Basic concepts of MR imaging, diffusion MR imaging, and diffusion tensor imaging.

Abstract

MR image contrast is based on intrinsic tissue properties and specific pulse sequences and parameter adjustments. A growing number of MRI imaging applications are based on diffusion properties of water. To better understand MRI diffusion-weighted imaging, a brief overview of MR physics is presented in this article followed by physics of the evolving techniques of diffusion MR imaging and diffusion tensor imaging.

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Basic Concepts of MR
Imaging, Diffusion MR
Imaging, and Diffusion
Tensor Imaging
Eduardo H.M.S.G. de Figueiredo, BSca,*,
Arthur F.N.G. Borgonovi, BScb,c, Thomas M. Doring, MScd,e
BASIC PHYSICS OF MAGNETIC RESONANCE
Magnetic resonance (MR) imaging stems from the
application of nuclear magnetic resonance (NMR)
to radiological imaging. The adjective “magnetic”
refers to the use of magnetic fields and “reso-
nance” refers to the need of matching the
frequency of an oscillating electromagnetic field
to the “precessional” frequency of the spin of
some nuclei in a tissue molecule.1
Interest in medical diagnostic possibilities of
NMR began in 1971, with the study by Damadian2
of the differences in relaxation times T1 and T2,
among different tissues, and between normal
and cancerous tissues.3In 1973, the imaging
area for MR started with pioneering articles pub-
lished by Lauterbur4and Mansfield and Grannell,5
when the idea of spatial varying magnetic fields to
give the localization
introduced.
The phenomenon of magnetic resonance is
basedon theinteraction
magnetic fields and nuclei, which have a nonzero
magnetic moment. According to classical theory
of electromagnetism, individual nuclear moments
information wasfirst
betweenexternal
called spins in a static magnetic field B0precess
with Larmor frequency w0about B0. A bulk of spins
forms the net magnetization vector pointing along
B0. When radiofrequency (RF) is applied in this
system at Larmor frequency, the spins absorb
the radiofrequency energy and the net magnetiza-
tion vector flips by a certain angle in relation to B0.
The net magnetization vector can be decomposed
into 2 components, a longitudinal component
paralleltoB0andatransversalcomponentperpen-
dicular to B0. As the transversal component
precesses around a receiver coil, it induces
a current in that coil, in accordance with Faraday’s
law of induction. This current becomes the MR
signal.6
After stopping sending RF to the spins system,
the MR signal decays mainly via 2 processes:
loss of phase between spins and energy release
to the environment. The loss of phase between
spins can occur due to interaction spin-spin, and
this process is described as T2 relaxation, or it
can occur due to B0inhomogeneities, described
as T2* relaxation. In both ways, the transverse
component of net magnetization vector decreases
and even though there is energy in the system, the
aGE Healthcare, Avenida das Nac ¸o ˜es Unidas, 8501, 3 andar, 05425-070, Sa ˜o Paulo, Sa ˜o Paulo, Brazil
bHospital das Clı ´nicas, Avenida Dr. Ene ´as de Carvalho Aguiar, 255, 3 andar, 05403-900, Sa ˜o Paulo, Sa ˜o Paulo,
Brazil
cHospital do Corac ¸a ˜o, Avenida Dr. Ene ´as de Carvalho Aguiar, 255, 3 andar, 05403-900, Sa ˜o Paulo, Sa ˜o Paulo,
Brazil
dFederal University of Rio de Janeiro, Avenida das Ame ´ricas, 4666, grupo302A, Barra da Tijuca, Rio de Janeiro,
Rio de Janeiro, Brazil
eClı ´nica de Diagno ´stico por Imagem, Avenida das Ame ´ricas, 4666, grupo302A, Barra da Tijuca, Rio de Janeiro,
Rio de Janeiro, Brazil
* Corresponding author.
E-mail address: eduardo.figueiredo@ge.com
KEYWORDS
?Magnetic resonance imaging?Diffusion-weighted imaging
?Diffusion tensor imaging
Magn Reson Imaging Clin N Am 19 (2011) 1e22
doi:10.1016/j.mric.2010.10.005
1064-9689/11/$ e see front matter ? 2011 Published by Elsevier Inc.
mri.theclinics.com

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MR signal decays. The interaction between spins
and the environment they insert, called spin-
lattice interaction and described as T1 relaxation,
causes the spins that form the net magnetization
vector to release their energy, and its longitudinal
component grows back along the B0direction.
The tissue relaxation characteristics are ex-
pressed in image contrast and are controlled by
the pulse sequence chosen. The “pulse sequence”
is a sequence of RF pulses, magnetic field gradient
pulses, signal sampling, and time periods between
them. RF pulses are basically responsible for exci-
tation of spins and their manipulation to obtain
a signal echo. Magnetic field gradients are respon-
sible for selecting the slice to be imaged, spatially
encoding the signal induced in receiver coils, using
frequency and phase as location information and
in some pulse sequences, gradients also control
the image contrast. Spin echo and gradient echo
are examples of pulse sequences that control
image contrast, by applying RF pulses or gradient
pulses, respectively.
Although tissue relaxation characteristics are
the main source of contrast information, MR
images can represent other aspects of the biologic
architecture. Random thermal motion of spins in
a gradient field causes a phase shift of their trans-
verse magnetization with respect to static spins,
and can be used as a source of contrast using
proper pulse sequence. To better understand
this mechanism of image acquisition, a brief over-
view of MR physics is presented in this article.
Magnetic Resonance Concepts
MR imaging works because, in the presence of an
external magnetic field, one can measure the inter-
action between protons in human body and the
external magnetic field itself. Protons interact
with the external magnetic field, due to an intrinsic
magnetic characteristic called spin. The classical
view of spin is the effect of one charged parti-
cleda proton, for exampledspinning around
itself, which creates a magnetic moment pointing
perpendicularly toward a spinning axis (Fig. 1). In
the presence of an external magnetic field, this
proton behaves similarly to a magnetic bar, tend-
ing to align to the field and precesses around it.
When we are dealing with individual protons, we
needtolookatthemthrough“quantummechanical
glasses,” and “not-so-intuitive” thoughts are
permitted under this treatment. A not-intuitive
behavior is that this proton can align in either
a parallel or an antiparallel orientation to the
external magnetic field. There is a difference of
energy between a parallel and an antiparallel state,
andifthisexactamountofenergyismatchedbyan
incident radiation and is delivered to the proton, it
absorbs this energy and changes from one state
to another. This phenomenon is called magnetic
resonance.Thedifference
dependsonthemagnitudeoftheexternalmagnetic
field B0where this proton is inserted, and is ex-
pressed in terms of the Larmor equation (1) as:
of energy states
w05yB0
(1)
In equation (1), w0is the frequency needed from
the incident electromagnetic field to match the
energy difference between states of nuclei, with
the gyromagnetic constant y in the presence of an
external magnetic field of amplitude B0. The gyro-
magneticconstantyisanimportantnucleuscharac-
teristic. Human body abundance and proper
gyromagnetic constant make hydrogen nuclei (H1)
the best choice for magnetic resonance imaging.
But in MR imaging, we are not dealing with
a single spin and instead of a quantum mechanics
view, a classical treatment can be used to under-
stand MR imagingphysics.
mechanics view, spins in the presence of an
external magnetic field B0start to precess around
B0, with an angular frequency w0 (Fig. 2). The
precessional frequency is given by the same rela-
tionship expressed in quantum mechanics treat-
ment, the Larmor equation.
The sum of many spins rotating around B0forms
the magnetization vector M that points in the same
direction as B0. When RF wave in resonance
frequency w0 irradiates the bulk of spins, the
quantum effect of states transition is analogous
to flip the vector M by a certain angle in relation
to B0. If RF in resonance frequency is applied by
a certain time that flips the vector M 90?, this RF
pulse is said to be a “90?RF pulse.” The same
nomenclature applies to a “180?RF pulse.”
Consider the following experiment: a bulk of
spins in the presence of external magnetic field
B0. At this time, the equilibrium time, vector M,
Inaclassical
Fig. 1. Intheclassicalview,achargedparticlewithspin
can be compared as a magnetic dipole such as a bar
magnet.(Courtesyof GEHealthcare,Sa ˜oPaulo,Brazil.)
de Figueiredo et al
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possesses only the longitudinal component Mz. A
90?RF pulse is applied. After turning RF off, vector
M has only the transversal component Mxyand
under the influence of B0starts to precess around
it. At this time, vector M induces current in the
receiver coil, according to Faraday’s law of induc-
tion, and generates the MR signal (Fig. 3). When it
passes through the receiver coil a maximum signal
is obtained, and when it is the farthest from the coil
a minimal signal is obtained. Plotting signal versus
time, a sinusoidal function is to be expected.
Butinsteadofasimplesinusoidalfunction,asinu-
soidalfunctionwithamplitudedecreaseovertimeis
obtained (Fig. 4). The signal represented in Fig. 4 is
called FID (free induction decay), which decays due
to a process known as relaxation.
Relaxation Effects
Signal relaxation is a result of loss of phase
between spins and energy release to the environ-
ment where spins are inserted. Immediately after
a 90?RF pulse, all spins that form the Mxyvector
are pointing in the same direction, and they are
said to be “in phase.” Turning RF off, they are
set free to precess around B0, and they precess
with an angular frequency according to the Larmor
equation. The first problem is that not all spins feel
the same static field, due to B0inhomogeneities,
and as they feel differently, they precess with
different frequencies. This difference
a loss of phase between them and after some
time carrying out vector addition, it can be under-
stood why the signal vanishes (Fig. 5). This
process is called T2* decay or T2* relaxation,
and effects of this nature can be restored once it
is not a random effect but a scanner characteristic.
The same loss of phase described here can
happen even though scanner B0homogeneity is
perfect. Spins that form vector Mxycan interact
with each other, feeling the tiny magnetic field of
neighboring spins, in addition to B0, and each
spinstartstoprecesswithadifferentfrequency,re-
sultinginthesamelossofphasecausedbyB0inho-
mogeneities. The central difference is that this loss
ofphasecannotberecoveredoncethermalmotion
of spins is random. This process is called T2 relax-
ation or T2 decay. T2* and T2 decay express expo-
nential behavior (see Fig. 4) and are modeled in
terms of transversal components of vector M.
Another source ofsignal decay is spins giving up
energy, which are absorbed by an RF pulse to the
causes
Fig. 3. In the transversal plane, rotating magnetiza-
tion vector M induces current in receiver coil accord-
ing to Faraday’s law of induction. (Courtesy of
Healthcare, Sa ˜o Paulo, Brazil.)
GE
Fig. 4. MR signal amplitude (proportional to M vector
magnitude) as function of time. The exponential
damped sine function is called FID and is character-
ized by relaxation effect. (Courtesy of GE Healthcare,
Sa ˜o Paulo, Brazil.)
Fig. 2. Precession of spin axis around B0. (Courtesy of
GE Healthcare, Sa ˜o Paulo, Brazil.)
Basic Concepts of MRI
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environment. In terms of magnetization vectors,
vector Mxy decreases and Mz increases over
time, until all energy has been released and the
vector M has only a longitudinal component
(Fig. 6), returning to equilibrium. The energy
exchange is governed by the interaction between
spins and the lattice, and it is called T1 decay or
T1 relaxation.
T1, T2, and T2* are tissue properties, and are
measured in seconds (T2* carries B0inhomogene-
ities information, caused by the tissue and the
scanner). T1 is the time needed for Mzto achieve
approximately 63% of its initial value, after a 90?
RF pulse. T2 and T2* are the times taken for Mxy
to achieve approximately 37% of its initial value,
after a 90?RF pulse, T2* being measured taking
into account inhomogeneity effects and T2 being
measured taking into account only spin-spin inter-
actions. Image contrast contains all relaxation
effects, but usually they are weighted in one effect,
meaning that differences in the gray scale mostly
representdifferencesin tissue relaxation
properties. Image weighting is controlled by pulse
sequences.
Pulse Sequences
In
magnetic field is not particularly uniform, dephas-
ing between spins caused by field inhomogenei-
ties arethemain source
Fortunately, this effect can be reversed by a well-
known RF pulse sequence called the “spin echo
method.”
The spin echo sequence is based on the appli-
cation of 2 RF pulses: a 90?RF pulse (or excitation
pulse) followed by a 180?RF pulse (or refocusing
pulse). The 90?RF pulse tips all spins that form
the vector M into a transversal plane and immedi-
ately after they reach the plane, they are in phase
and Mxyhas its maximum amplitude. After turning
the RF pulse off, spins start to precess and lose
phase by T2* relaxation effects. The 180?RF pulse
is applied after a time t, defined as t 5 TE/2 (t is set
equal to 0 at the time the first 90?RF pulse is
applied), rotating the spins by 180?in relation to
the position they were in. At the end of refocusing
the pulse application,
remains in the transversal plane, but “faster” spins
with higher precessional frequency are put behind
“slower” spins with lower precessional frequency.
The accumulated phase between spins caused by
field inhomogeneities through the time from t 5 0
seconds to t 5 TE/2 seconds is compensated after
a 180?RF pulse at the time t 5 TE seconds (called
echo time), TE/2 seconds after refocusing pulse.
Thewholepicture is
Figs. 7 and 8.
The realignment of spins is called spin echo. It is
possible to apply many refocusing pulses after
a 90?RF pulse, collecting many spin echoes as
shown in Fig. 8. B0inhomogeneities are canceled
in this method, because they are static in time, and
thecircumstance wherebythe external
of signalloss.
the spins localization
betterunderstoodin
Fig. 5. Signal loss due to dephasing between spins, caused by T2* and T2 relaxation. (Courtesy of GE Healthcare,
Sa ˜o Paulo, Brazil.)
Fig. 6. The decay of transverse magnetization occurs
by means of T2 and T2* effects, while recovery of
longitudinal magnetization reflects energy transfer-
ence to the environment. (Courtesy of GE Healthcare,
Sa ˜o Paulo, Brazil.)
de Figueiredo et al
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the same B0inhomogeneity that spins feel before
refocusing pulse they also experience after it.
But spin-spin interactions are not static in time
and T2 decay cannot be avoided. The longer the
time a spin echo is collected, the stronger a T2
relaxation effect is presented.
Another important figure in pulse sequence is
the repetition time TR, the time between 2 excita-
tion pulses. In the spin echo sequence, TR is the
time between 2 90?RF pulses applied in the
same location.
TR and TE are the main parameters in spin echo
sequence used to control contrast image weight-
ing. Signal is acquired in echo time t 5 TE, and ac-
cording to parameters set in the sequence, the
spin echo received in the receiver coil will express
differences among tissues, regarding proton
density (PD), and T2 or T1 relaxation.
Consider 2 different magnetization vectors, Ma
and Mb,possessing different PD, T1, and T2 prop-
erties. At equilibrium state, M vector magnitude
depends on PD available in the tissue. A 90?RF
pulse isapplied inthe system,and Maand Mbrelay
into the transversal plane. Dephase between spins
starts at different rates, and a refocusing pulse is
applied at TE/2 to provide a spin echo at TE. If
TE is short enough, T2 differences will not be rele-
vant and the signal induced in the receiver coil will
mostly represent PD properties. To fulfill this requi-
site, TR must be kept long enough, to allow M
vectors to recover all their longitudinal magnetiza-
tion and so that T1 differences will not influence
Mxyvector after the next excitation pulse (Fig. 9).
Therefore, PD-weighted images acquired with
a spin echo sequence are obtained by using long
TR and short TE.
If a T2-weighted image is desired, a long TR is
needed to avoid T1 influence through excitation
pulse repetitions, but TE must be adequate to
reflect T2 relaxation differences in signal ampli-
tudes acquired (Fig. 10). The optimal TE to opti-
mize contrast between Ma and Mb is the time
when the difference between transverse magneti-
zation of Ma and Mb is larger. Therefore, T2-
weighted images acquired with a spin echo
sequence are obtained by using long TR and
adequate TE.
To acquire T1-weighted images, a short TE is
needed for a signal echo not to reflect T2 differ-
ences. After applying the first excitation pulse,
Maand Mbremain in a transversal plane and de-
phasingbetween spinsiscausedby T2*
Fig. 7. Themechanismofspinechowithapplicationof180?RFpulse.(Courtesyof GEHealthcare,Sa ˜oPaulo,Brazil.)
Fig. 8. A spin echo experiment with acquisition of 2 echoes. (Courtesy of GE Healthcare, Sa ˜o Paulo, Brazil.)
Basic Concepts of MRI
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