Scanner Hardware

From Physics to Plumbing: The Machine That Makes the Signal

Every contrast mechanism, every pulse sequence and every reconstruction trick ultimately rides on three pieces of hardware: a main magnet that creates a strong, exquisitely uniform static field $B_0$; a set of three orthogonal gradient coils that impose controlled spatial variation on that field; and the radiofrequency (RF) chain that excites spins and detects their tiny induced signal. Understanding the engineering limits of these subsystems is not academic. The magnet's homogeneity dictates whether fat suppression works, the gradient slew rate sets your minimum echo spacing, and the receive array geometry determines how much you can accelerate with parallel imaging. This lesson treats the scanner as a quantitative system and connects each component to the numbers you see in protocols and on the console.

The Superconducting Main Magnet

Clinical MRI overwhelmingly uses superconducting solenoids wound from niobium-titanium (NbTi) wire embedded in a copper matrix. Below its critical temperature of about 9.2 K, NbTi carries current with essentially zero resistance, so the windings are cooled by immersion in liquid helium at 4.2 K. Once the magnet is energized ('ramped up'), the leads are joined by a superconducting persistent-mode switch and the power supply is disconnected. Current then circulates for years with a drift typically below 0.1 ppm per hour, which is why an MRI magnet is always on, even during a power outage.

The field a solenoid produces is set by the number of ampere-turns per unit length. For an idealized long solenoid the on-axis field is given below; real clinical magnets use carefully spaced discrete coil bundles to extend uniformity over a usable imaging volume.

Clinical field strengths cluster at 1.5 T and 3 T, with 7 T systems now approved for some neuro and musculoskeletal use. Higher $B_0$ raises the equilibrium magnetization (and thus SNR) approximately linearly, which is the engine behind the drive to higher fields. The trade-offs are cost, increased susceptibility artifact, faster $T_1$ lengthening, higher RF power deposition, and the $B_1$ inhomogeneity problems discussed below.

Homogeneity and Shimming

Spatial uniformity of $B_0$ is quoted in parts per million (ppm) over a defined diameter spherical volume (DSV). A typical specification is on the order of less than 1 ppm peak-to-peak over a 40 cm DSV. The reason such extreme uniformity matters is the Larmor relation: any field offset $\Delta B_0$ translates directly into a frequency offset $\Delta f = \gamma \Delta B_0 / 2\pi$, which corrupts spatial encoding and, critically, breaks frequency-selective fat saturation.

Two complementary techniques restore uniformity. Passive shimming places small ferromagnetic steel plates in trays inside the bore during installation to cancel large static inhomogeneities. Active shimming drives current through dedicated shim coils to generate corrective spherical-harmonic field terms; zeroth- and first-order shims can be adjusted per patient because the body itself distorts the field. The console performs an automatic shim over the prescribed volume before most sequences.

The Fringe Field

The magnetic field does not stop at the bore opening. The external fringe field extends meters into the room and is the basis for the 5 gauss (0.5 mT) line, the boundary inside which ferromagnetic projectile risk and pacemaker interference become significant. Modern magnets use active shielding (additional outer superconducting windings carrying opposing current) to collapse the 5 G line to a compact envelope, often within the magnet housing itself, which is why a 3 T scanner can sit in a modestly sized room.

Gradient Coils

Spatial encoding requires that the field vary with position. Three independent gradient coils superimpose linear field variations along x, y and z onto $B_0$. By convention the z gradient (along the bore) uses a Maxwell pair of opposed coils, while the transverse x and y gradients use saddle-shaped (Golay) windings. Each coil adds or subtracts a small, position-dependent field so the z-component of the total field becomes:

Two performance numbers dominate. Gradient amplitude $G_{max}$, typically 30 to 80 mT/m clinically, sets the maximum spatial resolution and the minimum slice thickness and field of view. Slew rate is how fast the gradient can switch, defined as the amplitude divided by the rise time:

Eddy Currents, Shielding and Duty Cycle

Rapidly switched gradients induce eddy currents in nearby conducting structures (the cryostat, RF shield), creating unwanted residual fields that distort images and shift phase. Actively shielded gradient coils add a second outer winding that cancels the field outside the imaging volume, dramatically reducing eddy currents; residual effects are corrected by pre-emphasis, a deliberate overshoot in the drive waveform. Because the coils dissipate kilowatts of resistive heat, sustained high-amplitude switching is limited by the duty cycle and gradient cooling.

Acoustic Noise and dB/dt Limits

The loud knocking of an MRI scan is pure physics. A current-carrying conductor in a magnetic field feels a Lorentz force $\mathbf{F} = I\,\mathbf{L}\times\mathbf{B}$. As the gradient current switches thousands of times per second, the coil windings are violently pushed against their mountings, radiating acoustic noise that can exceed 110 dB, mandating hearing protection. The same rapid field change is constrained physiologically: a time-varying field induces an electric field in tissue (Faraday's law), and above a threshold this causes peripheral nerve stimulation (PNS). Regulatory limits therefore cap the rate of change of field, dB/dt, which directly bounds usable slew rate.

The RF Subsystem

The RF chain both excites spins and listens for their signal. Transmit is usually performed by the large body coil built into the bore, most often a birdcage design that produces a uniform circularly polarized $B_1$ field across a large volume. Uniform $B_1$ matters because flip angle is proportional to the integral of $B_1$ over the pulse; spatial variation in $B_1$ produces spatially varying flip angle and therefore non-uniform contrast and signal.

Receive Phased Arrays and Parallel Imaging

Signal is detected by receive coils placed close to the anatomy. SNR improves the closer and smaller the coil, but a small coil sees only a small region. The solution is the phased array: many small overlapping elements, each with its own low-noise preamplifier and receive channel, combined to give both the high SNR of small coils and the coverage of a large one. Crucially, because each element has a distinct, known spatial sensitivity, the array enables parallel imaging (SENSE, GRAPPA), where undersampled k-space is reconstructed using coil sensitivity information to shorten scan time.

Parallel imaging accelerates by an acceleration factor R at a cost in SNR captured by the geometry factor:

The T/R Switch and Quadrature Detection

The transmitted RF pulse is on the order of kilowatts; the received signal is in the microvolt range. A transmit/receive (T/R) switch protects the sensitive receive preamplifiers by isolating them during transmit and connecting them during readout. The induced signal is then demodulated by quadrature detection: the signal is mixed against both a cosine (in-phase, I) and a sine (quadrature, Q) reference at the Larmor frequency, recovering both magnitude and phase. Quadrature detection improves SNR by a factor of $\sqrt{2}$ over single-channel detection and lets the system distinguish positive from negative frequency offsets, which is essential for correct spatial encoding.

Shielding: Keeping the Room Quiet and the Field In

Two distinct shielding problems must be solved. RF shielding uses a continuous conductive enclosure, the Faraday cage, lining the scan room (copper or aluminum sheet, with waveguide-filtered penetrations and RF-sealed doors and windows). It keeps external RF, especially from FM and broadcast sources near the 64 to 128 MHz Larmor band, from contaminating the receive chain, which would otherwise appear as the classic zipper artifact. Separately, magnetic shielding (the active shield windings discussed earlier, sometimes augmented by iron in the walls) confines the static fringe field. A broken or improperly sealed Faraday cage, for example a door left ajar, is a common real-world cause of RF interference artifacts.

Locating the Subsystems in the Bore

These subsystems are arranged as concentric shells around the patient. From the outside in: the cryostat and main magnet windings, then the actively shielded gradient coils, then the integrated transmit body coil and RF shield, and finally the patient with any receive array placed directly on the anatomy. Use the interactive model below to rotate the scanner, peel back each layer, and see how the magnet, gradients and RF coils nest together around isocenter.

Notice in the model how the gradient coils sit between the magnet and the bore: their proximity to the conductive cryostat is exactly why eddy currents and active gradient shielding matter, and their rigid mounting is what transmits the Lorentz-force vibration you hear as acoustic noise. The body transmit coil lies just inside the gradients and the RF shield, while the receive array hugs the anatomy at isocenter where homogeneity and SNR are best.

The Quench

A quench is the sudden, often catastrophic loss of superconductivity. If any part of the winding warms above its critical temperature, that segment becomes resistive; the persistent current then dissipates as ohmic heat, which warms adjacent windings and propagates the normal zone. The enormous stored magnetic energy (megajoules) is converted to heat in seconds, rapidly boiling off the liquid helium. Because one liter of liquid helium expands to roughly 750 liters of gas, a quench produces an explosive volume of cold gas.

| Subsystem | Key parameter | Representative value | What it limits | | --- | --- | --- | --- | | Main magnet | Field strength B0 | 1.5, 3, or 7 T | SNR, chemical shift, susceptibility | | Main magnet | Homogeneity | less than 1 ppm over 40 cm DSV | Fat suppression, encoding fidelity | | Gradients | Max amplitude | 30 to 80 mT per m | Spatial resolution, min FOV | | Gradients | Slew rate | 100 to 200 T per m per s | Echo spacing, EPI and DWI speed | | RF transmit | B1 uniformity | degrades at 3 T and above | Flip-angle and contrast uniformity | | RF receive | Channel count | 8 to 64 channels | Acceleration factor R, g-factor |