Magnetic resonance imaging (MRI) is amedical imaging technique mostly used inradiology andnuclear medicine in order to investigate the anatomy and physiology of the body, and to detect pathologies includingtumors,inflammation, neurological conditions such asstroke, disorders of muscles and joints, and abnormalities in the heart and blood vessels among other things.Contrast agents may be injectedintravenously or into a joint to enhance the image and facilitate diagnosis. UnlikeCT andX-ray, MRI uses noionizing radiation and is, therefore, a safe procedure suitable for diagnosis in children and repeated runs. Patients with specific non-ferromagnetic metal implants,cochlear implants, and cardiac pacemakers nowadays may also have an MRI in spite of effects of the strong magnetic fields. This does not apply on older devices, and details for medical professionals are provided by the device's manufacturer.

Certainatomic nuclei are able to absorb and emitradio frequency energy when placed in an externalmagnetic field. In clinical and research MRI,hydrogen atoms are most often used to generate a detectable radio-frequency signal that is received by antennas close to the anatomy being examined. Hydrogen atoms are naturally abundant in people and other biological organisms, particularly inwater andfat. For this reason, most MRI scans essentially map the location of water and fat in the body. Pulses of radio waves excite thenuclear spin energy transition, and magnetic field gradients localize the signal in space. By varying the parameters of thepulse sequence, different contrasts may be generated between tissues based on therelaxation properties of the hydrogen atoms therein.

When inside the magnetic field (B₀) of the scanner, themagnetic moments of the protons align to be either parallel or anti-parallel to the direction of the field. While each individual proton can only have one of two alignments, the collection of protons appear to behave as though they can have any alignment. Most protons align parallel toB₀ as this is a lower energy state. Aradio frequency pulse is then applied, which can excite protons from parallel to anti-parallel alignment, only the latter are relevant to the rest of the discussion. In response to the force bringing them back to their equilibrium orientation, the protons undergo a rotating motion (precession), much like a spun wheel under the effect of gravity. The protons will return to the low energy state by the process ofspin-lattice relaxation. This appears as amagnetic flux, which yields a changing voltage in the receiver coils to give a signal. The frequency at which a proton or group of protons in avoxel resonates depends on the strength of the local magnetic field around the proton or group of protons, a stronger field corresponds to a larger energy difference and higher frequency photons. By applying additional magnetic fields (gradients) that vary linearly over space, specific slices to be imaged can be selected, and an image is obtained by taking the 2-DFourier transform of the spatial frequencies of the signal (k-space). Due to the magneticLorentz force fromB₀ on the current flowing in the gradient coils, the gradient coils will try to move producing loud knocking sounds, for which patients require hearing protection.

The MRI scanner was developed from 1975 to 1977 at theUniversity of Nottingham by ProfRaymond Andrew FRS FRSE following from his research intonuclear magnetic resonance. The full body scanner was created in 1978.^[1]

Subatomic particles have thequantum mechanical property ofspin.^[2] Certain nuclei such as¹H (protons),²H,³He,²³Na or³¹P, have a non–zero spin and therefore amagnetic moment. In the case of the so-calledspin-1⁄2nuclei, such as¹H, there are twospin states, sometimes referred to asup anddown. Nuclei such as¹²C have no unpairedneutrons or protons, and no net spin; however, theisotope¹³C does.

When these spins are placed in a strong externalmagnetic field theyprecess around an axis along the direction of the field. Protons align in two energyeigenstates (theZeeman effect): one low-energy and one high-energy, which are separated by a very small splitting energy.

Quantum mechanics is required to accurately model the behaviour of a single proton. However,classical mechanics can be used to describe the behaviour of an ensemble of protons adequately. As with other spin${\displaystyle 1/2}$ particles, whenever the spin of a single proton is measured it can only have one of two results commonly calledparallel and anti-parallel. When we discuss the state of a proton or protons we are referring to thewave function of that proton which is a linear combination of the parallel and anti-parallel states.^[3]

In the presence of the magnetic field,B₀, the protons will appear to precess at theLarmor frequency determined by the particle's gyro-magnetic ratio and thestrength of the field. The static fields used most commonly in MRI cause precession which corresponds to aradiofrequency (RF)photon.^{[citation needed]}

The net longitudinal magnetization inthermodynamic equilibrium is due to a tiny excess of protons in the lower energy state. This gives a net polarization that is parallel to the external field. Application of an RF pulse can tip this net polarization vector sideways (with, i.e., a so-called 90° pulse), or even reverse it (with a so-called 180° pulse). The protons will come intophase with the RF pulse and therefore each other.^{[citation needed]}

The recovery of longitudinal magnetization is called longitudinal orT₁relaxation and occursexponentially with a time constantT₁. The loss of phase coherence in the transverse plane is called transverse orT₂ relaxation.T₁ is thus associated with theenthalpy of the spin system, or the number of nuclei with parallel versus anti-parallel spin.T₂ on the other hand is associated with theentropy of the system, or the number of nuclei in phase.

When the radio frequency pulse is turned off, the transverse vector component produces an oscillating magnetic field which induces a small current in the receiver coil. This signal is called thefree induction decay (FID). In an idealizednuclear magnetic resonance experiment, the FID decays approximately exponentially with a time constantT₂. However, in practical MRI there are small differences in thestatic magnetic field at different spatial locations ("inhomogeneities") that cause theLarmor frequency to vary across the body. This createsdestructive interference, which shortens the FID. The time constant for the observed decay of the FID is called theT^*
₂ relaxation time, and is always shorter thanT₂. At the same time, the longitudinal magnetization starts to recover exponentially with a time constantT₁ which is much larger thanT₂ (see below).

In MRI, the static magnetic field is augmented by afield gradient coil to vary across the scanned region, so that different spatial locations become associated with different precession frequencies. Only those regions where the field is such that the precession frequencies match the RF frequency will experience excitation. Usually, these field gradients are modulated to sweep across the region to be scanned, and it is the almostinfinite variety of RF and gradient pulse sequences that gives MRI its versatility. Change of field gradient spreads the responding FID signal in the frequency domain, but this can be recovered and measured by a refocusing gradient (to create a so-called "gradient echo"), or by a radio frequency pulse (to create a so-called "spin-echo"), or in digital post-processing of the spread signal. The whole process can be repeated when someT₁-relaxation has occurred and thethermal equilibrium of the spins has been more or less restored. Therepetition time (TR) is the time between two successive excitations of the same slice.^[4]

Typically, insoft tissuesT₁ is around one second whileT₂ andT^*
₂ are a few tens of milliseconds. However, these values can vary widely between different tissues, as well as between different external magnetic fields. This behavior is one factor giving MRI its tremendous soft tissue contrast.

MRI contrast agents, such as those containingGadolinium(III) work by altering (shortening) the relaxation parameters, especiallyT₁.

A number of schemes have been devised for combining field gradients and radio frequency excitation to create an image:

• 2D or3D reconstruction fromprojections, such as incomputed tomography.
• Building the image point-by-point or line-by-line.
• Gradients in the RF field rather than the static field.

Although each of these schemes is occasionally used in specialist applications, the majority of MR Images today are created either by the two-dimensionalFourier transform (2DFT) technique with slice selection, or by the three-dimensional Fourier transform (3DFT) technique. Another name for 2DFT is spin-warp. What follows here is a description of the 2DFT technique with slice selection.

The 3DFT technique is rather similar except that there is no slice selection and phase-encoding is performed in two separate directions.

Another scheme which is sometimes used, especially inbrain scanning or where images are needed very rapidly, is called echo-planar imaging (EPI):^[5] In this case, each RF excitation is followed by a train of gradient echoes with different spatial encoding. Multiplexed-EPI is even faster, e.g., for whole brainfunctional MRI (fMRI) ordiffusion MRI.^[6]

Imagecontrast is created by differences in the strength of the NMR signal recovered from different locations within the sample. This depends upon the relative density of excited nuclei (usuallywater protons), on differences inrelaxation times (T₁,T₂, andT^*
₂) of those nuclei after the pulse sequence, and often on other parameters discussed underspecialized MR scans. Contrast in most MR images is actually a mixture of all these effects, but careful design of the imaging pulse sequence allows one contrast mechanism to be emphasized while the others are minimized. The ability to choose different contrast mechanisms gives MRI tremendous flexibility. In the brain,T₁-weighting causes the nerve connections ofwhite matter to appear white, and the congregations ofneurons ofgray matter to appear gray, whilecerebrospinal fluid (CSF) appears dark. The contrast of white matter, gray matter and cerebrospinal fluid is reversed usingT₂ orT^*
₂ imaging, whereas proton-density-weighted imaging provides little contrast in healthy subjects. Additionally, functional parameters such ascerebral blood flow (CBF), cerebral blood volume (CBV) orblood oxygenation can affectT₁,T₂, andT^*
₂ and so can be encoded with suitable pulse sequences.

In some situations it is not possible to generate enough image contrast to adequately show theanatomy orpathology of interest by adjusting the imaging parameters alone, in which case acontrast agent may be administered. This can be as simple aswater, taken orally, for imaging the stomach and small bowel. However, mostcontrast agents used in MRI are selected for their specific magnetic properties. Most commonly, aparamagnetic contrast agent (usually agadolinium compound^[7][8]) is given. Gadolinium-enhanced tissues and fluids appear extremely bright onT₁-weighted images. This provides high sensitivity for detection of vascular tissues (e.g., tumors) and permits assessment of brain perfusion (e.g., in stroke).There have been concerns raised recently regarding the toxicity of gadolinium-based contrast agents and their impact on persons with impaired kidney function. (SeeSafety/Contrast agents below.)

More recently,superparamagnetic contrast agents, e.g.,iron oxidenanoparticles,^[9][10] have become available. These agents appear very dark onT^*
₂-weighted images and may be used for liver imaging, as normalliver tissue retains the agent, but abnormal areas (e.g., scars, tumors) do not. They can also be taken orally, to improve visualization of thegastrointestinal tract, and to prevent water in the gastrointestinal tract from obscuring other organs (e.g., thepancreas).Diamagnetic agents such asbarium sulfate have also been studied for potential use in thegastrointestinal tract, but are less frequently used.

In 1983, Ljunggren^[11] and Twieg^[12] independently introduced thek-space formalism, a technique that proved invaluable in unifying different MR imaging techniques. They showed that the demodulated MR signalS(t) generated by the interaction between an ensemble of freely precessing nuclear spins in the presence of a linear magnetic field gradientG and a receiver-coil equals the Fourier transform of the effective spin density,${\displaystyle \rho (\overrightarrow{x})}$. Fundamentally, the signal is derived fromFaraday's law of induction:

${\displaystyle S(t)={\tilde{\rho }}_{\mathrm{e}\mathrm{f}\mathrm{f}}\left(\overrightarrow{k}(t)\right)\equiv {\int }_{-\mathrm{\infty }}^{\mathrm{\infty }}\mathrm{d}\overrightarrow{x}\rho (\overrightarrow{x})\cdot e^{2\pi i\overrightarrow{k}(t)\cdot \overrightarrow{x}}}$

where:

${\displaystyle \overrightarrow{k}(t)\equiv {\int }_0^t\overrightarrow{G}(\tau )\mathrm{d}\tau }$

In other words, as time progresses the signal traces out a trajectory ink-space with thevelocity vector of the trajectory proportional to the vector of the applied magnetic field gradient.By the termeffective spin density we mean the true spin density${\displaystyle \rho (\overrightarrow{x})}$ corrected for the effects ofT₁ preparation,T₂ decay, dephasing due to field inhomogeneity, flow, diffusion, etc. and any other phenomena that affect that amount of transverse magnetization available to induce signal in the RF probe or its phase with respect to the receiving coil' s electromagnetic field.

From the basick-space formula, it follows immediately that we reconstruct an image${\displaystyle I(\overrightarrow{x})}$ by taking theinverse Fourier transform of the sampled data, viz.

${\displaystyle I\left(\overrightarrow{x}\right)={\int }_{-\mathrm{\infty }}^{\mathrm{\infty }}\mathrm{d}\overrightarrow{k}S\left(\overrightarrow{k}(t)\right)\cdot e^{-2\pi i\overrightarrow{k}(t)\cdot \overrightarrow{x}}}$

Using thek-space formalism, a number of seemingly complex ideas became simple. For example, it becomes very easy (forphysicists, in particular) to understand the role of phase encoding (the so-called spin-warp method). In a standard spin echo or gradient echo scan, where the readout (or view) gradient is constant (e.g.,G), a single line ofk-space is scanned per RF excitation. When the phase encoding gradient is zero, the line scanned is thek_x axis. When a non-zero phase-encoding pulse is added in between the RF excitation and the commencement of the readout gradient, this line moves up or down ink-space, i.e., we scan the linek_y = constant.

Thek-space formalism also makes it very easy to compare different scanning techniques. In single-shotEPI, all ofk-space is scanned in a single shot, following either a sinusoidal or zig-zag trajectory. Since alternating lines ofk-space are scanned in opposite directions, this must be taken into account in the reconstruction. Multi-shot EPI and fast spin echo techniques acquire only part ofk-space per excitation. In each shot, a different interleaved segment is acquired, and the shots are repeated untilk-space is sufficiently well-covered. Since the data at the center ofk-space represent lower spatial frequencies than the data at the edges ofk-space, theT_E value for the center ofk-space determines the image'sT₂ contrast.

The importance of the center ofk-space in determining image contrast can be exploited in more advanced imaging techniques. One such technique is spiral acquisition—arotating magnetic field gradient is applied, causing the trajectory ink-space to spiral out from the center to the edge. Due toT₂ andT^*
₂ decay the signal is greatest at the start of the acquisition, hence acquiring the center ofk-space first improvescontrast to noise ratio (CNR) when compared to conventional zig-zag acquisitions, especially in the presence of rapid movement.

Since${\displaystyle \overrightarrow{x}}$ and${\displaystyle \overrightarrow{k}}$ are conjugate variables (with respect to the Fourier transform) we can use theNyquist theorem to show that a step ink-space determines the field of view of the image (maximum frequency that is correctly sampled) and the maximum value of k sampled determines the resolution; i.e.,

${\displaystyle \mathrm{F}\mathrm{O}\mathrm{V}\propto \frac{1}{\mathrm{\Delta }k}\mathrm{R}\mathrm{e}\mathrm{s}\mathrm{o}\mathrm{l}\mathrm{u}\mathrm{t}\mathrm{i}\mathrm{o}\mathrm{n}\propto |k_{max}|.}$

(These relationships apply to each axis independently.)

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In thetiming diagram, the horizontal axis represents time. The vertical axis represents: (top row) amplitude of radio frequency pulses; (middle rows) amplitudes of the three orthogonal magnetic field gradient pulses; and (bottom row) receiver analog-to-digital converter (ADC). Radio frequencies are transmitted at the Larmor frequency of the nuclide to be imaged. For example, for¹H in a magnetic field of 1 T, a frequency of 42.5781 MHz would be employed. The three field gradients are labeledG_X (typically corresponding to a patient's left-to-right direction and colored red in diagram),G_Y (typically corresponding to a patient's front-to-back direction and colored green in diagram), andG_Z (typically corresponding to a patient's head-to-toe direction and colored blue in diagram). Where negative-going gradient pulses are shown, they represent reversal of the gradient direction, i.e., right-to-left, back-to-front or toe-to-head. For human scanning, gradient strengths of 1–100 mT/m are employed: Higher gradient strengths permit better resolution and faster imaging. The pulse sequence shown here would produce a transverse (axial) image.

The first part of the pulse sequence, SS, achieves "slice selection". A shaped pulse (shown here with asinc modulation) causes a 90°nutation of longitudinal nuclear magnetization within a slab, or slice, creating transverse magnetization. The second part of the pulse sequence, PE, imparts a phase shift upon the slice-selected nuclear magnetization, varying with its location in the Y direction. The third part of the pulse sequence, another slice selection (of the same slice) uses another shaped pulse to cause a 180° rotation of transverse nuclear magnetization within the slice. This transverse magnetisation refocuses to form a spin echo at a timeT_E. During the spin echo, a frequency-encoding (FE) or readout gradient is applied, making the resonant frequency of the nuclear magnetization vary with its location in the X direction. The signal is sampledn_FE times by the ADC during this period, as represented by the vertical lines. Typicallyn_FE of between 128 and 512 samples are taken.

The longitudinal magnetisation is then allowed to recover somewhat and after a timeT_R the whole sequence is repeatedn_PE times, but with the phase-encoding gradient incremented (indicated by the horizontal hatching in the green gradient block). Typicallyn_PE of between 128 and 512 repetitions are made.

The negative-going lobes inG_X andG_Z are imposed to ensure that, at timeT_E (the spin echo maximum), phase only encodes spatial location in the Y direction.

TypicallyT_E is between 5 ms and 100 ms, whileT_R is between 100 ms and 2000 ms.

After the two-dimensional matrix (typical dimension between 128 × 128 and 512 × 512) has been acquired, producing the so-calledk-space data, a two-dimensional inverse Fourier transform is performed to provide the familiar MR image. Either the magnitude or phase of the Fourier transform can be taken, the former being far more common.

edit
This table does not includeuncommon and experimental sequences.

Group	Sequence	Abbr.	Physics	Main clinical distinctions	Example
Spin echo	T1 weighted	T1	Measuringspin–lattice relaxation by using a shortrepetition time (TR) andecho time (TE).	Lower signal for more water content,[13] as inedema,tumor,infarction,inflammation,infection, hyperacute or chronichemorrhage.[14] High signal forfat[13][14] High signal forparamagnetic substances, such asMRI contrast agents[14] Standard foundation and comparison for other sequences
	T2 weighted	T2	Measuringspin–spin relaxation by using long TR and TE times	Higher signal for more water content[13] Low signal for fat in standard Spine Echo (SE),[13] though not with Fast Spin Echo/Turbo Spin Echo (FSE/TSE). FSE/TSE is thestandard of care in modern medicine because it is faster. With FSE/TSE, fat has high signal due to disruption ofhyperfineJ-coupling between adjacent fatprotons.[15] Low signal forparamagnetic substances[14] Standard foundation and comparison for other sequences
	Proton density weighted	PD	LongTR (to reduce T1) and shortTE (to minimize T2).[16]	Joint disease and injury.[17] High signal frommeniscus tears.[18] (pictured)
Gradient echo (GRE)	Steady-state free precession	SSFP	Maintenance of a steady, residual transverse magnetisation over successive cycles.[19]	Creation ofcardiac MRI videos (pictured).[19]
	Effective T2 or "T2-star"	T2*	Spoiled gradient recalled echo (GRE) with a long echo time and small flip angle[20]	Low signal fromhemosiderin deposits (pictured) and hemorrhages.[20]
	Susceptibility-weighted	SWI	Spoiled gradient recalled echo (GRE), fully flow compensated, long echo time, combines phase image with magnitude image[21]	Detecting small amounts of hemorrhage (diffuse axonal injury pictured) or calcium.[21]
Inversion recovery	Short tau inversion recovery	STIR	Fat suppression by setting aninversion time where the signal of fat is zero.[22]	High signal inedema, such as in more severestress fracture.[23]Shin splints pictured:
	Fluid-attenuated inversion recovery	FLAIR	Fluid suppression by setting an inversion time that nulls fluids	High signal inlacunar infarction,multiple sclerosis (MS) plaques,subarachnoid haemorrhage andmeningitis (pictured).[24]
	Double inversion recovery	DIR	Simultaneous suppression ofcerebrospinal fluid andwhite matter by two inversion times.[25]	High signal ofmultiple sclerosis plaques (pictured).[25]
Diffusion weighted (DWI)	Conventional	DWI	Measure ofBrownian motion of water molecules.[26]	High signal within minutes ofcerebral infarction (pictured).[27]
	Apparent diffusion coefficient	ADC	Reduced T2 weighting by taking multiple conventional DWI images with different DWI weighting, and the change corresponds to diffusion.[28]	Low signal minutes aftercerebral infarction (pictured).[29]
	Diffusion tensor	DTI	Mainlytractography (pictured) by an overall greaterBrownian motion of water molecules in the directions of nerve fibers.[30]	Evaluating white matter deformation by tumors[30] Reducedfractional anisotropy may indicatedementia.[31]
Perfusion weighted (PWI)	Dynamic susceptibility contrast	DSC	Measures changes over time in susceptibility-induced signal loss due togadolinium contrast injection.[32]	Provides measurements of blood flow Incerebral infarction, the infarcted core and thepenumbra have decreased perfusion and delayed contrast arrival (pictured).[33]
	Arterial spin labelling	ASL	Magnetic labeling of arterial blood below the imaging slab, which subsequently enters the region of interest.[34] It does not need gadolinium contrast.[35]
	Dynamic contrast enhanced	DCE	Measures changes over time in the shortening of thespin–lattice relaxation (T1) induced by agadolinium contrast bolus.[36]	Faster Gd contrast uptake along with other features is suggestive of malignancy (pictured).[37]
Functional MRI (fMRI)	Blood-oxygen-level dependent imaging	BOLD	Changes inoxygen saturation-dependent magnetism ofhemoglobin reflects tissue activity.[38]	Localizing brain activity from performing an assigned task (e.g. talking, moving fingers) before surgery, also used in research of cognition.[39]
Magnetic resonance angiography (MRA) and venography	Time-of-flight	TOF	Blood entering the imaged area is not yetmagnetically saturated, giving it a much higher signal when using short echo time and flow compensation.	Detection ofaneurysm,stenosis, ordissection[40]
	Phase-contrast magnetic resonance imaging	PC-MRA	Two gradients with equal magnitude, but opposite direction, are used to encode a phase shift, which is proportional to the velocity ofspins.[41]	Detection ofaneurysm,stenosis, ordissection (pictured).[40]	(VIPR)

The major components of anMRI scanner are: the main magnet, which polarizes the sample, the shim coils for correcting inhomogeneities in the main magnetic field, the gradient system which is used to localize the MR signal and the RF system, which excites the sample and detects the resulting NMR signal. The whole system is controlled by one or more computers.

The magnet is the largest and most expensive component of the scanner, and the remainder of the scanner is built around it. The strength of the magnet is measured inteslas (T). Clinical magnets generally have a field strength in the range 0.1–3.0 T, with research systems available up to 9.4 T for human use and 21 T for animal systems.^[42]In the United States, field strengths up to 7 T have been approved by the FDA for clinical use.^[43]

Just as important as the strength of the main magnet is its precision. The straightness of the magnetic lines within the center (or, as it is technically known, the iso-center) of the magnet needs to be near-perfect. This is known as homogeneity. Fluctuations (inhomogeneities in the field strength) within the scan region should be less than three parts per million (3 ppm). Three types of magnets have been used:

• Permanent magnet: Conventional magnets made from ferromagnetic materials (e.g., steel alloys containingrare-earth elements such asneodymium) can be used to provide the static magnetic field. A permanent magnet that is powerful enough to be used in an MRI will be extremely large and bulky; they can weigh over 100 tonnes. Permanent magnet MRIs are very inexpensive to maintain; this cannot be said of the other types of MRI magnets, but there are significant drawbacks to using permanent magnets. They are only capable of achieving weak field strengths compared to other MRI magnets (usually less than 0.4 T) and they are of limited precision and stability. Permanent magnets also present special safety issues; since their magnetic fields cannot be "turned off," ferromagnetic objects are virtually impossible to remove from them once they come into direct contact. Permanent magnets also require special care when they are being brought to their site of installation.
• Resistive electromagnet: Asolenoid wound from copper wire is an alternative to a permanent magnet. An advantage is low initial cost, but field strength and stability are limited. The electromagnet requires considerable electrical energy during operation which can make it expensive to operate. This design is essentially obsolete.
• Superconducting electromagnet: When aniobium-titanium orniobium-tin alloy is cooled byliquid helium to 4 K (−269 °C, −452 °F) it becomes asuperconductor, losing resistance to flow of electric current. An electromagnet constructed with superconductors can have extremely high field strengths, with very high stability. The construction of such magnets is extremely costly, and the cryogenic helium is expensive and difficult to handle. However, despite their cost, helium cooled superconducting magnets are the most common type found in MRI scanners today.

Most superconducting magnets have their coils of superconductive wire immersed in liquid helium, inside a vessel called acryostat. Despite thermal insulation, sometimes including a second cryostat containingliquid nitrogen, ambient heat causes the helium to slowly boil off. Such magnets, therefore, require regular topping-up with liquid helium. Generally acryocooler, also known as a coldhead, is used to recondense some helium vapor back into the liquid helium bath. Several manufacturers now offer 'cryogenless' scanners, where instead of being immersed in liquid helium the magnet wire is cooled directly by a cryocooler.^[44] Alternatively, the magnet may be cooled by carefully placing liquid helium in strategic spots, dramatically reducing the amount of liquid helium used,^[45] or,high temperature superconductors may be used instead.^[46][47]

Magnets are available in a variety of shapes. However, permanent magnets are most frequently C-shaped, and superconducting magnets most frequently cylindrical. C-shaped superconducting magnets and box-shaped permanent magnets have also been used.

Magnetic field strength is an important factor in determining image quality. Higher magnetic fields increasesignal-to-noise ratio, permitting higher resolution or faster scanning. However, higher field strengths require more costly magnets with higher maintenance costs, and have increased safety concerns. A field strength of 1.0–1.5 T is a good compromise between cost and performance for general medical use. However, for certain specialist uses (e.g., brain imaging) higher field strengths are desirable, with some hospitals now using 3.0 T scanners.

When the MR scanner is placed in the hospital or clinic, its main magnetic field is far from being homogeneous enough to be used for scanning. That is why before doing fine tuning of the field using a sample, the magnetic field of the magnet must be measured andshimmed.

After a sample is placed into the scanner, the main magnetic field is distorted bysusceptibility boundaries within that sample, causing signal dropout (regions showing no signal) and spatial distortions in acquired images. For humans or animals the effect is particularly pronounced at air-tissue boundaries such as thesinuses (due toparamagnetic oxygen in air) making, for example, the frontal lobes of the brain difficult to image. To restore field homogeneity a set of shim coils is included in the scanner. These are resistive coils, usually at room temperature, capable of producing field corrections distributed as several orders ofspherical harmonics.^[48]

After placing the sample in the scanner, theB₀ field is 'shimmed' by adjusting currents in the shim coils. Field homogeneity is measured by examining anFID signal in the absence of field gradients. The FID from a poorly shimmed sample will show a complex decay envelope, often with many humps. Shim currents are then adjusted to produce a large amplitude exponentially decaying FID, indicating a homogeneousB₀ field. The process is usually automated.^[49]

Gradient coils are used to spatially encode the positions of protons by varying the magnetic field linearly across the imaging volume. The Larmor frequency will then vary as a function of position in thex,y andz-axes.

Gradient coils are usually resistive electromagnets powered by sophisticatedamplifiers which permit rapid and precise adjustments to their field strength and direction. Typical gradient systems are capable of producing gradients from 20 to 100 mT/m (i.e., in a 1.5 T magnet, when a maximalz-axis gradient is applied, the field strength may be 1.45 T at one end of a 1 m long bore and 1.55 T at the other^[50]). It is the magnetic gradients that determine the plane of imaging—because the orthogonal gradients can be combined freely, any plane can be selected for imaging.

Scan speed is dependent on performance of the gradient system. Stronger gradients allow for faster imaging, or for higher resolution; similarly, gradient systems capable of faster switching can also permit faster scanning. However, gradient performance is limited by safety concerns over nerve stimulation.

Some important characteristics of gradient amplifiers and gradient coils are slew rate and gradient strength. As mentioned earlier, a gradient coil will create an additional, linearly varying magnetic field that adds or subtracts from the main magnetic field. This additional magnetic field will have components in all 3 directions, viz.x,y andz; however, only the component along the magnetic field (usually called thez-axis, hence denotedG_z) is useful for imaging. Along any given axis, the gradient will add to the magnetic field on one side of the zero position and subtract from it on the other side. Since the additional field is a gradient, it has units ofgauss per centimeter or millitesla per meter (mT/m). High performance gradient coils used in MRI are typically capable of producing a gradient magnetic field of approximate 30 mT/m or higher for a 1.5 T MRI. The slew rate of a gradient system is a measure of how quickly the gradients can be ramped on or off. Typical higher performance gradients have a slew rate of up to 100–200 T·m⁻¹·s⁻¹. The slew rate depends both on the gradient coil (it takes more time to ramp up or down a large coil than a small coil) and on the performance of the gradient amplifier (it takes a lot of voltage to overcome the inductance of the coil) and has significant influence on image quality.

Theradio frequency (RF)transmission system consists of an RF synthesizer,power amplifier andtransmitting coil. That coil is usually built into the body of the scanner. The power of the transmitter is variable, but high-end whole-body scanners may have a peak output power of up to 35 kW,^[51] and be capable of sustaining average power of 1 kW. Although theseelectromagnetic fields are in the RF range of tens ofmegahertz (often in theshortwave radio portion of theelectromagnetic spectrum) at powers usually exceeding the highest powers used byamateur radio, there is very little RF interference produced by the MRI machine. The reason for this is that the MRI is not a radio transmitter. The RF frequencyelectromagnetic field produced in the "transmitting coil" is a magneticnear-field with very little associated changingelectric field component (such as all conventional radio wave transmissions have). Thus, the high-powered electromagnetic field produced in the MRI transmitter coil does not produce muchelectromagnetic radiation at its RF frequency, and the power is confined to the coil space and not radiated as "radio waves." Thus, the transmitting coil is a good EMfield transmitter at radio frequency, but a poor EMradiation transmitter at radio frequency.

The receiver consists of the coil, pre-amplifier and signal processing system. The RFelectromagnetic radiation produced by nuclear relaxation inside the subject is true EM radiation (radio waves), and these leave the subject as RF radiation, but they are of such low power as to also not cause appreciable RF interference that can be picked up by nearby radio tuners (in addition, MRI scanners are generally situated in metal mesh lined rooms which act asFaraday cages.)

While it is possible to scan using the integrated coil for RF transmission and MR signal reception, if a small region is being imaged, then better image quality (i.e., higher signal-to-noise ratio) is obtained by using a close-fitting smaller coil. A variety of coils are available which fit closely around parts of the body such as the head, knee, wrist, breast, or internally, e.g., the rectum.

A recent development in MRI technology has been the development of sophisticated multi-element phased array^[52] coils which are capable of acquiring multiple channels of data in parallel. This 'parallel imaging' technique uses unique acquisition schemes that allow for accelerated imaging, by replacing some of the spatial coding originating from the magnetic gradients with the spatial sensitivity of the different coil elements. However, the increased acceleration also reduces the signal-to-noise ratio and can create residual artifacts in the image reconstruction. Two frequently used parallel acquisition and reconstruction schemes are known as SENSE^[53] and GRAPPA.^[54] A detailed review of parallel imaging techniques can be found here:^[55]

• Magnetic resonance imaging

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