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Diagnostic MRI in Dogs and Cats
Diagnostic MRI in Dogs and Cats WILFRIED MAI, Dr. Méd. Vét., MSc, PhD, Diplomate ACVR, Diplomate ECVDI Professor of Radiology Section Chief of Radiology University of Pennsylvania School of Veterinary Medicine Philadelphia, Pennsylvania USA
CRC Press Taylor & Francis Group 6000 Broken Sound Parkway NW, Suite 300 Boca Raton, FL 33487-2742 © 2018 by Taylor & Francis Group, LLC CRC Press is an imprint of Taylor & Francis Group, an Informa business No claim to original U.S. Government works Printed on acid-free paper International Standard Book Number-13: 978-1-4987-3770-8 (Hardback) This book contains information obtained from authentic and highly regarded sources. Reasonable efforts have been made to publish reliable data and information, but the author and publisher cannot assume responsibility for the validity of all materials or the consequences of their use. The authors and publishers have attempted to trace the copyright holders of all material reproduced in this publication and apologize to copyright holders if permission to publish in this form has not been obtained. If any copyright material has not been acknowledged please write and let us know so we may rectify in any future reprint. Except as permitted under U.S. Copyright Law, no part of this book may be reprinted, reproduced, transmitted, or utilized in any form by any electronic, mechanical, or other means, now known or hereafter invented, including photocopying, microfilming, and recording, or in any information storage or retrieval system, without written permission from the publishers. For permission to photocopy or use material electronically from this work, please access www.copyright.com (http://www.copyright.com/) or contact the Copyright Clearance Center, Inc. (CCC), 222 Rosewood Drive, Danvers, MA 01923, 978-750-8400. CCC is a not-for-profit organization that provides licenses and registration for a variety of users. For organizations that have been granted a photocopy license by the CCC, a separate system of payment has been arranged. Trademark Notice: Product or corporate names may be trademarks or registered trademarks, and are used only for identification and explanation without intent to infringe. Library of Congress Cataloging‑in‑Publication Data Names: Mai, Wilfried, editor. Title: Diagnostic MRI in dogs and cats / [edited by] Wilfried Mai. Description: Boca Raton : CRC Press,  | Includes bibliographical references and index. Identifiers: LCCN 2017061268 (print) | LCCN 2017061795 (ebook) | ISBN 9781315121055 (General eBook) | ISBN 9781498737708 (hardback : alk. paper) Subjects: LCSH: Veterinary radiography. | Magnetic resonance imaging. | MESH: Magnetic Resonance Imaging--veterinary | Dog Diseases--diagnostic imaging | Cat Diseases--diagnostic imaging Classification: LCC SF757.8 (ebook) | LCC SF757.8 .D46 2018 (print) | NLM SF 757.8 | DDC 636.089/607572--dc23 LC record available at https://lccn.loc.gov/2017061268 Visit the Taylor & Francis Web site at http://www.taylorandfrancis.com and the CRC Press Web site at http://www.crcpress.com
Preface Contributors Abbreviations
SeCtion 1 CHAPTER 1
ix xi xii
PHYSiCS AnD teCHniCAL ConSiDeRAtionS – iMAGe oPtiMiZAtion GeneRAL PRinCiPLeS oF MAGnetiC ReSonAnCe iMAGinG
Wilfried Mai CHAPTER 2
iMAGe CHARACteRiStiCS in MRi AnD PRinCiPAL PULSe SeQUenCeS
Wilfried Mai CHAPTER 3
J. Fraser McConnell CHAPTER 4.1
oPtiMiZeD teCHniQUe: BRAin
Silke Hecht CHAPTER 4.2
oPtiMiZeD teCHniQUe: SPine
Ruth Dennis CHAPTER 4.3
GeneRAL FeAtUReS AnD oPtiMiZeD teCHniQUe FoR tHe MUSCULoSKeLetAL SYSteM
Amy R. Zalcman, Cristi Cook, and Wilfried Mai CHAPTER 4.4
teCHniCAL PARtiCULARitieS WitH LoW-FieLD iMAGinG
SeCtion 2 CHAPTER 5.1
MRi oF tHe BRAin ConGenitAL AnD DeVeLoPMentAL DiSoRDeRS Silke Hecht
C on t e n t s
MetABoLiC AnD DeGeneRAtiVe enCePHALoPAtHieS
Silke Hecht CHAPTER 5.3
Benjamin Young CHAPTER 5.4
Silke Hecht CHAPTER 5.5
Silke Hecht CHAPTER 5.6
iSCHeMiC BRAin DiSeASe AnD VASCULAR AnoMALieS
J. Fraser McConnell CHAPTER 5.7
J. Fraser McConnell CHAPTER 5.8
Silke Hecht CHAPTER 5.9
AGinG CHAnGeS oF tHe BRAin
Silke Hecht and Amy Hodshon CHAPTER 5.10
CRAniAL neRVe DiSeASeS
SeCtion 3 CHAPTER 6.1
MRi oF non-neURoLoGiCAL StRUCtUReS oF tHe HeAD AnD neCK nASAL CAVitY AnD FRontAL SinUSeS
Jimmy H. Saunders and Susanne A. E. B. Boroffka CHAPTER 6.2
eYe AnD oRBit
Susanne A. E. B. Boroffka and Jimmy H. Saunders CHAPTER 6.3
eXteRnAL, MiDDLe, AnD inneR eAR Jimmy H. Saunders and Susanne A. E. B. Boroffka
C on t e n t s
non-neURoLoGiC ConDitionS oF tHe HeAD AnD neCK
Jimmy H. Saunders and Susanne A. E. B. Boroffka
SeCtion 4 CHAPTER 7.1
MRi oF tHe SPine noRMAL MRi SPinAL AnAtoMY, DeGeneRAtiVe DiSC DiSeASe, AnD DiSC HeRniAtion
Wilfried Mai CHAPTER 7.2
Ronaldo C. da Costa CHAPTER 7.3
DeGeneRAtiVe LUMBoSACRAL StenoSiS
Silke Hecht CHAPTER 7.4
ConGenitAL AnD DeVeLoPMentAL AnoMALieS AnD MALFoRMAtionS
Wilfried Mai CHAPTER 7.5
inFLAMMAtoRY AnD inFeCtioUS ConDitionS
Wilfried Mai CHAPTER 7.6
Wilfried Mai CHAPTER 7.7
iSCHeMiC MYeLoPAtHY, SPinAL CoRD HeMoRRHAGe, MYeLoMALACiA
Daniela Schweizer-Gorgas CHAPTER 7.8
eXtRAMeDULLARY CYSt-LiKe ConDitionS oF tHe SPine
Wilfried Mai CHAPTER 7.9
Wilfried Mai CHAPTER 7.10
MRi oF tHe BRACHiAL AnD LUMBoSACRAL PLeXUSeS
Wilfried Mai CHAPTER 7.11
DeGeneRAtiVe SPinAL DiSeASeS Wilfried Mai
C on t e n t s
MRi oF SPinAL tRAUMA
Wilfried Mai CHAPTER 7.13
MRi oF PARASPinAL SoFt tiSSUeS
SeCtion 5 CHAPTER 8
MRi oF tHe MUSCULoSKeLetAL SYSteM MRi oF MUSCULoSKeLetAL DiSeASeS
Amy R. Zalcman, Cristi Cook, and Wilfried Mai
SeCtion 6 CHAPTER 9.1
MRi oF tHe tHoRAX AnD ABDoMen CARDiAC MRi
Wilfried Mai CHAPTER 9.2
MRi oF non-CARDiAC tHoRACiC ConDitionS
Ruth Dennis CHAPTER 10
Wilfried Mai, Ruth Dennis, and Matthew Paek index
It is my great pleasure and honor to introduce this new textbook tackling the vast and increasingly complex topic of clinical magnetic resonance imaging (MRI) in the dog and cat. The use of clinical MRI in veterinary medicine is no longer the privilege of a few, and the vast majority of university veterinary hospitals worldwide have on-site MRI scanners. In addition, there has been a dramatic increase in the number of veterinary specialty practices that now have regular access to MRI capabilities. This has been made even easier with the development of veterinary-specific machines. With the increasing availability of the technology, there is also an increasing need for expertise in interpreting these images. Although there have been a large number of scientific publications on various aspects of clinical MRI in canine and feline practice, there is a paucity of reference textbooks that provide a well-illustrated and comprehensive overview of the current knowledge. Having been in the field of veterinary radiology education for quite a few years now, I saw a need for an updated and thorough text and pictorial review of the current stateof-the-art in small animal veterinary MRI. Although MRI now constitutes one of six individual examination sections at the American College of Veterinary Radiology (ACVR) Specialty board examination, trainees often comment that they are in dire need of a well-documented, evidence-based learning resource. It was to fill this void that I decided to embark on this adventure. I have been immersed in the field of MRI since my younger years when I completed a Masters and then PhD in Biomedical Engineering, during which I focused more on the physics and basic science aspects of the modality. I started practicing clinical MRI on a regular basis at the beginning of my career at the University of Pennsylvania as an Assistant Professor of Veterinary Radiology, and I sure would have loved, back then, to have had access to a good reference textbook in my learning process. The same struggle has been true for many of my earlier radiology trainees. With our gain in experience and comfort, the increasing number of quality scientific publications elucidating MRI patterns of various diseases, and the improvement in imaging technology, MRI has become much more integrated as a routine imaging modality. Still, I had been thinking about putting a group of experts together to share our experience and summarize current knowledge to hopefully facilitate the learning of
this mesmerizing imaging technology. My goal was to select a few collaborators that are recognized and trusted experts on the topics they would be writing about, either through their clinical research experience, publication track record, or general recognition in the veterinary radiology community. I was lucky that the people I had elected to collaborate all agreed to do so. I was even luckier that they all provided high-quality contributions, which I had the immense privilege and pleasure to edit and review. Even the chapters I wrote were submitted to the scrutiny of a careful editing process by Dr. Silke Hecht from the University of Tennessee (former president of the ACVR and former president of the CT & MRI Society). Dr. Hecht is a trusted name in the field of veterinary imaging in general and MRI in particular, with numerous publications and extensive experience on the topic. I cannot thank her enough for taking over this task, in addition to contributing many excellent chapters herself. Almost all the contributors are from the radiology specialty with a vast expertise in MRI, with the exception of Dr. Ronaldo da Costa, chief neurologist at the Ohio State University, whom I asked to contribute a chapter on cervical spondylomyelopathy, as he is regarded as a world expert on the topic. Examples of other experts I asked to participate in this endeavor include Dr. Ruth Dennis, an early pioneer in the development of clinical veterinary MRI, Drs. Fraser McConnell and Daniela Schweizer-Gorgas, for their experience and knowledge of MRI of ischemic and hemorrhagic neurologic disorders, Dr. Benjamin Young for his outstanding publication track record on MRI of inflammatory brain disease, Dr. Susanne Boroffka for her expertise in MRI of the orbit, and Dr. Jimmy Saunders for his well-published investigative work on cross-sectional imaging of nasal diseases. Dr. Cristi Cook from the University of Missouri contributed her vast knowledge on the multimodality imaging of canine orthopedic conditions, for which MRI is particularly suitable. Finally, I am indebted to one of my former radiology residents, Dr. Matt Paek, for contributing so many quality MR images that are the result of his exemplary and extensive use of the modality in clinical practice. Through the careful choice of these collaborators, and a thorough editing process and literature review, I believe that this textbook provides an accurate representation of the current evidence-based knowledge in veterinary MRI to this date. Regarding the contents, although countless general references exist on the topic, I did not want to pass on dedicating
P r e fac e
a few chapters to MRI physics and technology (Section 1), because I wanted this book to be a stand-alone reference for someone to use, whether it be a radiology/neurology resident in training or a practitioner with a need to learn about veterinary clinical MRI. I tried to keep these chapters simple enough, which is challenging because when it comes to MRI, once you pull at the thread there is no end to the unraveling. My goal was to keep it concise by using many visual aids and diagrams, and trying to avoid abstract concepts and equations whenever possible. Within that section, several chapters are dedicated to optimization of imaging technique and will hopefully be helpful for anyone who is getting started with MRI scanning, whether it be a veterinary technologist or a veterinarian. The following book sections are organized by anatomic regions (brain, head/neck, spine, musculoskeletal, thorax, and abdomen), and within each section, each chapter focuses on a disease category of that body region. I wanted the book to be easily searchable and this organization seemed to make the most sense to me. I adopted a ‘bullet-points format’ throughout the book, to keep the concepts concise and organized. For the most part, all the information presented in this book reflects knowledge that is supported by peer-reviewed scientific publications, which are referenced at the end of each chapter. I did not want to deliver a book of ‘opinions’. Such material already exists. I wanted to write a book of ‘facts’. Although this book is a thorough review of the current knowledge, it is also richly illustrated with annotated images showcasing the main features of the disease processes covered in each chapter. I decided to include images obtained at all magnet field strengths, so as to reflect the current reality of veterinary MRI, which uses low-, mid-, and high-field magnets. The magnetic field strength information, where known, is included at the end of each figure caption. I cannot conclude this preface without having a thought for all the mentors that have helped shape my passion for and knowledge of clinical MRI: • First and foremost, Dr. Dominique Begon, my radiology professor when I was a student and then intern at the École Nationale Vétérinaire d’Alfort in France, who recognized my genuine interest in diagnostic imaging and helped me ‘make it happen’. • Dr. Corinne Fournel, for giving me the unique opportunity to enroll into an alternative radiology residency training program under the mentorship of Dr. Paul Barthez.
Like Dominique, Corinne always trusted me, even when anyone else would have doubted; her constant and unconditional support have meant the world to me. • Dr. Paul Barthez, my residency mentor/director: he trusted I knew something about MRI physics, but he taught me all the rest. • Dr. Allan Johnson, who welcomed me at the Duke University Center for In Vivo Microscopy. This unbelievable and amazing human being has got to be the most humble and talented junior researchers’ mentor that I have ever met. His jovial and upbeat attitude, his ‘can-do’ approach, and his insatiable passion to teach, help, mentor, and push his trainees were a true inspiration to me. Of course, no such monumental task could ever be possible to achieve without the help and support of my colleagues at PennVet: Drs. Ana Cáceres, Jennifer Reetz, Yael Mosenco, and Jantra Suran; former colleagues Drs. Tobias Schwarz, Victoria Johnson, Gabriela Seiler, Jeff Wortman, and Darryl Biery; and the talented imaging technologists that have acquired many of the wonderful images presented in this book (chronologically, Amy Basatemur, Denise Priore, Russel White, Barb Kaminsky, and Lisa Chant). I also could not have done it without all the residents that I was lucky to train over the past 13 years, undoubtedly one of the most challenging but also most rewarding parts of my job. They have all in one way or the other shaped me as the instructor and mentor I am today, and have taught me so many invaluable lessons. Observing them struggling with a variety of concepts when it comes to clinical MRI has been an unparalleled resource for me to understand what works and what doesn’t in terms of teaching this topic, which I hope will translate in this book and help many other residents and practitioners down the road. Finally, I need to thank my mom and friends who have endured my ups and downs over the many years that took me to this point and particularly over the past 2 years while in the process of writing this seemingly never-ending book. A special thank you to my husband Minhtri Thach, who has been very patient when the priority of this writing and editing process took away from our time together and has pushed me so much to make sure it would be the best it could be. I am forever grateful for all these people. Wilfried Mai
Susanne A. E. B. Boroffka, Dr. med. vet., PhD, Diplomate ECVDI Boroffka Diagnostic Imaging Utrecht, The Netherlands Cristi Cook, DVM, MS, Diplomate ACVR Thompson Laboratory for Regenerative Orthopaedics Missouri Orthopedic Institute Columbia, Missouri, USA Ronaldo C. da Costa, DMV, MSc, PhD, Diplomate ACVIM (Neurology) Professor and Service Head – Neurology and Neurosurgery Department of Veterinary Clinical Studies The Ohio State University College of Veterinary Medicine Columbus, Ohio, USA Ruth Dennis, MA, VetMB, DVR, Diplomate ECVDI, FRCVS Head of the Diagnostic Imaging Unit Centre for Small Animal Studies Animal Health Trust Newmarket, Suffolk, United Kingdom Silke Hecht, Dr. med. vet., Diplomate ACVR, Diplomate ECVDI Professor in Radiology Department of Small Animal Clinical Sciences University of Tennessee College of Veterinary Medicine Knoxville, Tennessee, USA Amy Hodshon, DVM, Diplomate ACVIM (Neurology) Assistant Professor in Neurology and Neurosurgery University of Tennessee College of Veterinary Medicine Knoxville, Tennessee, USA Wilfried Mai, Dr. Méd. Vét., MSc, PhD, Diplomate ACVR, Diplomate ECVDI Professor and Chief of Radiology University of Pennsylvania School of Veterinary Medicine Philadelphia, Pennsylvania, USA
J. Fraser McConnell, BVM&S, DVR, Diplomate ECVDI, CertSAM, MRCVS Senior Lecturer in Veterinary Diagnostic Imaging Small Animal Teaching Hospital University of Liverpool (Leahurst Campus) Neston, Cheshire, United Kingdom Matthew Paek, VMD, MS, Diplomate ACVR Synergy Veterinary Imaging Partners Frederick, Maryland, USA Jimmy H. Saunders, Dr. med. vet., PhD, CertVR, Diplomate ECVDI Professor in Medical Imaging Department of Medical Imaging of Domestic Animals and Orthopedics of Small Animals Ghent University Faculty of Veterinary Medicine Merelbeke, Belgium Daniela Schweizer-Gorgas, Prof. Dr. med. vet., Diplomate ECVDI Head of Division of Clinical Radiology Department of Clinical Veterinary Medicine Vetsuisse-faculty, University of Bern Bern, Switzerland Benjamin D. Young, DVM, MS, Diplomate ACVR VCA Alameda East Veterinary Hospital Denver, Colorado, USA Amy R. Zalcman, DVM Resident in Radiology Veterinary Health Center University of Missouri Columbia, Missouri, USA
ADC ADM AHNPE ATP AVM bFFE CAA CBF CKCS CME cMRI CNS CPP CSF CSM CVR CVT dGEMRIC DISH DSH DTI DWI ECG EPI FCD FCE FCEM FE FFE FID FISP FISS FLAIR FLASH FMPGR fMRI FOV FROMS FSE GE FISP GFE GME GPE GRE GRE-EPI HASTE HPE IR IVDD
apparent diffusion coefficient ascending/descending myelomalacia acute hydrated nucleus pulposus extrusion adenosine triphosphate arteriovenous malformation balanced fast field echo ceroid amyloid angiopathy cerebral blood flow Cavalier King Charles Spaniel canine monocytotropic (or monocytic) ehrlichiosis cardiac MRI central nervous system cerebral perfusion pressure cerebrospinal fluid cervical spondylomyelopathy cerebrovascular resistance cerebral vein thrombosis delayed gadolinium enhanced MRI of cartilage diffuse idiopathic skeletal hyperostosis domestic short-haired (cat) diffusion tensor imaging diffusion-weighted imaging electrocardiogram/electrocardiography echoplanar imaging focal cortical dysplasia fibrocartilaginous embolism fibrocartilaginous embolic myelopathy frequency-encoding fast field echo free induction decay fast imaging with steady-state precession feline injection site sarcoma fluid attenuated inversion recovery fast low angle shot fast multiplanar gradient recalled acquisition in the steady state functional MR imaging field of view feline restrictive orbital myofibroblastic sarcoma fast spin echo gradient echo fast imaging with steady-state precession frequency-encoding gradient strength granulomatous meningoencephalomyelitis phase-encoding gradient strength gradient echo gradient echo EPI half Fourier acquisition single shot turbo spin echo holoprosencephaly inversion recovery intervertebral disc disease
MHz MPS MRA MRCP MRI MRS MRV MTT NCL NE NEX NLE NME NMR N FE N PE NSA OA-CSM PC VIPR PD PDW PE PNET PNST PRF PWI rBW rCBF RF ROI SCC SE SE-EPI SNR SPAIR SPIR SSFP SS-FSE STIR SWI T1W T2W TBI TE TI TIA TOF TR TSE TTP VOX
megahertz mucopolysaccharidosis magnetic resonance angiography magnetic resonance cholangiopancreatography magnetic resonance imaging magnetic resonance spectroscopy magnetic resonance venography mean transit time neuronal ceroid lipofuscinosis necrotizing encephalitis number of excitations necrotizing leukoencephalitis necrotizing meningoencephalitis nuclear magnetic resonance number of pixels in the FE direction number of pixels in the PE direction number of signal averages osseous-associated cervical spondylomyelopathy phase-contrast vastly undersampled isotropic projection reconstruction proton density proton density-weighted phase-encoding primitive neuroectodermal tumor peripheral nerve sheath tumor pulse repetition frequency perfusion-weighted imaging receive bandwidth relative cerebral blood flow radiofrequency region of interest squamous cell carcinoma spin echo spin echo EPI signal-to-noise ratio spectral attenuated inversion recovery spectral presaturation with inversion recovery steady state free precession single shot fast spin echo short tau inversion recovery susceptibility-weighted imaging T1-weighted T2-weighted traumatic brain injury time of echo (or echo time) time from inversion/inversion time (tau) transient ischemic attack time of flight repetition time turbo spin echo time to peak signal loss volume of the voxels
PHYSICS AND TECHNICAL CONSIDERATIONS – IMAGE OPTIMIZATION CHAPTER 1
General principles of magnetic resonance imaging
Image characteristics in MRI and principal pulse sequences
CHAPTER 4.1 Optimized technique: brain CHAPTER 4.2 Optimized technique: spine CHAPTER 4.3 General features and optimized technique for the musculoskeletal system CHAPTER 4.4 Technical particularities with low-field imaging
GENERAL PRINCIPLES OF MAGNETIC RESONANCE IMAGING
CONTENTS Atomic and nuclear structure.....................................................................................................................................................................................3 What is magnetism? ..................................................................................................................................................................................................4 Spins – magnetic moment ........................................................................................................................................................................................4 Spin angular momentum ...........................................................................................................................................................................................5 The net magnetization (M0): spin-up and spin-down.................................................................................................................................................6 Transverse (Mxy) and longitudinal (Mz) components of the net magnetization ...........................................................................................................8 Measuring the net magnetization: magnetic resonance .............................................................................................................................................9 The return to equilibrium: spin-lattice and spin-spin relaxation .............................................................................................................................. 11 The longitudinal relaxation time: T1 ........................................................................................................................................................................ 12 The transverse relaxation time: T2 ........................................................................................................................................................................... 13 The free induction decay ......................................................................................................................................................................................... 13 T2* versus T2.......................................................................................................................................................................................................... 15 A basic pulse sequence: the spin echo, or how to measure the true T2 ................................................................................................................... 15 What about the repetition time? ............................................................................................................................................................................... 17 The MR image: field of view, matrix, pixel, and voxel ............................................................................................................................................. 18 From image to spatial frequencies: notion of Fourier transform .............................................................................................................................. 19 Spatial encoding: slice selection .............................................................................................................................................................................21 Spatial encoding: notions of frequency-encoding and phase-encoding .................................................................................................................. 24 A bit more about phase-encoding ...........................................................................................................................................................................26 The frequency domain, or k-space, and its relationship to the MR image ...............................................................................................................28 Further reading........................................................................................................................................................................................................35
It is beyond the scope of this textbook to provide in-depth details regarding the physics of magnetic resonance imaging (MRI). There are a large number of textbooks, scientific articles, and free online resources available for readers who are interested in a more detailed description. However, we will explain the basics and go over the material that is important for radiologists to understand: • The basic principles of nuclear magnetic resonance and the phenomenon of relaxation (longitudinal and transverse). • The concepts of image formation, and how they influence image quality such as signal-to-noise ratio, image contrast, and spatial resolution. • The mechanisms and pros/cons of the essential pulse sequences often used in diagnostic imaging. • Common image artifacts and prevention/correction.
ATOMIC AND NUCLEAR STRUCTURE • Atoms are made of a nucleus and orbiting electrons. • The nucleus itself consists of protons and neutrons. • The electrons have a negative charge and weigh about 2,000 times less than protons and neutrons; protons have a positive charge while the neutrons have no charge (Table 1.1). • The number of protons in a nucleus defines the identity of the element: for example, all hydrogen atoms contain 1 proton, all carbon atoms contain 6 protons. This number is called the atomic number, Z. • The total number of protons and neutrons in an atom is called mass number, A. • Two atoms with the same number of protons but a different number of neutrons are called isotopes.
Table 1.1 Mass and charges of atomic particles. PARtiCLe
1.6727 × 10
1.6750 × 10−24
9.110 × 10−28
• It is customary to represent an element X as A X: X defines the element (i.e., its number of protons) and A defines which isotope of the element is considered. For example, 13C (six protons) is the isotope of carbon that contains seven neutrons; the most abundant isotope of carbon in nature is 12C (six protons and six neutrons).
WHAT IS MAGNETISM? • Magnetism and displacement of an electric charge can be related to each other: an electrical current running through a cable generates a magnetic field that one can detect by placing a compass in the vicinity. Conversely, an electrical current can be induced by placing a bar magnet in the center of a solenoid cable. • Magnetism is a fundamental property of matter, caused by the orbiting electrons at the atomic level. These orbiting electrons can cause atoms to possess a small magnetic field called ‘magnetic moment’. • In a nucleus, the nucleons (protons and neutrons) rotate around their axis and, given the existence of electric charges, you can already see how the nucleus of an atom also has the potential of intrinsic magnetic properties. • The units to measure magnetic fields are the Gauss (G) and Tesla (T), the official SI unit being the Tesla. The Tesla is defined as such: a particle carrying a charge of 1 Coulomb and passing through a magnetic field of 1 Tesla at a speed of 1 meter per second perpendicular to said field experiences a force with magnitude 1 Newton. 1 Tesla ≈ 10,000 Gauss • The Earth’s magnetic field is about 0.5 Gauss or 5 × 10 −5 Tesla; compare this with the magnetic field strengths used in high-field MRI, typically 1.5 to 3 Tesla. This is about the strength of electromagnets used to pick up cars in junk yards. • Magnetic properties of materials vary depending on their composition, and in MRI four different types of magnetic properties are typically encountered: • Ferromagnetism: materials that have a large positive magnetic susceptibility. When they are placed in a magnetic field, the field strength is much stronger inside the material than outside; such materials typically contain iron, nickel, or cobalt, and include magnets and various objects that can be found in
veterinary patients such as identification microchips, surgical implants, and ballistic projectiles. • Paramagnetism: refers to materials with some ions with unpaired electrons such as ions of various metals like Fe (iron), Mg (magnesium), and Gd (gadolinium). The electronic magnetic moment due to the unpaired electrons confers on these ions a positive magnetic susceptibility capable of effecting magnetization of other structures around them. This magnetic susceptibility is much less than that of ferromagnetic materials, but sufficient to cause effects on MR images. This is exploited with some of these ions being used as MRI contrast agents (gadolinium). • Super-paramagnetism: these are materials made of discrete individual domains of elements that, when in bulk, have ferromagnetic properties. This results in a positive magnetic susceptibility that falls somewhere between that of ferromagnetic and paramagnetic materials; some of these can be used as contrast agents in MRI, such as super-paramagnetic iron particles. • Diamagnetism: refers to materials that do not possess intrinsic atomic magnetic moment but, when placed in a strong external magnetic field, slightly repel the field, resulting in a negative magnetic susceptibility. Water and most biologic tissues are diamagnetic.
SPINS – MAGNETIC MOMENT • All fundamental particles (electrons, protons, neutrons) spin around their own axis and, therefore, as a result of electromagnetism, they have an associated magnetic field called ‘magnetic moment’ or ‘magnetic dipole moment’ (Fig. 1.1). This property is called ‘spin’. • Spin is a fundamental property of nature, like mass or charge. It comes in multiples of ½ and can be positive or negative. Individual unpaired electrons, neutrons, and protons have a spin of +½ or −½. • Note that neutrons, although devoid of a net charge, do possess a magnetic moment (non-zero spin). This is because at the sub-particle level, they are made (like protons) of quarks, which are electrically charged; the total charge of a neutron’s quarks is zero, but their spatial distribution within the neutron generates a magnetic moment when the neutron spins. The magnetic moment of a neutron is equal to about two-thirds that of a proton. • Particles with similar spin can pair up, with their magnetic moments facing opposite directions, in the same way that two identical little magnets would. This in turn eliminates the observable manifestations of their individual spins. In a nucleus, protons (as well as neutrons) form pairs and their individual magnetic moments cancel each other out (Fig. 1.2). • As a result, all isotopes that contain an odd number of protons and/or neutrons possess an intrinsic ‘nuclear magnetic moment’ and ‘nuclear spin angular momentum’ (see below),
G e n e r a l P r i nc i pl e s of M ag n e t ic R e son a nc e I m agi ng
Fig. 1.1 The hydrogen atom contains one single proton; as this proton spins around its axis, it generates a small magnetic field, and behaves as a little bar magnet with a magnetic moment vector pointing south to north.
and can be used for nuclear magnetic resonance (NMR) experiments. They are said to have a ‘non-zero spin’; spin is a quantum mechanical feature of the nucleus. In contrast, all isotopes with an even number of protons and neutrons have an intrinsic nuclear magnetic moment of zero (Fig. 1.2). • For example, carbon 13C (seven neutrons [i.e., three pairs of neutrons + one unpaired neutron] and six protons [three pairs]) has magnetic properties that make it usable in NMR, while 12C (three pairs of neutrons and three pairs of protons) does not. • In MRI, the signal is derived from hydrogen protons (1H, one unpaired proton). In NMR spectroscopy, other elements can be used to generate an NMR signal, such as certain isotopes of phosphorus (31P), carbon (13C), sodium (23Na), or fluorine (19F).
• There are two reasons why hydrogen protons (1H) are very good for MRI: • They are abundant in biologic tissues due to their richness in water molecules; • They have the ability to generate a relatively strong NMR signal compared with other elements that have a non-zero spin. • ‘Nuclear magnetism’ refers to the fact that a nucleus with a non-zero spin will behave as a little bar magnet, with magnetic field lines around it (Fig. 1.1). This magnetic field can be described by a vector, where the length of the vector describes the magnitude of the magnetic field generated, and the direction of the vector describes the orientation of the magnetic field generated. Much like the needle of a compass, when that nucleus is placed in a strong external magnetic field B0, its magnetization vector will try to align itself parallel to the direction of the external magnetic field vector B0 .
SPIN ANGULAR MOMENTUM • If you place a spinning top on a table and spin it, it will quickly deviate from the vertical axis due to interactions between the rotational force and the gravitational force, and will start to wobble around the vertical axis. In the absence of friction (which will eventually make it stop spinning and fall), it would continue to spin and wobble around the vertical axis at a specific angle to that axis due to its specific ‘angular momentum’, which depends on its mass (Fig. 1.3). • There is an analogous property of small nuclear particles with a non-zero spin that is proportional to their mass, and this is called ‘spin angular momentum’. Because
PAIRED MAGNETS Their opposite poles attract each other
HELIUM NUCLEUS 2 paired protons (+) 2 paired neutrons (N) → No magnetic moment
HYDROGEN NUCLEUS 1 unpaired proton → Non-zero magnetic moment
Fig. 1.2 Left: identical bar magnets pair up with their opposite poles, matching up so that their magnetic fields cancel each other. Middle: the same phenomenon happens in a nucleus with an even number of neutrons and protons; in this example, the helium nucleus contains two paired protons and two paired neutrons, and therefore has no net magnetic moment. Right: the hydrogen nucleus contains a single (thus unpaired) proton and therefore possesses a magnetic moment.
Fig. 1.3 The top is spinning around its own axis (red arrow). Due to the interaction of the mass of the spinning top with the gravitational force, the rotating axis of the spinning top has a gyration pattern (blue arrow) around the axis of the force of gravity (green arrow). That gyration motion is called ‘precession’.
Spinning motion of the proton
of the spin angular momentum, when a nucleus with a non-zero spin is placed in an externally applied magnetic field, it is not perfectly aligned parallel to the external magnetic field; instead, the angular momentum forces it to wobble around the magnetic field. The envelope of the trajectory of the spin magnetic vector is therefore coneshaped; this particular motion due to the spin angular momentum is called ‘precession’ (Fig. 1.4). • Note that the spinning motion of the particle exists even in the absence of an external magnetic field, but there is no precession in that case. Spin angular momentum exists even in the absence of an external magnetic field as an intrinsic property of spins; it is the interaction between the spin angular momentum and the external magnetic field that causes the motion of precession. • The frequency (ω0, in radians per second [= angular frequency], or f 0 in megahertz [number or rotations per second]) of that precession motion is a function of the particle and the strength B0 (Tesla) of the externally applied magnetic field:
ω 0 rad.s−1 = γ rad.s−1.Tesla−1 × B0 ( Tesla )
Fig. 1.4 The spin angular momentum of a particle with a non-zero spin causes it to precess around the axis of the strong external magnetic field B0.
f0 ( MHz ) = γ MHz.Tesla −1 × B0 ( Tesla ) ω 0 = 2 × π × f0 • γ is the ‘gyromagnetic ratio’, and is unique for an element; for hydrogen protons: • γ = 2.67 × 108 rad.s−1.T−1 • γ = 42.57 MHz.T−1 • For example, in a magnetic field of 1 Tesla, the precession motion of hydrogen protons will be at a frequency of 42.57 MHz, or 42.57 million rotations per second. • The equation ω0 = γ × B0 is called the ‘Larmor equation’. • ω0 is called the ‘Larmor frequency’.
THE NET MAGNETIZATION (M0): SPIN-UP AND SPIN-DOWN • Regardless of the body part being imaged, some principles always apply in MRI. The basic principle relies on placing a patient in the bore of a magnet and generating a signal: the NMR signal. • Consider a sample of protons (hydrogen nuclei). Outside of an externally applied magnetic field, at the equilibrium
G e n e r a l P r i nc i pl e s of M ag n e t ic R e son a nc e I m agi ng
M0 = 0
• Fig. 1.5 A theoretical sample of a few protons (hydrogen nuclei). In the absence of a strong external magnetic field, they are oriented in a random fashion so that the sum of their individual magnetization vectors, called the net magnetization vector (M0), is equal to 0.
state, the magnetic field vectors of protons are oriented randomly, so that the vector sum M 0 (macroscopic magnetization, or ‘net magnetization vector’) of the magnetic fields of all individual protons would equal zero (Fig. 1.5). This is the reason why a biologic tissue sample at rest has no measurable magnetization. • When placed in a strong external magnetic field (B0), a non-zero magnetization is created in the sample of pro tons, the direction of which is parallel to that of B0 . We say that the sample of protons is now ‘magnetized’. • That net magnetization is due to the fact that, like the needle of a compass, protons will become parallel to
the direction of the external magnetic field vector B0 . However, due to quantum mechanics constraints, they will not all have the same orientation; some spin vectors will be pointing in the direction of B0 (‘spin-up’ or parallel), and some in the opposite direction (‘spin-down’ or anti-parallel). This is different from what physical bar magnets would do in a magnetic field where they would all be aligned in the direction of the external magnetic field; this is because protons obey the rules dictated by ‘quantum physics’, as opposed to bar magnets, which obey the rules of ‘classic mechanics’. Note that, in reality, the individual spin vectors are pre cessing around the direction of B0 , with the tip of the vectors describing a circular trajectory, the plane of the circle being perpendicular to the direction of B0 (Fig. 1.6). The reason why the protons, when placed in an external field, can exist in one of these two states is a result of quantum mechanics physics, which dictate that a particle with a spin value of s can exist in [2s + 1] orientations, or ‘states’, when placed in an external magnetic field. Since protons have a spin of ½, they can adopt two states, or orientations, when placed in B0 . The population of protons is distributed about equally between the spin-up and spin-down groups, but because the energy required to remain at the spin-down level (anti-parallel) is slightly higher, there is a slight excess of protons in the spin-up orientation than in the spin-down. This results in the overall sample of protons gaining a net magnetization vector M 0 that is oriented parallel to, and pointing in the same direction as, B0 (Fig. 1.6).
z y Mz
Mxy = 0
DOWN DOWN M0
UP NUP = 5 NDOWN = 4
Fig. 1.6 When placed in a strong external magnetic field B0 , the individual spin vectors are precessing around the direction of B0 , with the tip of the vectors describing a circular trajectory, the plane of the circle being perpendicular to the direction of B0 . There is a slight excess of spin-ups versus spin-downs (in this case 5 versus 4), resulting in a non-zero net magnetization vector M0.
• According to Boltzmann’s statistics, the ratio between the number of spins up (Nup) and spins down (Ndown) depends on the energy difference (ΔE) between the two states and the temperature (T) as follows, where kb is the Boltzmann constant:
N up = exp k b .T N down
z component y component
• The energy difference ΔE between the two states depends on the strength of the magnetic field B0. At the temperature of the human body, and in a 1.5 T magnetic field, the ratio would be equal to 1.000004, which means that for every 1 million protons in the anti-parallel direction, there are one-million-and-four protons in the parallel direction. • As you may imagine, the ‘net magnetization’ created by such a minor difference is pretty small. One could calculate that, assuming the human head has a volume of 1,500 mL and is made of 80% water, the net magnetization at B0 = 1.5 T would be about M0 = 20 μT. This is 75,000 times weaker than the main magnetic field B0. Thus, specific strategies will be needed to extract such a weak signal from the strong background created by B0. • The magnitude of M 0 depends on the amplitude of the magnetic field B0, the density of protons in the sample (proton density [PD]), the temperature T, and the gyro γ magnetic ratio γ = of the element at hand (h is the
Fig. 1.7 Schematic representing a proton oriented parallel to B0 and precessing around it, in a Cartesian coordinates system (x,y,z). One can consider two vectorial components to the magnetic moment vector of that spin. The z component is the projection onto the z-axis and the xy component is the projection onto the (x,y) plane. Note that as the proton precesses around B0, the z component remains constant and static, while the (x,y) component (red vector) will rotate in the plane (x,y) at the precession frequency.
Planck constant and kb the Boltzmann constant):
( γ .h ) 4.k b .T 2
M0 = B0 × PD ×
• This relationship shows that: • Higher magnetic field B0 yields higher net magnetization M0 in a given sample of tissue, hence the MRI signal will be better in a 1.5 T MRI (high-field magnet) machine than in a 0.2 T MRI (low-field magnet) machine. • Higher PD yields higher net magnetization: tissues rich in protons, such as highly hydrated tissues, will be more magnetized and produce more signal than tissues with poor PD. You can already start to understand why MRI has high sensitivity to pathology; for example, brain edema, richer in water than normal brain parenchyma, will be readily detectable with MRI.
TRANSVERSE (Mxy) AND LONGITUDINAL (Mz) COMPONENTS OF THE NET MAGNETIZATION (Fig. 1.7) • Let us consider an (x,y,z) Cartesian coordinate system, with the z-axis parallel to the main magnetic field B0 , and the (x,y) plane perpendicular to the direction of B0 .
• In that system, once placed in the magnetic field B0 , the protons will all individually precess around the direction of the z-axis. Although they all have the same precession frequency ω0, they are doing so incoherently: they are said to not be ‘in phase’, so that the projection of their magnetic vectors onto the (x,y) plane makes up a group of vectors with random directions and equal magnitude. • As a result, the sum of the projections of these vectors in the (x,y) plane equals zero; in other words, the total component of the net magnetization in the (x,y) plane at equilibrium, called the ‘transverse magnetization’, is zero: M xy = 0. • On the other hand, the projection of the individual precessing spin vectors onto the z-axis is either a vector pointing in the direction of B0 (spin-up) or 180° opposite (spin-down), and all of these projections have the same amplitude, so that the sum of an up-projection and a down-projection equals zero; since there is a slight excess of spin-ups in the sample, the resulting vector sum along the z-axis, also called ‘longitudinal magnetization’, Mz, is a positive vector pointing in the same direction as B0 . • From this it becomes clear that, at equilibrium, there exists a net magnetization vector M 0 , which is aligned exactly with the main magnetic field B0 . The components of that net magnetization vector in the (x,y) plane
G e n e r a l P r i nc i pl e s of M ag n e t ic R e son a nc e I m agi ng (transverse magnetization) and along the z-axis (longitudinal magnetization) are such that (Fig. 1.6): M xy = 0
Mz = M0
MEASURING THE NET MAGNETIZATION: MAGNETIC RESONANCE • As you have probably guessed, the aim of MRI will be to measure the net magnetization in individual voxels, made of packets of tissue with different densities of protons, and protons with different magnetic behaviors depending on their physicochemical environment. • The net magnetization vector M 0 that we defined in the previous paragraphs is of much smaller magnitude than the main magnetic field B0 it is parallel to, and therefore it is virtually impossible to measure that magnetization as this tiny signal is so much smaller than B0. • The strategy to measure that magnetization is to flip the net magnetization vector M 0 into a plane perpendicular to B0 . Indeed, the projection of B0 in a plane perpendicular to it will be null, and therefore if one can selectively shift the net magnetization from its original longitudinal orientation to a plane perpendicular to it, it will become measurable without being ‘hidden’ by the intense mag netic field B0 . For measurement, we will need to use a detector that only measures magnetic fields in the transverse plane. • The property used to achieve this goal relies on ‘magnetic resonance’. As you remember, when placed in a strong magnetic field, the protons align themselves along the direction of B0 , with about half parallel to (spin-up) and half anti-parallel (spin-down) to the direction of B0 . There is a slight excess of protons in the ‘spin-up’ orientation, because this energy level is slightly less than the ‘spindown’ orientation. The energy difference, ΔE, between the two states is directly proportional to the strength of the magnetic field B0. • One can force protons to transition between the energy levels by providing energy to the sample of protons, equal to the energy difference ΔE between the two states. This energy is provided in the form of an electromagnetic radiation, which is a rotating magnetic field B1. This rotating (or oscillating) magnetic field is: • Generated by a radiofrequency (RF) pulse within the transmit coil of the MR system. • Perpendicular to B0 (i.e., in the (x,y) plane). • Rotating around the axis of B0 (= the z-axis), at the Larmor frequency of protons. • The principle of resonance is that the protons will only transition between the energy levels if the RF pulse is applied at the proton’s Larmor, or precessional, frequency
ω0 = γ × B0. This frequency is called the ‘resonance frequency’. For the experiment to work, the RF pulse and the protons have to be ‘in resonance’. To understand the concept of ‘resonance’, think about pushing a child on a playground swing; the swing has an intrinsic oscillating frequency that depends on its length and load. If you swing the child at the same frequency as the natural frequency of the swing, you are very efficient in transferring energy to the swing; however, if you try to do it ‘out-of-tune’ with regard to the frequency of the swinging pattern, it will be less efficient, more difficult or even counterproductive. The energy transfer will be maximum when you swing the child at the exact same frequency as the natural frequency of the swing (i.e., when you are in ‘resonance’ with the swing). When the RF pulse is applied, two things happen (quantum mechanics interpretation): 1. The system is provided with energy, allowing more and more protons to transition from the spin-up (lower energy) to the spin-down (higher energy) orientation; the net result of this is that the longitudinal component Mz of the net magnetization M0 is progressively decreasing, as there are fewer and fewer protons in the spin-up orientation. When there is an equal number of protons in the spin-up as in the spin-down orientation, the longitudinal component of the net magnetization becomes null. 2. The precession motions of individual protons are more and more ‘in-phase’ with each other; as a result, the transverse components of the spin vectors are not oriented randomly in the (x,y) plane anymore, and therefore their sum, the transverse magnetization M xy, is not equal to 0 and progressively increases. When all protons in the sample are ‘in-phase’, M xy reaches its maximum value equal to the original magnitude of the net magnetization, M0. The net result of these phenomena is that the net magnetization vector will progressively tilt away from the initial orientation, with its tip describing a spiral motion that fits in a sphere, from the north pole of that sphere to its equator (Fig. 1.8). After a 90° pulse it will end up being completely in the (x,y) plane, perpendicular to both B0 and B1 and, like B1, rotating at the Larmor frequency in that plane (Figs. 1.8, 1.9). The vector M 0 is said to be ‘flipped’ away from B0 . As this happens, the longitudinal component of the net magnetization vector M 0 (Mz) progressively decreases, while its transverse component (M xy) progressively increases. M xy reaches a maximum value when M 0 is completely in the transverse (x,y) plane, and its amplitude is then equal to the amplitude of the longitudinal magnetization at equilibrium: M0. At that point in time, the longitudinal component Mz of the net magnetization vector M 0 is equal to 0 (Fig. 1.9).
During the application of the RF pulse, the extremity of the magnetization vector M0 is describing a spiral along the surface of a sphere, and progressively shifting away from B0
After a certain time, the magnetization vector M0 is completely shifted in the transverse plane, the equatorial plane of the sphere: this is a 90° pulse
M0 If the RF pulse is kept on, the magnetization vector M0 continues its path along the southern hemisphere until it reaches a direction completely opposite to the equilibrium and to B0: this is a 180° pulse
Fig. 1.8 Trajectory described by the net magnetization vector M0 during the application of the RF pulse.
Mxymax = M0 x
Mxy = M0
Mz = 0
Fig. 1.9 During the application of the RF pulse (using a rotating magnetic field B1 ), the longitudinal component of the net magnetization vector M 0 (Mz) progressively decreases, while its transverse component (M xy) progressively increases. M xy reaches a maximum value when M 0 is completely in the transverse (x,y) plane, and its amplitude is then equal to the amplitude of the longitudinal magnetization at equilibrium: M0. At that point in time, the longitudinal component of the net magnetization vector M 0 is equal to 0, and both M 0 and B1 (perpendicular to each other) are rotating in the transverse (x,y) plane, at the Larmor frequency.
• If one measures the angle α between M 0 and B0 (z-axis) during this process, one can observe a steady increase in the value of α from 0 to 90° as M 0 transitions from being parallel to B0 into the (x,y) plane; the net value of
α is called the ‘flip angle’ and depends on the gyromagnetic ratio, the magnitude of the rotating magnetic field B1, and the time t that B1 is applied for: α = γ × B1 × t
G e n e r a l P r i nc i pl e s of M ag n e t ic R e son a nc e I m agi ng
Table 1.2 Changes in amplitude of the longitudinal and transverse components of the net magnetization vector M0 , depending on the flip angle. eQUiLiBRiUM
−M0 (maximum, negative)
• If the RF pulse continues to be applied after M 0 has been flipped into the transverse plane (a ‘90° pulse’), the magnetization vector will continue its course, with its tip describing a spiral along the surface of the southern hemisphere (Fig. 1.8), down to a point where it will be oriented completely parallel to B0 but in a direction completely opposite. This corresponds to a flip angle of α = 180° (a ‘180° pulse’, also called an ‘inversion pulse’). • The duration of the RF pulse is very short, in the order of a few milliseconds. The longer the RF pulse is applied, the larger the flip angle α. • Table 1.2 describes the changes in amplitude of the longitudinal and transverse components of the net magnetiza tion vector M 0 , depending on the flip angle. • Note that, given the relationship between the ‘flip angle’, the amplitude of the RF pulse (B1), and the duration of the pulse (t) [α = γ × B1 × t], the same flip angle can be obtained with various combinations of amplitude of B1 and application time t. In practice, stronger pulses are used to increase flip angles, in order to maintain minimal image acquisition times. • The principal benefit of the magnetic resonance experiment is that it shifts the net magnetization into a plane that is now perpendicular to B0 and, since B0 has no vector component in that plane, the net magnetization becomes measurable; therefore, now, the magnetic properties of every single voxel of the patient become, in theory, measurable.
THE RETURN TO EQUILIBRIUM: SPINLATTICE AND SPIN-SPIN RELAXATION • In the previous paragraph we saw that, during a 90° RF pulse, the net magnetization is shifted into the transverse (x,y) plane, and that this is the result of two phenomena: 1. Energy absorption (from the RF pulse) allowing spinups to transition to the spin-down state until there is an equal number of spins ‘up’ and ‘down’, leading to nulling of the longitudinal magnetization Mz. 2. Synchronization of the precession motion of the protons, which now precess ‘in-phase’, leading to an
increase in the transverse magnetization M xy until a maximal amplitude, equal to M0. • This state is unstable, so that as soon as the RF pulse stops, these two separate processes are reversed, and the system returns to equilibrium (stable state): • There are transitions between the two energy levels, which restore a slight excess of spin-ups in the sample (the spin-up state is more stable [lower energy] than the spin-down) and therefore allow the longitudinal magnetization to progressively grow back to its initial and maximum value, M0; this phenomenon is called ‘longitudinal, or spin-lattice, relaxation’ because it relies on energy exchange between the protons and the molecules in their environment (lattice). • The protons undergo progressive dephasing due to local magnetic field inhomogeneities induced by the movement of adjacent spins due to molecular vibrations, rotations, or collisions (Brownian motion). This results in a rapid decrease in amplitude of the transverse magnetization M xy; this phenomenon is called ‘transverse, or spin-spin, relaxation’. The term ‘spinspin’ conveys the fact that this relaxation is because of interactions between protons, due to the different molecular environments they are in, causing local variations in the magnetic field the protons experience (B0 +/− delta); these microvariations in the local magnetic field change the precession frequency of individual protons due to the Larmor equation (ω = γ × B0), which in turn causes rapid dephasing of the protons in any given voxel. These local inhomogeneities are the same reason why, at equilibrium, the protons are precessing out-of-phase, resulting in a null value of the transverse magnetization. • Due to the relaxation phenomena, after a 90° RF pulse there is an exponential regrowth of the longitudinal magnetization Mz and an exponential decrease of the transverse magnetization M xy (Figs. 1.10, 1.11). These phenomena are due to two separate processes, respectively: (1) energy exchange between protons and microenvironment (lattice), and (2) rapid dephasing of the precession motion of the protons. Because these processes are independent, the rate of regrowth of the longitudinal magnetization Mz is not the same as the rate at which the transverse magnetization M xy decreases after the end of the RF pulse. • The rate of Mz regrowth is characterized by a specific time, called T1 (‘longitudinal relaxation time’), which is the time it takes for the longitudinal magnetization to reach 63% of its original amplitude before the RF pulse was applied (which, as you remember, equals the amplitude of the equilibrium net magnetization M0) (Fig. 1.10). After a 90° pulse, Mz regrows as a function of time (t) according to the following equation: t − Mz ( t ) = M0 × 1 − e T1
Mz = 100% = M0 95% 87% 63%
T1 500–1,000 ms in biologic tissues
Fig. 1.10 T1 relaxation curve showing the exponential regrowth of the longitudinal magnetization after a 90° RF pulse. The T1 relaxation time is a characteristic of specific biologic tissues and corresponds to the time after which 63% of the maximum value at equilibrium, M0, is reached.
Mxy= 100% = M0
37% 13% 5% T2 50–100 ms in biologic tissues
• The rate of M xy decay is characterized by another specific time, called T2 (‘transverse relaxation time’), which is the time it takes for the transverse magnetization to decrease by 63% of the amplitude it had at the end of the RF pulse (which equals the amplitude of the net magnetization M0) (Fig. 1.11). In other words, at a time of T2 after the end of the RF pulse, M xy will be equal to 37% of the maximum value it had at the end of the 90° pulse (M0). At the end of the 90° pulse, the transverse magnetization M xy decreases as a function of time according to the following equation: M xy ( t ) = M xy _ max × e
= M0 × e
THE LONGITUDINAL RELAXATION TIME: T1 • T1 is an intrinsic characteristic of the biologic tissue containing protons. In biologic tissues, T1 is in the order
Fig. 1.11 T2 relaxation curve showing the exponential decrease of the longitudinal magnetization after a 90° RF pulse. The T2 relaxation time is a characteristic of specific biologic tissues and corresponds to the time after which 37% of the maximum value at equilibrium, M0, is reached.
of 500–1,000 ms. T1 depends on molecular structure as well a solid or liquid state of tissues; for example, T1 tends to be longer in fluids than in solids and is shorter for fatty versus non-fatty tissues. • Tissues have specific T1 because, due to their molecular structures or state (liquid or solid), the Brownian motion of the molecules (translations, rotations, collisions) varies significantly between different tissues. The frequency of these molecular collisions has an influence on the ability of the protons to exchange energy with the lattice during the relaxation period: • When the natural frequency of molecular collisions in the tissue is close to the Larmor frequency of protons (condition of ‘resonance’, which maximizes energy transfer), the energy exchange during relaxation is maximal and T1 is short. This is, for example, the case in fatty tissue, which is why fat has a short T1. • Conversely, when the collisional frequency is very low (e.g., in water molecules in ice, which have
G e n e r a l P r i nc i pl e s of M ag n e t ic R e son a nc e I m agi ng poor mobility), the energy exchange during relaxation is slow and T1 is long. • Another example of long T1 is water, in which the collisional frequency of the highly mobile, small-sized, water molecules is much higher than the protons’ Larmor frequency. This is a reason why when a tissue undergoes pathologic change with increased water content, the T1 of that tissue will increase; an example of that situation is brain edema, and as we will see later, this will translate to a darker (hypointense) brain parenchyma on T1-weighted images. • There is a relationship between the T1 of a given tissue and the magnetic field B0, due to the Larmor equation (ω = γ × B0). When B0 increases, so does the Larmor frequency of protons in that field. This in turn increases the difference between Larmor frequency and Brownian collisional frequencies (which are constant and independent of B0). This is the reason why, for example, T1 of fat will be longer at 3 Tesla than at 1.5 Tesla.
THE TRANSVERSE RELAXATION TIME: T2 • Like T1, T2 is a characteristic of each biologic tissue. • T2 in biologic tissues is a lot shorter than T1 (typically 10 times shorter, about 50–100 ms). The shorter the T2, the faster the transverse magnetization will disappear after the end of the RF pulse. • Like T1, T2 depends on the molecular structure and state (solid/liquid) of tissues. This again can be explained by the Brownian motion differences in tissues: • In tissues where there is high mobility of small molecules such as free water, the fast motion of the small molecules
averages out the local inhomogeneities of the magnetic field induced by said molecules, so that individual protons see, on average, a more homogeneous magnetic field. This, in turns, decreases spin-spin interactions and lengthens the transverse relaxation process: T2 is therefore longer. This is the reason why fluids, such as cerebrospinal fluid (CSF) in neuro-MRI, have a long T2. For the same reason, accumulation of free water in tissues, such as seen with brain edema, will also induce an increase in T2 of these tissues. As we will see later, this is why fluids like the CSF, or edematous brain tissue, will appear bright (‘hyperintense’) on T2-weighted images. • Conversely, in tissues where there is slow motion of molecules, such as solid tissues, or tissues containing large molecules, the T2 will be shorter.
THE FREE INDUCTION DECAY • As we just saw, during relaxation, Mz regrows and M xy decreases. Due to the differences in T2 (tens of milliseconds) versus T1 (hundreds of milliseconds) the decrease in M xy is much faster than the regrowth of Mz. As a result, if we now combine M xy + M z = M 0 , we can see that the tip of the vector M 0 during relaxation will have an envelope that resembles the bell of a trumpet, somewhat conical (Fig. 1.12). • This is different from the spherical envelope that was observed during the application of the RF pulse, and that difference is inherently due to the specific relaxation times T1 and T2, with T2 T1 (i.e., for tissues with short T1, such as fat), nulling occurs at a time equal to about T1*(ln 2) = 0.69*T1, as the second term in between the brackets becomes closer to 0. This does not, however, hold true for tissues with long T1, such as fluids. • The TI determines the amount of longitudinal magnetization that is available from each tissue to be flipped into the transverse plane by the 90° RF pulse. • Obviously, depending on the TI value, some tissues will still have a negative longitudinal magnetization by the time the 90° RF pulse is applied. In most image reconstruction schemes, the sign of the net magnetization of a given tissue does not matter, as signal is displayed as ‘magnitude’ (i.e., the absolute value of the net magnetization vector). This is the ‘magnitude reconstruction’ where signal will vary from 0 (black) to maximum (white) (Figs. 2.14, 2.16). • Some vendors have an option to use ‘phase-corrected’ reconstruction as opposed to ‘magnitude reconstruction’. This is called ‘true IR’ or ‘real IR’, and in this scheme, the information about the polarity of the net magnetization during recovery is preserved, with negative values darker while positive values are made brighter (hyperintense). On such images, air (zero signal) is a mid-shade of gray, midway between completely black and completely white. • IR sequences tend to be T1W as the 180° RF preparatory pulse exaggerates the differences in T1 of tissues (i.e., the inversion pulse increases the ‘T1-dynamic range’ in the image). • The main use of IR sequences today is through their ability to suppress the signal from specific tissues. Knowing the T1 relaxation time of specific tissues at specific field strength, one can use IR sequences to suppress selectively the signal from these tissues, such as fat (short tau inversion recovery [STIR]) or fluids (fluid attenuated inversion recovery [FLAIR]). This is achieved by using a TI equal to the nulling time of the longitudinal magnetization of the tissue that is targeted (e.g., for tissues with short T1:0.69*T1tissue).
[…] RF pulses STIR
Fig. 2.14 Principle of a STIR pulse sequence. A preparatory 180° inversion RF pulse is used first and flips all protons in the opposite direction. Protons then start to recover their longitudinal magnetization in the direction of B0 at variable rates, depending on the T1 relaxation time of the tissue they are in. At variable time points the longitudinal magnetization of all tissues would cross the 0 line, which defines their nulling point. At an inversion time (TISTIR) equal to the nulling time of the longitudinal magnetization of fat, the classic fast spin echo sequence is started with the 90° excitation pulse followed by multiple 180° pulses. The gray scale on the right illustrates the expected signal intensity obtained from tissues as a result of this imaging scheme when images are displayed using the ‘magnitude reconstruction’. Fat will be black as it does not have any longitudinal magnetization, therefore will not generate any transverse magnetization or MR signal. Pathology will tend to be quite intense, as it usually contains a large amount of free water, which lengthens its T1 time, thereby causing a significant amount of residual negative longitudinal magnetization by the time the 90° RF pulse is applied. Fluids will tend to be very bright (hyperintense) due to the long relaxation time.
Short tau inversion recovery
• In a STIR pulse sequence (Fig. 2.14), the TI is short, matching the nulling time of fat; for example, at 1.5T, T1 for fat is equal to about 220 ms, so using a TI of about 150 ms (0.69*220) will null the signal from fat when the 90° pulse is applied. • Nulling of fat signal with STIR is more efficient than with the fat saturation techniques, as STIR is not sensitive to magnetic field inhomogeneities, which make saturation techniques less suitable, especially when imaging large FOVs, and in newer 1.5T systems, where the extremely short bore limits the homogeneity of spectral fat suppression at the ends of the magnet. • STIR pulse sequences cannot be used with gadolinium contrast agents, because the T1 of enhancing tissue is shortened and closer to that of fat, leading to suppression of signal from enhancing tissues on STIR images. Therefore, if fat suppression is needed on post-contrast scans, fat saturation techniques are used.
• The benefits and uses of the STIR pulse sequence are multiple: • It is widely used in musculoskeletal imaging, because the suppression of the bright signal from the fatty bone marrow enhances the conspicuity of pathologic changes to the bone, such as edema, bruising, or neoplasia (Fig. 2.15). • Pathologic lesions (e.g., tumors, infection) are often rich in free water, which lengthens the T1 of the lesion. When short to medium TIs are used, this T1 lengthening creates a substantial negative longitudinal magnetization of these tissues at the TI time, causing these lesions to look brighter (hyperintense) on magnitude reconstructed images. Note that for the same reason, fluids (e.g., CSF, synovial fluid, cysts) tend to look quite hyperintense on STIR images, like ‘T2-pathology’ scans. • Because hyperintense signal from pathology is enhanced on STIR images while bright signal from
I m age C h a r ac t e r is t ic s i n M R I a n d P r i nc i pa l P u l s e Se qu e nc e s
Fig. 2.15 Comparison between a T2W-TSE sagittal image of the shoulder (left) and a STIR image (right) in a dog with shoulder osteochondrosis. The subchondral bone marrow lesions are more obvious in the STIR image (arrows) and the joint effusion in the caudal pouch (dotted arrow) is also more conspicuous as its signal is enhanced compared with the suppressed fat around it (which was hyperintense on the T2W image, arrowhead).
background or surrounding fat is reduced, the lesionto-background contrast tends to be enhanced, which makes STIR images suitable for fast ‘screening’ of large FOVs looking for bright lesions.
Fluid attenuated inversion recovery
• In FLAIR pulse sequences, the TI is calculated so that signal from fluid is nulled (Fig. 2.16). • Relatively pure fluids, such as CSF, have long intrinsic T1 relaxation times, therefore their TI null is not simply a function of 0.69*T1, as for fat, but it more strongly depends on TR as shown in the TI null equation above, and also depends on TElastecho. • FLAIR images can be made with different types of weighting: • T1-FLAIR images are obtained with relatively short TR and TE values in order to minimize T2-weighting. With these sequences, TI null for fluids is usually around 800–1,000 ms. • T2-FLAIR images are obtained with very long TR and TE values to maximize T2-weighting and hence sensitivity to pathology. With these sequences, TI null for fluids is usually around 2,000–2,500 ms. • The benefits and uses of a T2-FLAIR pulse sequence are due to its high sensitivity to pathology owing to the heavy T2-weighting. Water-rich lesions are usually very bright and conspicuous on T2-FLAIR images. T2-FLAIR also increases the conspicuity of lesions that are located close to the bright signal of CSF on T2W images such as
intra- or periventricular lesions in the brain (Fig. 2.17), meningeal lesions, or peripheral spinal cord lesions; for example, this can be useful for a diagnosis of meningitis. This pulse sequence is now a standard part of most clinical protocols in neuroimaging. • A T1-FLAIR pulse sequence is not typically part of standard imaging protocols, but it can occasionally be useful to image lesions surrounded by CSF such as meningiomas in the subarachnoid space (Fig. 2.18) or intraventricular tumors, especially after the application of contrast material (gadolinium). The contrast enhancement of the lesion and suppression of the signal from the CSF can lead to an increase in lesion-to-background contrast that can be beneficial.
Gradient echo pulse sequences General principles of gradient echo imaging
• Up until now, we have studied pulse sequences of the spin echo family. In this type of pulse sequence, the inhomogeneities of the magnetic field that cause ultrafast decay of the transverse magnetization after the 90° RF pulse (according to the T2* time) are corrected for by the application of a 180° RF pulse, which rephases protons and generates an echo at the TE time. • Note that the generation of an echo is essential to the way that signal is recorded and spatially encoded before being stored in k-space, since the center of the echo, containing the higher amplitude signals with
[…] RF pulses FLAIR
Fig. 2.16 Principle of a FLAIR pulse sequence. A preparatory 180° inversion RF pulse is used first and flips all protons in the opposite direction. Protons then start to recover their longitudinal magnetization in the direction of B0 at variable rates, depending on the T1 relaxation time of the tissue they are in. At variable time points the longitudinal magnetization of all tissues would cross the 0 line, which defines their nulling point. At an inversion time (TI FLAIR) equal to the nulling time of the longitudinal magnetization of (pure) fluid, the classic fast spin echo sequence is started with the 90° excitation pulse followed by multiple 180° pulses. The gray scale on the right illustrates the expected signal intensity obtained from tissues as a result of this imaging scheme when images are displayed using the ‘magnitude reconstruction’. Fluid (such as CSF) will be black as it does not have any longitudinal magnetization, therefore will not generate any transverse magnetization or MR signal.
information on contrast, must correspond to the 0 values of k x (as you remember, the center of k-space contains the higher amplitude frequencies with information on contrast). That is why in spin echo we let the magnetization dephase naturally, before rephasing it with a 180° RF pulse, which will generate an echo, the center
Fig. 2.17 Transverse images of the brain at the same level in a dog with a choroid plexus tumor. Left is a T2W image and right is a T2-FLAIR image. There is dilation of the ventricular system (*). The outline of the tumor (T) is more conspicuous on the T2-FLAIR image where the signal from the hyperintense CSF around it has been suppressed.
of which will be aligned with the center of the acquisition window while the FE gradient is on. • A relatively long TR is necessary in spin echo pulse sequences to allow for enough recovery of the longitudinal magnetization prior to the application of the next 90° RF pulse. However, long TRs equate to long acquisition
I m age C h a r ac t e r is t ic s i n M R I a n d P r i nc i pa l P u l s e Se qu e nc e s
Fig. 2.18 Transverse T1-FLAIR pre-contrast (top) and postcontrast (bottom) images at the level of C1 in a dog with a meningioma. Even on the pre-contrast image, the outline of the meningeal tumor (arrows) along the left side is well seen due to the suppression of the CSF signal (*), which appears very dark on these images. C, spinal cord.
times, and it would be beneficial to use significantly shorter TRs to decrease acquisition times. As shown in Fig. 2.19, when short TRs are used (TR < T1), there is insufficient recovery of the longitudinal magnetization between successive 90° RF pulses, a phenomenon known as ‘saturation’; this results in low amplitude transverse magnetization induced by each 90° RF pulse, and therefore low signal. • There is another strategy to generate an echo that does not rely on a 180° RF pulse, but on the use of a ‘reversal gradient’, and for that reason is called ‘gradient echo’ pulse sequence. • As you can see in Figs. 2.20 and 2.21, we can use two successive gradients of opposite polarities to dephase and then rephase the protons, thereby generating a regrowth in transverse magnetization signal in the form of an echo. This causes the free induction decay (FID) signal to rapidly decrease as protons are dephased, to then regrow in
Mzmax = M0 Mz_eff
90° RF pulse
Large transverse magnetization
Mzmax = M0 Small MZ_EFF
90° RF pulse
Small transverse magnetization
Fig. 2.19 Effect of the TR on the available longitudinal magnetization to generate signal in spin echo. If complete recovery were allowed during TR, the longitudinal magnetization would fully regrow to reach the maximum value Mzmax = M0 (see Chapter 1), and therefore each 90° RF pulse would generate a maximum transverse magnetization for signal generation. In practice, the longitudinal magnetization is not allowed to recover completely as this would generate very long acquisition times. Top: When TR is long and on the order of or superior to T1, there is a large amount of longitudinal magnetization recovered between each TR (Mz_eff = effective longitudinal magnetization), which will generate a substantial transverse magnetization when the 90° RF pulse is applied. Bottom: When TR is short (