MR hardware I
MRI system
MRI system
Main magnet (properties)
• We learned that the Larmor precessional frequency depends on the static magnetic field applied: ω0 = γB0 where the gyromagnetic ratio γ is a large scaling factor, so that small variations in the static magnetic field produce large undesired frequency variations.
• It is very important therefore, that the homogeneity of the static magnetic field generated by the main magnet be extremely good (< 5 ppm) and very stable with time (< 0.1 ppm/hour).
• For good SNR performance, the field generated by the clinical system must be very intense (0.2- 3T, and above for research and animal applications).
• The main magnet in an MRI system is one of the most critical components to the performance of the instrument.
• In general there are three classes of MR magnets:
• Permanent magnets
have a weak field strength and are not very practical.
• Electromagnets
are generally more suited to low field applications
due to difficulties of power dissipation and stability • Superconducting magnets
above ~0.3T, superconducting systems are used almost exclusively, and form the basis of clinical MRI systems
•Superconductivity is a physical phenomenon whereby certain metals and alloys when cooled below few degrees Kelvin behave as though they are ‘perfect’ conductors.
• A typical superconductor used for MRI is a Niobium Titanium (NbTi, type II superconductor) alloy in a cooper matrix. Anywhere from 50 to 400 km of the wire may be required for a modern MR magnet.
• Right: 0.7mm diameter superconducting wire with ~6,000 filaments of 5μm NbTi.
Main magnet (wire material)
Main magnet (limit of superconductivity)
• Unfortunately, superconductivity is not without its limitation. The current density (A/m2) that conductor can safely carry is dependent on the temperature of the conductor and static field strength it resides in.
• In general, the lower the temperature and the lower the immersed field, the greater the ability to carry current.
• NbTi coils are often kept at ~4.2K inside liquid helium.
• The critical field strength, above which superconductivity ceases is ~10T for NbTi.
• In a conventional 1.5T system, for example, the coils would reside in a maximum field of about 6.5T.
• The implication of using superconducting coils are that once they are energized to the desired current, the complete loop may be closed, the power supply removed and the current still circulates! (provided the coil structure is maintained at a low enough temperature)
• Surrounding the coils and radially outwards is a range of chambers designed to provide a thermal gradient to the outside world and prevent liquid helium from evaporating.
•This is very important as liquid helium is expensive and the longer the time between refills the better.
Main magnet (cooling system)
Main magnet (coil geometry)
•Most clinical magnets use coaxial superconducting coil – the exact structure (position, dimensions, number of turns per coil, type of wire in each coil etc) of which requires very careful design.
• In most designs, a very pure magnetic field is generated at the centre of the magnet structure.
•The entire system has generally a rather large inductance.
Main magnet (stray field)
• Magnet designs are further complicated by the need to restrict stray or fringe fields that emanate from the system.
• People working in the nearby environment must be working within statutory working exposure limits to static magnetic fields.
Main magnet (shielding)
• In order to prevent flux leakages from magnet systems, the general methods of magnetic shielding have been developed: passive and active shielding.
• Passive shielding involves the use of a 3D iron or steel box being constructed around the magnet – I.e. walls, floor, ceiling of the magnet room. This results in a significant structure of often considerable weight (100-200 tonnes!), having implications for the floor structure on which the magnet is placed.
• Active shielding involves the use of an additional set of superconducting coils at a larger radius than the primary coils, but where the current density is opposite in sign to the primary. This method has the advantage of enabling compact magnet designs.
• Why high-field magnets?
•Higher signal intensity (SI) and signal to noise ratio (SNR)
• Decreased image acquisition time
- minimize motion, kinetics and breath holds
• Higher spatial resolution
• Improved contrast and exploration of contrast agents
Main magnet (high field)
• Although routine clinical magnets are currently limited to an upper field strength of 3T, magnets for research purposes have been installed for human imaging up to 8T.
• Opens new doors for anatomic, metabolic, functional and molecular MR imaging
- Improved spectral separation
- Increased BOLD effect
• There are still a number of technical issues to be solved for very high field clinical imaging.
• able to be pulsed very rapidly (100-200μs rise-time).
• low resistance to minimize power dissipation
• gradient coils need to be shielded so that eddy currents are not induced in the cold main magnet
• torques/forces and accompanying acoustic output and vibration should be minimal during current switching.
• cooled (water-cooled) to maintain a suitable temperature
Gradient coils (properties)
• The role of gradient coils is to encode the MR signal with its position of origin in the patient and thereby permit the formation of images.
•Gradient performance is very important in producing fast and accurate imaging examinations.
• very linear gradients (<5% deviation) over whole DSV.
Gradient coils (simple designs)
• A Gz gradient can be generated by two circular coils (z-gradient coil) separated along the z- axis by a distance of 3a (a is the coil radius) and where the current in the coils is flowing in the oppositr direction. This is termed a Maxwell pair.
• A similar process to the design of Gz gradient coils may be followed to design transverse gradient coils Gx and Gy. The resultant designs tend to use a ‘4-loop’ configuration around the circumference of the cylinder and are called Golay coils.
Gz Gx and Gy
Gradient coils (distributed designs)
• Modern gradient coils are designed with multiple turns to increase the sensitivity (mT/m/A) of the coil and much effort is included in the design process so that the coil may still be rapidly switched.
Gz
Gx and Gy
• As gradient fields are pulsed, the fields that they generate outside the coil, link with other conductors and induce currents in them (Faraday’s Law of Electromagnetism).
•The induced eddy currents generate their own fields that generally oppose the initial field generated by the coils and subsequently distort the intended spatial and temporal properties of the gradient fields in the DSV.
• These effects are undesirable and continue to be a problem since the advent of MRI.
Gradient coils (eddy currents)
• Just as magnets need to be shielded to reduce stray fields, so do gradient coil, although the reasoning behind this is somewhat different.
Gradient coils (shielding)
• Gradient coils are typically actively shielded, which involves placing an additional layer of coils radially outwards from the primary coils and which, while connected in series, are wound in the opposite direction to the primary coils.
• This effectively reduces the leakage fields to the main magnet and consequently the undesired eddy currents.
• imperfections in the magnet manufacturing process, as very small displacements in the coils during manufacture and cooling to cryogenic temperatures can result in significant magnetic field impurities.
• the installation site, as large ferrous installations near the magnet can alter the homogeneity of the magnetic field over the DSV.
• the patient and radiofrequency coils in the DSV, which alter the magnetic field due to their magnetic properties.
MR image distortion due to ferromagnetic particles in the eyeliner and foundation make- up of a female patient
Shims
(B0 inhomogeneity)
• We have learned that the static magnetic field B0, must be as pure as possible over the DSV. If it is not, FID from a simple experiment will decay more rapidly than it otherwise would. This is undesirable as the SNR decreases and spectral line-widths increase.
• The static field inhomogeneities may be caused by:
Shims
(B0 inhomogeneity)
• The magnetic field in the imaging volume is measured and analysed in spherical harmonics and small thin pieces of ferromagnetic iron are placed at appropriate locations on a cylindrical surface (within the magnet bore) to cancel undesired spatial field variations. This process is called passive shimming of the magnet and is engaged after the magnet has been energized. Active shimming can also be used.
Illustration of a ferromagnetic iron shim matrix, where light greys denote larger shim thickness. Some iron pieces can be axially and/or radially longer than others.
B0 distortion caused by the head (a) B0 map
(b) Normal image