RF coils: A practical guide for nonphysicists

Abstract

Radiofrequency (RF) coils are an essential MRI hardware component. They directly impact the spatial and temporal resolution, sensitivity, and uniformity in MRI. Advances in RF hardware have resulted in a variety of designs optimized for specific clinical applications. RF coils are the "antennas" of the MRI system and have two functions: first, to excite the magnetization by broadcasting the RF power (Tx-Coil) and second to receive the signal from the excited spins (Rx-Coil). Transmit RF Coils emit magnetic field pulses ( B1+) to rotate the net magnetization away from its alignment with the main magnetic field (B₀ ), resulting in a transverse precessing magnetization. Due to the precession around the static main magnetic field, the magnetic flux in the receive RF Coil ( B1-) changes, which generates a current I. This signal is "picked-up" by an antenna and preamplified, usually mixed down to a lower frequency, digitized, and processed by a computer to finally reconstruct an image or a spectrum. Transmit and receive functionality can be combined in one RF Coil (Tx/Rx Coils). This review looks at the fundamental principles of an MRI RF coil from the perspective of clinicians and MR technicians and summarizes the current advances and developments in technology.

Level of evidence: 1 Technical Efficacy: Stage 6.

Keywords: RF coil; arrays; clinical application; coil array design; novel RF coils; signal-to-noise ratio.

Figure 1

The basic components of any MRI system: The main magnet produces the B₀ field, necessary to align the spins and achieve equilibrium. Gradient coils enable image encoding in the x, y, and z direction (ie, the frequency, phase, and slice‐encoding directions). The RF coil is the part of the MRI system that excites the aligned spins and receives an RF signal back from the sample. All the components are controlled and interfaced with the user via a console.

Figure 2

(B) To excite the spins, the transmit coil receives a signal from the controller/computer via a digital‐to‐analog converter (DAC). (A) The receive coil takes up the response from the excitation, amplifies, and digitizes (ADC) it. (C) The schematic of a transmit‐receive RF coil: the T/R‐switch controls the transmission and reception of RF signals.

Figure 3

Noise in the image appears as a grainy random pattern similar to snow on a TV screen. It represents statistical fluctuations in signal intensity that do not contribute to image information, and have two basic sources: Brownian motion of molecules in the human body and electronic noise of the receiver, which both add up. If the signal from a slice is too weak, it may be “washed over” by noise (Courtesy of Ref. 93).

Figure 4

The rotating magnetization induces a voltage at the terminals of the loop, which can cause a current in a single loop or coil. Linearly polarized RF coils detect the rotating magnetization (MR signal) along a single direction. Quadrature (circularly polarized) coil arrangements detect the MR signal in orthogonal directions.

Figure 5

The optimal combination of signals from an array coil requires knowledge of the spatial variation of each element's sensitivity. A pixel at point 1 of a 3‐element (channel) array, has contributions from Coil 3 and 2, but very little from Coil 1. Coil 1 will contribute primarily noise leading to a suboptimal combination of the pixel at point 1. If the contribution is weighted by the sensitivity at this location, the contribution of Coil 1 will be weighted close to zero, reducing the combined noise and hence increasing the SNR.

Figure 6

Different types of array coil combinations: (A) parallel array, (B) decoupled array, (C) phased array, (D) representation of various loop‐shaped surface coil designs.

Figure 7

SNR along a 1D profile as a function of channel count for a few representative arrays (courtesy of Ref. 94).

Figure 8

A typical circuit schematic for a receive‐only coil element. In this case, the loop comprises a conductive wire with two tuning capacitors (C_Tune and C) to fine‐tune the coil frequency, a detuning trap (L and pin diode D) to actively deactivate/detune (turn off) the loop while RF excitation by the transmit coil, and a matching capacitor (C_Match) to transform the element impedance to 50 Ω of the preamplifier, which finally amplifies the MR signal. The pin diode D is powered by a certain DC bias. The passive detuning circuit, consisting of an adjustable capacitor CV and crossed diodes and serves, like the RF fuse F, as a second stage of safety to the Rx loop during transmission by the Tx coil, and as well for the patient. This example coil thus has three redundant safety features to prevent interactions between the Tx coil and the receive loop that could cause heating or other safety concerns.

Figure 9

The SNR is influenced by several factors: The effective temperature factor T or T_eff at room temperature which includes the losses of all components; the performance of the preamplifier to amplify the EMF, which is described through the noise figure; the coil quality factor Q, which is the ratio of stored energy to dissipated energy; the filling factor η, which is the ratio of magnetic field energy stored inside the sample volume versus the total magnetic energy stored by the loop; the geometry or g‐factor, which is simply the ratio in noise between accelerated and unaccelerated imaging important in parallel imaging; and the volume of interest and the field strength B₀.

Figure 10

Table a,b display the optimal coil radii (mm) of a surface coil loop for target depths of interest, determined with full‐wave simulations and SNR loss of the lossy coils of radii r _R measured and calculated (courtesy of Ref. 90). The graph below table b shows the optimal coil radius r _R (mm), including coil losses, as a function of field strength (T) for various target depths, as determined by full‐wave numerical method of moments (MoM). The quasistatic optimum coil radii r ₀, without coil losses, are indicated by horizontal bars in the center of the plot (courtesy of Ref. 55).

Figure 11

The behavior of two loops with resonance frequency f ₁ = 1/[2π√(L ₁ C ₁)] and f ₂ = 1/[2π√(L ₂ C ₂)] can be described with the impedance curve. The mutual inductance increases as the distance between the coils decrease, meaning that moving the loops further the two peaks will fuse to one. If the coupling between two coils gets stronger, they over‐couple and the resonance frequency splits into two current modes: a co‐rotating (+) and a counter‐rotating (–) mode. The optimal distance for geometrical decoupling depends on the loop dimensions. The ideal distances for maximal passive/inductive decoupling depend on the loop shape and size (a), (b).