Lecture Notes MRI(Part2)

8/13/2019 Lecture Notes MRI(Part2)

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Medical Applications of MRI 2011 handout 1

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Medical Applications of MRI

Dr John Thornton

e-mail: [email protected]

6 Lectures:

Review of nuclear magnetisation & the magnetic resonance signal

Conventional MRI:

Spin echo and gradient echo imaging

Proton density, T1, T2 and T2* image contrast

Fast Imaging Methods

Magnetization Transfer Contrast

Diffusion-weighted MRI

MRI Contrast Agents

MR angiography

MR Perfusion ImagingFunctional MRI

MRI Safety

Suggested Textbooks:

MRI From Picture to Proton, 2nd

Edition (McRobbie, Moore, Graves and Prince, 2007)

Magnetic Resonance Imaging (Kuperman, 2000)

MRI Basic Principles and Applications, 4th

Edition, (Brown & Semelka, 2010)

Terminology:

In these lectures:

Spins = nuclear dipoles = nuclei with magnetic moment = nuclear magnets

MR = (nuclear) magnetic resonance = NMR

Contrast = difference in image intensity for different tissues

= difference in numerical values for relevant pixels

difference in MR signal intensity originating from the each tissue

Overall Learning Objectives

By the end of these 6 lectures you should be able to understand and describe

the physical origin of image contrast in MRI

how scanner parameters may be selected to provide different types of medically useful contrast

How we can exploit the physics of magnetic resonance and tissue-water behaviour to providesophisticated structural and functional information


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Magnetization, RF pulses and the MR signal

Some nuclei (in particular the hydrogen nucleus, or proton) have the quantum mechanical property of spin and behave

as microscopic magnets.

In the presence of a magnetic field Bo, on average, more nuclear spins align with the field than against and an

equilibrium total magnetisation (Mo) parallel to Bo is produced.

By convention the direction of Bo (and therefore Mo) defines thez axis

z

B1

M0

B0

B0

z

At equilibrium: M0

z

M0

B0

x

B1

Apply a second magnetic field, B1, perpendicular to thez

axis and rotating at the Larmor frequency: =.Bo

To simplify things, we take a coordinate system

(the rotating frame) which rotates at the same

rate as B1:

B1 now appears stationary


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B1 exerts a torque on Mo causing it to rotate down towards thex-yplane:

A B1 field applied for a time just sufficient to rotate Mo through 900

is termed a 900

pulse and generates a transverse

magnetization Mxy

Immediately following a 900

pulse, the longitudinal component ofM, i.e. Mz, is zero.

Viewed in the laboratory frame, Mxy now rotates about the z axis with an angular frequency = -B0 and therefore

induces a signal in a receiver coil.

The process of creating longitudinal magnetization (Mxy) e.g. by means of a 900

RF pulse, is known as excitation.

Relaxation

after the 90o

RF pulse the magnetization returns towards equilibrium (relaxes):

M0

B1After 90

opulse |Mxy| = |Mo|

Mz = 0

After a short time (typically a

few ms in MRI), Mo is rotated by

90o

onto thex-yplane

B1 is then removed

NB The detected signal magnitude (and hence final image intensity) is proportional to |Mxy|.

No |Mxy|: no signal!

(Mz is stationary both in the rotating frame and the laboratory frame, and therefore induces no

signal voltage in the receiver coil)

Immediately

following 90o

pulse:

Mxy = |Mo|

Some time

later:

Mxy < |Mo|

Mz

Some more

time later,

t >> T1:Mz = Mo

Magnetization has returned to

equilibrium state

x


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The components of the magnetization before and after a 90o

pulse may be represented graphically:

The longitudinal magnetization (Mz) recovers to equilibrium exponentially with a time constant T1:

Mz = Mo.(1-exp(-t/T1)) (T1or spin-lattice relaxation) [eqn 1]

where t represents time from the end of the 900

pulse

The transverse magnetization (Mxy) decays to zero with a time constant T2:

Mxy = Mo.exp(-t/T2) (T2 or spin-spin relaxation) [eqn 2]

T1 relaxation involves an exchange of energy between the spins and their environment (the lattice), causing the

relative populations in the up and down states to return towards their equilibrium distribution so that Mz recovers

towards M0

T2 relaxation involves a gradual reduction in Mxy by the above mechanism, and additionally because the spins interact

magnetically amongst each other (hence spin-spin) the net result of which is that they gradually lose phase coherence

(i.e. their individual magnetizations no longer all point in the same direction in thex-yplane) and hence their vector sum

(Mxy) decays in magnitude with time

Because of this extra contribution to the decay of Mxy, in biological tissue, T2 < T1

Inversion

Starting again from the equilibrium situation (|Mz| = | Mo|), the rotating B1 field is applied for twice as long as for a 900

RF pulse.

M therefore continues to rotate past thex-yplane until it is oriented along the z direction.

This is known as an applying an 180o

inversion RF pulse: immediately following this the magnetization is said to be

inverted.

The longitudinal magnetization then recovers exponentially towards equilibrium with time constant T1:

SOLUTIONS OF THE

BLOCH EQUATIONS

dM/dt = M x B

after modification to

include relaxation

effects

Exponential recovery: time constant T1

90opulse applied here

Mo

Mxy

Mz

Exponential decay: time constant T2

time

time

signal Mxy


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Following inversion, Mz is described by the equation:

Mz = Mo (1-2exp(-t/T1)) inversion recovery(IR)equation

Here t represents time after the inversion pulse

Note: Mxy is zero at all times in this case

Field inhomogeneities and the spin echo

In reality the main magnetic field, B0, is never perfectly uniform its value varies slightly with position. This B0inhomogeneity accelerates the decay of Mxy: spins in different positions experience slightly different magnetic fields,

and thus precess at slightly different frequencies.

This frequency variation causes them to get out of step with each other , i.e. lose phase coherence, more rapidly than

they otherwise would such that their vector sum decreases more rapidly than would be caused by spin-spin (T2)

relaxation alone.

The signal decay is therefore described by a new time constant: T2* (< T2)

Such dephasing and consequent accelerated signal loss is usually a problem, but can be useful as it may reflect local

physiological changes - e.g. as in fMRI (see later).

Spin Echo Formation

This accelerated (T2*) decay due to field inhomogeneities is reversible by the applications of an 180o refocusing RF

pulse to form a spin echo:

Immediately

after inversion

pulse:

M0

Mz

-M0

M0

Before

inversion

pulse:

Exponential recovery: time constant T1

more time

later:

Some time

later:

time

x


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Signal at spin-echo peak xy(TE)= Mo exp(-TE/T2)

Refocusing vs. inversion

Both refocusing and inversion involve an 1800

RF pulse, but the effect is different because the magnetization is

orientated differently at the start of each pulse:

Inversion:

Start with magnetization along the positive z direction, Mz = Mo; Mxy = 0. Apply B1 pulse along+yin this case (1800

y)

M0

B1

z

180o

refocusing pulse

(pancake flipper)

flips spins about x-axis

Mxy

Time constant

T2*

180o rf

pulse

Time constant T2

180o

MxyMxy

Individual nuclei precess

at slightly different

frequencies

immediately after a 900

pulse

x

timet=0 t=TE/2 t=TE

Signal maximum occurs at the Echo Time (TE)

Signal,

S Mxy


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Refocusing:

Starts with magnetization in the x-y plane, i.e. Mz =0; Individual spins have become out of phase: magnetization is

fanned-out in thex-yplane. Apply 1800

B1 pulse in this case along-x(1800

x), spins then tend to rephase.

Gradient echo formation

An echo signal may also be obtainedwithoutan 180o

refocusing RF pulse by performing a field gradient reversal.

We deliberately introduce strong magnetic field inhomogeneities by applying a (negative) linear gradient. This greatly

increases the effective rate of T2* signal decay.

If the field gradient is then reversed, after an appropriate time the phase changes caused by the negative gradient are

exactly cancelled by the (opposite) phase changes caused by the positive gradient: a gradient echo is formed.

Note only phase differences due to the applied gradients are reversed: signal loss due to field inhomogeneities intrinsic

to the tissue (or scanner Bo field) is not recovered and the amplitude of the gradient echo is determined by the natural

T2* decay rate.

Signal at gradient-echo peak Mxy(TE)= Mo exp(-TE/T2*)

TE

time

time

Gradient,

G

negative Gx positive Gx

Mxy

Time constant

T2*

Area A

Area Bgradient-echooccurs when

Area A = Area B

x

B

x

B

z

Dephasing occurs much more

rapidly than for the usual T2*

decay due to large additional field

inhomogeneities caused by the

applied gradient

Just after 180opulseJust before 180

opulse

Mxy

B1

x

Mxy

B1

x

Rotate by 180 about

x axis


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The Spin-Echo Imaging Sequence

Both TE and TR are under software control and may be chosen by the operator to change the image contrast.

In general,

the magnitude and timing of the RF (B1) pulses determines the image contrast.

The field gradients Gx, Gy and Gz provide spatial discrimination

These two aspects of image formation (contrast and localization) can be to a certain extent considered independently.

Image

contrast

Spatial

ocalisation


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MRI in medicine

Most clinical MRI is performed to provide structural & anatomical information: e.g. to detect atrophy, tumours, the

presence of haemorrhage, evidence of ischaemia (stroke) etc.

MRI examinations may also provide functional information relating to e.g. blood flow in vessels, cerebral activation

(fMRI), organ perfusion, uptake of tracer agents etc. giving additional information regarding tissue status(viable/compromised/activated etc.)

Since the hydrogen nucleus (proton) has both the highest NMR sensitivity, and a very high concentration in the human

body (primarily as H2O), clinical MRI is concerned almost exclusively with imaging this nucleus, and therefore MRI scans

usually represent maps of the distribution of water in the body (there is also some contribution from protons in lipid

(fat) molecules ).

A primary source of contrast in MRI therefore relates to the concentration of hydrogen nuclei (proton density) in

particular tissue types. (NB the equilibrium magnetisation Mo is proportional to the proton density).

As a consequence of differences in their microscopic environment protons in different tissue types exhibit different T1and T2 relaxation times, e.g. in head imaging at 1.5Tesla approximate values are:

Protons associated with solid structures (e.g. compact bone) have T2 relaxation times too short for them to be

detected in standard MR images and hence appear dark.

Because different tissues exhibit different relaxation times, and since pathology (tissue injury) can change these

relaxation times, MRI is said to have excellent soft tissue contrast.

By choosing suitable scan parameters, the image contrast may be manipulated to reflect these proton density andrelaxation time differences.

Contrast in conventional MRI

The signal (i.e. numerical value in each pixel) in a spin-echo image is given by:

Signal, S Mxy = Mo[1-exp(-TR/T1)].exp(-TE/T2)

By manipulating TR and TE we can change the signal dependency (weighting) with respect to M o, T1, and T2. This

generates different contrasts for tissues with different relaxation times and proton densities. We will now consider how

to generate images whose contrast is weighted in terms of the proton density, T2, T2* and T1 of the tissues beingimaged.

TISSUE T1(ms) T2(ms)

white matter 600-800 60

grey matter 800-1000 80

fat 150 50

cerebrospinal 2700 500fluid


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Signal Intensity in spin-echo imaging

The signal intensity in each pixel for a spin-echo image is given by:

S Mxy = Mo[1-exp(-TR/T1)].exp(-TE/T2)

S, the signal = numerical value of image pixel

or how dark or light the tissue appears in image

Mo = the equilibrium magnetization

number of nuclei per unit volume the concentration of water or fat molecules i.e. the proton density (PD)

Spin-echo imaging I: Proton density-weighted contrast

The proton density (PD) is proportional to the number of hydrogen nuclei (and hence water and lipid

molecules) in a given volume of tissue. is proportional to the PD. A proton-density-weighted imagegives contrast (relative signal intensity) dependent almost exclusively upon with very small or negligibledependence upon T1 or T2.

S PD.[1-exp(-TR/T1)].exp(-TE/T2)

T1 = longitudinal (spin-lattice) relaxation time constant

T2 = transverse (spin-spin) relaxation time constant

PD = the proton density

TE = the echo time

TR = the repetition time

Properties of the tissue

Scanner parameters

under control of the

operator

We can choose TE and TR to vary the amount by which S

depends upon T1, T2 or PD


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For PD weighting:

Signal Mxy = Mo[1-exp(-TR/T1)].exp(-TE/T2)

= Mo

Requires long scan times if TR is sufficiently long to completely remove T1 dependence (for which TR should

be T1 so that the spins are fully relaxed between 90o pulses) modern MRI systems use fast-spinecho techniques to reduce scan time (see later).

N.B. PD weighted contrast may also be obtained using a gradient-echo pulse sequence with low flip angle

also see later.

Spin-echo imaging II: T2weighted contrast

We use a spin-echo sequence, this time with both long TR and long TE:

90oSpin-echo signal

Etc.

NB Abbreviated version of pulse-sequence on page 8 not showing

the imaging gradients

Typical values:

TE 10 msTR 4000 ms

TR

TE

180oTE

TR long TE short

0 1

TE

TR

90o 180

o

Spin-echo signal

Etc.

NB Abbreviated version of pulse-sequence on page 8

T

90o

180o

Typical values:

TE 100 ms

TR 4000 msN.B sequence identical to previous figure,

except TE longer!

TR Etc.


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For T2 weighting:

Signal Mxy = Mo[1-exp(-TR/T1)].exp(-TE/T2)

Therefore to a good approximation:

Signal Mxy(TE) = Mo.exp(-TE/T2)

and 2 tissue types (e.g. white and grey matter in the brain) having different T2 values will give different

image signal intensities even if they have identical PDs.

Increasing the echo time (TE) increases the degree of T2 signal dependence (i.e. increases the "T2 weighting"

or T2 contrast).

Note the image signal intensity in the above equation is still dependent upon M0 (i.e. PD), but the image is

additionally T2 weighted. Since the PD (and hence M0) may not vary much between different soft tissues,

with sufficiently long TE theT2 effect predominates.

Like PD-weighted spin-echo imaging, T2-weighted imaging requires long scan times if we wait for complete

relaxation to completely eliminate T1 influences (i.e. TR T1 as above), but scan times are againcommonly decreased by using the fast spin-echo method (see later).

Spin-echo imaging III: T1weighted contrast by progressive saturation

We use a spin-echo sequence this time with a short TR (


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In the steady state (once the sequence has been running long enough for an equilibrium situation to havebeen developed) the signal intensity is again given by the spin-echo signal equation. This time the echo time

TE is made as short as possible, so the signal has very little dependence upon T2:

Signal, S Mxy Mo[1-exp(-TR/T1)].exp(-TE/T2) Mo[1-exp(-TR/T1)]

So image contrast now depends on M0 and T1, and for sufficiently short TR, the T1 dependence (contrast)

dominates. In this case signal the intensity of longer T1 tissues is preferentially decreased.

Advantage of this approach: short TR means T1-weighted images can be produced in a relatively short

acquisition time.

Spin-echo imaging IV: inversion recovery imaging

We can add an extra 180o inversion pulse to the start of each repetition of a spin-echo imaging sequence.

We then wait for a time TI (the inversion time) after this pulse before continuing with the rest of the spin-

echo imaging sequence. This introduces a different kind of T1 signal dependence.

Each 900

pulse shown is in fact

the first RF pulse for eachrepetition of the spin-echo

imaging sequence

As usual the phase encode

gradient in incremented for

each repetition

Repeat until the required number of phase

encoding steps have been acquiredShort T1

tissueLong T1

tissue

time

Also known as a progressive saturation pulse sequence

90o 90o 90o 90o

TR

90o 90oMz

(v. short TE


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During the inversion time TI there is partial recovery of Mz towards M0 by T1 relaxation. Again we consider

behaviour of Mz:

Repeat e.g. 256 times with increase phase encoding gradient

TI

slice

select

Phase

encode

read

90o

180o Spin echo

180

inversion

pulse

Repeat until the required

number of phase encoding steps

have been acquired

180

T contrast

90oM0

Mz

-M0

TR

TI

lon T1 tissue

shortT1 tissue

time

T contrast

90o

Again follow each 90 pulsewith an 180

0refocusing pulse

and apply imaging gradients

Phase encode gradient

incremented for each repetition


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The full expression for the signal intensity is:

S Mxy = Mo [1-2.exp(-TI/T1)+ exp(-TR/T1)].exp(-TE/T2)

For long TR ( T1) and very short TE ( T2), this simplifies to

S(TI) Mxy = Mo (1-2.exp(-TI/T1))

S depends heavily on T1 - inversion recovery imaging can give strong T1 contrast

Selective Nulling

Note that the IR method forces the longitudinal magnetization to pass through zero. If the image data is

acquired at the exact time when Mz for a particular tissue passes through zero we get no signal for this

tissue it has been nulled and will appear black in the final image.

This occurs at TI = ln(2).T1 = 0.693T1

We can selectively eliminate signals from specific tissues in this way:

e.g. STIR (Short TI inversion recovery): reduces the signal from fat tissue (fat suppression)

Fat has a T1 of approx. 200 ms at 1.5T. Therefore the inversion recovery curve for fat tissue

passes through zero at 0.693x200 = 139 ms

FLAIR (Fluid Attenuated Inversion Recovery) suppresses long T1 fluids e.g. cerebrospinal

fluid (CSF)

180

inversion

pulse

90o 90o

Mz

time

Short T1 tissuelong T1 tissue

Acquire image here:

null short T1 componentAcquire image here:

null long T1 component


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Advantages of inversion recovery sequence for T1-weighted imaging:

There a greater dynamic range of potential T1 contrasts available due to relaxation from Mo to + Mo(compared with progressive saturation)

Selective nulling is possible (see above)

Disadvantage: long scan times due to long TR requirement (usually mitigated using fast-spin echo

methods).

(In addition to progressive saturation and inversion recovery techniques, T1-weighted imaging may also be

obtained using a gradient-echo imaging sequence - see later).

CONTRAST AND NOISE

An objective in medical is to maximize contrast between different tissues, e.g. between healthy and

diseased tissue, while minimizing the deleterious effects of image noise.

To separate features in our object we must have image contrast, here defined as a difference in image

signal intensity between tissue types or between normal/abnormal tissue in the same organ.

As already mentioned, the image signal in MRI is always dependent to a certain extent upon Mo (i.e.

proportional to the proton density), but usually the image siganl is also made to depend upon (weighted in

terms of) other factors e.g. T1 or T2 (and other types of contrast to be described later).

Our aim is always to maximise signal contrast, but we must also take into account the influence of noise inthe imaging system.

Consider the case of a T2-weighted sequence for 2 tissues with identical M0 but different T2s. (Similar

considerations apply for other contrasts e.g. T1-weighted contrast etc.). The graph shows how the tissue

signal levels depend on the choice of echo-time, TE.

There is a TE which gives a maximum difference between the curves and therefore images collected withthis TE will show the greatest contrast.

Maximum

Signal difference (contrast)

Time (TE)

Signal

Signal from tissue 2:short T2

Signal from tissue 1:

long T2

Noise level 1

Noise level 2


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However, we have to consider the effects of image noise. Noise in MRI originates largely from the random

thermal motion of ions within tissue, which induce random electrical signals in the MRI receiver coil. The

precise mean noise level in an image depends upon a number of factors, including the volume of tissue

within the receiver coil and the receiver bandwidth.

At the echo-time giving the maximum signal difference the signal-to-noise ratio may be poor: e.g. if the

noise were at Noise Level 2 in the example in the diagram, the image would be largely obscured by noise

and diagnostically useless. Thus in order to obtain the most diagnostically sensitive images, a compromise

between contrast and SNR may be required, and we aim to maximise the contrast-to-noise ratio (CNR)

defined as

CNR = (S1-S2)/Snoise

where S1 and S2 are the signals from tissues 1 and 2, and Snoise is the mean noise level.

S1/Snoise and S2/Snoise are the signal to noise ratios (SNRs) for the 2 tissues.

Image noise affects imaging performance in 2 ways:

1. it limits the minimum signal levels which may be detected

2. it may also limit the maximum image contrast which may be obtained

For a particular pulse sequence, the objective is always to select pulse sequence parameters which

maximize the contrast-to-noise ratio for the tissues/pathologies in which we wish to discriminate.


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T1-weighted:

We use a shorter TR to increase the T1 weighting in the final image

TR is too short to allow complete T1 relaxation, so

Mxy| after 90o =Mz | before 90o |M0| and the image signal is much reduced

However, scanning time is now faster: 256 x 0.5s 2 minutes

To scan even faster:

We might try drastically reducing TR to decrease the scan time even further.

However, this would cause the remaining signal to be too small too be useful (remember the importance of

maintaining an adequate contrast-to-noise ratio):

In this case, |Mxy| after 90o (= |Mz| before 90

o) is so small that signal would be below the noise level,

and therefore useful images may not be obtained in this way.

Mz immediately prior to each RF pulse, and hence the subsequent signal, may be increased by reducing the

disturbance from equilibrium at each repetition, i.e. by using an excitation pulse with a flip angle < 90o

:

M0

Mz90

o90

o90

o90

o

timeVery short TR (e.g. < 0.1s)

short TR (e.g. 0.5s)

M0

Mz

90o 90o 90o 90o

time

Smaller Mz < Motherefore smaller

signal


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e.g. immediately following a 30o

RF pulse

Mz = MZstart.cos(30o) = 86.6 % of |MZstart|

Mxy = MZstart.sin(30o) = 50 % of |MZstart|

i.e. we can generate significant Mxy without reducing

Mz by very much.

Considering again for comparison the case of 900

RF pulses applied with short TR:

This is effectively equivalent to a very heavily T1-weighted progressive saturation sequence (as discussed

earlier), where, assuming a TE sufficiently short to ignore T2 effects has been chosen,

Signal Mo[1-exp(-TR/T1)]

For low flip-angle RF pulses, i.e. flip angle < 90o, solution of the Bloch equations shows that theequilibrium value of Mz immediately prior to each RF pulse is greater than for the case of 90

opulses, and

hence the available signal is also greater.

M0

Mz o o


MZstart

B1

z

Mz

Mxy

= 30o

M0

Mz90

o90

o90

o90

o



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The signal is now given by a more general equation:

S Mo[1-exp(-TR/T1)].sin() (again assuming v. short TE)1-cos() exp(-TR/T1)

A maximum signal for a tissue with given T1 is obtained at a specified TR by using a flip angle

satisfies the equation:

cos( expTRT and is in this case known as the Ernst angle.

In order to further decrease the total acquisition time, the signal is acquired using a gradient echo

(saving the time required for a refocusing pulse and its associated field gradients), so that a shorter

TR may be used in order to minimize the time taken to acquire the image:

Example: for a tissue with T1 = 800ms; TR = 100ms

= 90o: Signal Mo[1-exp(-TR/T1)] = 0.12 x M0

= 30o: Signal Mo[1-exp(-TR/T1)].sin(30o)1-cos(30o) exp(-TR/T1)

= 0.25 x M0 i.e. signal strength doubled !

Also assumes sequence is

spoilt see below


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Because the TR is short compared with T2, immediately before each0 RF pulse there may remain asubstantial transverse magnetization, Mxy (in a spin-echo sequence, this would normally have

decayed to zero). Left alone, this magnetization might interfere with the transverse magnetization

produced by the subsequent RF pulses, causing errors in the image formation process. The purpose

of the crusher gradients is to avoid this by completely dephasing the transverse magnetization, such

that the vector sum of all the individual nuclear moments contributing to M xy is zero before the next

RF pulse is applied.

TE

0, slice selectiveRF pulse gradient echo

acquire

0

slice select

gradient

readgradient

Phase encodinggradient

spoiler gradientsensure Mxy = 0 before

next0 pulse

TR (short !)

Phase encodegradient

incremented foreach repetition

Repeat number

of phase encodingsteps

RF/signal


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Digression: T2*-weighted imaging

For simplicity, the above treatment assumed that the echo time TE was very short. More generally,

for a gradient-echo pulse sequence, explicitly including the TE:

S Mo[1-exp(-TR/T1)].sin() . exp(-TE/T2* )1-cos() exp(-TR/T1)

If TR is long, and small, this simplifies to: S Mo. exp(-TE/T2* )

For longer TEs (say 40ms) we can obtain contrast that depends upon T2* which is sensitive to local

disturbances of the magnetic field homogeneity, e.g. a local build up of iron in tissue as may occur in

blood and blood-breakdown products following bleeding (haemorrhage).

[Fast T2*-weighted images may also be obtained using echo-planar imaging see later]

Advantages of gradient echo imaging:

- Fast: Imaging times less than 1 minute possible

- can also obtain true 3D scans in reasonable examination times ( < 10 minutes)

- versatile: image contrast can be selected by choice of, TR and TE (the sequencemay additionally be preceded by an 180

oinversion RF pulse to provide additional T1-

weighted contrast)

Disadvantages of gradient echo imaging:

- SNR may be low (but can be improved by signal averaging)

- lack of 180o

refocusing pulse may result in susceptibility artefacts.

Gradient echo imaging: summary

Signal , S Mo[1-exp(-TR/T1)].sin() . exp(-TE/T2* )1-cos() exp(-TR/T1)

Appropriate choice of, TR and TE determines contrast:

Short TR ( 100ms), Short TE, and/or large ( 800) T1 weighting

Moderate TR ( 500ms), short TE and small (10-200) PD weighting

long TE ( 40ms), moderate TR ( 500ms) and small (10-200) T2* weighting


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Fast Imaging II: Echo Planar Imaging (EPI)

In both the gradient- and spin-echo imaging pulse sequences, a series ofn (= number of phase

encoding steps) gradient- or spin-echoes are acquired one at a time. For each acquisition the phase-

encoding gradient is increased by a fixed amount. Total imaging acquisition time isn TR.

In EPI, the equivalent dataset is acquired in a single shot following a single RF excitation pulse.

A slice-selective 90o

RF pulse is applied, followed by reversal of the read gradient to form a gradient

echo in the usual way. After the acquisition (digitization) of this gradient echo, the read gradient is

reversed again causing a second gradient echo to form. This process is repeated to form a train of,

say, 64 gradient echoes which are all separately digitized.

Before the first of these gradient echoes is acquired, a large negative phase-encoding gradient is

applied. A small positive phase-encoding gradient (blip) is switched on in between subsequent

gradient echoes.

At the time of each gradient echo, the nuclear spins have experienced the cumulative sum (i.e. time

integral) of the preceding phase encoding gradient pulses. Therefore the first gradient echo has

experienced a large negative phase-encoding gradient, the second effectively a smaller negative

phase-encoding gradient (the initial large negative excursion + the first positive blip), the third a

smaller again negative phase-encoding gradient (the initial large negative excursion + the first

positive blip + the second positive blip) and so on.

By the time of the final gradient-echo acquisition, the effective cumulative phase-encoding gradient

has stepped from the initial large negative value to a large positive value.

From the point of view of phase-encoding, the data set is equivalent to that which would beobtained if each gradient-echo were acquired in completely separate acquisitions, with the phase

encoding stepped between each one, as happens in a standard spin- or gradient-echo imaging

sequence.

Thus complete 2-dimensional image information is built up in a single shot following a single 90o

RF excitation pulse. This may be achieved in times of the order of 100 ms.


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An echo-planar imaging pulse sequence

The EPI sequence as shown is based entirely on gradient echoes, and therefore produces images

with T2*-weighted contrast. Alternatively, the addition of a single 1800

RF refocusing pulse before

the read gradient is switched on for the first time gives an image with T2-weighted contrast.

The EPI sequence may be preceded by a contrast preparation module (e.g. an 1800

inversion pulse

to generate T1-weighting, or diffusion-weighting gradients (- see later)) to generate specific tissue

contrasts.

900

R.F

Slice

gradient

read

gradient

Phase-

encoding

gradient

signal

Gradient

echoes

Equal

areas

Equal

areas Keep repeating read

gradient reversal until

all phase encode steps

acquired, e.g. 64 times

Equivalent to in other sequences


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Advantages of EPI:

- can "freeze" subject motion

- high temporal resolution : useful for fMRI or perfusion imaging (see later)- high signal-to-noise ratio per unit time

Disadvantages of EPI:

- rapid gradient switching requires high-performance system hardware

(however, modern all MRI systems are generally EPI-capable)

- rapid gradient switching generates large amount of acoustic noise

(although not necessarily worse than other high-speed MRI methods)

- gradient switching rate limits resolution (typically 64 x 64 pixels or 96x96)

- magnetic susceptibility variations can cause image distortion and signal loss

Fast Imaging III: Fast Spin-Echo Imaging (FSE)

(Also known as Turbo Spin-Echo imaging (TSE)). The principle here is similar to EPI: a series

of separately phase-encoded echoes is generated following a single 900

excitation pulse, but this

time a train of 1800

refocusing pulses are added to form a series of spin echoes rather than the

gradient echoes used in EPI (see diagram on next page).


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FSE sequence: schematic diagram for 3 1800

refocusing pulses

90o

180o

refocusing

RF ulse

90o RF

pulse

Signal Mxy

MxMx

180o

Mx

180o

Mx

180o

180o

refocusing

RF ulse

Spin Echo

180o

refocusing

RF ulse

Spin Echo Spin Echo

Behaviour of Mxy

Phase encoding gradient Equal areas, opposite polarity

[positive pulses cancel out phase

changes caused by original negativephase encoding gradient, so that thephase encoding is reset to zero

before each 1800pulse].

Phase encoding gradient incrementing for each

acquired echo as in previous sequences

Equal areas, opposite polarity


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In a FSE sequence, a number (typically 8) of 1800

RF refocusing pulses are applied after each

900

pulse, producing 8 spin-echoes per repetition period. Each spin echo is generated and digitized in

the presence of slice select and read gradients as described previously. Crucially each spin echo also

experiences a phase-encoding gradient pulse, the magnitude of this gradient being unique for each

echo. In this way, 8 phase encoding steps are collected for each repetition. After a recovery delay TR,

a second set of 8 spin-echoes is acquired in exactly the same way except that in this case the value ofeach of the phase-encoding gradients are different from those in the first set. This process is

repeated until all the phase encoding steps (typically 256) have been acquired. In the case of a

256x256 pixel image acquisition, for a standard spin-echo imaging sequence, the total imaging time is

256 x TR. For a FSE sequence however, the total imaging time will be (256/8) x TR = 32 x TR, i.e. the

imaging time is reduced by a factor of 8. For each of the 32 individual acquisitions, different phase

encoding gradient magnitudes are used, so that, at the end 256 spin echoes have been acquired,

each having experienced a unique phase encoding gradient.

Important difference between EPI and FSE imaging: In FSE imaging, each echo is followed by a phase-

encoding gradient pulse equal in area but of opposite polarity to that preceding it. This effectively

resets the phase encoding after each spin echo, so that the phase encoding gradient affecting aparticular echo is independent of the phase encoding for previous echoes; in EPI the phase encoding

affecting each echo is the cumulative sum of the preceding phase encoding gradient pulses.

T2-weighted FSE, sometimes with an additional initial 1800

inversion RF pulse for FLAIR suppression

of the CSF signal, are commonly used for clinical investigations in the brain.

Advantages of FSE imaging

High resolution (c.f. EPI) 256x256 or 512x512 pixels per image possible

Good T2 or PD contrast depending upon order of phase-encoding steps Good resistance to magnetic susceptibility problems

Disadvantages of FSE imaging

Not as fast as EPI FSE is a multi-shot sequence Lots of 180o RF pulses


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Summary: conventional MRI images with PD, T1,T2 or T2* based contrast

T1 Contrast PD Contrast

Progressive saturation Standard spin-echo sequence

(short TR/TE spin echo sequence) (long TR, short TE)

Inversion recovery Gradient echo sequence

(Inversion pulse preceding short (typically short TR, short TE, low )

TE/long TR spin-echo sequence)

Gradient echo sequence FSE sequence

(with short TR, short TE, high ) (long TR, short effective TE)

T2 Contrast T2* Contrast

Standard spin-echo sequence Gradient echo sequence

(long TR, long TE) (typically moderate TR and , long TE)

FSE sequence EPI without 180o refocusing pulse

(long TR, long effective TE) (long TE gradient echo)

EPI with 180o

refocusing pulse (long TE)

Beyond Conventional MRI Other Contrast Mechanisms

T1, T2, T2* and PD provide excellent soft-tissue contrast, and these quantities change with many

pathologies, and so provide image contrast which is diagnostically useful. The use of sequences toprovide this type of contrast has been referred to as conventional MRI

However, changes in these quantities may be non-specific, and tend reflect long-term tissue injury

rather than recently occurring (acute) changes.

Methods have therefore been developed which exploit other contrast mechanisms that more

directly reflect the physiological status of tissue, we will now explore some examples of theses.

1. Magnetization Transfer Contrast (MTC)

The spin-spin (T2) relaxation of water molecules is influenced by random magnetic fieldfluctuations, largely due to the tumbling motion of neighbouring (magnetic) molecules. Populations

of molecules which are free to tumble havelong T2 relaxation times, whereas water molecules which

are bound to e.g. membranes or macromolecules are less free to tumble and hence haveshort T2relaxation times.

In tissue (e.g. brain or muscle), water molecules may be considered to exist in one of two

microscopic environments: either bound to macromolecules or cell membranes such that their

rotational motion is restricted, or free to undergo rotational motion, e.g. water in the cytosol. The

T2 of bound molecules is very short (< 1 ms) so that they are invisible to conventional MRI.

MTC imaging is a way of observing the effects of these bound moleculesindirectly.


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30

Fourier theory predicts an inverse relationship between the relaxation time (decay) constant

(i.e. T2) in the time domain, and resonance line width in the frequency domain:

A series of RF pulses applied "off-resonance" (i.e. B1 rotating with angular frequency different from

that of the rotating proton transverse magnetizations Mxy) have no direct effect on the free water,

because their frequency lies outside the resonant frequency range for these spins. However, because

the RF frequency lies within the line width of the bound water, the magnetization of this water is

reduced to close to zero (saturation). Under certain circumstances, e.g. due to physical exchange,

chemical exchange or magnetic interactions (dipolar coupling) between spins in the bound-water and

free-water populations, (i.e. magnetization transfer) the magnetization of the free water may also

become reduced, and hence the MRI signal intensity is decreased:

Therefore if off-resonance (1-5 kHz) saturation pulses are added to conventional imaging sequences,

tissues which have a significant fraction of bound water molecules will show a reduced image

intensity. Thus MTC provides additional contrast between tissues containing complex microstructure

(larger bound water fraction, e.g. brain tissue), and those which dont (e.g. blood).

respo

nse

frequency

Fourier

Transformation

bound water molecules

- short T2

time

Signa

l

free watermolecules

- long T2

free watermolecules

- narrow resonance

bound water molecules

- broad resonance

CASE 1: No transfer of magnetizationbetween free and bound water

The off-resonancepulse has no effect on

the free water pool

Apply off-resonance

saturation pulses

The bound pool

ma netization is reduced

RF pulses

CASE 2: No transfer of magnetizationbetween free and bound water

Free water magnetization

reduced due to exchangeof magnetization between

free and bound poolsRF pulses


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31

Applications of MTC:

1. Selective suppression of "solid" tissue, e.g. for magnetic resonance angiography where signal

from brain tissue is suppressed in order to highlight that from free water in the blood (see

later).

2. Investigation of conditions in which tissue micro-structure is disturbed, e.g. in demyelinating

diseases such as multiple sclerosis, the solid/free water dynamics are altered such that areas

of diseased brain tissue show changes in MTR.

In this type of application it is common to produce maps of the magnetization transfer ratio

(MTR): A control image is acquired without off-resonance saturation pulses, and a second

image is acquiredwith the off-resonance saturation pulses, and for each pixel the ratio of the

value from each of the acquired images is calculated.


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32

Beyond Conventional MRI Contrast Mechanisms

2: Diffusion Weighted Imaging (DWI)

Water diffusion in tissue

Water molecules undergo random thermal motion or self diffusion, familiar as "Brownian Motion".

After a time t an individual water molecule following a random walk will have achieved a total

displacement from its start position which may be represented by a vector r.

Considering a large ensemble of water molecules, after a certain time t, adding all the displacements

together the mean vector displacement is zero molecules are equally as likely to have experienced

a positive as a negative displacement.

However, the mean squared displacement (for 3 dimensions) is given by:

= 6Dt Einstein relation

where t is the observation time and D is the diffusion coefficient.

For free water, D = 2.0 x 10-3

mm2.s

-1, therefore if t = 100 ms, the root mean square displacement

(()) 35 m. In tissue, cellular structures have dimensions < 100m, so that over thistimescale, the free motion of water molecules is restricted. This causes the measured diffusion

coefficient to be reduced compared with free water. Therefore in tissue instead of D we refer to the

Apparent Diffusion Coefficient (ADC < D). The degree to which diffusion is reduced in tissue is a

reflection of tissue microstructure, and hence if MRI can be made sensitive to diffusion effects,

diffusion may provide a useful source of contrast, changing with tissue type and pathology.

Generating diffusion-weighted contrast in MRI

A standard spin-echo sequence is modified with the addition of diffusion sensitising gradient pulses:

(cont. on next page)

r

r

Position at time 0

Position at time t

After a time t individual molecules in a particularvolume have undergone displacements (r) with

random directions and lengths

Therefore summing over all molecules in the

volume, the mean displacement, = 0


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33

final

21

1

1x x x

yyyy

x

Period 1Consider a water molecule at position x1 during

period 1. During this interval,

=.Gdiff.x1

At the end of period 1, the phase acquired by a

spin at position x1 during Period 1 is,

1 =.tPeriod 1 =.Gdiff.x1.

time

Gdiff

Period 1

Gdiff

180090

0

Spin echo

Period 2

x1 x

B

x2 x

B

Period 2Imagine that the water molecule now moves so

that it is at positionx2 during period 2. During

this interval, =.Gdiff.x2

the additional phase acquired by a spin at

position x2 during Period 2 is,

2 =.tPeriod 2 =.Gdiff.x2.

The total phase for a particular spin at the peak of the spin echo must be calculated taking into account

the effect of the 1800

refocusing pulse, which rotates the spins by 1800

about thexaxis:

After 900

pulse:

At the end

of period 1:

After 1800

pulse:

At the end

of period 2:

final = 2 -1

= .Gdiff. x2x1)

During periods 1 and 2, if the gradient Gdiff is applied in the x direction, in the rotating frame, B = Gdiff.xand

therefore the angular frequency of spins depends upon position: =.Gdiff.x

The presence of the diffusion gradients affects the amplitude of the spin-echo obtained. To see how this is so,

consider the following highly-simplified model of water molecule motion. Remember:

i) Magnetization in the x-y plane, Mxy, rotates about the z axis with angular frequency =B0i i) By the phase of Mxy at any instant, we mean the angle between Mxy and (say) thexaxis at that point in time

iii) This phase angle, =t, where t is the time from the start, and


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34

For stationary spins, =2, and as a result of the 180o

pulse, the phase advance acquired during the

first diffusion gradient pulse (Period 1) is cancelled out by that produced by the second diffusion

gradient pulse during Period 2, and the total phase gained (final) is zero (c.f. spin echo formation)

For moving spins, the phase gained during the second diffusion gradient pulse is not equal to that

acquired during the first, and a net phase difference,final, remains at the echo time.

In reality, the diffusion motion of spins is more complex, but the same basic principle holds. Since the

diffusive motion of spins is random, the phase advance,final, gained by individual spins is randomlydistributed (because e.g. the values ofx1 andx2 in each case are randomly distributed), and hence at

the echo-time the individual magnetization vectors point in different random directions in the x-y

plane, and their vector sum, and hence the final spin-echo signal magnitude, is reduced:

The more rapidly the spins diffuse, the greater the range of the individual x1s and x2s, the larger thephase dispersal between individual spins, and the smaller their final vector sum Mxy.

Quantitatively, for general simple self-diffusion, the spin-echo signal magnitude is given by:

S = So.exp(-b.D)

b =Gdiff2(

where So is the signal magnitude obtained without applying the diffusion gradients is themagnetogyric ratio, Gdiffis the diffusion gradient strength and and are respectively the durationand separation of the leading edges of the diffusion gradient pulses (see diagram at top of previous

page).

b is known as the "diffusion weighting" (units s.m-2

) or b factor

(typically diffusion weighting = 1000x106

s.m-2

, sometimes expressed as a b factor of 1000)

Increasing b (by increasing Gdiff, or) increases the diffusion-weighting of the image.

Incorporating the diffusion-weighting gradients into a spin-echo imaging sequence in this way

produces diffusion-weighted images in which water which is free to diffuse gives a low signal, and

water which is less mobile gives a high signal. Repeating the image sequence with increasing values

of b allows the actual ADC in each pixel to be estimated and the calculation of maps of ADC. For

example, a diffusion-weighted image is collected with diffusion weighting b = 1000x106

s.m-2

(b1000)

and a second which identical except the diffusion weighting is zero (b0). If, for a given pixel, the

Resultant

magnetization < |M0|

Slow diffusion: Rapid diffusion:

Resultant

magnetization


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35

image intensities for the two images are given by S1000 and S0 respectively, then the ADC for that pixel

is given by:

ADC = -1/(b1000-b0) . ln(S1000/S0)

= -1/1000x106

. ln(S1000/S0)

ADC maps are useful because they eliminate the T2 and PD weighting present in the source images,

and provide quantitative information directly comparable between patients.

Clinical Application of DWI: Cerebral ischaemic injury (stroke):

Cerebral ischaemia means loss of blood supply to the brain. This means the cells will receive

insufficient oxygen etc. and are unable to maintain osmotic homeostasis i.e. they swell up. This

causes a bigger fraction of the tissue water to be located inside the cells, where diffusion is lower,

and therefore the average ADC is reduced. ADC reduction occurs within minutes of ischaemia (c.f. T2changes which may take hours or days to occur following ischaemia). DWI is widely used in clinical

MRI to assess acute stroke and other ischaemic conditions.

Diffusion Anisotropy and Diffusion Tensor Imaging

The diffusion of free water is said to be anisotropic: there is no preferential direction for the

molecular motion: motion in all directions is equally probable.

However, the diffusion of water in tissue is impeded by cell membranes and other structures.

Therefore, taking as an example white matter nerve bundles in the brain, the ADC measured withdiffusion-gradient directions applied parallel to the nerve fibres main axis appears higher than that

measured with diffusion-gradient directions orthogonal to the fibre directions.

Therefore the measured ADC depends on the relative geometric orientation of the tissue and the

measurement gradient direction.

neuron

High

diffusion

(ADCparallel)

Low diffusion

(ADCorthogonal)


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36

If diffusion is anisotropic, it can no longer be adequately described by a simple scalar diffusion

coefficient (i.e. the ADC). In the most general case the diffusion behaviour is described by a tensor. A

tensor is a matrix of values. The diffusion tensor is 3 x 3 matrix, containing 9 values which

characterise the interactions between the diffusion gradient orientations and the 3-dimensional

directional dependence of the molecular diffusion. The diffusion tensor (DT) is written as:

The details are complicated and beyond the scope of these lectures, but if at least 6 diffusion-

weighted images are acquired, each with the diffusion-sensitizing gradients applied in a different

direction, together with one image with no diffusion weighting (b=0), then it is possible to calculateall the elements of the DT for each image pixel, and calculate a vector representing the principle

direction of the underlying diffusion. The direction of this vector for each pixel in the image indicates

the direction of the white matter fibres in this volume of tissue. By comparing the direction of this

vector in neighbouring pixels it is possible, e.g. by calculating streamlines, to reconstruct the path of

underlying white matted tracts: this is the principle of white matter tractography.

seed voxel

Alternatively, to obtain a measure of diffusion which is independent of direction, a simple scalar

quantity analogous to the ADC discussed above, then we can calculate the mean diffusivity (MD),

given by:

MD = (Dxx + Dyy + Dzz)/3

Diffusion Weighted Imaging Pulse Sequences

In practice DWI (and DTI) are performed by adding the diffusion-weighing gradients to a spin-echo

echo-planar imaging pulse sequence. A fast (single-shot) EPI imaging sequence is required i) to

avoid errors caused by bulk motion of the subject and ii) to allow the collection of a large number of

images with different diffusion-gradient directions, for DTI acquisitions, within an acceptable total

scan time.

DT =

Dxx Dxy Dxz

Dyx Dyy Dyz

Dzx Dzy Dzz

A streamline is detected where there the change in

direction of the principle diffusion vector is smallbetween neighbouring pixels. In the brain this

would indicate a probable white matter tract


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37

position position

Region of

magnetic

field

gradient

Beyond Conventional MRI Contrast Mechanisms

3. BOLD contrast and fMRI

Blood Oxygenation Level Dependant (BOLD) contrast

Deoxygenated haemoglobin is paramagnetic (i.e. has unpaired electrons) - and therefore has a

relatively high magnetic susceptibility.

Oxygenated haemoglobin is diamagnetic - with only a relatively low magnetic susceptibility (close to

that of brain tissue)

Therefore the magnetic susceptibility of blood depends upon the relative proportions of oxy- and

deoxy-haemoglobin it contains and therefore to the blood oxygenation level

The presence of paramagnetic deoxyhaemoglobin in blood vessels causes the magnetic field within

the vessels to be higher than in the surrounding tissue, resulting in local magnetic field gradients,and hence a reduction in T2

*in the tissue:

Thus T2*-weighted images, e.g. gradient-echo based EPI, can be used to detect changes in the local

blood oxygenation.

Functional Magnetic Resonance Imaging (fMRI)

fMRI is primarily applied to studies of the brain.

In the brain, specific "processing" functions are located in well defined anatomic regions, e.g. the

visual cortex, auditory cortex or motor cortex.

During "functional activation", as a result of increased metabolic demand, these regions demonstrate

focally increased blood flow.

B

tissue tissuecapillary

B

Low oxygenation

-> more deoxyhaemoglobin

-> blood has high magnetic susceptibility

-> high magnetic field gradients around vessels

-> more rapid spin dephasing => short T2*

high oxygenation

-> less deoxyhaemoglobin

-> blood has low magnetic susceptibility

-> smaller magnetic field gradients around vessels

-> less rapid spin dephasing => long T2*


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38

Time (seconds)

Sudden stimulus

e.g. sound,

flashing light etc,

The haemodynamic response of the brain produces an apparent "over-supply" of blood to the

activated region. Blood volume and blood flow increases, and although more oxygen is being

consumed, the net result is a decrease in the concentration of deoxyhaemoglobin. This causes the

magnetic susceptibility of the blood to decrease. Local microscopic field gradients are reduced and

T2*

in the tissue is lengthened. Therefore on T2*-weighted images, activated regions demonstrateincreased intensity, or "light up":

Typically a gradient-echo EPI sequence is used in fMRI so that images of the whole brain may be

obtained every few seconds.

Changes in the local value of T2* due to neuronal activation can therefore be rapidly followed,

allowing the location of specific functional areas to be determined to improve understanding of how

the brain performs certain tasks.

T2*-weighted

Signal

intensity

Deoxy-

haemog

lobin

Blood

flowMetabolic

demand

At rest: On activation:

Signal

intensityBlood

flow

Metabolicdemand T2*

0 2 3 6 8 10 12

Blood

suscepti

bilityT2*

Bold Response to a stimulus


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39

Regional activation detected with BOLD contrast can be demonstrated e.g. in the visual cortex (in

response to patterns and lights), in the motor cortex (in response to finger movement) and in the

noun and verb mediation speech centres of the brain responsible for forming responses to word

presentation tasks.

fMRI is becoming clinically useful for planning neurosurgical procedures where it is important toknow where certain critical functional regions (e.g. speech centres) are located so that they may be

avoided during the surgical approach.

Practical Issues

a) A high temporal resolution is required and therefore EPI is usually used, so the spatial resolution is

limited (for this reason, for display purposes, fMRI data are frequently overlaid upon high resolution

anatomical reference images)

b) The absolute signal change is dependent upon experimental parameters, so is not a direct

measure of blood flow, c.f. PET.

c) The signal changes upon activation are small (< 5%) so averaging between experiments in the same

subject, and even across subjects, is often performed

d) Complex image processing and statistical analysis is required.

-----------------------------------------------------------------------------------------------------------------

Contrast Agents in MRI

Contrast agents provide additional image contrast in order to improve diagnosis yielding:

- increased sensitivity

- increased specificity

- functional information

MRI contrast agents themselves are not seen directly in MRI images, but their presence causes

changes in the relaxation times of the surrounding tissue water.

Paramagnetic contrast agents are commonly used; they contain a metal ion having unpaired

electrons. The most efficient elements are:

Element Symbol Number of unpaired electrons

Gadolidium Gd3+

7

Manganese Mn2+

5

Dysprosium Dy3+

5

Iron Fe2+

4


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These ions are toxic, and are therefore bound in stable, biochemically-inert complexes (known as

chelates) e.g. DTPA (diethylene-triamine-penta-acetic acid) which form low molecular weight water-

soluble agents which are in general safely excreted by the kidneys in a few hours.

Gadolidium compounds are commonly used (high paramagnetic moment) e.g. Gd-DTPA (e.g.

"Magnevist"):

Such contrast agents may affect the MRI signal in 2 ways:

1) Positive Contrast Effect

As the paramagnetic molecules tumble due to thermal motion, they produce local random rotating

magnetic fields in the vicinity of the water protons. If these time-varying magnetic fields have

components at the correct frequency (e.g. the Larmor frequency) they interact with tissue water

spins causing a shortening of the T1 and T2 relaxation times. Under normal imaging conditions, the

dominant effect is T1 shortening and regions taking up the agent appear bright (positive contrast) on

T1-weighted imaging sequences such as progressive saturation.

After injection these agents distribute into the intravascular and extracellular space of the body, but

the relatively large molecules cannot cross the intact blood-brain barrier (BBB). Pathologicalbreakdown or absence of the BBB allows contrast agents to cross into the extracellular space of the

brain and alter T1 values locally.

Pathologies in which this occurs include tumours, infarctions, infection and acute demyelination. In

cancer sometimes living tumour tissue can be distinguished from the necrotic core and surrounding

oedema in tissue outside the tumour.

Dynamic studies can be used to assess organ function, such as filtration rate in the kidneys, liver

function or membrane permeability in the brain.


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2) Negative Contrast Effect

In healthy brain these types of contrast agents are restricted to the blood vessels and cant pass

directly into the brain tissue. They make the blood more paramagnetic and hence microscopic field

gradients are produced around the blood vessels (capillaries) causing a reduction in T2*. Therefore

signal intensity in T2*

-weighted images is reduced as the contrast agent passes through the tissue(this is very similar to T2* shortening in BOLD contrast).

This effect can be used to assess local blood flow through the brain. If a bolus of agent is injected into

a vein it will eventually pass through the brain via the heart. If rapid T 2*-weighted images are

repeatedly acquired (using an EPI sequence to produce an image every second or so) we can track

the passage of the contrast agent bolus through the brain as a function of time:

Such signal intensity curves provide information about tissue perfusion.

Cerebral perfusion is a general term used to describe the delivery of oxygen and other metabolic

substrates to the brain tissue via the blood. It is measured in terms of the Cerebral Blood Flow (CBF).

CBF quantifies the rate of blood supply via the capillaries to a unit mass of brain tissue. It is non-

directional and is approximately 100ml/100g/min in healthy brain. If CBF decreases to less than

approx 20 ml/100g/min for any substantial period, cell death will occur. This might occur acutely

following the rupture or blockage of a blood vessel (i.e. a stroke) or chronically after long-termdisruption of the normal blood supply to the brain. Assessment of perfusion using MRI contrast

agents in this way is therefore an important clinical tool.

By analyzing the shape of the curve of T2*-weighted signal intensity against time as the bolus of

contrast agent passes through the brain (e.g. the duration, width and depth of the signal dip) it is

possible estimate CBF directly.

From shape of curve (time-to-

minimum, overall time-integral

etc) perfusion can be quantified

Time (seconds)0 5 10 15 20 25

T2*-weighted

image

intensit

Bolus

Injection


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42

The Safety of Gadolinium Contrast Agents

Until recently gadolinium MRI contrast agents were considered relatively safe. However, in 2006

several studies suggested an association between the disease nephrogenic systemic fibrosis (NSF)

and MRI contrast agent administration. NSF is an uncommon but serious acquired systemic disorder

which affects patients with impaired kidney function. NSF causes swelling and tightening of the skin,

usually limited to the extremities, but 5% or less of patients have exceedingly rapid disease course

that may result in death.

It is hypothesised that NSF may be caused by the gadolinium ion becoming dechelated, i.e. separated

from the chelate complex, allowing the toxic ion to circulate unshielded in the blood stream.

Therefore before gadolinium contrast agents are administered, a blood test is usually performed to

ensure that the patients kidney function is adequate to allow the contrast agent to be excreted

sufficiently rapidly, before dechelation can occur.

We also now prefer to use contrast agents with a cyclic molecular structure in preference to those

with a linear structure (such as Magnevist) as these form chelates which are considered to be more

stable.

Other Contrast Agents

Other types of contrast agent are in use or under development, many of them organ-specific

targeting e.g. the liver, specific tumours and the heart. Some, such as magnetite, a

Time

S.I.

Low perfusion e.g. cerebral white matter

High perfusion (e.g. cerebral greymatter)

Gadolinium Contrast Agents Summary

positive contrast: T1-weighted progressive saturation imaging

Shortened T1signal increases

negative contrast: T2*-weighted (gradient-echo) imaging

Shortened T2*signal decreases


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43

superparamagnetic particle which can be coated with an inert resin, may be taken orally (as well as

intravenously), enhancing image contrast in the gut.

Exciting recent developments include the development of smart MRI contrast agents, which

produce paramagnetic contrast only if certain specific important biochemical/molecular species are

expressed in tissue, and iron-oxide agents which may be used to magnetically label macrophages(and also, potentially, stem cells), so that their distribution within tissue may be tracked.


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- 4 4 -

Magnetic Resonance Angiography

Angiography is the direct visualization of flowing blood in arteries and veins.

We will consider 2 methods of magnetic resonance angiography (MRA):

(i) Time-of-flight (TOF) MRA

(ii) Phase-contrast (PC) MRA

Both generate contrast caused by the movement (flow) of blood, and both are "bright-blood"

methods where blood vessels appear bright compared with background tissue.

Time-of-Flight Angiography

A gradient echo imaging sequence is used with short repetition time (TR


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- 4 5 -

compared with static tissue. Thus arteries containing rapidly flowing blood appear bright, and the

static tissue appears dark.

Phase-Contrast Angiography

Phase-contrast (PC) MRA uses an imaging pulse sequence with additional motion sensitizing

gradients in a manner similar to diffusion-weighted imaging. The difference is that here we are

interested in uniform coherent fluid flow, where all molecules in a pixel move with the same velocity

v, whereas for DWI, the molecules within a pixel move with velocities having random directions and

magnitudes.

PC MRA relies on the creation of a uniform phase shift for all spins within a pixel as the blood moves

in a gradient applied parallel to the direction of flow.

between eacho pulse a fresh volume of blood flows into the imaging slice

Imaging

z

v = 0

v = z/(2TR)

v z/(2TR)

image

Some, or all, of thepartially saturatedspins in the blood

vessel are replaced,increasing the vessel

signal

Partially saturated spins unsaturated spins


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- 4 6 -

After a RF excitation pulse, spins precess (rotate) in the rotating frame about the z axis, and gain a

phase in a time t:

The degree of velocity encoding is defined in terms of the velocity encoding factor, or venc:

Frequency, phase, dt Gx..x(t) = dt

= Gx..(x0 + v.t)

For static spins: x = x0 For flowing spins: x = xo + v.t

After the application of velocity sensitizing gradients:

Phase, T

dt2

0

=T

T

xx

T

xx tvGtxGtvGtxG

2

2

0

0

2

0 ...2

1......

2

1...

= - Gx . v.T2

Gx

-Gx

Equal magnitude and durations

STATIC SPINS

FLOWING SPINS

Ph

ase

,

Ph

ase

,

time

time

time

0 T 2T

(2T) = 0

(2T) = -.Gx.v.T2

is a phase increment due toblood velocity

Velocity sensitizing

gradients


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Venc =/GT2 (units are m.s-1

)

The venc represents the maximum velocity which can be encoded unambiguously, i.e. the

velocity which produces a of 180o.

PC MRA is usually performed using a gradient echo sequence. Scanner imperfections (e.g. B0 non-

uniformity) can cause phase differences between neighbouring pixels which have nothing to do with

blood flow. In order to cancel out these unwanted phase differences, in the simplest form of PC

MRA, the image is acquired twice: once with velocity encoding gradients applied in a positive

direction, and once with the same velocity encoding gradients applied in a negative direction. It is

possible to reconstruct each MRI image such that a value forthe phase ofthe MR signal in each pixelis obtained. For each pixel, the phase angle obtained with the positive velocity encoding gradients

applied is subtracted from the phase angle obtained for the same image pixel acquired with negative

flow-sensitization gradients. For static tissue, the phase is the same in each case and subtraction

results in a phase value of zero. However, for pixels containing flowing blood there is a phase

difference between each image (proportional to the flow velocity) and the final subtraction image, in

which the pixel intensity is proportional to the phase difference, shows only flowing blood.

In order to produce images without distracting signal from outside the tissue (i.e. from air) where the

signal phase is random, the phase subtraction image is usually multiplied, pixel-by-pixel, with the

conventional magnitude image.

Occasionally, instead of calculating phase differences for each pixel, PC MRA data is processed by

treating the signal from each pixel as a complex number (i.e. the x axis corresponds to the real axis

and y to the imaginary axis), and performing a complex subtraction between the two images for each

pixel. This can provide better flow-contrast in certain circumstances.

PC MRA has the advantages that it may be made sensitive to vessels containing slowly flowing blood

(the signal from which is likely to be completely suppressed in the TOF technique), and by judicious

choice of the velocity encoding gradients, flow velocity and direction may be estimated. A

disadvantage of PC MRA compared with TOF MRA is that at least 2 acquisitions must be performed

(one with and one without velocity encoding) and so imaging times are longer

MRA image processing/presentation

For both methods, 3-dimensional data sets are acquired (either by using multi-slicing or true 3D

sequences) and algorithms such as the Maximum Intensity Projection (MIP) method used to highlight

image 1:(+ve velocity

encoding)

Resultantvector for

flowing spins

image 2:(negative velocity

encoding)

-

static spins

Flowing spins

static spins

Flowing spins

image 1 image 2:(complex subtraction)

=

Resultant forStatic spins is zero


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the 3D vascular structure by selecting only the brightest pixels for display and thereby further

suppressing the static background. In this way interactive display of the 3D distribution of blood

vessels viewed from any direction is possible.

MR angiography can be improved with the use of:

a) contrast agent adminsitatration to shorten blood T1 and hence increase its signal relative to

tissue (as in contrast enhanced magnetic resonance angiography or CEMRA)

b) an additional MTC off-resonance pre-saturation pulse applied to decrease the intensity of

background signal relative to that of blood.

These valuesbecome pixelvalues for first

row of MIPdirectionof blood

flow

stationarytissue

MRangiogram

Maximum Intensit ro ection for MRA

maximumintensity

projection

3D stackof ima es

Take series of rays across image:get maximum value along each ray

one slicethrough

3D ima e-

blood


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Safety in MRI

MRI does not use ionizing radiation and is therefore generally considered a very safe imaging

modality.

However various factors involved in MRI represent hazards to both patients and staff. These include:

(i) The static magnetic field, B0.

(ii) Time varying magnetic fields i.e. the switched imaging gradients

(iii) The radio frequency (B1) fields.

(iv) The cryogens (liquid nitrogen and liquid helium) required to maintain the

superconductivity of the magnet

(v) Intravenous contrast agents (see discussion in earlier lecture)

Safety guidelines for the management of these risks are laid down (in the UK) by the Department of

Health (via the MHRA) (DB 2007(03) Safety Guidelines for Magnetic Resonance Imaging Equipment

in Clinical Use

http://www.mhra.gov.uk/Publications/Safetyguidance/DeviceBulletins/CON2033018 )

Hazards due to the MRI scanner environment

These are principally hazards created by the static B0 magnetic field which extends beyond the

confines of the scanner over an ellipsoidal region centred on the centre of the magnet. The stray

fields of whole-body MRI systems may extend over a number of metres.

To reduce the stray fields, 2 types of shielding are used:

1. Passive Shielding iron plates are attached either to the outside of the cryostat, or within

the walls, floor and ceiling of the scanner room

2. Active Shielding The superconducting coil winding is continued in the opposite direction

outside the inner main magnet winding. This self-shielding partially cancels the field

outside the main magnet coils thereby reducing the magnetic field exterior to the scanner.

The safety guidelines state that the general public should not be exposed to fields of more than 5

Gauss (0.5 mT) since cardiac pacemakers and other active medical implants may be affected abovethis field strength.

A Controlled Area is defined to enclose the 5 Gauss field contour lines. Signs and physical barriers

must be used to restrict access to areas inside the 5 Gauss contour line.

Projectiles: The most imminent danger to patients and personnel is from ferromagnetic objects such

as pens, scissors etc. which may be attracted to the magnet with great force and act as projectiles.

Ferrous objects experience a displacement force if the field is varying in space, and also a rotational

torque even if the magnetic field is uniform.

All staff and patients are required to empty their pockets and remove jewellery etc. beforeapproaching the scanner. Access to the scanner room is restricted by locked doors to prevent the


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unauthorized introduction of ferromagnetic material. Metal detectors may be used at the entrance of

scanner suites to reduce this hazard.

Items introduced into the Controlled Area should be correctly labelled as

1. MR Safean item which poses no known hazards in all MR environments

2. MR Conditional

An item which has been demonstrated to pose no known hazards in a specified MR

environment with specified conditions of use. Field conditions that define the specified MR

environment include field strength, spatial gradient, dB/dt (time rate of change of the

magnetic field), radio frequency (RF) fields, and specific absorption rate (SAR). Additional

conditions, including specific configurations of the item, may be required.

3. MR Unsafe

an item which is known to pose hazards in all MR environments

Implants: surgical implants e.g. surgical clips, pins, plates, prostheses, neurostimulators, implantedinfusion pumps or pacemakers pose a danger: implants can shift position due to magnetic forces with

a risk of haemorrhage or other injury. They may also get dangerously hot during scanning, and active

electronic implants may malfunction.

Established implants such as pins and plates attached to bones should not move, and teeth fillings

and false teeth are usually not affected.

Patient notes are examined and patients asked to complete a safety questionnaire before scan. A

particular danger may be posed from any particles of metal dust in eyes: often a plane x-ray may be

required to exclude this possibility.

Pacemakers: These may be affected by fields of 17 Gauss or more, and for safety a 5 Gauss threshold

is specified for public exposure (e.g. in corridors surrounding the scanner). Patients with pacemakers

should not be scanned (although MRI Conditional pacemakers are now coming on to the market).

Patient notes are examined and patients asked to complete a safety questionnaire before scan.

Cooling gases (cryogens): In superconducting magnets a "quench" (a very rare event) involves the

magnet windings loosing their superconductivity and becoming resistive, when due to the large

current they carry they heat up rapidly causing all the liquid helium and nitrogen in the magnet to

evaporate very rapidly: large volumes of gas are produced which may completely fill the scanner

suite.

Because of the risks of asphyxia, provision is made to vent these gases to the outside via suitable

piping, and oxygen sensors trigger emergency extraction systems if the fraction of air in the scanner

room drops below safe levels.

Bio-effects: Hazards due to electromagnetic field interactions during scanning

Exposure to MRI is considered safe for patients and staff`. There is no evidence of the initiation of

cancer or other harmful effects.


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Static magnetic fields (B0): There is no evidence of risk due to short or long term exposure. Ethical

approval is required to scan human subjects at magnetic fields greater than 4Tesla. Research MRI

scanners use fields up to 8 Tesla for human studies.

Time varying magnetic fields (Gx, Gy, Gz imaging gradients): As the magnetic field gradients are

switched on and off during scanning, the resulting fluctuating magnetic fields (dB/dt) can induceelectrical currents in tissue which may exceed the nerve depolarization threshold and cause

peripheral nerve stimulation (PNS). While this unlikely to cause permanent injury, it may cause

discomfort to patients or cause them to move ruining the scan. To prevent PNS, safety guidelines

specify that dB/dt should not exceed 20T/s for times greater than 120 microseconds.

Software checks and hardware interlocks prevent the scanner from accidentally exceeding this dB/dt

limit.

(Between 2 and 5 T/s, magnetic stimulation of the optic nerve or retina can occur, producing a

harmless flashing sensation in the eyes (magnetic phosphenes)).

Gradient switching acoustic noise: A second hazard resulting from the switched imaging gradients

concerns acoustic noise. This arises due to movement of the gradient coil windings against their

mountings caused by the Lorenz force arising as the current through them is rapidly changed in the

presence of the strong static magnetic field. This acoustic noise may be stressful for patients or even

hazardous. Noise levels increase with field strength and gradient switching speed.

Hearing protection (ear plugs or ear defenders) is commonly provided.

Radio frequency (B1) fields: Rapidly oscillating RF electro-magnetic fields can induce currents in

electrically conductive tissue of sufficient magnitude to cause significant heating. The higher thefrequency of the RF used (i.e. the higher the magnetic field strength), the larger the amount of heat

deposited in the tissue by a given RF pulse. Tissues with a high concentration of ions will absorb the

greatest RF power. Parts of the body with a low blood supply are at greatest risk (e.g. the eyes). The

specific absorption rate (SAR), measured in watts per kilogram, is used to quantify the rate of tissue

heating. For a particular imaging sequence the SAR depends upon the field strength, the number of

RF pulses per unit time, their magnitude, the RF coil design and the mass of tissue exposed. Safety

guidelines state that in normal uncontrolled operation the whole-body SAR should not exceed 2

W/kg for exposure times up to 15 minutes. The guidelines are designed to limit the temperature rise

of the tissue to less than 10C.

MRI systems have software checks and hardware interlock systems to prevent the SAR guidelines

from being exceeded. In order to accurately estimate the mass of tissue exposed, it is necessary to

know the weight of the patient, and this is therefore entered at the scanner console prior to scanning.

Lecture Notes MRI(Part2)

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