2.
MRI HARDWARE
The main
components of an MRI system are the superconducting magnet, the gradient
system, the RF system and the computer system. The magnet produces a strong,
static field and the radiofrequency transmit and receive coils excite and
detect the MR signal. The magnetic field gradients localise the MR signal and
the computer system facilitates scanner control, image display and archiving.
This chapter will describe each of these components in turn. Figure 2-1 below
shows the general layout of an MRI system.
Figure
2- 1:
Typical Architecture of an MRI System
2.1. Magnet Types
The magnet is the main component of
any MR system and there are four different types of magnets capable of MRI:
air-cored resistive magnets, iron-cored electromagnets, permanent magnets and
superconducting magnets. Superconducting magnets are by far the most common
type used in the NHS and the three main manufacturers are Siemens, GE and
Philips. I had the opportunity to use two magnets at Ninewells, one of which
was a Siemens Symphony 1.5T and the other a Siemens Impact 1.0T magnet. Table
2-1 below outlines some characteristics of a typical superconducting magnet.
The static magnetic field B0 is non-uniform and its homogeneity is
increased for clinical use through shimming. The homogeneity in parts per
million for a typical 1.5T magnet is approximately 5 ppm 1.
2.1.1. Superconducting Magnets
Magnet Type
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Characteristics
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Manufacturers
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Example
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Superconducting
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Flux lines horizontal.
Low power requirements.
Expensive to purchase.
Large fringe fields.
High field strength.
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Siemens SYMPHONY
Philips INTERA
GE Signa Excite
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Table
2- 1:
Characteristics of a Typical Superconducting Magnet
The way a superconducting magnet works
is as follows. As resistance decreases, current dissipation also decreases.
Resistance depends on the material from which the loops of wire are made, the
length of wire in the loop, the cross-sectional area of the wire itself and the
temperature of the wires. Superconductors have effectively zero resistance
below a certain very low temperature called the critical temperature. The coils
are wound from niobium-titanium (NbTi) filaments embedded in a copper matrix.
The copper serves to protect the NbTi wires in the event of a quench 1. The NbTi filaments become
superconducting at around 7.7K. Figure 2-2 below shows the cross-section of a
typical superconducting magnet.
Figure 2- 2: Cross-section of a
Typical Superconducting Magnet
Initially, the current is passed
through supercooled wires to create the magnetic field. Then the wires are
supercooled with cryogens (liquid Helium). I observed this procedure being
performed on a GE Signa Excite 1.5T scanner at Perth Royal Infirmary. It
involved a gradual power input accompanied by gradual cryogenic cooling. Such
magnets produce high magnetic fields with low power requirements. With
resistance at zero, there is no current dissipation and no additional power
input is required to maintain the field 2.
2.1.2. Fringe Fields and Shielding
Walls, floors and ceilings cannot contain static magnetic
fields. The stray magnetic field outside the bore of the magnet is known as the
fringe field. All magnets have a fringe field to some extent and these fields
must be taken into account when installing a magnet.
Fringe fields can be compensated for
by the use of magnetic field shielding which may be active or passive. Passive
shielding is the more expensive alternative using iron plates to restrict the
field lines. Some manufacturers offer actively shielded magnets that reduce the
fringe field to about 30m2. Active shielding partially cancels the
field outside the main magnet coils thus reducing the magnitude of the fringe
field. The 0.5mT isomagnetic line is taken as the critical cut-off limit.
During my MRI training I performed measurements to determine the extent of the
0.5mT line around a Siemens
Impact 1.0T Scanner. The results of this are presented in chapter four – MRI
Safety.
2.1.3. Quenching
A quench is basically a sudden loss of
the main magnetic field and may occur intentionally or unintentionally. Stored
electrical energy in the NbTi winding will be dissipated as heat if the
superconducting process fails. This heat will cause other parts of the windings
to rise above their maximum rated temperature and this will propagate the
effect throughout the magnet. This will result in a sudden loss of the B0
field and the loss of the cryogen. A safety feature of any MRI system is the
presence of quench-pipes that are vented outside the building and this prevents
cold burns and suffocation in the event of a quench 1.
2.1.4. Shim Coils
Most modern MRI techniques (e.g. EPI,
chemical shift imaging) require magnetic fields homogeneous to less than 3.5
parts per million (ppm) over the imaging volume. The raw field produced by a
superconducting magnet is approximately 1000 ppm or worse, thus the magnetic
field has to be corrected or shimmed. Usually this is accomplished by a
combination of current loops (active or dynamic shims) and ferromagnetic
material (passive or fixed shims). Gradient coils are used to provide a
first-order shim. The patient distorts the magnetic field when put into the
scanner and so an active shim correction must be made before scanning. The
operator will perform active shimming to improve the homogeneity on an
individual patient basis 3.
2.2. Gradients and Gradient Coils
In order to localize (spatially) the
MR signal we three orthogonal linear magnetic field gradients in the x, y and z
directions. There are three gradient coils situated within the bore of the
magnet and these are named according to the axis along which they act when
switched on.
The z-gradient alters the magnetic field strength along
the z-axis, (long axis of magnet bore). The y-gradient alters the field along
the y-axis (vertical axis of the magnet bore). The x-gradient alters the field
along the x-axis (horizontal axis of the magnet bore). Figure 2-3 below shows
the relative positions of the x, y and z-axes.
Figure 2- 3:
Magnet and Gradient Orientations
Gradients are alterations to the main
magnetic field and are generated by coils of wire located within the bore of
the magnet through which current is passed. Passage of current through a
gradient coil induces a gradient (magnetic) field around it that either
subtracts from or adds to the main static magnetic field strength, B0.
Gradients have many purposes including
slice selection, spatial encoding, flow compensation, spoiling, rewinding and
pre-saturation.
The magnitude of B0 is altered in a linear fashion
by the gradient coils, so that the magnetic field strength and therefore the
precessional frequency experienced by nuclei situated along the axis of the
gradient can be predicted. This is called spatial encoding. Nuclei that
experience an increased magnetic field strength due to the gradient speed up
(precessional frequency increases), whereas nuclei that experience a lower
magnetic field strength slow down (precessional frequency decreases).
Therefore, the position of a nucleus along a gradient can be identified
according to its precessional frequency.
The “magnetic isocentre” is the centre
point of the axis of all three gradients and the bore of the magnet. The field
strength here remains unaltered even when gradients are applied as shown in
figure 2-4 below.
Figure 2- 4: Magnetic Isocentre
2.2.1. Z-Axis Gradient (Maxwell Pair)
The z-gradient (denoted Gz)
is produced by a single pair of coils with currents flowing through them in
opposite directions. This is called a Maxwell Pair. The optimum gradient
linearity occurs when the coil separation is approximately R√3 where R is the
coil radius. Figure 2-5 below left shows the configuration of a Maxwell Pair of
gradient coils.
Figure 2- 5: (L) Maxwell Pair. (R)
Golay Set
2.2.2. X- and Y-Axis Gradients (Golay Set)
The y-gradient (denoted GY)
is generated using a Golay Set of gradient coils. A Golay Set is shown in
figure 2-5 above right. It consists of four coils on the surface of a
cylindrical drum with the currents producing a quadrupolar magnetic field (two
north and two south poles). The x-gradient axis (denoted GX), is
generated using an identical set of Golay coils orthogonal to the first set.
2.2.3. Eddy Currents and Pre-emphasis
Rapid switching of the gradients
induces eddy currents in conducting components in close proximity to the
magnet. The field generated by eddy currents combines with the true gradient
field and results in waveform distortions. These distortions in turn lead to
image distortion artefacts and signal loss. The effects of eddy currents can be
reduced by pre-emphasis of the gradient waveforms, so that after interference
from the eddy current fields, the resulting waveform has approximately the
ideal shape.
2.2.4. Slew Rate
The strength
of the gradient is expressed in mTm-1 with typical clinical values
of about 30-50 mTm-1. Figure 2-6 below illustrates the parameters
used to calculate the slew rate.
Figure 2- 6: Gradient Slew Rate
The ratio of the maximum gradient, (Gmax) to the RISE TIME, (tR) is called the SLEW RATE,
(SR).
SR = Gmax / tR in mT/m. msec.
Some modern MRI scanners (Siemens
SYMPHONY, GE SIGNA, Philips INTERA) have Gmax as high as 50 mT/m and a tR of 180 µsec giving a
typical SR as high as 120 mT/m. msec, (compared to an SR ~ 15 in the 1980’s).
2.3. The Radio Frequency System
A radiofrequency (RF) system in its
simplest form consists of a transmitter, coil and receiver 1. The transmitter produces suitably shaped pulses of
current at the Larmor frequency and when this is applied to the coil, results
in an alternating magnetic field. The coil also serves to detect the signal
from the patient. Frequency encoding results in a narrow range of frequencies
(± 16 kHz), centered on the Larmor frequency. The receiver demodulates this
range of frequencies from the higher carrier signal. Figure 2-7 below
illustrates a typical RF system.
Figure
2- 7:
The Radio Frequency System
2.3.1. Transmitter
The transmitter
produces RF pulses with suitable center frequencies, amplitudes and phases in
order to excite nuclei within the selected slice. The center frequency is
determined by the slice position and slice select gradient, the bandwidth
controls the thickness of the excited slice and the amplitude of the RF pulse
controls how much magnetization is flipped by the pulse.
2.3.2. Transmit Coils
The coils that
excite the MR signal must produce a uniform field B1 at right angles
to the B0 field. In order to maximise their uniformity, transmit
coils are usually large. The main
transmitting coil is usually the body coil surrounding the entire patient. It
is built into the scanner bore and is not visible. Body or volume coils are
usually of the saddle or birdcage design as shown in figure 2-8 below.
Figure 2- 8: (L) Saddle Coil. (R) Birdcage
Coil
To obtain a homogenous B1 field, the
current around the surface of the cylinder should vary sinusoidally. For a
saddle coil this is achieved by six wires arranged at 60º intervals. The
birdcage coil improves homogeneity over the saddle coil by increasing the
number of conductors around the circumference.
2.3.3. Receiver Coils
The purpose of a receiver coil is to maximise signal
detection and minimise noise. To minimise noise and maximise the SNR, the coil
dimensions must be of a comparable size to the patient’s anatomy. The two
classes of receiver coils are volume and surface types. Volume coils completely
surround the anatomy of interest and are often dual-function transmit/receive
coils. Surface coils are receive-only because of their inhomogeneous reception
field. They are useful for detecting signal near the surface of the patient.
Receiver coils can also operate in quadrature, resulting in a √2 improvement in
SNR.
2.3.4. RF Coil Types (Surface and Phased Array)
Surface
coils are placed over the anatomical ROI for example when imaging the
spine. The signal-response of a surface
coil in non-linear with increasing depth and results in an intensity fall-off
in the patient. Some surface coils are flexible and are useful when they can be
wrapped around the ROI. It must be ensured that the surface coil is
perpendicular to B0 otherwise no signal will be detected.
Phased
array coils consist of a number of small coils simultaneously and individually
receiving a MR signal. The mutual inductance must therefore be kept to a
minimum. Phased array coils produce the SNR of a small coil but have a much
larger FoV. If each coil is connected to a separate receiver then the
inter-coil noise is uncorrelated and this produces a higher SNR than if the
coils were connected to only one receiver.
Several
types of coils are used and are shown below in figure 2-9 including: Saddle
(volume) coils, Birdcage (volume) coils, Surface coils, Linearly Polarised
coils, Circularly Polarised coils and Phased Array coils.
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Figure
2- 9: (A)
Head Coil - Quadrature transmit and receive. (B) Peripheral Vascular Phased
Array Coil. (C) Phased Array Coil. (D) Breast Array Coil. (E) General Purpose
Coil. (F) Knee Coil.
2.3.5. Computer Systems
The complexity of modern MRI systems necessitates the
presence of several sub-systems under their own microprocessor control but
linked to the main host computer.
The
computer system generally consists of an operator console enabling operator
input, scan selection and sends imaging sequence to scan computer. The scan
computer contains a real time operating system; controls gradients and controls
RF transmit and receive. An array processor handles image reconstruction,
multi-planar reconstructions and maximum intensity projections. The flat panel
display allows the operator to supervise data storage, recalling, copying,
archiving etc…Images may be viewed slice-by-slice or as a cine loop. A two-way intercom
is controlled from the operator console and facilitates contact with the
patient. The remote drug delivery system (e.g. contrast agents) is also
controlled from here.
2.4. RF Shielding – Faraday Cage
Unwanted
external RF signals from television channels cause RF noise, so too do
flickering fluorescent lights, monitoring equipment). It occurs at the specific
frequency of the unwanted RF pulse and can be removed by improved RF shielding
using copper sheets. It is recommended to remove monitoring devices if possible
and ensure the door to the magnet room is closed. A Faraday age surrounds the
entire magnet room as shown in figure 2-10 below left. A penetration panel
shown in figure 2-10 below right is the entrance point for all electrical
cables, hydraulic, air and water into the scan room. The penetration panel
ensures that RF cannot enter the room from the outside and RF from inside
cannot escape.
Figure 2- 10: (L) Newly built Scan Room
at Perth Royal Infirmary showing RF shielding on walls and ceiling. (R)
Penetration panel in Perth Royal Infirmary MRI Scan Room.
2.5. References
[1]
D. W. McRobbie, E. A. Moore, M. J.
Graves, M. R. Prince. MRI – From Picture to Proton
(2003). Cambridge University Press. Pp 164-188.
[2]
J. P. Hornak. The
Basics of MRI (2004). http://www.cis.rit.edu/htbooks/mri/
[3]
R. H. Hashemi, W. G. Bradley. MRI
The Basics (1998). Lippincott Williams and Wilkins.
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