The development of magnetic resonance imaging (MRI) has vastly altered the field of radiology, particularly neuroradiology and by extension to a certain degree the practice of neurology and neurosurgery. Unlike the rapid implementation of X-rays after they were discovered by Wilhelm Roentgen in 1895, it took more than 40 yr before the clinical application of nuclear magnetic resonance (NMR) principles was realized. The fundamentals of NMR were first outlined by a Dutch physicist named G. J. Gorter in 1936 (1,2) and refined by Bloch and Purcell in 1946 (3,4) but it was not until 1973 when Lauterbur (5) suggested using magnetic field gradients and NMR to encode position information that the current clinical use of NMR medical imaging was established. Still, it required many years of contributions from the basic science fields of physics, chemistry, and engineering before the development and subsequent implementation of the clinical scanners that we use today.
Every clinical imaging MR magnet requires a minimum of five coils for operation. The most important is the main gradient (in the z direction, along the length of the bore) with a field of B0 that lines up all the hydrogen proton spins at Larmor frequency, which is defined as the product of field strength (B) and the gyromagnetic constant of the proton (y). The Larmor frequency of the water proton (chosen because of its abundance in body tissue) is dependent on the magnetic field strength, calculated as 42.57 MHz at 1 T and 63.86 MHz at 1.5 T. Magnetic field strength is invariably described in Gauss (G, smaller field) and Tesla (T, larger field) where 1 T =10,000 G. To give a perspective on the scale, the earth's magnetic field is measured at approx 0.5 G. The goal of the magnet is for B0 to give a spatial homogeneity of approx 10 ppm/40-cm sphere and temporal stability of 0.1 ppm/h.
The other four coils include a set of three orthogonal (z or slice-select, y or phase-encoding, x or frequency-encoding) gradients that can be turned on and off, which gives each proton spin a different time dependence that is essential for position encoding. The last coil is the radiofrequency (RF) coil, which
From: Minimally Invasive Neurosurgery, edited by: M.R. Proctor and P.M. Black © Humana Press Inc., Totowa, NJ
provides a uniform low-amplitude magnetic field near the Larmor frequency that can be switched on and off for excitation (short burst of RF at 0.1-10 kW for a few milliseconds) and for receiving or readout of signals (free induction decay 10-1000 ms) (6-8).
Other varieties of more specialized RF coils include surface coils (placed close to the anatomical site), which are often smaller receive-only coils tailored to specific anatomical regions such as the orbit, temporal mandibular joint, extremities, and spine for increased signal-to-noise ratio (SNR) and at the same time using the larger body RF coil as transmitter. Phase-array coils are in essence multiple surface coils assembled to give an improved SNR but able to cover a larger area or field of view often used in the setting of total spine imaging (9). The design of the orthogonal gradient coils is such that they are perpendicular to B0 (except the z gradient), the most common being the quadrature configuration that yields an increase in SNR of the square root of 2.
During the early years, the permanent and resistive magnetic materials used were limited by their tremendous weight (tons) and inability to sustain a homogeneous magnetic field. Most, if not all, of the current mid- to high-field magnets (0.5-4 T) are made from superconducting alloys (niobium-titanium) that are wound into a coil configuration and whereby the magnetic field is generated by an electric current that passes through these coils. To maintain a temperature below the transition temperature of the superconducting alloy (10°K), liquid helium at 4°K is circulated around the coil while at the same time further buffering with the ambient temperature is accomplished by circulating liquid nitrogen at 77°K within the outer layer (10) (Fig. 1).
All magnetic fields are generated from electric currents and electric fields secondary to either changing electric current or charge accumulation. The magnetic field that a proton experiences depends on the cumulative sum of fields from all five coils. Coil design and configuration, the materials used in the manufacture of the magnet, electrical conductivity, and heat generated (by gradient coils) are a few of the variables that contribute to the imperfection of the field homogeneity and magnet inefficiency. In addition, eddy currents are often generated from the support material, RF coils, and at times even the mere presence of the patient within the magnet. These eddy currents in turn can be corrected by shimming or shielding via smaller coils that have opposite flowing electric current (11,12).
Unlike conventional X-rays, MRI is not a transmission technique but rather a technology whereby signal is generated from proton spins (primarily in cellular water and lipid and not DNA, protein, or bone) that have been perturbed from equilibrium. The voltage generated is secondary to the spin magnetization that is proportional to the static magnetic field strength and spin density. The tissue contrast thus generated results from small differences in tissue water concentration under nonequilibrium conditions. The term spin relaxation consists of both the time necessary for the spins to return to the equilibrium state and the transfer of energy among themselves and the surrounding environment.
X.Y.Z.Bo shield windings
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