INTRODUCTION TO M.R.I. HARDWARE
Senior Radiographer MRI
Royal Adelaide Hospital
These notes outline the major components of a Magnetic Resonance Imaging
(MRI) machine. The intention is to familiarise the reader with the primary
functional components common to all MRI machines by discussing design variations,
key specifications, and the role of each component in the process of image
Image Creation & Functional Overview
Primary System Components
Console & Workstations
Patient Communication and Monitoring
Image Output Devices
THE PROCESS OF MR IMAGE CREATION
When a patient is placed in a strong magnetic field their tissue becomes
magnetised primarily due to the behaviour of certain magnetic nuclei. This
nuclear magnetic field has a component which rotates at a specific frequency
due to a phenomenon known as nuclear magnetic resonance (NMR). The resonant
frequency is proportional to the strength of the applied magnetic field,
and known as the Larmor Frequency
Virtually all clinical MR imaging relies of manipulating the nuclear
magnetic field of hydrogen which has a Larmor frequency of 63.87 Mhz. at
a field strength of 1.5 Tesla. The MRI image is a digital map of NMR signal
intensity. To arrive at a useful image, the MRI scanner must create detectable
NMR signals from tissues so that different tissues will give different
signal intensity (contrast), and spatially localise the NMR signals.
FUNCTIONAL OVERVIEW OF THE MRI SYSTEM
The MRI Magnet provides the external magnetic field(B0) which brings
the intrinsic magnetic fields of certain nuclei into specific alignments,
setting the conditions for NMR to occur. Spatial uniformity of the B0 field
is maintained by the shim system.
The RF Transmitter produces shaped pulses of alternating current
at appropriate frequencies which, when passing through the transmit
coils, create a magnetic field rotating at the Larmor frequency (B1).
This excitation field realigns and organises the intrinsic nuclear magnetic
fields of tissues so they coherently rotate in the X-Y plane (Mxy). The
rotating magnetic fields can induce an RF alternating current (MR signal)
in the receive coils (detection). The receiver systems demodulates
MR signal to extract information that is digitised to create the raw data.
The raw data is used by the array processor to calculate digital
Image contrast is manipulated by the method and timing of MR signal excitation
and detection through the use of pulse sequences set up by scanning interface
software in the host computer, and controlled by sequence control
software in the host and distributed computers.
Spatial localisation is achieved by the Gradient System which alters
the magnetic field strength in different locations for the period that
the gradient field is applied, creating frequency shift, phase shifts and
creating the precondition for frequency selective excitation.
The MR Radiographer handles the images and other equipment maintenance
functions at an operating console or work-station via image
manipulation software within a User Interface.
The Host Computer is used to manage image data display, storage,
transmission, and transfer to hard copy devices.
The magnet of an MRI scanner provides a stable, strong and spatially uniform
magnetic field within a structure that allows adequate patient access.
The unit of magnetic field strength is the Tesla (T). Another, less
frequently used, unit of magnetic field strength is the Gauss (G).
1 Gauss = 10-4 Tesla 1 Tesla = 10,000 Gauss.
MRI scanners use a magnetic field strength of 0.04 T to 2 T with most
in the 0.5T - 1.5T range. By comparison, the earth's magnetic field is
about 0.5 x 10-4 T, and a fridge magnet is about 0.2 Tesla.
Three types of magnets are available:-
Permanent magnets are constructed from magnetised ferromagnetic material
such as alloys of iron and cobalt (ALNICO), or more efficiently with rare
earth alloys such as Neodymium Iron Boron (Ne-Fe-B) or Samarium Cobalt
(Sm-Co) ALNICO permanent magnets are extremely heavy considering the achievable
field strengths. A 0.2 T magnet using ALNICO weighs 23 Tonnes, using Ne-Fe-B
the weight is reduced to 4 tonne unit, but the material is very expensive.
Virtually all permanent magnets use a vertical main field alignment with
the patient sliding in horizontally between two close poles. This geometry
require different RF coil and gradient coil designs to that used in horizontal
field magnets. The field of a permanent magnet is always "on", yet they
require no applied power and have a stable field as long as the room temperature
is well controlled. Field homogeneity is relatively poor compared to that
of superconductive magnets, and is generally achieved over a smaller volume.
Older commercial MR systems have offered permanent magnets between
0.064 T and 0.3T but they constitute a very small portion of the installed
base. More recently GE and Hitachi released systems based on a 0.2 Tesla
permanent magnet. Marketing of these scanners concentrates on their potential
in interventional, biopsy, and joint motion work as well as helping to
claustrophobia in some patients.
In a resistive electromagnet the field is generated by the flow of current
through resistive electrical windings. Resistive systems require the constant
application of large amounts of electrical power to maintain the field.
A 0.15T magnet uses 50 kilowatts, and a 1.5T magnet would theoretically
consume 5 Megawatts. Most of this power is dissipated as heat necessitating
elaborate cooling systems.
Air core resistive magnets can be designed with 2 or 4 sets of solenoidal
windings arranged co-linearly around the magnet bore to yield a horizontal
magnetic field. Iron core resistive designs generally use two solenoid
windings around the limbs of an iron rectangular yoke fitted with large
poles in the horizontal limbs and yielding a vertical field at the patient
bore. The yoke can augment the field of the electromagnetic windings by
up to 40%.
Resistive windings can also be used to augment the fields of permanent
magnets in a hybrid configuration that offers less mass than the permanent
approach and less power usage than the resistive approach for a given field
strength, although the heat produced by the resistive windings and the
temperature dependence of the permanent magnetic material can create design
Field homogeneity is relatively poor compared to that of superconductive
magnets, and is generally achieved over a smaller volume restricting high
quality imaging with gradient echo sequences and large fields of view.
General purpose scanners using resistive magnets have been offered
at between 0.1T and 0.3T but occupy less than 1% of the world-wide installed
base of about 5000 scanners. This figure has increased with the release
by Siemens in 1993 of the Magnetom OPEN, an interventional MRI unit based
on a 0.2 Tesla resistive magnet by Oxford Magnet technology. Picker based
its OUTLOOK system on the same magnet but runs at slightly higher field
strength. Other new resistive designs are aimed at the niche market of
the small parts scanners, where they offer suitable field strengths at
low cost for the user with a low patient throughput.
The most common MR magnet type is an air core electromagnet made with windings
made of superconductive material. The superconductive windings show virtually
zero electrical resistance when they are cooled to very low temperatures
and so, as long as the material is kept below its critical temperature,
the current flow producing the magnetic field continues without application
of electrical power.
MR magnets use multi-filament windings of Niobium-Titanium alloy (Nb-Ti)
embedded in a copper core. Nb-Ti becomes superconductive at about 10oK
(-263o C.), so the windings are bathed in liquid Helium (4oK)
within a structure called the cryostat. The cryostat is constructed of
low thermally conductive material and a series of refrigerated zones and
extreme vacuums to thermally shield the cryogenic zones from the relatively
high temperature outside environment. Liquid Helium (LHe), and in older
units liquid Nitrogen (LN2), is constantly boiling away and
the cryogen supply needs to be topped up regularly. This is expensive and
time consuming. High cryogen usage was a feature of older superconductive
magnets, but designs from the past 3 years use significantly less liquid
Helium (approx 0.1 litres per hour), and no liquid Nitrogen.
Once energised the magnetic field is always "on", but it can be turned
off (quenched) in about 30 minutes by discharging the current flow through
a resistive shunt made by heating a small section of superconductive material
in parallel with the main windings. In older designs section this process
consumes a large percentage of the cryogens and take a day to restore a
stable field (ramping), but new magnet designs effect a controlled quench
with little cryogen loss, and can be ramped up within a few hours.
General purpose scanners use superconductive magnets are available from
a variety of manufacturers, operating at fields of 0.5 T to 2 T. Over 90%
of the world's installed MRI systems.
The magnetic fields generated by high strength superconductive magnets
extend over a large area, presenting a number of siting difficulties. This
"fringe field" can disrupt the operation of computer storage media, colour
video monitors, CT scanners, and Gamma cameras. Dynamic intrusions of ferro-magnetic
material into the fringe field will distort the path of the magnetic field
lines (flux lines), induce current variations in the superconductive windings,
and create reciprocal field distortions within the magnet bore that degrade
homogeneity. These transient field distortions will oscillate in intensity,
taking as long as 20 minutes to decay effectively. Siting of high field
MRI scanners is made more practical if the magnetic field can be shielded
or contained to reduce the fringe field area.
Magnetic Field Shileding
This approach uses ferromagnetic material (usually iron) to create a structure
that will contain the magnetic flux lines more effectively than air. The
design can place iron close to the magnet, restricting fringe field intensity
within the scan room, or incorporate the shield in the walls of the room
to contain it at that point. The closer the shield is to the magnet, the
more critical is its design, material, and symmetry. In a derivation of
high energy linear accelerator magnet designs, the shield can also be design
to augment the field strength. Wether the shield is placed close to the
magnet or in the walls of the scan room, the mass of a given iron alloy
required to achieve a specific degree of fringe field containment is constant.
About 20 tonnes of iron will effectively shield magnets of 1 to 1.5 Tesla,
sufficient to require careful design of the magnet room floor. By containing
the flux lines in the iron the fringe field is greatly contracted, so passive
shielded magnets are less able to be affected by moving ferromagnetic objects
outside the scan room. Close passive shields, like the Siemens Magnashield
are waning in popularity due to their mass and construction expense, although
Toshiba uses the technique to achieve part of the shielding offered by
its Gemini Shield magnets. Room wall shielding is still popular, particularly
for screening high traffic areas, but active shielded magnets have become
the norm for most manufacturers.
In this design approach the fringe field is suppressed, rather than contained.
Current flow in a set of extra superconductive windings generate a field
that oppose a proportion of the field produced by the main windings in
a manner that balances out the dipole, or ovoid, pattern of the fringe
field, leaving the quadrupole fringe field pattern. The quadrupole field
intensity decreases much more rapidly with distance form the magnet centre.
(Dipole B µ 1/d3 Quadrupole
B µ 1/d5).
The main windings of a 1 T active magnet generates a 1.5 T field opposed
by a 0.5 T shield winding, thus active shield magnets are considerably
more expensive than unshielded models. This extra expense is usually offset
by avoiding building modifications and the cost of passive shields. Active
shield magnets weigh only a little more than unshielded magnets.
The prime disadvantage of simple active shield magnets is that at high
fields they can be more sensitive to moving ferrous objects than an unshielded
magnet. This sensitivity results from the dynamic nature of the balance
of the shielding field and the main field and their interaction in the
centre of the magnet bore to create a homogeneous field for imaging. Transient
disturbances in the ferromagnetic environment of the fringe field alter
the current flow in the main magnet windings and the shield magnet windings
differently, disturbing the balance between their fields and varying their
interaction in the imaging region, thus creating a larger inhomogeneity
than would have been the case for an unshielded magnet. The creation of
a significant field disturbance depends on the magnetic susceptibility
of the object, its mass, and the distance form the magnet centre (or more
accurately the strength of the fringe field it encounters). Thus small
steel objects such as beds, and wheelchairs close to the magnet can create
transient problems of field inhomogeneity as well as heavy trucks or traffic
much further away.
Oxford Magnet Technologies has addressed this problem in its recent
shielded magnets by adding extra superconductive windings to the outside
of the main and shield windings. These "B0 Shield" windings
are shorted together but isolated from the magnet windings. In the normal
state do have no current flowing. They are constructed so that when the
fringe field is distorted the extra currents induced flow in the B0
shield windings rather than the main or shield windings, thus preserving
their balance. The current in the B0 shield is discharge automatically
at a time when the magnet is not in use. This style of magnet is supplied
to Siemens, Picker, and Toshiba for a variety of field strengths using
the commercial name of external interference shield (EIS).
Considerable debate has flourished about the optimum field strength for
MRI scanners. Selection discussions which centre purely on the issue of
field strength are frequently ideologically or commercially motivated,
or indicate a poor understanding of the role other components of an MRI
scanner. Field strength will dictate to a certain extent the level of imaging
speed, image quality, and capability of the scanner, however the combination
of magnet homogeneity, gradient system performance, scan sequences, and
coil inventory play a greater role in determining the overall capacity
of the machine. The required capability of a unit should be determined
relative to its planned use, with potential purchase choices rated against
this. The final significant determinant of imaging performance remains
the competence of the MRI Radiographer, as displayed through appropriate
patient management and technique selection, working to extract the optimum
image and examination in each circumstance.
Does Field Strength Equal System Capability ?
The spatial uniformity of magnets are designed as near to perfect as
possible, but manufacturing tolerances and ferro-magnetic material near
the installation site will introduce significant spatial variations in
the magnetic field of the sensitive volume that will make imaging impossible
or sub-optimal. In addition to these "permanent" influences on field homogeneity,
the ferromagnetic environment of a site may change as buildings, car-parks,
or structures surrounding the MRI installation are altered through the
life of the equipment. These variations must be compensated for by a process
called "shimming". Shimming has the aim of creating a large volume highly
uniform magnetic field. The resultant field is called the sensitive volume,
and is defined as field variations in parts per million, over a specific
volume diameter ie. +/- 2 ppm 50 cm. DSV. On a shorter time scale, ferromagnetic
patient trolleys, large vehicle traffic around the scan site, and the individual
patient, can introduce transient field homogeneity variations.
Shimming levels are not as critical for spin echo imaging where 180o
pulses recall the echo in a manner that compensates for up to 50 ppm of
local field homogeneity, but for gradient echo and multi-shot sequences
(EPI, GRASE, segmented GE) a high order of field homogeneity ensures a
long T2* and the preservation of the Mxy for as long as possible. Fat saturation
techniques, offset field of view imaging of joints, particularly with gradient
echo sequences, and long field of view spinal imaging with phased array
coils also place extra value on the production of a large well shimmed
volume. These sequences are more frequently being used in standard protocols,
so the homogeneity figures of the magnet and the capability of the shim
system have become more critical issues in system design. The uniform field
achieved by shimming also has a role in ensuring accurate slice thickness
and profile, slice location, and geometric fidelity for all imaging.
Shimming is typically achieved by some combination of four methods:-
Gradient Offset Shimming
Small ferromagnetic plates are fitted within the energised magnet at specific
locations to regionally correct field distortions. The quality of shim
available form passive systems depends on the range of plates and locations
possible and the quality of the homogeneity correction programmes used
to calculate the required plates and positioning.. Older systems could
only compensate for gross field inhomogeneities but current systems can
achieve very high order homogeneity. Passive shimming is perfectly stable
over time with no power consumption, but it cannot compensate for short
or long term changes in the external field environment, or for patient
Active Resistive Shimming
A collection of electromagnetic coils (the shim set) are fitted within
the magnet bore. These coils are energised with suitable current levels.
A typical shim set has 12 to 18 sets of coils which operate to distort
the field in the primary axes of the magnet (X,Y,Z). For each plane there
can be coils that provide a linear field distortion (1st order shim) a
field that varies with the square of distance (2nd order shim) and so on,
as well as inverse 1st, 2nd and 3rd order shim sets. Shimming is achieved
by choosing the appropriate current for each coil MR systems using resistive
shimming are subject to image degradation if there is any drift in the
shim power supplies and will typically requires a stabilisation period
after daily power up
Each patient in the bore caused unique field homogeneity variations
due to their body shape and particular spread of tissue and magnetic susceptibility.
As the set up of a resistive shim system can be rapidly changed it can
adjust for these individually induced inhomogeneities in a process known
as patient shimming. Patient shimming is essential for MR spectroscopy
and for achieving optimum results for frequency based fat or water suppression
techniques, MTC techniques, where signal loss due to T2* relaxation times
becomes significant such as high order fast spin echo, GRASE, segmented
k-space, and EPI
Patient shimming is typically automatic, and can compensate to some
degree for stable changes in external environment. At least one installation
has used servo linked alteration to the shim set to automatically compensate
for nearby elevator traffic.
Active Superconductive Shimming
This approach is similar to resistive shim system, but using a shim coil
set of superconductive wire fitted within the cryostat. The result is a
highly stable shim with no need for constant power application. By design
the superconductive shim system usually cannot be quickly adjusted to suit
Gradient Off-Set Shimming
In many systems the gradient coils are used to create the linear or first
order shim corrections by carrying a small current flow at all times (offset
bias). This simplifies the shim set construction and saves space in the
patient bore by removing the need for a first order shim coil. The level
of bias current is kept small so as not to compromise the peak performance
of the gradient coils and amplifiers. Most current equipment carry out
patient shimming by adjusting gradient offsets.
To select the orientation of the MR scan plane, and to spatially localise
the MR signals within each slice, a series of "gradient" magnetic fields
are applied during the imaging sequence. These fields are called gradients
because they vary linearly in space. Their effect is to strengthen or weaken
the main magnetic field in a spatially predictable way.
The application of a gradient field in conjunction with an RF pulse
of controlled frequency bandwidth allows control of the location and thickness
of the slice of tissues that will experience NMR (slice selective excitation).
Localisation within the slice in one dimension within the excited slice
is achieved by applying a gradient during echo formation and signal readout,
which frequency encodes the component of the MR signal depending on spatial
location (frequency encoding). Localisation in the final dimensions is
achieved by analysing the effect of spatially dependant phase shifts created
by transient gradient applications before signal readout (phase encoding).
The Gradient system is also used to restore the phase coherence of transverse
magnetisation in gradient echo and spin echo sequences.
MR uses three gradient fields, each made to vary the field in one of
the three prime axes of the machine. The fields are created by passing
DC pulses (gradient current), through the gradient coils. The gradient
currents are very high(200 - 600 Amps) and need to be switched on and off
quickly (<< 1 ms) and accurately. This is achieved by digitally controlled
high power amplifiers, referred to as the gradient amplifiers. Each of
the three gradient coil sets has a dedicated gradient amplifier, the combination
is controlled by a computer called the gradient control unit. Pulse shapes
used to switch the gradient amplifiers are created by the host computer
or by dedicated pulse control systems. Sophisticated gradient pulse control
makes possible Oblique slice selection techniques to reduce motion artefacts
(gradient moment nulling) and the localisation of saturation slabs.
Due to the high current loads and heavy duty cycle, gradient systems
dissipate a lot of heat. Gradient amplifiers need air cooling, some older
designs require water cooling as well. Standard gradient coil sets are
cooled by forced air, but most newer high powered systems use closed loop
The gradient coil set is secured firmly within the magnet just inside
the shim coils As the current flows in the coil the interaction of the
gradient field and the main field tries to distort the gradient coil or
thrust it out of the bore. The flexing of the gradient coil structure is
heard as the beating noise of the MRI scan. Some manufacturers entirely
encase the gradient coil set in epoxy resin to quieten the system, however
this approach may impedes cooling. Stronger gradients will create higher
noise levels experienced in the bore, although there is some hope that
better gradient coil designs will reduce noise levels. Current typical
noise levels of >85 dB are already sufficient to require that patients
must be given hearing protection during MR procedures.
Radio waves are not involved in the MRI process. The electromagnetic waves
used for radio transmission and reception are a cross polarised combination
of electrical and magnetic fields with the electric field component about
120 times stronger than the magnetic field. Nuclei virtually never emit
radio waves, and the design of MR coils aims at produce a rotating magnetic
field and little or no electrical field. For MR the "Radio" tag comes simply
because the Larmor frequency of NMR nuclei fall within the section of the
electromagnetic spectrum known as Radio Frequencies (10kHz-2GHz). While
this discrimination may appear pedantic, it has two redeeming features.
First it is a more accurate description of the process, and secondly it
clarifies the description of the role of the RF system in NMR and MRI signal
production by emphasising the role of a simple concept, electromagnetic
Radio-Frequency (RF) Systems
All excitation energy is provided by magnetic fields rotating in the
X-Y plane, at the Larmor frequency of Hydrogen (42.58 Mhz/T). These rotating
magnetic fields are created by passing RF alternating current, generated
by the RF transmitter, through suitably shaped and orientated conductors
known as the RF coils. This energy realigns the nuclear magnetic fields
of the subject into the transverse plane and bring the nuclei fields into
phase coherence by acting as a fixed magnetic field in the rotating frame.
The nuclear magnetic fields rotating coherently in the transverse plane
induce RF currents (MR signal) in the receive coil which are detected and
analysed in the receiver system.
RF DESIGN ISSUES
The RF shield is a Faraday cage, ie. a closed conductive structure, that
usually encloses the entire MRI examination room. This structure excludes
all RF fields from the environment of the patient and the receiver coils.
This is necessary as the magnitude of the received MR signal (10-9 watts)
is 1000 times weaker than RF field energies typically encountered in a
hospital environment. Without the RF shield these environmental fields
could swamp the MR signal and imaging would be difficult. The shield can
be constructed of any conductive material but non ferro-magnetic materials
are preferred. Doors must have special seals fitted, windows are usually
double glass laminated with a fine conductive mesh, and all cables into
the room must be fitted with RF filters. Non-conductive lines can pass
through tubes with an appropriate length to diameter ratio (L=5D), similar
tubes are used in air-conditioning ducts and cryogen venting pipes.
RF Transmitter Design
The RF transmitter provides the excitation pulses for the MRI pulse sequences.
It must be capable of accurately delivering pulses of RF current of controlled
amplitude, frequency, and phase. Sophisticated pulse prescription components
create the low frequency modulation envelopes of the required RF output.
These pulses are used to modulate the output of a high quality RF power
amplifier (RFPA). Typical MRI amplifiers have maximum power capabilities
of between 5 and 20 Kilowatts. Solid state or valve final amplifier stages
may be used with no obvious advantage in one system over the other. Amplifier
control can be analogue or digital, but RF pulse shaping and timing must
be digitally controlled for best results.
RF Receiver Design
The receive system handles the extremely low energy signals induced by
the patient in the RF coils. Conventional receive systems have sufficiently
high sensitivity and low noise figures for this task. The key issues for
MRI receivers are stability and control of amplification, the accuracy
of frequency and phase shift detection, and reference to scanner wide time
base. The received signals are amplified and then demodulated in a phase
sensitive detector to yield real (sine) and imaginary (cosine) signals
that represent the phase and amplitude data of the MR signal.. The signals
are digitised to become one line of the raw data array of an MR image (k-space).
The timing of the received signal relative to the phase encoding regime
of the gradient system dictates which line of Ky the data represents. Current
MR systems have digitally controlled receiver systems to allow more accurate
timing of operation. "Digital" receive systems vary from traditional analogue
receivers by digitising the MR signal before demodulation, virtually all
MR receivers have digital control of amplifiers and attenuators. Several
manufacturers are equipping scanners with multiple receivers operating
simultaneously to allow the use of phased array receive coils.
The coils produce a rotating magnetic field from the transmit (Tx) current
and have RF current induced in them by the rotating nuclear magnetic fields
of the patient during the receive phase. Coils are designed to create a
magnetic field (B1) within a specific region, by the principle of reciprocity
this is also the area within which time varying magnetic fields will be
able to induce current flow in the coils, and the region is known as the
sensitive volume. Coils design dictates that they respond preferentially
to a range of frequencies, thus coils manufactured for a 0.5T machine will
not work in a 1.5T scanner. Designs which respond more sensitively to a
very narrow range of frequencies are known as high Q (quality) coils while
lower Q coils will respond less sensitively to a broader frequency range.
RF designers generally choose to pursue a high Q or low Q approach and
this affects the need for accurate patient tuning of the RF coils (see
coil matching below). Overall the choice of a high Q or low Q design has
little discernible affect of image quality of the scanner.
Coils and Signal to Noise Ratio
While coils are to a degree frequency discriminating, they do not discriminate
between currents induced by the transverse magnetisation representing tissue
NMR behaviour (signal) and the portion of black body radiation generated
by the motion of molecules (thermal noise) that happens to be aligned in
the transverse plane and at the MR frequency. This noise degrades image
quality in all situations. The slices selective approach to MRI causes
signal to be generated only by part of the tissue in the sensitive volume
of the coil but noise comes from all of the coil volume. Therefore it is
best to select a coil with a sensitive volume that matches the target volume
(useful field of view in 3 dimension) as closely as possible. As long as
the target volume is included in the coil's sensitive volume, the smaller
the coil the higher the signal to noise ratio. Typical MRI machines are
available with body, head, and extremity coils and a range of anatomy specific
surface coils, although the range and quality of coils varies between manufacturers
and MR systems. The MR radiographers must choose the most appropriate coils
for the examination to ensure optimum results. In day to day
operation, the selection of an appropriate sized coil is the primary means
of assuring the best image quality.
CP versus LP Transmission
In all alternating current systems the simple electrical resistance of
a circuit is complicated by a frequency dependant component known as reactance.
The reactance of a circuit is derived from inductors (coils or turns of
conductors) and capacitors (gaps, spaces, or components). and impedance.
Reactance and resistance combined together represent the impedance of the
circuit which is its resistance to current flow alternating at a given
frequency. To allow for loss free transfer of current from the transmitter
to the coil, or more crucially the coil to the receiver the impedances
of these components must be matched. This is usually done by adjusting
capacitors in the coil tuning circuit, which may be incorporated into each
coil, or built into the scanner where coils are plugged in. systems as
different patients will have varying natural Larmor frequencies which may
Coil Impedance Matching
The transmit coil in particular must have a uniform sensitive field which
covers the entire field of view as the strength of the B1 field must be
equal throughout the volume to ensure all nuclei magnetic vectors are flipped
the same degree. Many systems use all of the uniform volume coils as transmit
coils, but some equipment uses the body coil as the transmit coil in all
applications and optimises receive coil sensitivity at the expense of minor
volume uniformity variations.
Uniform Volume Coils
In many applications (head, abdomen, limb joints) whole axial sections
of the body need to be imaged, so a coil that delivers uniform sensitivity
over its volume is desirable. The size of the coil is chosen to achieve
the highest filing factor. This is usually achieved with cylindrical coil
designs placed around the patient.
The body coil is the largest uniform volume coil, fitted within the
magnet bore immediately under the bore covers. Smaller diameter coils are
used for extremity work and head imaging. Coil sensitivity is related to
coil volume. A 30 cm. head coil will have nearly 4 times the sensitivity
of the 50 cm. diameter body coil, while the smaller extremity coil can
have 6 times the sensitivity. The smallest coil that can accommodate the
body part to be imaged will usually give the best result.
In many applications the transmit coil is the same as the receive coil
(head, body, knee, ankle) but for a range of common examinations where
the anatomy of interest is near the surface (spine, TMJ, shoulder, orbit)
specialised receiver coils called surface coils provide large advantages.
A receive coil can have a spatially varying field sensitivity. Most surface
coils are flat and receive signals more sensitively from the regions close
to them, so surface coil images appear brighter near the coil plane. This
is used to advantage by excluding unwanted tissues while maximising the
signal from target structures. Within its useful region a flat surface
coil will exhibit sensitivity 1 to 5 times higher than a cylindrical coil
that might cover the region. This increased sensitivity can be used to
increase image quality or traded for increased resolution.
Phased array receive systems attempt to apply the SNR advantages of small
field of view coils to relatively large fields of view by aiming a number
of small receiver coils at the area of interest and combining their signals
in a manner which cancels noise. They offer increases in SNR of 40% to
200% in a range of clinical applications, and should radically alter the
imaging approach used in spine, neck, pelvis, breast, paediatric body,
thorax and upper abdomen examinations.
In conventional MRI systems the receiver is typically connected to
one coil that is of sufficient size to cover the target field of view.
With a phased array system there are multiple (typically 4) receivers.
Each receiver is connected to one of a number of a small volume receiver
coils constructed in an array to surround a specific body part. The coils
must be designed and positioned so as not to constructively interfere in
any area of their sensitive volumes, so that the noise each coil receives
is not correlated to noise received by other elements of the array. Coils
in this configuration are said to be decoupled. The MR signals from each
receiver are combined in a way that adds the MR signal from the patient
but allows the random unrelated noise components to mutually cancel. The
result is the coverage of a large area with the performance of the small
coils, and significant noise reduction.
G.E. researchers P. Roemer, W. Edelstein et al first described this
concept of simultaneous multiple reception as a method to extend the performance
of conventional surface coils. The same workers later coined the phrase
"phased array" when describing the flat style of array, and detailed techniques
for collecting and combining the received signals of their four independent
receive systems. Shortly thereafter a related group described a volume
phased array. These three papers effectively defined the configurations
of the first wave of commercially available phased array systems released
by G.E. in late 1991.
The initial phased array systems were linearly polarised (LP). Subsequently
circular polarised (CP or Quad) PA has been developed. Quad PA is technically
more difficult to implement than LP but can offers a further 20 to 40 percent
increase over a similar LP array depending on the dimensions of the coil
elements employed. Because of its high hardware component, (new receiver
chains significant extra raw data memory, new coils, and extra demands
on array processing) subsequently upgrading to phased array is expensive
(approx $100,000 to $150,000 plus coils), but it should be clear that phased
array capability is essential to any MR user seeking high quality imaging.
More recent work has aimed at switching the received signal of 4 de-coupled
coils into a single receiver in a shared time concept, but this approach
seems to place severe restrictions on sequence options.
Phased Array Coil Configurations
The spine arrays are developed from the initial flat array described by
Roemer et al in 1988 and incorporating the switched array methods described
by Siemens researchers at the same SMRM meeting. The individual coil elements
of a switched array can be selected or de-selected via software which can
even be incorporated into the scan sequence. The arrays are typically 60
cm. long and contain 6 to 9 small surface coil elements of which up to
four are switched on at a time to yield two of three 40 cm. fields. To
obtain good resolution, a rectangular 512 matrix is often used.
These coils are designed to allow full spine imaging without patient
handling. They are especially well suited to multiple region exams for
trauma, suspected spinal block by metastatic disease, identification of
spinal multiple sclerosis plaques, and the vaguely localised symptoms of
spinal cord ischaemia.
Body Array (Liver Array, Pelvic Array)
An array of 4 surface coils, two fitted posteriorly and two anteriorly
around the body for fields of view about 200 mm to 400 mm and offering
at least 200% SNR increase over the conventional 60 cm. body coil. There
is some inhomogeneity across the imaged volume with these variable spacing
arrays, but in practice this should not be a problem. Body arrays can be
used to image the liver, pelvis, thorax or as a paediatric whole abdomen
coil. This type of coil should largely replace the body coil as a receiver
device and may see the end of endorectal and vaginal coil usage.
Designed for head and neck examinations particularly MRA. Coverage extends
from the aortic arch to the top of the head and performance is significantly
better that the conventional head neck array which suffers low SNR due
to its high coil volume.
A body array incorporating an disposable endocavity coil as one of its
elements to allow high resolution at the prostate or cervix with extra
detail in the surrounding field of view. This product (only offered by
one manufacturer) may provide a more workable image than a conventional
endocavity exam, but the question will be wether adding the endocavity
coil to the array supplies sufficient additional useful data to justify
The array processor converts the raw data into image data using a mathematical
algorithm known as Fourier transformation. It is a dedicated computer system
which communicates with the host computer drawing raw data from a temporary
storage device or RAM, and transferring the completed images to the
main archiving device. The time taken to process an MRI image greatly affects
the speed of the MR system in practical applications. Typical image processors
take 0.05 to 0.1 second to process a 256 x 256 matrix image. As a modern
scanner operating at full capacity may produce between 1000 and 2000 images
per day even small differences in image processor speeds become significant.
This aspect of performance has become more significant with the increasing
use of high 5122 imaging matrix (which takes 4 times as long
to process as a 2562 image), and phased array images which takes
4 to 5 times as long.
IMAGE PROCESSOR (Array Processor)
The operators console allows the radiographer total control over scanning,
image display, archive, filming, photography and patient monitoring functions.
Many systems are also fitted with ancillary console for the use of radiologists,
a second radiographer filming, or other clinicians. These second consoles
cannot control the scanning functions and operate without reducing the
speed of the main system functions. This lack of interference depends largely
on the design of software and the computer network. Second consoles are
commonly used for prescribing and viewing reconstructions and angiograms,
and should be connected to the main image output devices.
Increasingly MRI systems are employing stand-alone image work-stations.
These devices have an independent computer system but communicate with
the host computer and image storage devices. They allow complex image manipulating
software to be used virtually without interference to the main system.
Work-stations are typically used to generate MR projection angiograms,
and display these dynamically as rotating 3D perspectives, generate full
volume models of the region imaged for interactive cut-aways, and arbitrary
plane reconstructions form 3D data sets.
The host computer co-ordinates all functions of the MRI system. It handles
all instructions from the MR operator via the MR operating software, all
image manipulation and display functions, image and file managements tasks,
and the creation of control pulses throughout the scanning period. Fault
finding, service, quality control and external communication software is
included on many current systems. The capacity to handle several processing
tasks simultaneously is determined by the architecture and capacity of
the host computer in conjunction with its operating system software. Simultaneous
or parallel tasking can greatly increase the speed and useability of the
Currently available MRI scanners use host computers with a wide range
of capability and speed. Recently manufacturers are acknowledging that
image handling systems require large computing and storage capacity and
improving this aspect of their design. It is essential to choose a machine
with some redundant computer capacity and an active path for upgrades.
During the examination the patient is isolated in the bore of the MRI magnet
for 45 minutes to 2 hours. Claustrophobia and boredom will affect many
patients, so some amusement and good communication with the Radiographer
throughout the examination is essential Virtually all systems have a hand
held device (usually pneumatic) for allowing the patient to attract the
radiographer's attention. Most systems are supplied with a room audio system
carrying a voice channel and music from radio (FM), CD or tape players.
The audio should also be directed to a variety of sound reducing headsets
compatible with the range of RF coils used. If any of these facilities
are not available on a unit, they should be retro-fitted. A video camera
is often used so that the radiographer can see the patient in the bore.
Devices also exist to display video (Closed circuit, TV programmes, diagnostic
images) to the patient and other personnel in the room.
PATIENT COMMUNICATION & MONITORING
Most MR scanners are equipped with MR compatible ECG sets, a pneumatic
bellows fitted around the abdomen or a nasal thermistor for providing a
trace of respiratory motion, and a finger probe for displaying peripheral
blood pulsation. These devices are primarily intended for triggering or
gating scan sequences to reduce artefacts from physiological motion, or
to better image that motion. The ECG and pulse probes particularly can
be used in a secondary role for patient condition monitoring, but most
manufacturers warn against this. MRI compatible patient
monitors are available although there is no standard code on what constitutes
an MRI compatable device. The monitors should not interfere
with the MRI signal, and must not present any hazard to the patient or
staff. Care and dilligence are needed to asses the suitability
of any monitoring devices that have conductive components attached
to the patient as burns have been caused at many sites.
Issues of patient monitoring
are further discussed in an associated paper on this website.
MR images are purely digital, after production they exist as data stored
on computer disc. This data is frequently archived to tape or optical disc.
Virtually all diagnosis and clinical use is based on these images presented
on large format transparent film (hard copy). As the image data contains
a greater dynamic range than any visual system can reproduce the images
are photographed by the Radiographer to best display the relevant information.
Photography can make or break an examination, and is a basic skill of every
radiographer working with a digital modality. Hard copy is produced on
a laser camera, or multi-format video camera.
IMAGE OUTPUT DEVICES
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