. For instance, a very sensitive liposome based PARACEST agent, or LIPOCEST agent, has been proposed and demonstrated in vitro. 11 The dendrimer system also has been recently demonstrated with the ability to improve the efficiency of PARACEST contrast system

156 However, considering the expertise in chemistry in our laboratory, polymer scaffolds have been chosen to improve the sensitivity of Yb-DO3AoAA. A HPMA co-polymer platform is currently under development, which will allow the covalently attachments of the DO3-oAA to the HPMA polymer structure (Figure 5.2). The expected enhancement of sensitivity will be 10- to 100-fold for Yb-DO3oAA. Moreover, this polymer platform can be easily extended to other PARACEST contrast agents, just by simply following the same procedure. CH 3 CH 3 H 2 C C x H 2 C C y C O C O NH CH 2 NH CH 2 HO O O N N Yb 3+ N N OH O O HC OH HC HN H 2 N NH CH 3 CH 3 Figure 5.2. The schematic representation of an Yb-DO3AoAA HPMA co-polymer system as a highly sensitive NO detecting agent Enzyme Based Amplification Another effective strategy to boost sensitivity of MRI contrast agent for biomarker detection is based on enzymatic catalysis. The smart or responsive MRI contrast agents

157 can be designed to respond enzymes as biomarkers. These enzymes are able to continuously convert substrates (e.g. responsive contrast agents) to products as long as the substrates are presented and the turn-over of enzyme is relatively quick. Consequently, the products (with either the turned on or turned off MRI signals) will accumulate to a significantly high level by many orders of magnitude compared to the concentration of the functional enzyme. The functional enzyme can be either a target or a reporter. This strategy has been quickly adapted to MRI molecular imaging for the studies of gene expressions and enzyme disease biomarkers based on a variety of enzyme systems, such as β-galatosidases, 27 proteases, 28, 29 and peroxidase. 30 The NO responsive PARACEST agent,yb-do3oaa, that has been described in the Chapter 2 has limited ability for in vivo detection of NO mainly due to its ultra low concentration (nm μm) and fleeting lifetime. Although a nanoparticle system can be used to deliver sufficient agent with a long life time, alternatively, such a limitation can also be circumvented by applying the proposed PARACEST agent agent to monitor appropriate NO synthase activity. For humans, there are three major NO synthases, i.e. inducible NO synthase (inos), epithelial nitric oxide synthase (enos) and neuronal nitric oxide synthase (nnos) (Figure 5.3). inos and nnos are predominantly located in the cytosol, while enos is membrane-associated and releases the NO to the extracellular space. While enos and nnos constitutively produce NO at ultralow concentration (e.g nm in vascular wall 31 ) corresponding to the maintaining and modulating the physiological processes, the inos can produce up to 10 μm NO during the pathophysiological processes. 32 Moreover,

158 the inos is a very important biomarker involving in pathological processes of many diseases such as cancer 33, arthritis, 34 and cystic fibrosis (CF). 35 Upon the accumulation of product by continuously reacting with a low concentration of NO, Yb-DO3AoAA is able to detect inos activity based on the enzyme amplification strategy. This leads to the potential in vivo applications of the proposed NO detecting contrast agent. To test the feasibility and applicability of in vivo inos activity detection by Yb DO3AoAA, an appropriate transgenic animal model with significant changes in inos activity therefore needs to be obtained to validate the feasibility and quantification of this approach. The first trial has been conducted to detect inos in the nasal epithelium of a live CF mouse because CF mice have less or no inos than healthy mice. However, the preliminary study was not successful mainly due to the technical hurdles in PARACEST agent administering and the low water content in the nasal epithelium. Our future works therefore include exploration of other appropriate in vivo models and hence expanding Yb-DO3oAAA project to realistic in vivo applications. Figure The three major isomers of nitric oxide synthase and their functions

159 5.2. IMPROVEMENT OF CELLULAR INTERNALIZATION Most biomarkers such as genes, RNAs, enzymes and metabolites are located in intracellular space. Therefore the molecular imaging of these biomarkers requires the cross-membrane transportation of imaging probes into the cytosol (cytoplasm), which is commonly referred to as cellular internalization. Compared to methods that target biomarkers located either in extracellular space or on the cell membrane surface, intracellular imaging methods can be developed for highly diverse biomarkers and therefore are more informative to diagnose pathophysiological processes. In addition, the probes are entrapped and accumulated in the intracellular space after cellular internalization. Consequently, imaging signals can be amplified with relatively high temporal stability (which overcomes the low temporal resolution problems of MRI). Unfortunately, most of the currently available MRI contrast agents only have limited cellular internalization ability. The further modification of contrast agents has to be taken under consideration in order to realize efficient intracellular delivery. Cells can uptake nutritional materials from the surrounding environment through endocytosis and transportation (or active transportation, ATP-driven transportation, or membrane protein-mediated transportation). The former mechanism includes pinocytosis, phagocytosis and receptor mediated endocytosis (Figure 5.4). Pinocytosis, or celldrinking, is a process of non-specific uptake of fluid that contains solutes (soluble small molecules) through the formation of small vesicles ( 150 μm). Phagocytosis, or celleating, conversely carries large solid particles cross the cell membrane by the formation of large endocytic vesicles known as a phagosome. Phagocytosis is specifically performed by phagocytes such as macrophages, monocytes, dendritic cells, and

160 granulocytes. 36 The fluid phase endocytoses have been successfully employed to intracellularly internalize imaging probes including small molecular weight Gd(III) chelates, 37 and nanoparticles (MION). 38, 39 However, the strategies based on fluid phase endocytosis generally lack specificity and suffer from low efficiency of cell internalization for non-dividing cells, 40 which leads to a limited applicability for many in vivo applications. Figure 5.4. The schematic representations of a) pinocytosis b) phagocytosis c) receptor mediated endocytosis and d) transporter delivery mechanism for cell to uptake extracellular molecules

161 Receptor-mediated endocytosis is a more specific process to carry materials across membranes by receptor reorganization and facilitation. Receptor-medicated endocytosis is very promising for the in vivo delivery of imaging probes owing to high efficiency and the ability to specifically deliver contrast agents to targeted cells. Human transferrin receptor (htfr) 41 and folate receptor 24 have been primarily studied for internalization of contrast agents so far. Other receptors such as β-d-galactose receptors 42 and LDL receptors 13 are also reported as efficient internalization systems for MR molecular imaging. The targeting property of repceptor-mediated endocytosis is very desirable in molecular imaging studies. Therefore a number of nanoparticle delivery platforms, such as dendrimers 24, 25 and SPIO, 41 have been reported to conjugate with receptor ligands to further improve efficiency of the targeted delivery. Recently, many studies have shown that the intracellular delivery of imaging probes can be more efficient by conjugating to a cell penetrating peptide (CPP). 40 The CPPs are referred to as a group of small peptides (10-15 amino acids) with cell membrane translocation properties. The Tat peptide from the HIV-1 Tat protein and the antennapedia-derived peptide from the drosophila antennapedia homeodomain are the two most widely used CPPs, 43 while other CPPs including poly-cations, Tranportan and VP22 have also been studied for a variety of applications. The mechanism of a CPP to facilitate the cross-membrane delivery is still not very clear but likely through endocytosic processes. 43,44 CPP (e.g. Tat) has been demonstrated with markedly improvement of intracellular delivery of MRI contrast agents including small molecular weight agents 45 and nanoparticles. 46 To conjugate with CPP is often considered as one of

162 most practical ways to improve the efficiency of cellular internalization, especially for small molecule probes. Another growing field of cellular internalization is the use of transporter systems (Figure 5.2.d). A specialized protein transporter can facilitate the cross-membrane delivery of large amount of specific nutritional materials with very high efficiency. Highly efficient transporter-based internalization systems for molecular imaging have been explored by using an organic anion transport protein (OATP) for hepatocytes 47 and a glutamine transporter for tumor cells. 48 This strategy is a worthwhile alternative to receptor mediated endocytosis with a highly targeting ability for in vivo applications. The understanding of potential cellular internalization mechanisms is extremely important for our future studies to allow the further improvement of detecting PARACEST agents. For instance, we have developed a NO responsive PARACEST agent as shown in Chapter 2. One of the important applications of this agent is to investigate the inos activity as discussed in Section 5.1. However, both the inos and produced NO are located in intracellular spaces. The detection of inos therefore requires the cross-membrane delivery of Yb-DO3AoAA by one of the mechanisms that have been introduced above. Considering the feasibility of chemistry in our laboratory and potential future applications (e.g. for tumor angiogenesis), the strategy that employs cell penetrating peptide (CPP, i.e. Tat) appears the best choice for the delivery Yb-DO3oAA. Moreover, Tat can also be conjugated with different payload systems (i.e. polymer, dendrimer or liposome) as a technique platform to combine high loads of PARACEST agents and a CPP intracellular delivery system. Such a delivery system can be used for the detection of intracellular inos activity with high sensitivity in the near future

163 The efficient internalization of Yb-DO3oAA also endows the intracellular ph (ph i ) measurement ability. By conjugating with a Tat peptide, the new contrast agent Yb- DO3oAA-Tat is ready to enter the cytosol and therefore the same protocol as described in the Chapter 3 can be used to measure intracellular ph. This will lead to another potential project, the simultaneous detection of extracellular ph and intracellular ph. To realize it, two PARACEST agents, Tm-DO3AoAA and Yb-DO3AoAA-Tat, may be administrated at the same time and selective imaging of the PARACEST effects of the Tm agent and Yb agent may be performed. The Yb-DO3AoAA-Tat is the intracellular ph indicator while Tm-DO3AoAA is only able to report the extracellular ph. Because the PARACEST offsets of Tm-DO3AoAA and Yb-DO3AoAA are distinct, the selective detection of each PARACEST signal is feasible, which enables the simultaneous intracellular and extracellular ph determination. In practice, the extracellular Yb- DO3AoAA-Tat has been to be considered in order to accurately measure the intracellular ph. A dynamic study would be useful to distinguish the signals from intracellular and extracellular Yb-DO3oAA due to the lengthened retention time after the agents are entrapped in the intracellular space. A more accurate measurement can be endeavored by the use of a CPP conjugated ph-responsive polymer delivery system. The encapsulated Yb-DO3oAA in a polymer system will not show PARACEST contrast because of the limited access to the outside bulky water environment. Only after the ph-responsive polymer is degraded in the endosomes (~ ph 5) through the CPP facilitated internalization, the encapsulated Yb-DO3oAA will be released to the cytosol and be exposed to bulky water to trigger the PARACEST contrast

164 5.3. IMPROVEMENT OF IN VIVO IMPLEMENTATION BY USING A CONTROL AGENT The in vivo application of a responsive PARACEST agent may need an unresponsive control agent due to the need to distinguish pharmacokinetics from the presence of the biomarkers. For example, the absence of a detectable PARACEST effect after administration Yb-DO3oAA may be due to insufficient concentrations of this imaging agent in the tissue of interest, rather than the response to NO. Owing to the ability of selective detection in a PAARCEST study, an unresponsive contrast agent with distinct PARACEST offsets can be combined with the responsive PARACEST agent by a means of direct coupling of the agent or conjugations of the agents to a polymer scaffold. Therefore the pharmacokinetics of the responsive PARACEST agent can be indirectly monitored by the PARACEST signal of the control agent, which enables the tracking of the pharmacokinetics and biodistribution of responsive agent when applied to live subjects. 49 This approach has been demonstrated in vitro in Chapter 2, where the NOunresponsive Tm-DOTAMGly was subjected to the same reaction conditions as Yb- DO3oAAto obtain MR images with selective saturation at -51 ppm corresponding to the PARACEST effect of Tm-DOTAMGly. The result showed that no difference was detected between Tm-DOTAMGly before and after exposure to NO, which proved that this unresponsive PARACEST MRI contrast agent can be used to monitor the local concentration of the imaging agent during in vivo applications. In spite of the feasibility showed within this in vitro study, the in vivo implementation of such an approach may still be challenging. Differential pharmacokinetics and biodistributions can complicate the evaluation of responsive contrast agents during in

165 vivo molecular imaging applications. 50 To ensure that the in vivo pharmacokinetics of responsive and unresponsive PARACEST agents are identical, this strategy must be extended by covalently linking both agents or by packaging both agents in the same delivery vehicle. In practice, the approach of using of a delivery vehicle is preferable because high sensitivity can be achieved at the same time. The potential payload systems that can be used to conjugate or encapsulate both agents are liposomes, dendrimers and polymers. Since we have proposed that a polymer system can be used to improve the sensitivity of Yb-DO3oAA in the preceding section, the same method can be used to conjugate the control agent to the same copolymer system (Figure 5.5). CH 3 CH 3 CH 3 C H 2 C x H 2 C C y H 2 C C z O C C O C O COOH NH CH 2 NH CH 2 NH CH 2 HO O O N N + Yb 3 N N OH O O HN O O N N + Tm 3 N N NH O O CH CH 3 HC OH CH 3 HC CH 3 HN H 2 N NH HN NH COOH HOOC Figure 5.5. The schematic representation of the proposed Yb-DO3AoAA and Tm- DOTAMGly bifunctional HPMA co-polymer. This system can be used to detect NO with high sensitivity by the PARACEST effects of Yb-DO3AoAA, with traceable pharmacokinetics by the PARACEST effect of the conjugated NO-unresponsive Tm- DOTAMGly

166 5.4. IMPROVEMENT OF SELECTIVITY Reduction of Direct Water Saturation Although PARACEST is considered as a very promising MRI technique, most of the PARACEST studies reported so far still have been within in vitro models. Based on the pioneering works in the last decade, in vivo studies have been recently endeavored by Vinogradov A. and her coworkers with an on-resonance approach, 51 and by our group (unpublished data) with an off-resonance approach as shown in Chapter 4 (The onresonance approach refers to the application of Waltz decoupling RF pulses at the water resonance offset in order to obtain higher sensitivity in an RF power-efficient way, while the off-resonance approach refers to the application of a saturation RF pulse at the PARACEST offset and has been currently used for most the PARACEST studies due to the ability of selective detection. However, the in vivo application of PARACEST is still a formidable challenge. One of the major concerns is the direct saturation of the bulk water pool. The PARACEST signals from the direct saturation of the agent may be overlapped with the broad signals generated from direct saturation of the water and may become difficult to distinguish. This effect is demonstrated in Figure 5.6, where the broad water notch (±30 ppm with respect to water resonance offset, for a 10-μT RF saturation) was caused by the direct water saturation in a mouse tumor region of interest (ROI). The significant broader water spillover within an in vivo model is mainly due to the high content of macromolecules. Therefore the in vivo application of Yb-DO3AoAA will be seriously limited by the saturation power (<5 μt) in order to avoid the interferences from direct water saturation. One of the practical solutions is to use other Ln(III) with higher PARACEST frequency offsets. Tm(III) therefore should be used instead when higher

167 saturation power is desired. Another solution is to use more selective pulse (e.g. sinc and eburp) and a smaller bandwidth (<0.5 ppm) to avoid the extra power spillover. If the asymmetric analysis is not required (e.g. DCE PARACEST), an adiabatic Delay Alternating with Nutations for Tailored Excitation (DANTE) pulse train 52 also can be used for saturation with a number of advantages, such as the higher efficiency, insensitivity to B 1 inhomogeneity and sharp saturation profiles. 53 Figure 5.6. CEST spectra of a mouse tumor ROI with 10 μt and 5 μt RF saturation power. The selective saturation pulse used was a train of 1000 Gaussian pulses with a bandwidth of 3 ppm. For the display purpose, the CEST spectrum obtained by 10 μt was slightly shifted up by 5% Reduction of the Interferences of Multiple PARACEST Sites Yb-DO3AoAA possesses two PARACEST offsets, -11 ppm and +8 ppm, corresponding to the amide proton and amine protons respectively. If an asymmetric analysis is used,

168 significant interferences can be caused due to the spillover of saturation of one PARACEST offset to its neighbors, i.e. -11ppm to -8 ppm and +8 ppm to +11 ppm. The in vitro NMR data can be corrected by fitting a three-lorentzian-line model to account for the interference. However, such an approach may take too long to obtain and process data and therefore limits potential in vivo applications. One of the alternatives is to use another Ln(III) complex with distinguishable and highly shifted PARACEST offsets. Our study has demonstrated that Tm has 3 PARACEST offsets at -52 ppm, -36 ppm and +16 ppm (Figure 5.7). Figure 5.7. The CEST spectra of Tm-DO3AoAA at ph 6.93 and ph 7.4. The spectra were obtained by using a 3 second CW pulse (12 μt) at 19 o C on 600 MHz Varian NMR spectrometer. The low sensitivity of Tm-DO3AoAA can be greatly improved by polymerization with the same procedure described in the preceding part (Section 5.1 and Figure 5.2)

169 Moreover, the polymerized Tm-DO3AoAA will be more favorable for the intravenous administration owing to the improved sensitivity and broader distribution of PARACEST offsets. In spite of the success of subcutaneous injection of Yb-DOA3oAA in our in vivo ph study (Chapter 3), the intravenous administration of polymerized Tm-DO3AoAA will be ultimately more favorable for pre-clinical and clinical studies IMPLEMENTATION OF HIGH TEMPORAL RESOLUTION METHODS A PARACEST MRI acquisition method must be tailored to the properties of PARACEST contrast agents, in addition to other properties of the MRI experiment such as T 1 MRI relaxation times of the tissue of interest. In addition to the RARE and FLASH sequences that have been demonstrated in Chapter 4 with great improvements in temporal resolution for PARACEST studies, other MRI pulse sequences with fast acquisition schemes or potential short saturation periods can be employed for the high temporal resolution PARACEST MRI under different circumstances (e.g. field strength and scanner ) (Table 5.2). These pulse sequences include Echo-planar Imaging (EPI), Fast Imaging with Steady state Precession (FISP)/ Gradient-Recalled Acquisition in the Steady State (GRASS), and parallel imaging. These methods may be more applicable than RARE and FLASH sequences for particular biomedical applications that require higher temporal resolution. The developments in implementing these new MRI pulse sequence methods to PARACEST study can further boost temporal resolution of dynamic PARACEST methods. A preliminary result has demonstrated that a higher temporal resolution (~3s per CEST image) can be achieved by using a FISP-CEST method developed by Flask and

170 his colleagues (Figure 5.8). The higher temporal resolution PARACEST MRI methods will greatly facilitate the exploration of in vivo applications and translation to clinic. Table 5.2. MRI pulse sequence candidates for high temporal resolution PARACEST MRI methods and their preferable main magnetic field strength (B 0 ) Multiple echo approach Fast steady-state approach Typical TR and TE Favorable B 0 RARE TR: ms No limit TE:<15 ms TR: FLASH ms No limit TE:<15 ms EPI TR: ms 1.5T TE:<15 ms True-FISP TR<50 ms TE:<15 ms 1.5T FISP/GRASS TR<50 ms TE:<15 ms No limit Parallel imaging Varied No limit

171 Ms/Mo Saturation Frequency (Hz) EuDOTAMGly PBS Figure 5.8. The CEST spectra of 40 mm Eu-DOTAM-Gly and PBS by the FISP-CEST method. A modified FISP pulse sequence (TR/TE=4/2ms, res=230 μmx230μmx1mm tip angle=60 o ) that included a CEST preparation scheme has been used with a Bruker Biospec 9.4T small animal scanner. The CEST preparation scheme consisted of 1000 selective Gaussian pulses (2.25ms, 20 μt) REDUCTION OF RF POWER DEPOSITION The translation of established methodologies to clinical scanners will be the ultimate future goal for preclinical PARACEST studies. In addition to other technical hurdles, the high RF power deposition involved in PARACEST may be beyond the clinical criteria and therefore hinder the translation process. The reduction of RF power thus will be an essential part of future studies in order to expand PARACEST to clinical studies

172 The Specific Absorption Rate (SAR) is used in the clinic to describe the RF power deposition in a mass of tissue during an imaging experiment. The FDA has established guidelines to regulate the clinical allowable SAR in MRI studies. 54 When a RF pulse train is used in the saturation scheme, the SAR is proportional to the product of the number of RF pulses (n), square of saturation power (B 2 1 ) and pulse duration (pw) as shown in Eq. [5.1]. 55 SAR n B pw [5.1] 2 1 Because a long saturation scheme is often required to obtain sufficient sensitivity, the SAR may be significantly high and beyond the FDA guidelines for most PARACEST studies according to Eq. [5.1]. For the applications in higher magnetic fields, the higher SAR can be more problematic due to quadratic relationship with SAR and field strength. 56 Therefore the current preclinical PARACEST methods have to be modified to improve the efficiency of RF saturation without compromising the sensitivity/contrast. There are two potential strategies that can be used improve the RF efficiency. The first is to optimize the RF pulse and imaging parameters. In practice, shaped pulses are always used for a preclinical and clinical MRI scanner due to the hardware limitation to produce CW pulses. 57 Different shapes of RF pulses have been demonstrated with markedly different efficiencies with respect to the magnetization transfer contrast. 56 It has been suggested that tradeoffs exist between contrast, selectivity and RF efficiency for different shapes of RF pulses. In our preceding studies, the Gaussian pulses have been chosen because of the reported good saturation profile (i.e. strong with modest selectivity). 56 However, other pulse shapes can be used to improve the RF efficiency, especially when the selectivity can be further sacrificed for the PARACEST offsets that are far away from

173 the water resonance frequency (>50 ppm). For example, the square pulse has been reported to achieve 1.67 times higher MT contrast than the Gaussian pulse with the same peak B 1 power and pulse duration (pw). 56 Additionally, other parameters also should be optimized for a given pulse shape. One effective way is to decrease the saturation power (B 1 ) while increasing the pulse duration (pw) to maintain the contrast. 55 Nevertheless, the extension of imaging time (TR) also can effectively reduce SAR because less energy deposits in tissues per unit of time although the price of low temporal resolution has to be paid. The second approach is to apply the RF saturation only to the parts of acquisition in k- space. One of the obvious ways to reduce the RF power while mostly maintaining the contrast is to collect partial or half of k space data as same as the other partial Fourier transform MRI techniques. 58 Consequently, up to half of the RF energy deposit in imaging subjects can be reduced at the price of loss of SNR. In addition to the asymmetric partial Fourier transform approach, a keyhole approach, which symmetrically collects the low frequency components around the DC center in k space, can be used because it has been suggested that the low spatial frequency domain data contribute to the majority of contrast. 59 To illustrate the applicability of such a keyhole strategy in PARACEST, a set of pre contrast and post-contrast data have been processed both in the frequency domain and in the image domain (Figure 5.9). The contrast image produced by a keyhole approach (Figure 5.9h) demonstrated a very similar result as the original contrast image (Figure 5.9g)

174 Figure 5.9. The demonstration of keyhole approach for the reduction of RF power in a PARACEST study. Images a d are frequency domain (k-space) data and images e-h are image domain data by Fourier transforming from corresponding k-space dataa a-d respectively. The a and e are precontrast images, b and f are postcontrast images, c and g are contrast images without any manipulation, d and h are contrast images after low pass filtering around the DC center in k space (red line block). The images were mouse liver before and after intravenous administration of Eu(III)DOTA-OBS2Gly2COOH by using preset-flash method same as described in Chapter 4. However, the partial data collection in k space may lead to a poor image quality and reduced spatial resolution. Alternatively, the imaging scheme can still be kept in a full Fourier transform fashion while different weighting factors will be added to the parameters in CEST preparation schemes, such as saturation power, pulse number and pulse duration with respect to the different phase encoding step or k-space position. 60,