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: a solenoid 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 a niobium-titanium or niobium-tin alloy is cooled by liquid helium to 4 k (−269 °c, −452 °f) it becomes a superconductor, 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 a cryostat. despite thermal insulation, sometimes including a second cryostat containing liquid nitrogen, ambient heat causes the helium to slowly boil off. such magnets, therefore, require regular topping-up with liquid helium. generally a cryocooler, 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. alternatively, the magnet may be cooled by carefully placing liquid helium in strategic spots, dramatically reducing the amount of liquid helium used, or, high temperature superconductors may be used instead. 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 increase signal-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. fid signal from a badly shimmed sample has a complex envelope. fid signal from a well shimmed sample, showing a pure exponential decay. shims 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 and shimmed. after a sample is placed into the scanner, the main magnetic field is distorted by susceptibility 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 the sinuses (due to paramagnetic 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 of spherical harmonics. after placing the sample in the scanner, the b0 field is 'shimmed' by adjusting currents in the shim coils. field homogeneity is measured by examining an fid 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 homogeneous b0 field. the process is usually automated. gradients 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 the x, y and z-axes. gradient coils are usually resistive electromagnets powered by sophisticated amplifiers 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 maximal z-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). 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 and z; however, only the component along the magnetic field (usually called the z-axis, hence denoted gz) 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 of gauss 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. radio frequency system the radio frequency (rf) transmission system consists of an rf synthesizer, power amplifier and transmitting 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, and be capable of sustaining average power of 1 kw. although these electromagnetic fields are in the rf range of tens of megahertz (often in the shortwave radio portion of the electromagnetic spectrum) at powers usually exceeding the highest powers used by amateur 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 frequency electromagnetic field produced in the "transmitting coil" is a magnetic near-field with very little associated changing electric 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 much electromagnetic 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 em field transmitter at radio frequency, but a poor em radiation transmitter at radio frequency. the receiver consists of the coil, pre-amplifier and signal processing system. the rf electromagnetic 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 as faraday 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 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 and grappa. a detailed review of parallel imaging techniques can be found here: