High Field MRI and Safety I Installation1

A. MAioRAnA, T. ScARAbino, V. d'Alesio, M. Tosetti, M. ARmillottA, U. SAlvolini

High-field magnetic resonance (MR), originally developed in the framework of spectroscopy and functional neuroradiology, is set to become an important diagnostic tool not only in research but also in advanced clinical practice.

High magnetic fields afford a better signal/noise ratio (SNR) and consequently better spatial resolution in a shorter acquisition time, even though the diagnostic outcome is then subject to the dependence on the magnetic field of other factors that variously contribute to image quality.

The rationale for the utilization of high magnetic fields in MR diagnostic imaging is obvious.

The distribution of the population into two spin levels is statistically determined:

-f = e ne where AE = h • y ■ B0 depends on the static magnetic field, h and y are constants, nf is the spin population in the fundamental state and ne is the population in excited state.

An increase in B0 values results in an inversion of the states of the two populations, and therefore in a stronger MR signal.

In fact, the signal intensity is proportional to the square of the static magnetic field, since:

where the number of Nspin and Vspin and the voltage induced by each spin both depend linearly on B0 [1, 2].

However, if the signal is proportional to the square of the static field B0 and the noise is proportional to B0, then from 1.5 T to 3.0 T the SNR doubles. This allows an acceptable image quality to be obtained even with increased spatial resolution or reduced time of acquisition.

Clearly, achieving greater spatial resolution while minimizing undesired partial volume effects requires increasing the gradient steepness Gz and reducing the encoding frequency bandwidth , since the slice thickness is defined as:

In any case, shimming requires very strong field homogeneity. For instance, the first 3.0 T system installed in Italy achieves a field homogeneity of less than 1 ppm in a spherical volume of 33 cm, and of less than 0.3 ppm in a spherical volume of 24 cm, with a peak gradient ramp of 50 mT/m and a slew rate of 150 mT/m/ms.

Another key principle is that the resonance frequency <w0 = y ■ B0 depends on the static magnetic field. It therefore is 43 MHz at 1.0 T and 128 MHz at 3.0 T, resulting in greater radiofrequency (RF) absorption by biological tissues, whose conductivity increases with frequency. This poses problems when designing coils suitable for the greater power applied as well as for patient safety, as increased tissue temperature is one of the risks associated with RF electromagnetic fields [3, 4]. Finally, the increased susceptibility of tissues exposed to the magnetic field can result in local inhomo-geneities that cannot be corrected, and eventually present as artefacts. This effect can be exploited in functional MR BOLD studies, which are based on the changes in blood oxygenation generated by magnetic susceptibility inhomogeneities [2, 4, 5].

Notwithstanding technical and cost problems, high-field MR offers very significant advantages, and recent human imaging studies at 340 MHz have demonstrated that safety margins still exist above 3.0 T and 4.0 T, while spectroscopic analysis has overcome the 1 GHz threshold [6].

The growing interest in 3.0 T and higher magnetic fields and their expected increasing diffusion in clinical practice have brought the safety issues back to the forefront.

When planning the installation of a high-field MR unit, the strength of the static magnetic field is one of the major problems to be addressed by those responsible for safety and technicians alike. In fact, this is but one, albeit the most apparent, element to be considered when estimating the associated risks and benefits. Patients in the tunnel of a high-strength imager are exposed to a magnetic field many thousands of times greater than the earth's, even though no special patient or operator safety precautions are required compared with low or medium magnetic fields. However, during scanning patients are also exposed to gradient switch-

on and to RF impulses for signal decoding and spin excitation, both of which are related to the intensity of the static magnetic field and carry different though acceptable patient risk.

Approval of high-field MR tomographs for diagnostic purposes dates from 1997, when the US Department of Health and Human Services Food and Drug Administration, Center for Devices and Radiological Health (FDA), classified as carrying a significant risk, and therefore subjected to specific authorization, all MR systems having: static magnetic fields exceeding 4.0 T; a specific absorption rate (SAR) exceeding 4 W/kg averaged over the whole body for any 15-min period, 3 W/ kg averaged over the head for any 10-min period, 8 W/ kg for any gram of body or chest tissue in any 15-min period, or 15 W/kgfor any gram oftissue in the extremities in any 15-min period; field gradients sufficient to induce patient discomfort or pain; and acoustic noise reaching sound pressure levels of 99 dB(A) (with reference to the response curve of the human ear, A curve) or peak values exceeding 140 dB(A). Based on these conditions, a 3.0 T MR unit meeting SAR, gradient and noise requirements is substantially equivalent to a 1.5 T unit and does not carry a significantly greater risk, hence further restrictions [7].

At that time, the components of the early high-field tomographs available on the market were substantially similar to those of the 1.5 T machines from which they had been derived. They thus presented a number of drawbacks, such as poor homogeneity of the static magnetic field, insufficient gradient slew rate, inadequate coil structure in relation to the greater resonance frequency, excessive fringe field extension, noise and weight, and were also extremely expensive. All such factors hampered the diffusion of the technique, which remained confined to specific research fields for a long time.

By contrast, last-generation units have been conceived as high-field diagnostic machines and are equipped with technical features that allow operation with a degree of safety comparable to the conventional low- and medium-field scanners [3, 8]. The passive-shield 3.0 T prototype was bulky and heavy and its 0.5 mT line (or 5 gauss limit) was an ellipsoid measuring 8.5 x 6.5 m in an axial radial direction. Adequate additional shielding of the fringe field in the magnet room would have required hundreds of tons of iron. The recent introduction of active shielding has significantly reduced magnet weight. The 3.0 T unit installed at the San Giovanni Rotondo Scientific Institution is still bulkier and heavier than the corresponding 1.5 T machine, but the 0.5 mT line occupies a 4.4 x 2.7 m area in the scan room compared with 3.9x2.4 m of the same type 1.5 T magnet installed 10 years previously [4]. In particular, the 200 mT line, which is crucial towards meeting current operator safety regulations, is practi cally contained in the tomograph's volume. This feature protects staff from excessive whole-body exposure even for long stays in the scan room, while also due to the presence of passive ferromagnetic barriers around the scan room perimeter, the 0.1 mT line barely exceeds the boundary of the magnet room

Due to the intrinsic weakness of the MR signal and to the high sensitivity of image reconstruction systems, the magnet room needs to be protected by a Faraday cage capable of attenuating the outside electromagnetic noise. The shielding must be able to attenuate signals by at least 80 dB, and it obviously also works the other way around, i.e. by shielding the operators from any type of exposure to the RF electromagnetic fields generated by the coils.

In 1999, the FDA approved for sale some high-field tomographs for clinical diagnostic imaging [4].

In July 2003 it replaced its 1997 guidelines and laid down new upper static magnetic field limits for MR diagnostic units. These new limits are 8.0 T for adults and 4.0 T for neonates less than 1 month old [9].

In Europe, particularlyin Italy, the earliest reference technical regulation, CEI EN 60601-2-33/A11 of 1998 [10], was superseded by IEC 60601 1-2-33 Ed. 2.0 of 2002 [11], which later became CEI EN 60601-2-3 of 2004 [12].

Recently, the International Commission for Non-Ionizing Radiation Protection (ICNIRP) has issued patient safety guidelines for MR scanning [13].

In Italy, MR units with a static field strength of or exceeding 2.0 T are not approved for clinical use and are restricted to documented research applications [14]. Further legislation [15, 16] lays down SAR and field gradient limits substantially in line with the indications of the Comitato Elettrotecnico Italiano (CEI) [10]. In addition, MR machines with static magnetic fields exceeding 2.0 T, classified by the law as group B systems, can be installed only at major research institutions subject to ministerial authorization [14]. Such units also need to be involved in scientific or clinical research projects mandating the use of such high field strengths, whereas units with magnetic fields exceeding 4.0 T may be authorized only for specific, documented needs of scientific or clinical experimental research limited to the limbs. Authorization of group B tomographs by the Health Ministry is currently also subject to the prior technical opinion of the Istituto per la Sicurezza e la Prevenzione degli Incidenti sul Lavoro (ISPESL), the Istituto Superiore di Sanita (ISS) and the Consiglio Su-periore di Sanita (CSS). In practice, meeting the requirements for the authorization of MR imagers, especially high-field ones, is hampered by an intricate and fragmentary legislation [17].

As regards staff safety and protection, limits for static magnetic fields were issued by the ICNIRP in 1994 [18]. Whereas the limits adopted in Italy [15] are sub stantially similar (save for those regarding the limbs), there are considerable differences in the time periods to which some of these limits are to be applied.

With high-field systems, correct dimensioning or upgrading of ventilation and helium venting pipes in relation to the type of magnet being installed and room size should be addressed at the design stage. Indeed, in the event of a quench, i.e. the sudden inactivation of the magnet, the liquid helium contained in a superconducting magnet (which is more abundant in a high-strength unit) rapidly turns to gas and may saturate the magnet room atmosphere.

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