Introduction

In MRI a very large magnet generates a strong static magnetic field (B₀) (>10,000 x B_earth) to align the nuclear magnetic moments in the body, which are otherwise in random orientation. This creates a bulk nuclear magnetisation (M₀) proportional to the static field and nuclear spin density. The precession of M₀ around the static field is determined by the gyromagnetic ratio (γ) which is a measure of the magnetic moment-to-angular momentum ratio of the nucleus. To detect a particular nucleus, a high-power RF pulse excitation with angular frequency equal γB₀ (Larmor frequency) is applied by a Tx coil with axis (in x-y plane) perpendicular to B₀ (z-axis). This generates a small magnetic field B₁ that rotates at the same speed as the spins around B₀ perturbing M₀ from its equilibrium state (precessing around B₀) to precesses, only temporarily, around B₁ field. After this pulsed excitation, M₀ returns back to its equilibrium state aligned with B₀. This free precession back to equilibrium induces a time variant voltage (ε ) (MR signal) on a tuned receive (Rx) coil with axis also perpendicular to B₀ (could be same Tx coil). The amplitude and transient response (oscillating at the Larmor Frequency) of the magnetization depends on tissue relaxation constants (T1, T2, T2*) and nuclei spin density (ρ). These parameters can differentiate healthy from diseased tissue as well as to provide insights about biological processes like blood perfusion. Since the magnetization precession frequency is proportional to the static field, the signal generated can be spatially encoded by well-defined varying magnetic field gradients (𝐺⃗). Field gradients in three directions (x, y, z) can be generated by three sets of coil arrangements that are concentrically located inside the magnet coil. A linear relationship between precession frequency and a spatial coordinate makes possible the generation of an image by a simple Fourier transform. All events involved in the generation of this image (RF excitation, gradient encoding and signal acquisition) are controlled by a pulse sequence executed in the scanner console which communicates with all system hardware.

Superconducting Magnets

Most MRI magnets today are built with several kilometers of superconducting wire, mostly niobium-titanium (Nb-Ti) and niobium-tin (Nb3-Sn). Because in MRI we want image contrast to be associated to tissue properties, we need extremely high magnetic field homogeneity ( ΔB₀ < 1 ppm). This stringent homogeneity specification differentiates these magnets from those with similar strength used in particle accelerators and nuclear fusion. In search of higher spatial resolution, higher signal to noise ratio (SNR) and higher signal contrast (CNR), stronger magnets are being developed for both human and animal studies. Today we have field strengths for human MRI systems up to 11.7 T and for NMR systems up to 32 T. Up to 21 T superconducting magnets are built with Nb₃-Sn or Nb-Ti or a combination of both. These low temperature superconductors (LTS) need to be kept at temperatures as low as 2 °K. This requires a cryogen system with large amount of liquid helium (He), and liquid nitrogen (N). Today, thanks to advanced engineering of the cryogenic system, we can find almost helium-free (few liters) commercial low field systems. This is important to overcome the decreasing supply and increasing price of He. Pushing MRI to higher field strengths requires new superconductors materials. There is active research on building magnets with high temperature superconductors (HTS), the most promising so far being magnesium diboride (MgB₂). These HTS are kept to temperature close to or above liquid nitrogen (> 77 °K) eliminating the need of liquid He in the cryogenic system. In addition to the magnet coils, an MRI system can have active shielding coils, also built from superconducting wire, located at both ends of the magnet coil to reduce stray fields outside the bore. To correct for unavoidable bulk B₀ inhomogeneities passive shimming can be performed in site by the arrangement of ferromagnetic small parts. To compensate for B₀ inhomogeneities introduced in the MRI experiment, dynamic electrical shimming can be performed through additional wired loops driven independently to generate different field harmonics.

Gradient Coils

MRI scanners have 3 sets of gradient coil arrangements integrated into an annular layer inside the bore of the main magnet winding. Each of these coils generate an independent spatial field change along x, y and z (Gx, Gy, Gz). The field change is linear in the imaging volume. The z-gradient is generated by sets of circular loop pairs coaxially arranged along z-axis and symmetrically spaced from the z=0 plane. The currents on each side of the pairs run in opposite directions such as to increase (parallel field) or decrease (anti-parallel field) the total magnetic field. In most MRI scanner today, the x-and y-gradients are generated by finger-print coils laid on a cylindrical surface concentrically to the magnet coil. Both gradient coil layouts are similar but with axis rotated 90 degrees. As in the z-gradient coil, loops are spatially symmetric driven by anti-symmetric currents such as to generate fields in parallel and antiparallel with B0 across x and y direction. These coils have an additional coil layer that minimizes eddy currents in surrounding conductive surfaces specially the cryostat. Each of these gradient coils is independently driven by a power amplifier. In order to generate enough field change and spatially encode the nuclear spin resonance, gradient coils can be driven by currents as high as 900 A. Gradient encoding only happens in a portion of the imaging cycle, which can be as short as few milliseconds, and these coils are switched on-off at frequencies in the kHz well inside the audio range. Because these large coils have inductances on the order of thousands microhenries, the fast switching of the currents induces voltages on the order of thousands of volts. Gradient hardware in the MRI scanner handles the highest power of all other hardware (in the order of MW). At kHz frequency range amplifier losses can be minimized by ultra-high-power efficient switch-mode amplifiers using Pulse-Width- Modulation (PWM). Different from magnet coils, gradient coils are not built with superconductors so at these high currents power dissipation in the wire can be in the order of several kW. Therefore, water cooling is required. This is usually implemented by chilled water pipes in close contact with gradient coil wires.

RF Hardware

Of all the MRI hardware, the RF coil is the most accessible and customizable hardware to improve image quality for a specific application. In MRI scanners with field strengths up to 3 T there is a RF transmitter integrated in the scanner bore known as body coil. This coil has a cylindrical shape and it is built from multiple rungs distributed along the cylindrical surface and aligned in the direction of the static field (z-axis). Both end-ring segments connecting each rung are capacitively or inductively coupled to the next rung. This “birdcage” structure has 2 ports that are located 90 degrees azimuthally from each other and receive high voltages with same amplitude but with a 90 degrees phase delay. These voltages generate a sinusoidal current distribution along the rungs that creates a time variant magnetic field (B₁), on resonance, rotating with the nuclear magnetization Therefore this Tx hardware can generate a very homogenous excitation of the magnetization in a well-defined volume. The body coil can be used as a Tx only or as a Tx-Rx coil. It has been shown that receiving the MR signal with an array of surface loops, instead of a volume coil, provides higher SNR and image acceleration by skipping acquisition lines and taking advantage of coil spatial information. This approach is known as parallel imaging. Most imaging protocols today use the body coil as a Tx only coil and external Rx arrays with as many as 128 channels. At higher field strengths the shorter excitation wavelength compromises the homogeneity achieved with this coil architecture, and other Tx coil topologies need to be considered. One approach to achieve higher homogeneity of the excitation field is to use an array of Tx coil independently controlled. This method is known as parallel transmission (pTx). Some ultra-high field scanners have already a pTx interface, up to 16 channels, implemented in their control. The technology for these ultra-high field systems continues to evolve to compensate for field homogeneities caused by tissue properties of the body and to automate methods for obtaining high quality images while ensuring patient safety.