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This chapter explains some of the basic principles that are used in the construction of a homogeneous superconducting magnet for magnetic resonance, including the basic physics of magnetic field generation, as well as superconducting materials and their properties. Concentrating on magnet design for human MRI, the chapter describes how the stray-field of a magnet can be controlled and how the magnetic field is stabilised. The cryogenic design and practical considerations for the operation of a magnet system are included with a number of key safety recommendations. Finally, the article discusses future developments, in particular operation of human systems at field strengths in excess of 10 T, and cryogen-free systems.

For whole-body MRI, an ultra-high field (UHF) magnet is currently defined as a system operating at 7 T or above. Over 70 UHF magnets have been built, all with the same technical approach originally developed by Magnex Scientific Ltd. The preferred coil configuration is a compensated solenoid. In this case, the majority of the field is generated by a simple long solenoid that stretches the entire length of the magnet. Additional coils are wound on a separate former outside the main windings with the purpose of balancing the homogeneity. Most of the magnets currently in operation are passively shielded systems where the magnet is surrounded by a steel box of 200–870 tonnes of carbon steel. More recently actively shielded magnets have been built for operation at 7 T; in this case the stray field is controlled by with reverse turns wound on a separate former outside the primary coils. Protection against quench damage is much more complex with an actively shielded magnet design due to the requirement to prevent the stray field from increasing during a quench. In the case of the 7 T 900 magnet this controlled by combining some of the screening coils into each section of the protection circuit. Correction of the field variations caused by manufacturing tolerances and environmental effects are made with a combination of superconducting shims and passive shims. Modern UHF magnets operate in zero boil-off mode with the use of cryocoolers with cooling capacity at 4.2 K. Although there are no cryogen costs associated with normal operation UHF magnets require a significant volume (10 000–20 000 l) of liquid helium for the cool-down. Liquid helium is expensive therefore new methods of cool-down using high-power cryocoolers are being implemented to reduce the requirement.

Superconducting magnets have been used in magnetic resonance imaging (MRI) since the early 1980s when the first low-field 0.35 and 0.5 T systems were introduced. Since then, field strengths have gradually increased and, today, 3.0 T is recognized as a standard field strength for clinical diagnostic imaging while magnets are also now available with field strengths up to 11.74 T for research in biomedical imaging applications. Superconductivity is the physical phenomenon in which materials have zero electrical resistance below a critical temperature and a critical field. Most MRI magnets are constructed from niobium–titanium (NbTi) superconductor and sit in a bath of liquid helium boiling at 4.2 K. The critical current and magnetic field characteristics of the superconductor determine the maximum field strength a magnet can achieve, although operating at temperatures below 4.2 K provides enhancement of the superconductor properties allowing for higher field strengths to be achieved. Controlling the extent of the stray fields around MRI magnets has always been an issue, and various techniques have been employed utilizing both steel shielding (passive) and active shielding (reverse wound coils). Today, virtually all MRI magnets up to 7.0 T use active shielding technology to reduce the stray fields and help with magnet-siting issues. The liquid helium reservoir containing the magnet is housed in a cryostat that is designed to minimize the heat loads reaching the reservoir. The cryostat is an evacuated vessel with one or more radiation shields cooled by a cryorefrigerator. Today, most systems use a cryorefrigerator that provides sufficient cooling power at 4.2 K to allow the evaporated gas to be recondensed and operate with no loss of liquid helium. This article describes the current state of superconducting magnet technology used in MRI systems. Keywords: superconducting; magnet; cryostat; shimming; shielding; high field

Purpose: The purpose of this report is to describe the design and construction of an 8 T/80 cm whole-body MRI system operating at 340 MHz. Method: The 8 T/80 cm magnet was constructed from 414 km of niobium titanium superconducting wire. The winding of this wire on four aluminum formers resulted in a total inductance of 4,155 H. Gradient subsystems included either a body gradient or a head gradient along with a removable shim insert. The magnet and gradient subsystems were interfaced to two spectrometers. These provided the control of the gradient amplifiers and the two sets of four RF power amplifiers. The latter provide in excess of 8 kW of RF power from 10 to 140 MHz and 10 kW of RF power from 245 to 345 MHz. A dedicated computer-controlled patient table was designed and assembled. The entire system is located in a clinical setting, facilitating patient-based studies. Results: The 8 T/80 cm magnet was energized without complication and achieved persistent operation using 198.9 A of current, thereby storing 81.5 MJ of magnetic energy. Exceptional performance was observed for nearly all components both in isolation and when combined within the complete system. Conclusion: An 8 T/80 cm MRI system has been assembled. The magnet subsystem is extremely stable and is characterized by good homogeneity and acceptable boil-off rates. Index Terms: Magnetic resonance imaging, apparatus and equipment-Magnetic resonance imaging, techniques-Tesla-Image quality.

A 3.0 Tesla, 0.80-m bore magnet replaced our previous 1.9 Tesla, 0.76-m magnet. The 3.0 Tesla replacement magnet had a dipole moment of 1.7 with respect to that of the 1.9 Tesla magnet. The pre-existing cylindrically symmetric passive steel shielding was modified to confine the fringe fields of the replacement magnet to match those of the previous magnet. A cylindrically symmetric inner shield insert of about 20,000 kg was designed, fabricated, and installed. Upon energization of the magnet, the combined shielding met all design criteria. Alternative designs, optimization of cylindrically symmetric designs, and costs of fabrication, are presented herein.