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AAPM/RSNA PHYSICS TUTORIAL 275
AAPM/RSNA PhysicsTutorial for ResidentsMR Artifacts, Safety, and QualityControl1
Jiachen Zhuo, MS ● Rao P. Gullapalli, PhD
Artifacts in magnetic resonance (MR) imaging result from the complexinteraction of contemporary imager subsystems, including the mainmagnet, gradient coils, radiofrequency (RF) transmitter and receiver,and reconstruction algorithm used. An understanding of the sources ofartifacts enables optimization of the MR imaging system performance.The increasing clinical use of very high magnetic field strengths, high-performance gradients, and multiple RF channels also mandates re-newed attention to the biologic effects and physical safety of MR imag-ing. Radiologists should be aware of the potential physiologic effects ofprolonged exposure to magnetic fields, acoustic noise, and RF energyduring MR imaging and should use all the available methods for avoid-ing accidents and adverse effects. Imaging equipment should be regu-larly tested and monitored to ensure its stability and the uniformity ofits functioning. Newly installed or upgraded MR systems should betested by a physicist or qualified engineer before use. In addition, theauthors recommend participation in the MR imaging accreditationprogram of the American College of Radiology to establish the initialframework for an adequate quality assurance program, which then canbe further developed to fulfill local institutional needs.©RSNA, 2006
Abbreviations: ACR � American College of Radiology, FDA � Food and Drug Administration, GRE � gradient-recalled echo, RF � radiofre-quency, SE � spin echo, STIR � short inversion time inversion recovery, 3D � three-dimensional
RadioGraphics 2006; 26:275–297 ● Published online 10.1148/rg.261055134 ● Content Codes:
1From the Department of Radiology, University of Maryland School of Medicine, 22 S Greene St, Baltimore, MD 21201. From the AAPM/RSNAPhysics Tutorial for Residents at the 2004 RSNA Annual Meeting. Received June 27, 2005; revision requested September 16 and received October13; accepted October 14. All authors have no financial relationships to disclose. Address correspondence to R.P.G. (e-mail: [email protected]).
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IntroductionSince the invention of magnetic resonance (MR)imaging in 1972, the field has grown by leaps andbounds (1). Today, MR imaging is available insmall community hospitals, in the far corners ofthe country, and in many private outpatient clin-ics. The large growth in this field is attributable torapid technologic advances in several areas, in-cluding magnet technology, gradient coil design,radiofrequency (RF) technology, and computerengineering. In stride with the rapid technologicadvances, there has been phenomenal growth inthe number of applications for MR imaging. Weno longer look to MR imaging to provide onlystructural information, but also functional infor-mation of various kinds. Information about bloodflow, cardiac function, biochemical processes,tumor kinetics, and blood oxygen levels (for map-ping of brain function) are just a few examples ofthe data that can be obtained with MR imagingtoday (2–6).
Since the first U.S. Food and Drug Adminis-tration (FDA)-approved MR imaging system ar-rived in the marketplace around 30 years ago, theavailable magnetic field strengths have increased20-fold and more, the gradient capabilities haveincreased 100- to 200-fold, and the RF chain hasbeen completely revamped with phased-array coilimaging and parallel imaging (7,8). We have gonefrom the single receiver to the quadrature coil, tomultielement fast receivers. The growing demandfor new applications has helped to fuel rapidgrowth in the number and type of pulse se-quences available to radiologists. Imaging tech-niques have progressed from the simple spin-echo(SE) sequence to include steady-state, fast SE,echo-planar, and parallel imaging sequences(9,10). At the same time, radiologists and clini-cians have had to adjust to the pace of change.They have had to learn to apply these new tech-niques and to distinguish between clinical featuresand potential artifacts on the resultant images.
Demands for higher spatial and temporal reso-lution for both structural and functional imaginghave led to the proliferation of higher magneticfield strengths. Over the past 5 years, ultra-high-field-strength magnets (�1.5 T) have taken themarket by storm and have quickly moved from aresearch environment to a clinical environment.The high magnetic field strengths and the high-performance gradients have brought a new aware-ness of the issue of safety for both clinicians andpatients. The effects of exposure to magneticfields and the compatibility of the many so-calledMR-compatible or MR-safe surgical implants andother tools are being reinvestigated at high fieldstrengths. The use of rapid imaging techniquessuch as echo-planar imaging and half-Fourierrapid acquisition with relaxation enhancement,or RARE, along with the use of ultra-high-speedgradients, has raised awareness and concern forthe biologic effects of exposure.
With the expansion of the array of applicationsand increased number of users, it has becomenecessary to institute measures for quality controlover the daily operations of MR imaging centers.The effort to control quality has been spear-headed mainly by the American College of Radi-ology (ACR), which has instituted a program forcertification in MR imaging. Although certifica-tion is voluntary, more and more payers of clinicalMR imaging costs are basing their payments onwhether a site is ACR accredited or not.
The different components of an MR imagingsystem are shown in Figure 1. The main sub-systems are the magnet, gradient coil, RF genera-tor, and computer, the last of which controls theinterplay between subsystems and the reconstruc-tion, storage, and display of the images. There aremany sources of artifacts on MR images. Thesecould broadly be classified as image reconstruc-tion–related, system-related, and physiology-re-lated sources. Typically, image reconstruction–related artifacts occur because of limitations in-trinsic to the reconstruction algorithm used by the
Figure 1. Schematic showsthe basic components and ar-chitecture of an MR imagingsystem. T/R � transmit-receive.
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Teaching PointWe no longer look to MR imaging to provide only structural information, but also functional information of various kinds. Information about blood glow, cardiac function, biochemical processes, tumor kinetics, and blood oxygen levels (for mapping of brain function) are just a few examples of the data that can be obtained with MR imaging today (2-6).
be due to transient effects generated within one ormore of the subsystems or could be a sign of deg-radation of some of the electronic components inthe subsystem. System-related artifacts can beminimized through the adoption of a good qualityassurance or quality control program, such as theone recommended by the ACR. Artifacts relatedto physiology are dependent on complex interac-tions between the subject and the MR imagingsystem. These include artifacts due to breathing-related or other motion of the subject, as well asflow-related artifacts. Physiology-related artifactsmay be minimized with an understanding of the
basic anatomy and physiology involved and theappropriate use of specific pulse sequences.
The next section, a brief overview of k-space, isfollowed by a discussion of the types of commonartifacts seen on MR images. Safety concerns re-lated to MR imaging are discussed according tosubsystem, and guidelines for practice safety inthe MR imaging workplace are provided. Finally,the authors suggest some basic quality assuranceprocedures that may help to minimize artifactsand increase system uptime while maintaininghigh quality.
k-SpaceMR imaging systems collect data over time. Thedata that are captured during MR imaging arecalled k-space data or, simply, raw data. Typi-cally, the data are collected by using quadraturedetection, which provides both real and imaginaryk-space data. k-Space data include useful infor-mation but can be interpreted only after they aretranslated into images with the Fourier transformmethod (Fig 2). Several types of motion may oc-cur during the collection of k-space data. Themotion may cause artifacts, which represent thecomplex interaction between the motion and thedata acquisition time. Hence, an understandingof k-space is a critical element in the understand-ing and possible elimination of motion artifacts.
The process of k-space detection and imageacquisition is shown in Figure 2, and the typicalcharacteristics of k-space are shown in Figure 3.
Figure 2. Schematic ofquadrature coil detectionshows the collection and com-bination of real and imaginaryMR signal data to produce acomplex map of k-space,which is then subjected toFourier analysis to obtain theMR image.
Figure 3. Schematic shows the location oflow-frequency information about image con-trast resolution at the center of k-space, and, atthe edges of k-space, high-frequency informa-tion about spatial resolution and fine structure.
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particular vendor. System-related artifacts might
Teaching PointSystem-related artifacts might be due to transient effects generated within one or more of the subsystems or could be a sign of degradation of some of the electronic components in the subsystem. System-related artifacts can be minimized through the adoption of a good quality assurance or quality control program, such as the one recommended by the ACR. Artifacts related to physiology are dependent on complex interactions between the subject and the MR imaging system.
As mentioned earlier, k-space is the rawest formof data obtained at MR imaging. An acquisitionwith a 256 � 256 matrix contains 256 lines ofdata, and each of those lines contains 256 datapoints. The y-dimension in this 256 � 256 arrayis called the phase encoding direction, and thex-dimension is called the frequency encoding di-rection. The distance between neighboring pointsin k-space determines the field of view of the ob-ject imaged, and the extent of k-space determinesthe resolution of the image. For a matrix withthe same resolution, sparse sampling of the pointswould require less gradient strength and wouldproduce images with a larger field of view, whereasfaster sampling of points would require a highergradient strength and would produce images witha smaller field of view. Filling of k-space is accom-plished by acquiring frequency encoded datasamples for a given phase encoding step. Sam-pling is much quicker in the frequency encodingdirection than in the phase encoding direction,where the time between adjacent samples isgreater than or equal to the repetition time of theparticular sequence (except for fast SE and echo-planar imaging sequences). Low spatial frequen-cies are encoded in the center of k-space and
provide contrast resolution to the image, whereashigh spatial frequencies are encoded toward theedges of k-space and contribute to the spatialresolution of the image (11,12) (Fig 4). Fouriertransform of k-space is then applied to convertthe data into an image. Each pixel in the resultantimage is the weighted sum of all the individualpoints in k-space. Therefore, the information ineach pixel is derived from a fraction of every pointin k-space. On the basis of these facts, we canconclude that any disruption of k-space, whetherby motion, extraneous frequencies, or frequencyspikes, has the potential to corrupt the entire im-age. In the next section, we describe the types ofartifacts that might appear on MR images.
Spike (Herringbone) ArtifactGradients applied at a very high duty cycle (eg,those in echo-planar imaging) may produce baddata points, or a spike of noise, in k-space. Thebad data might be a single point or a few pointsin k-space that have a very high or low intensitycompared with the intensity of the rest of k-space.The convolution of this spike with all the otherimage information during the Fourier transformresults in dark stripes overlaid on the image
Figure 4. Maps of k-space data (top row) and corresponding MR images (bot-tom row) obtained in a phantom illustrate the relationship between raw k-spacedata and the reconstructed MR image after Fourier transform. (a) Image recon-structed only with data from the center of k-space has high contrast but low spa-tial resolution (poor definition). (b) Image reconstructed only with data from theperiphery of k-space shows well-defined edges but has poor contrast resolution.(c) Image reconstructed with all the k-space data has both good contrast andgood spatial resolution.
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Teaching PointFourier transform of k-space is then applied to convert the data into an image. Each pixal in the resultant image is the weighted sum of all the individual points in k-space. Therefore, the information in each pixel is derived from a fraction of every point in k-space. On the basis of these facts, we can conclude that nay disruption of k-space, whether by motion, extraneous frequencies, or frequency spikes, has the potential to corrupt the entire image.
(Fig 5). The displacement of the spike of noisefrom the center of k-space determines the spacingbetween the stripes and the angulation of thestripes with respect to the readout direction.Spike noise usually is transient, but if it is not at-tended to, it can become chronic. Spike noiseusually occurs because of loose electrical connec-tions that produce arcs or because of the break-down of interconnections in an RF coil, and it ismore evident with the use of high-duty-cycle se-quences (13).
On the other hand, the pattern of stripes pro-duced by a spike in k-space can be used for tag-ging, an important technique in cardiac imaging.In this scheme, preparation pulses are appliedprior to the imaging sequence. The preparationpulse has an excitation profile such that it formsechoes in different parts of the k-space. Fouriertransform of such images produces tags in a grid-like pattern, with known spacing between the
stripes. In cardiac imaging, these tags are appliedat the start of each cardiac phase, and cardiac im-age acquisition then is performed at multiplephases of the cardiac cycle. Following the changesin tag position during the cardiac cycle from thepoint of initial placement can be useful for assess-ing cardiac motion (14).
Zipper ArtifactThe zipper artifact is a common equipment-related artifact caused by the leakage of electro-magnetic energy into the magnet room. It appearsas a region of increased noise with a width of 1or 2 pixels that extends in the frequency encodingdirection, throughout the image series (Fig 6).All magnet rooms are shielded to eliminate inter-ference from local RF broadcasting stationsor from electronic equipment that emits an
Figure 5. Spike artifact. Bad data points in k-space (arrow in b) re-sult in band artifacts on the MR image in a. The location of the baddata points, and their distance from the center of k-space, determinethe angulation of the bands and the distance between them. The inten-sity of the spike determines the severity of the artifact.
Figure 6. Zipper artifact (RF leakage artifact). Images show constant-frequency artifacts (arrows) pro-duced by RF leakage from electronic components brought into the magnet room.
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electromagnetic signal that could interfere withthe MR signal. Leakage is usually caused by elec-tronic equipment brought into the MR imagingroom and the frequency generated by this equip-ment, which is picked up by the receive chain ofthe imager subsystem. The persistence of theproblem even after the removal or electrical dis-connection of all electronic equipment in the im-ager room might indicate that the RF shield hasbeen compromised.
Motion-related ArtifactsPatient motion during image acquisition pro-duces a commonly seen artifact. The artifact ap-pears as blurring of the image as well as ghostingin the phase encoding direction. The time differ-ence in the acquisition of adjacent points in thefrequency encoding direction is relatively short(order of microseconds) and is dependent on thesampling frequency or the bandwidth used. Thetime difference in the acquisition of adjacentpoints in the phase encoding direction is muchlonger and is equal to the repetition time used forthe sequence. The positional difference becauseof motion introduces a phase difference betweenthe views in k-space that appears as a ghost on theimage.
Respiratory and cardiac motion also can causemovement-related artifacts in the phase encodingdirection. In general, periodic movements causecoherent ghosts, whereas nonperiodic movementscause a smearing of the image (15). Motion arti-facts are generally caused by phase differences
between adjacent k-space lines that are encodedat different phases of cardiac pulsation and respi-ration. Figure 7a shows an image affected by arti-facts due to respiration. The best way to eliminaterespiration-related motion is to perform breath-hold imaging (Fig 7b). However, the number ofsections that can be obtained during a 20-secondbreath hold is limited. Multiple breath holds maybe performed to increase coverage and obtain im-ages in batches, but this requires cooperationfrom the patient. Some patients might have diffi-culty with breath holding. For such patients, re-spiratory gating may work better, with image ac-quisition performed at a certain phase of the re-spiratory cycle. The drawback of respiratorygating is that the acquisition time is increased.Another way to effectively reduce respiratory mo-tion artifacts is to use respiratory compensation orphase reordering (also called respiratory-orderedphase encoding, or ROPE), whereby the phaseencoding steps are ordered on the basis of thephase of the respiratory cycle (16). This tech-nique allows a smooth phase variation across k-space and the limitation of variation to a singlecycle of respiration instead of several cycles. Fig-ure 8 shows an example of images acquired withand without respiratory motion compensation byusing a fast SE sequence. Respiratory-orderedphase encoding works well in patients with regu-lar respiratory cycles but not in those with irregu-lar breathing.
Real-time navigator echo gating is an elegantnon–breath-hold technique that can be used tocompensate for several different types of motion(17) (Fig 9). With this technique, an echo from
Figure 7. Motion-related artifacts. (a) Abdominal image obtained without breath holdshows breathing-related motion artifacts. (b) Abdominal image obtained with breath holdshows a minimal artifact due to respiration and several motion artifacts due to cardiac pulsa-tion.
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Figure 8. Images acquired without and with compensation for respiratory motion. (a, c) Images ac-quired without compensation show respiration-induced artifact. (b, d) Images acquired with compensa-tion are unaffected by respiratory motion.
Figure 9. Placement of the navigator section for respiratory motion compensation. (a) Image withaqua overlay shows the navigator section from which the displacement information is obtained to deter-mine the diaphragmatic position. (b) Graph shows diagphragmatic movement, indicated by the whitewave and green line. The yellow boxes represent the best time to image (“window of opportunity”).
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the diaphragm is obtained to determine the dia-phragmatic position, and the timing of the acqui-sition is adjusted so that data are acquired onlyduring a specific range of diaphragmatic motion.The navigator echo part of the acquisition is in-terleaved with the actual imaging sequence to fa-cilitate real-time monitoring.
Cardiac pulsation can also lead to motion arti-facts, especially during imaging of the heart. Car-diac motion–related artifacts are avoided by usingelectrocardiographic gating to time image acquisi-tion so that it occurs at the same phase of eachcardiac cycle. Further artifact suppression can beachieved while imaging the heart if the acquisitionis performed during breath holding.
N/2 Ghost and Seg-mented-k-Space ArtifactsOver the past few years, single-shot and multishotimaging sequences have been used increasingly inthe clinical environment. Echo-planar imaging isbased on the continuous reversal of the echoes byusing a gradient pulse (commonly called gradient-recalled echo [GRE] imaging) after a single exci-tation for the acquisition of all lines in k-space toform a single image. In this case, every alternateline in k-space is read in the opposite directionbecause of the reversed polarity of the readoutgradient. Prior to Fourier transform of thek-space data, these lines must be reversed. Thisreversal process may result in the introduction ofphase errors in every alternate line of k-space.Phase errors can arise from minor deviations in
the linear course of the gradient as it traversesfrom maximum positive polarity to maximumnegative polarity, the existence of eddy currents inthe system, or poor shimming. The mismatch inphase causes the appearance of ghosts on the re-constructed images. On images acquired with asingle-shot echo-planar imaging sequence, theghost appears as an additional image with re-duced intensity that is shifted by half the field ofview, as shown in Figure 10b, since half the linesof k-space are different from the other half. Thistype of ghost artifact is commonly referred to asan N/2 ghost. The minimization of phase errorshelps reduce the intensity of the ghost imagewhile enhancing the main image (18). A similarphenomenon occurs in multishot imaging se-quences in which there is a phase discrepancybetween the segments of k-space. The number ofghosts on the image increases with the number ofdiscontinuities in the k-space, as illustrated byFigure 10c. On images acquired with a multishotsequence, errors could result from the number ofechoes per shot and the number of segments inthe k-space. To accurately correct for or minimizethese artifacts, it may be necessary to obtain anadditional navigator echo (18,19).
Images acquired with fast SE sequences alsoare susceptible to segmented-k-space artifacts.Here the discontinuities may occur because of aminor discrepancy in the timing of the sequence(eg, between the multiple RF pulses or betweenthe data collection windows) or because of eddycurrents within the system. Any of these itemscould easily introduce phase errors that result inghosting. In most situations, a user can do noth-ing but call service support to address the prob-lem.
Figure 10. Ghosts caused by phase errors at MR imaging in a phantom. (a) Initial image,acquired by using a fast SE sequence, shows no ghost artifact. (b) Image acquired withsingle-shot echo-planar imaging shows an N/2 ghost, a typical result of phase error (differ-ence between odd- and even-numbered echoes). (c) Image acquired with a segmented-k-space MR sequence with eight phase-encoding lines per segment shows a coherent ghost.
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Flow ArtifactsFlowing blood is another source of motion-re-lated artifacts, as shown in Figure 11. The bloodflow is manifested as ghosting in the phase encod-ing direction. GRE sequences are much moresusceptible to flow artifacts than are SE se-quences. In SE sequences, the flow usually ap-pears dark (produces no signal) because the flow-ing blood that is exposed to the 90° excitationpulse moves out of the imaging section before therefocusing 180° pulse is applied, while the bloodthat moved into the section at the same time wasnever exposed to the excitation pulse. In GREimaging, the in-flow effect produces the bright-blood phenomenon. A common way to reducethe motion artifact caused by through-plane flow
is to apply a saturation band adjacent to the imag-ing section. With this method, all spins within theslab are tilted toward the axial plane by a 90° RFpulse and then spoiled with the application ofstrong gradient crusher pulses before image ac-quisition starts. The spins, thus saturated, exhibitno signal when they move into the imaging vol-ume (Fig 11b).
Flow artifacts could also be minimized by us-ing flow compensation or gradient momentnulling. In this technique, flowing spins arerephased by using motion-compensating gradientpulses. As illustrated in Figure 12, if no motion
Figure 11. Control of flow-related artifacts. (a) Image shows a cardiac pulsation artifact. (b, c) Images obtainedwith a saturation band superior to the imaging section (b) and with first-order motion compensation (c) show lesssevere flow-related artifacts.
Figure 12. Schematics show the MR signal effects of flow compensation with gradient mo-ment nulling. Top: Typical readout gradient waveform. Bottom: Phase of cumulative sta-tionary spins (solid line) and constant-velocity spins (dotted and dashed lines). (a) Duringimaging without flow compensation, the moving spins are not refocused at the desired echotime (TE), and this leads to the loss of signal from flowing spins. (b) During imaging withgradient moment nulling, all the spins are refocused, and flow velocity is compensated for bythe 1:2:1 ratio of the gradient lobe areas. GGMR � gradient moment nulling gradient, GR �readout gradient.
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compensation is applied with the gradients, theflowing spins are not in phase with the static spinswhen the echo forms. With motion compensa-tion, the flowing spins with constant velocity (ir-respective of what the velocity is) are broughtback into phase, with no effect on static spins.Figure 11c illustrates the decrease in the severityof flow artifacts with the use of first-order motioncompensation, in comparison with flow artifactsin Figure 11a, which was acquired without mo-tion compensation. The penalty for using motioncompensation is increased echo time. Higher-order motion compensation also can be obtainedfor spins with constant acceleration and for spinswith varying acceleration (wobbling) by applyingadditional gradient pulses, but this increases theecho time even further (20).
Susceptibility EffectsWhen placed in a large magnetic field, tissues aretemporarily magnetized, with the extent of themagnetization depending on the magnetic suscep-tibility of the tissue. The effect of tissue magneti-zation slightly alters the local magnetic field. Thedifference in tissue susceptibility causes field in-homogeneity between tissue boundaries, whichmakes spins dephase faster, producing signals oflow intensity. This signal loss is especially severeat air-tissue or bone–soft tissue boundaries, be-cause air and bone have much lower magneticsusceptibility than do most tissues. Local mag-netic fields tend to take on different configura-tions around these interfaces, and the difference
introduces geometric distortion into the resultantimages, especially when sequences with long echotimes are used. SE sequences are less affected bylocal field inhomogeneity because of the 180° re-focusing pulse, which cancels the susceptibilitygradients. Echo-planar images generally are sub-ject to more severe susceptibility artifacts becausethe echoes are refocused by using gradient refo-cusing over a long period. Figure 13a shows asusceptibility artifact that occurred at echo-planarimaging. Note the distortion, near the air-tissueinterface, that was caused by the susceptibilitydifference. One way to minimize this distortion isto orient the phase encoding gradient along thesame axis as the susceptibility gradients. Figure13a shows echo-planar images acquired with thephase encoding gradient in a left-right directionwith regard to the image, whereas Figure 13bshows the same echo-planar images acquiredwith the phase encoding gradient in the anterior-posterior direction. Although the severity of thedistortion is markedly reduced, there is still someresidual distortion and some stretching in the di-rection of the phase encoding gradient. A goodunderstanding of the anatomy and the types oftissues involved may help reduce susceptibility-related artifacts. The best way to minimize sus-ceptibility artifacts is to use an SE sequence orreduce the echo time and increase the acquisitionmatrix. Proper shimming over the volume of in-terest before image acquisition is also critical toimprove the local field homogeneity.
The artifact that is caused by the presence ofmetallic implants is of the same nature, but it ismore severe because most metals have muchhigher magnetic susceptibility than does body
Figure 13. Echo-planar images show magnetic susceptibility artifacts. (a) Left-right phase encodingcauses severe distortion of the signal (arrows). (b) Anterior-posterior phase encoding minimizes the dis-tortion, since the phase axis is symmetric around the susceptibility gradients.
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tissue. As shown in Figure 14, metal-related arti-facts typically are manifested as areas of completesignal loss because the local magnetic field is sostrong that the spins are almost immediatelydephased. Sometimes, depending on the strengthof the resultant distortion of the magnetic field,part of a section may appear on images acquiredin a completely different section because of thefrequency difference between the spins in themetal and those next to the metal. The visual dis-tortion created by metallic implants on MR im-ages cannot be completely eliminated, but its ef-fect can be minimized by a large receiver band-width and a decreased echo time. Fast SEacquisitions with a high bandwidth typically workwell in such cases. However, care should be takento avoid the heating of tissue that is adjacent tothe metal (21).
Chemical Shift ArtifactsProtons in water and protons in fat have a signifi-cantly different chemical environment, whichcauses their resonance frequencies to differ. At1.5 T, protons from fat resonate at a point ap-proximately 220 Hz downfield from the waterproton resonance frequency, and this differencein frequency is linearly related to magnetic fieldstrength. The shift in Larmor frequency betweenwater protons and fat protons is referred to aschemical shift. The suppression of fat protons atMR imaging is very useful for improving contraston images of the breast, abdomen, and opticnerve (22).
The chemical shift between water and fat cancause artifacts in the frequency encoding direc-tion. If this occurs, there will be a slight misregis-tration of the fat content on images, because ofthe slight shift in the frequency of the fat protons.
The number of pixels involved in this slight shiftdepends on the receiver bandwidth and the num-ber of data points used to encode the frequencydirection. In mathematic terms, the result may beexpressed as follows: CSA � �� � Nfreq/BWrec,where CSA is a chemical shift artifact, �� is thefrequency difference between fat and water, Nfreqis the number of samples in the frequency encod-ing direction, and BWrec is the receiver band-width. This implies that the chemical shift can beminimized by using a higher receiver bandwidth.
Chemical shift artifacts were common on T2-weighted images obtained before the invention offast SE imaging. The overall signal intensity on anormal T2-weighted image was low, and in orderto increase the signal-to-noise ratio, it was a com-mon practice to use a low receiver bandwidth forthe T2-weighted echo. With the use of fast SEsequences, in contrast, the goals are to minimizethe echo spacing and to maximize the decayingT2 signal. These goals are usually achieved byusing a shorter RF pulse and a higher-bandwidthreceiver window.
The signal from fat can dominate the dynamicrange of an image and suppress the contrast be-tween lesions and normal tissue, especially onabdominal MR images. Echo-planar images arevery susceptible to artifacts from off-resonancefrequencies, and unless the fat frequency is sup-pressed, the images will be degraded by fat-re-lated artifacts. Any off-resonance signal (such asthat of fat) will be affected by a position shiftalong the image because of the long duration ofdata sampling (50–100 msec). Such a shift couldaffect approximately 12 pixels at a readout time of50 msec, as shown in Figure 15a. A commonmethod for minimizing fat-related artifacts is to
Figure 14. Sagittal MR image shows a mag-netic susceptibility artifact that resulted from thepresence of metallic dental fillings.
Figure 15. Chemical shift artifact at echo-planar im-aging. (a) Image shows severe chemical shift artifactfrom insufficient fat suppression. (b) Image obtainedwith fat saturation shows minimization of the chemicalshift artifact or off-resonance effect.
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apply a frequency-selective RF pulse to null thefat signal before applying the imaging pulse se-quence (23). Successful fat saturation will resultonly if the magnetic field homogeneity is suffi-cient throughout the imaging region. Figure 16illustrates the difference between a poorly shimmedobject and a well-shimmed object at MR imagingwith fat saturation. The result, in the case of thepoorly shimmed object, is saturation not only offat but also of a significant amount of water con-tent; the resultant images are nondiagnostic.Therefore, it is necessary to properly shim theregion before applying fat saturation. Propershimming results in the depiction of two distinctsignals in fat and in water. As shown in Figure15b, good shimming, followed by fat saturation,enables the avoidance of off-resonance effectsfrom fat. Another method is to use a short inver-sion time inversion recovery (STIR) sequence(24), as shown in Figure 17. In this method, theshort T1 of the fat signal is used to suppress thesignal from fat through the use of an inversionrecovery sequence. If the T1 of any tissue isknown, the inversion time needed to null the sig-nal from that tissue in an inversion recovery se-quence is given by 0.7 � T1. In the case of fat, theinversion time is around 150 msec at 1.5 T. Thedecision to use fat saturation or a STIR sequencedepends on the specific application; both meth-ods have advantages and disadvantages (25).
In-phase versusOpposed-phase ImagingChemical shift can be used to advantage, as isroutinely done in liver imaging for the detectionof fatty infiltration. Because the precessional fre-quencies of fat and water are different, they go inand out of phase from each other after excitation.Since the difference in frequency between fat andwater is about 220 Hz, the fat and water signals
are in phase every 1/220 Hz. In other words, theyare in phase at 4.45 msec, 8.9 msec, and so on(these time points correspond to the first cycle,second cycle, and later cycles). Within each cycle,the fat and water signals would be 180° out ofphase with one another at 2.23 msec, 6.69 msec,and so on. Figure 18 shows the difference be-tween in-phase and opposed-phase abdominalimages obtained with echo times of 2.2 msec and4.4 msec, respectively, at 1.5 T. At the in-phaseecho times, the water and fat signals are summedwithin any pixel that contains a mixture of thetwo components, whereas at opposed-phase echotimes, the signal in these pixels is canceled out,leading to the appearance of a dark band at thefat-water interface. In-phase and opposed-phaseimaging sequences are sometimes also referred toas chemical shift sequences (26).
Figure 16. Schematics show the effect of fat saturation with an RF pulse appliedat different frequencies for differentiation of water from fat. Left: Good magneticfield homogeneity, which results from a good shim, ensures effective fat suppres-sion on images. Right: Poor homogeneity, or a bad shim, may produce nonuniformfat suppression on images and compromise the clinical reading.
Figure 17. STIR sequence. The inversion time(TI) is set so that the available net magnetizationof fat is zero at 0.7 � T1 of fat. ADC � analog-to-digital converter, Mz � longitudinal magnetiza-tion, RF � RF pulse.
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Aliasing (Wraparound) ArtifactsA common artifact routinely encountered in MRimaging is aliasing. Aliasing occurs whenever theimaging field of view is smaller than the anatomybeing imaged. It should be noted that the field ofview in both the readout and the phase encoding
directions is inversely proportional to the incre-mental gradient step required from one point tothe next. This means that the field of view is thedistance along the gradient that one must move toexperience one complete cycle. Therefore, when afield of view is selected, it is expected that the gra-dient will move one cycle (0–2�) from one end ofthe field to the other. If the sensitivity of the RFtransmission coil extends beyond this field ofview, the spins outside the field will be excited,but they will be part of the next cycle in the phaseencoding direction (eg, 2–4�, 4–6�, and so on).However, the phase angles of the spins outsidethe field of view are essentially equivalent to thoseof the spins within the selected field of view buton the opposite side of the image. For the Fouriertransform, spins at 10° are equivalent to spins at2� � 10°, and the result is an overlap of signalsoutside the intended field of view with signalswithin that field of view. Figure 19 shows aliasingof signals from outside the field of view in the im-aging volume of interest. Aliasing is always en-countered in the phase encoding direction be-cause the frequency encoding direction is typi-cally oversampled within the system. Aliasing alsocan occur in the section direction in a three-di-mensional (3D) acquisition. In this case, the sidelobes of the RF pulse that excites the 3D slab pro-duce a signal from outside the field of view, andinformation in the first section may be aliased inthe last section of a 3D acquisition.
Especially in the coronal plane, large fields ofview that are smaller than the object will not onlycause aliasing but also create an interference pat-tern called the moiré or fringe artifact. Homoge-neity of the main field over large fields of viewdegrades toward the edges of the field, causingphase differences between the two edges. Whenaliasing occurs, the overlap of signals from oneside of the body to the other side, with mismatchedphases, produces a moiré artifact (Fig 20).
Figure 18. (a) In-phase MR image acquired with an echo time of 2.2 msec. (b) Opposed-phase MRimage acquired with an echo time of 4.4 msec. Both images were acquired at 1.5 T.
Figure 19. Aliasing artifact caused by imaging withtoo small a field of view. (a) Image shows aliasing arti-fact. (b) Image obtained with the phase encoding andfrequency encoding axes exchanged shows no aliasing.
Figure 20. Aliasing artifact caused by phase errors atboth sides of the magnet. Image shows a moiré artifactproduced by the addition and cancellation of signals.
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An easy way to eliminate aliasing is to exchangethe readout direction with the phase encodingdirection (swap the imaging axes) so that theanatomy fits into the phase field of view. How-ever, this method may result in motion artifactsalong the phase encoding direction that obscure apathologic entity. Other ways of decreasing alias-ing artifacts are to increase the phase field of viewor to apply spatial saturation pulses outside thefield of view so that the signals from outside thefield of view will not give rise to artifacts withinthe field.
The interference pattern also occurs when dif-ferent echoes from different excitation modes arereceived within the same acquisition window butwith slightly different echo times. The secondecho is often a stimulated echo. The spacing ofthe fringes in this case is inversely proportional tothe difference in echo times. The only way to ad-dress this problem is either to adjust the timing ofthe sequence so that the stimulated echo is super-imposed on the main echo or to use selectivesaturation to null the signal from the stimulatedecho.
Gibbs EffectGibbs or truncation artifacts are bright or darklines that appear parallel with and adjacent tothe borders of an area of abrupt signal intensitychange on MR images. Insufficient collection ofsamples in either the phase encoding direction orthe readout direction leads to Gibbs rings as aresult of the Fourier transform (27) (Fig 21). Oneusually sees Gibbs rings in the phase encodingdirection because a phase encoding matrix that issmaller than the readout matrix is often selectedto decrease the acquisition time. Gibbs rings canbe minimized by increasing the acquisition matrixsize and maintaining the same field of view, butthe improvement comes with a penalty of in-creased acquisition time and a reduced per-pixelsignal-to-noise ratio.
Section Cross TalkCross talk between sections occurs at imaging ofcontiguous sections (with no gap) in a multisec-tion two-dimensional acquisition. Typically, inthe presence of cross talk, all sections except theedge sections in a multisection acquisition havereduced signal intensity. The RF pulses used forexcitation are not perfect in that they do notproduce a section profile with a straight edge. In
Figure 21. Gibbs ring artifact. (a) Axial image obtained with a low spatial resolu-tion (128 � 128) in a cylinder shows a Gibbs ring artifact at the edges of the cylin-der. (b) Image obtained with a higher spatial resolution (256 � 256) shows mini-mization of the artifact. The dotted line indicates the desired object profile, and thered line indicates the object profile with two different resolution parameters.
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addition, any side lobes to the RF pulse mightexcite the neighboring section by a few degrees.In other words, the neighboring section is sub-jected to an RF pulse more than once during asingle repetition time, which causes partial satura-tion of the signal in that section and leads to alower signal intensity. Possible remedies for crosstalk between sections include RF pulses withsharper section profiles, an increased gap betweensections, and/or multiple sections imaged in sepa-rate batches (one batch consisting of all the odd-numbered sections; the other, all the even-num-bered sections). Three-dimensional images arenot vulnerable to this effect because the wholevolume undergoes excitation and because sec-tions within the volume are acquired by using thegradients, as is the case with normal frequencyand phase encoding.
Imaging System Cross TalkMany imaging centers have more than one MRimaging system with the same field strength. Ifmultiple MR imaging systems are operated atabout the same frequency and at the same time, it
is very important that each of the MR imagingrooms be surrounded by a tight shield to avoidcross talk between the imagers. It is probably bestto ground the RF shield separately and keep theRF and gradient amplifiers for the individual im-agers away from each other. To ensure that crosstalk is not occurring, regular testing should beperformed by applying intensive RF pulses andhigh-bandwidth reception on both imagers simul-taneously.
Gradient-related DistortionWhen an electrical current is applied to a gradientcoil, a varying magnetic field is produced that islinear through the isocenter of the magnet buttapers toward the sides of the magnet. As illus-trated in Figure 22, the gradient linearity getsworse with the distance from the isocenter. Theeffect of this distortion is to compress the images,thereby mismapping the spins from their true lo-cation at the edges of images obtained with a largefield of view. Most MR imagers are equipped withdistortion-correction algorithms to compensatefor the gradient nonlinearity. Figure 23a showsthe distortion on a spinal image with a large fieldof view, and Figure 23b shows the effect of thedistortion-correction algorithm on the image. Be-cause magnet sizes have become smaller in recentyears, the possibility of geometric distortionshould be given greater consideration, and theaccuracy of correction algorithms should be as-sessed (28).
Parallel Imaging ArtifactNew signal acquisition schemes such as SMASH(simultaneous acquisition of spatial harmonics),SENSE (sensitivity encoding), and GRAPPA(generalized autocalibrating partially parallel ac-quisitions), which are categorized as parallel im-aging techniques (29–31), have spawned a wholenew avenue for growth in MR imaging methodsand applications. Techniques such as echo-planarimaging and fast SE sequences may benefit fromthe new reliance on spatial information from
Figure 22. Schematic shows thegeometric distortion of a typical gra-dient profile along the x-axis, withdecreasing linearity (solid line) as thedistance from the magnet isocenterincreases. The red dotted line showsthe desired linear gradient profile.Bz � magnetic field strength, MaxFOV � maximum field of view.
Figure 23. (a) SE image obtained with a large fieldof view shows the result of gradient geometric distor-tion. (b) Image obtained with a vendor-supplied cor-rection algorithm shows correction of the geometricdistortion.
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phased-array coils to replace some of the spatialencoding performed with the usual gradient steps.Linear combinations of the sensitivities of the in-dividual coils in an array are used to form spatialharmonics. These harmonics can be used to re-place data obtained from some of the traditionalphase encoding steps, thereby reducing the acqui-sition time. The reduced acquisition time can beused to improve spatial or temporal resolution.However, care must be taken to avoid the use ofoveraggressive acceleration factors or of fields ofview smaller than the object being imaged. Paral-lel imaging artifacts differ from the routinely seenMR imaging artifacts. When the field of view issmaller than the object, aliasing artifacts can still
occur, but additional artifacts appear in the centerof the image (32). Figure 24 shows increasedghosting and noise in the center of the image withprogressive reductions in the field of view. In ad-dition, while temporal resolution can be improvedwith the use of acceleration factors greater thanone, the overall magnitude and inhomogeneity ofbackground noise increase with the accelerationfactor, as shown in Figure 25 (33,34).
There are other artifacts that may be encoun-tered in MR imaging, and it is impossible to listand discuss each one in this article. The artifactsdescribed are those most frequently encountered,and a good understanding of the sources of theseartifacts will lead to a better understanding of sys-tem performance and to more effective correctiveaction.
Figure 24. Parallel imaging artifacts resulting from excessively small fields of view. A comparison ofimages acquired with a field of view of 22 cm (a), 18 cm (b), and 16 cm (c) and with a constant accelera-tion factor of two shows increasing severity of aliasing artifacts, as well as increasing noise at the imagecenter, in b and c.
Figure 25. Parallel imaging artifacts resulting from excessive acceleration factors. A comparison of im-ages acquired at a constant field of view of 22 cm and with no acceleration (a), with an acceleration fac-tor of two (b), and with an acceleration factor of three (c) shows a marked increase in noise and loss ofresolution with increases in the acceleration factor.
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MR SafetyMR safety issues are related to the three mainsubsystems within the MR imaging system. Theseare the static magnetic field, the time-varied mag-netic field from the gradient subsystem, and theRF field from the RF subsystem. The magneticfield strengths of clinical MR imaging systemsvary from 0.2 to 3.0 T. Static magnetic fieldstrengths greater than 1.5 T may induce mild sen-sory effects when a person moves within the field.Time-varied gradient field strength (ie, the rate ofchange in the magnetic field, rendered in math-ematic terms as dB/dt), which is measured ingauss per centimeter per second, may cause pe-ripheral nerve stimulation. In addition, the RFfield has the potential to heat tissue because theconductivity of the tissues supports the transfer ofRF energy into conductive electricity within thetissue. Because of these potential biologic effects,the FDA has published guidelines for the safe op-eration of MR imaging systems. Manufacturersare mandated to ensure that their equipment op-erates in accordance with these guidelines, andproof of compliance is required by the FDA be-fore any MR imaging system is approved (35).
Static Magnetic FieldsMost epidemiologic studies about the effect ofstatic magnetic fields on the human body havebeen performed in populations of MR imagingtechnologists and other workers in a controlledenvironment. Investigators have focused on thelong-term exposure of workers with electrolyticcells in chemical separation plants to a static mag-netic field of very low strength (15.0-mT) orworkers in various national research labs withshort-term exposure to fields of up to 2.0 T (36–38). None have demonstrated a significant in-crease in any disease rate. Some creatures, such asbees and pigeons, are especially sensitive to varia-tions in the magnetic field of the earth because ofthe magnetite content in their brains, whichserves as an internal compass. Although no suchstructures have been identified in higher organ-isms or in humans, it is known that moving bloodcan induce electric potential on the order of milli-volts within a static magnetic field. The so-calledmagnetohydrodynamic effect is not consideredhazardous at the magnetic field strengths that areapproved for clinical use. However, at higher fieldstrengths, the probability exists that induced po-tential might exceed 40 mV, the threshold for de-polarization of cardiac muscle. Despite very lowor nonexistent risk from high-strength static mag-netic fields, there are sporadic reports of head-ache, vomiting, hiccupping, numbness, and tinni-tus, but they have not been substantiated. It
should be noted that these sensory effects alsowere reported when the electrical current throughthe magnets was turned off. However, investiga-tors in other studies have shown significant sen-sory effects (nausea, vertigo, and a metallic taste)at 4.0 T compared with 1.5 T. There also havebeen reports of an experience of flashing lightsresulting from rapid movement in large magneticfields. These lights, which are called magneto-phosphenes, are thought to be caused by retinalstimulation by induced currents. In 1996, theFDA designated all field strengths of less than 4.0T as posing a nonsignificant risk. Based on dataobtained in humans at strengths of more than 4.0T and in animals at as much as 16.0 T, this limi-tation for clinical magnets provides a substantialthreshold for any adverse events that might arisefrom static fields (39,40).
Although exposure to magnetic fields may besafe, unsafe practices around the magnet have ledto many accidents, some of them fatal. Access tothe magnet room must be strictly controlled toprevent unintended subjects and personnel frombeing exposed to the high-strength magnetic field.Large signs should be posted to warn bystandersof the presence of a large magnetic field and thehazards associated with entering the magnetroom. In addition, the nursing staff, technolo-gists, radiologists, and other clinicians should beeducated about the hazards of approaching amagnet. Precautions should be taken so that noperson is able to take any metallic objects insidethe magnet room, because such objects could be-come projectiles that might harm the patient onthe imaging table. If the 5-gauss line falls outsidethe shielded magnet room, the area should beclearly marked and barricaded so that people withcardiac pacemakers do not enter this zone. A vis-ible indicator of the 5-gauss line within the mag-net room also is helpful for planning the place-ment of monitoring equipment and other appara-tuses in the MR imaging room.
Prior to MR imaging examinations, technolo-gists and clinicians also should carefully questionpatients about whether they have ferromagneticimplants such as surgical clips, coils, stents, andpacemakers. Because such implants are ferromag-netic, the static magnetic field may torque theobject and exert a translational force. Most of thefatal accidents in MR imaging were caused by thefailure to identify subjects with a cardiac pace-maker. The FDA recently received several reportsof serious injury, including coma and permanentneurologic impairment in patients with neuro-logic stimulators, that occurred during an MR
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Teaching PointThe magnetic field strengths of clinical MR imaging systems vary from 0.2 to 3.0 T. Static magnetic field strengths grater than 1.5 T may induce mild sensory effects when a person moves within the field. Time-varied gradient field strength (ie, the rate of change in the magnetic field, rendered in mathematic terms as dB/dt), which is measured in gauss per centimeter per second, may cause peripheral nerve stimulation. In addition, the RF field has the potential to heat tissue because the conductivity of the tissues supports the transfer of RF energy into conductive electricity within the tissue.
imaging procedure. The most likely cause of suchinjury was heating of the electrodes at the end ofthe lead wires, which damaged the surroundingtissue. According to the FDA report, “Althoughthese reports involved deep brain stimulators andvagus nerve stimulators, similar injuries could becaused by any type of implanted neurologicstimulator, such as spinal cord stimulators, pe-ripheral nerve stimulators, and neuromuscularstimulators.” If MR imaging is to be performed ina patient with an implanted neurologic stimula-tor, the documentation for the specific model thatis implanted in the patient should be reviewedwith particular attention to warnings and precau-tions (41). The radiologist may need to consultwith the implanting or monitoring physician toobtain this information. In general, any instruc-tions for MR imaging that are in the labeling forthe implant, including information on typesand/or strengths of MR imaging equipment thatwas tested for interaction with the particular im-planted device, should be noted and followed ex-actly. The radiologist may need to consult withthe device implant manufacturer for this informa-tion. The force fields generated by implants aremore and more relevant as MR imagers with fieldstrength of 3.0 T take a stronger hold in the mar-ketplace. MR imaging personnel, including radi-ologists, therefore need to have the appropriateprocedures in place to track patients with suchimplants before approving them for MR imaging.Shellock (42) provides an excellent reference listof implants for which MR imaging is contraindi-cated.
Time-varied Magnetic FieldsSpatial localization and encoding are achieved byusing gradient coils that are inside the main mag-net. The magnetic field generated by a gradientcoil is extremely small, compared with the mainmagnetic field. The rate of change in the gradientfield strength is expressed in milliteslas per meterper millisecond or in gauss per centimeter persecond, and gradient field strengths as high as 45mT/m with very rapid switching capabilities arenow available for clinical use. Each orthogonalaxis has a pair of gradient coils, and by applyingelectrical current in opposite directions to the coilloops, varying linear magnetic fields can be gener-ated within the main magnetic field.
Just as the magnetic signal from nuclear pre-cession induces a current that is picked up by thereceiver coil, an electrical current and field may
be induced in a patient with changing magneticfields, as described by the Faraday law of induc-tion (41). Normal imaging sequences induce cur-rents of a few tens of milliamperes per squaremeter, a level that is far below that present in thenormal brain and heart tissue. However, the rapidswitching of magnetic field gradients that is pro-duced with single-shot techniques such as echo-planar imaging is capable of generating currentsthat exceed the nerve depolarization thresholdand cause peripheral nerve stimulation. The den-sity of the induced current depends on the num-ber of gradient switching operations per unit oftime and on the rate of change of the magneticfield during each switching operation. The num-ber of gradient switching operations per unit oftime determines the duration of the current in-duced, while the rate of change in the magneticfield during each switching operation determinesthe maximum current induced (also referred to asthe maximum slew rate) with a given gradientsystem (43).
The rate of change in the magnetic field (dB/dt)in MR imaging is classified as an extremely lowfrequency electromagnetic field. Biologic effectsof extremely low frequencies emitted by transmis-sion lines have received a lot of attention latelybecause the results of a number of epidemiologicstudies suggested a weak negative link betweenextremely low frequency emissions and health.While many have questioned the design of thosestudies (because of the lack of matched controls,the absence of a dose-response relationship, orthe absence of control for the effects of carcino-gens), it should be remembered that they wereperformed in human subjects who were exposedto extremely low frequencies over very long peri-ods. In comparison, the exposure to extremelylow frequencies that is incurred in a single MRimaging examination is negligible (it is equal to orless than the background exposure). Other ef-fects, such as magnetophosphenes, also have beenreported for exposure to time-varied gradientfields (44–46).
The likelihood that peripheral nerve stimula-tion will occur is greatest during echo-planar im-aging. This is especially true of acquisitions inoblique planes of section, since a higher slew rateresults from the combined contributions of gradi-ents from more than one axis. The likelihood ofperipheral nerve stimulation is also greatest whenthe readout gradient at echo-planar imaging is inthe craniocaudal direction.
Given that peripheral nerve stimulation islikely to occur with the use of techniques such as
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echo-planar imaging, the FDA, adapting the stan-dards of the International Electrotechnical Com-mission (technical report no. 60601–2-33), speci-fied three modes of operation for time-variedmagnetic fields. The normal mode, which is usedfor routine imaging of patients, is limited to 20T/sec for gradient ramp time exceeding 120 �sec,for a total of less than 2400 T/sec (with ramp timebetween 12 and 120 �sec). The upper limit forthe first controlled mode is 60,000 T/sec to pre-vent cardiac stimulation. This limit is unlikely tobe reached by any of the commercial scanners.The second controlled mode can be used only forresearch and with the approval of an institutionalreview board. Current safety standards do ad-equately address the safety of patients and per-sonnel. All MR imaging systems automaticallymonitor the rate of change in the magnetic fieldand provide warnings about the likelihood of pe-ripheral nerve stimulation. Proper communica-tion with patients should help alleviate further anyconcerns that the patient undergoing MR imagingmay have (47,48).
Acoustic NoiseThe pulsed currents that are passed through thegradient coils produce a torque on the coil itselfand result in the movement of the gradient coilsagainst their mountings. This movement, muchlike that of the diaphragm of a loudspeaker, cre-ates the loud banging noises heard by the subjectduring MR imaging. The FDA follows the adviceof the International Electrotechnical Commis-sion, which essentially asks the manufacturer toprovide a warning if the acoustic noise from agiven pulse sequence is likely to exceed 99 dB(A),which is sufficiently removed from 140 dB(A),the level at which damage to the hearing may oc-cur. Wearing proper headgear for hearing protec-tion is extremely important during MR imagingand can lower noise levels by about 30 dB(A).Manufacturers of MR imaging systems are ac-tively working to further decrease noise levels byusing methods such as active noise cancellationfilters and vacuum packing (49).
RF EnergyProlonged exposure of biologic tissues to RF en-ergy may result in tissue heating. The conductiv-ity of the tissue allows the absorption of RF en-ergy, which results in heating. This energy ab-sorption is described in terms of the specificabsorption rate. Such heating, if not controlled,can have physiologic effects such as changes inmental function and cardiac output. A rise in core
body temperature by less than 1°C generally is nocause for concern, because cell death does notoccur until the temperature exceeds 42°C. Attemperatures of less than 42°C, autonomic re-sponse mechanisms in the body reverse any of thechanges that occur with a temperature increasebeyond normal levels (50). However, care shouldbe taken when imaging patients with metallic ob-jects (ferromagnetic or otherwise), which mayresult in increased heating compared with heatingin normal biologic tissues. Failure to use MR-compatible leads during electrocardiographicallygated cardiac MR imaging could result in skinburns.
The specific absorption rate is measured inwatts per kilogram. Because the imaging systemcomputer calculates the specific absorption rateon the basis of the weight value entered in thesystem, the correct weight of the subject must beentered during the patient registration process.The FDA limit for a permissible increase in bodycore temperature during MR imaging is 1°C. Inaddition, the FDA considers a significant risk tobe incurred if the specific absorption rate exceeds3 W/kg averaged over the head for 10 minutes, 4W/kg averaged over the whole body for 15 min-utes, 8 W/kg per gram of tissue averaged over thehead or torso for 5 minutes, or 12 W/kg averagedover the extremities for 15 minutes (51).
It should be noted that the specific absorptionrate increases with the square of the frequencyand, hence, depends on magnetic field strength.With MR imaging at 3.0 T, a subject will experi-ence nine times the specific absorption rate expe-rienced with MR imaging at 1.0 T. The specificabsorption rate increases with the square of theflip angle, the size of the patient, and the dutycycle of the RF pulses. The MR imager operatorcan minimize heat deposition by controlling theflip angle of the pulse, increasing the repetitiontime, decreasing the number of sections for agiven repetition time, increasing the echo spacingin a fast SE sequence, decreasing the number ofechoes in a fast SE sequence, and reducing therefocusing flip angle in a fast SE sequence. Ingeneral, the key to decreasing the specific absorp-tion rate is to decrease both the power and theduty cycle of the RF pulses. This implies that pro-tocols based on RF energy–intensive sequences orprotocols for higher-field-strength magnets mustbe optimized to remain within the specific absorp-tion rate limits while producing images with thedesired level of contrast resolution.
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Quality ControlWhen an institution or a clinic makes a substan-tial capital investment in an MR imaging system,it counts on reaping benefits from that investmentfor a long time. To obtain those dividends, a goodquality-control or quality-assurance program isessential. Dividends are directly related to patientreferrals, and referrals in turn depend on the reli-ability of MR imaging and the image quality atthe particular site. In a complex system such asthe MR imager, many things can go wrong. Therecould be a catastrophic failure of the entire systemor a problem caused by one of the subsystems.Electrical and electronic components couldgradually deteriorate, and if the deterioration goesundetected, it could lead to substantial imagerdowntime. Even a seemingly benign undertakingsuch as upgrading of software might change thecharacteristics of the imaging system in ways thatcould go undetected unless appropriate testing isperformed before and after the upgrade.
Ideally, after the installation of a new systemor upgrading of an existing system, the systemshould be evaluated by a physicist or a qualifiedengineer prior to use. The qualified engineer isusually the service engineer from the vendor ofthe MR imaging system who performs all the testsand provides a report to the customer (52,53).Acceptance testing is performed with MR imag-ing of phantoms to obtain information about vari-ous imaging parameters. Qualified personnelfrom within the respective MR imaging centerscan use the test reports as a guideline in institut-ing a quality-control plan that tracks the variousparameters over time. The parameters that aretracked should provide an indication of the over-all stability of the system as well as of the variousRF coils and should point toward any deviationsfrom the norm in any subsystem. System stabilitycould be measured by tracking the homogeneityof the magnetic field, signal-to-noise ratio, resolu-tion, and contrast-to-noise ratio in a standardphantom. Such data should be collected on a pe-riodic basis. The length of the interval betweensuch tests may vary from site to site; testingshould be accommodated realistically in the over-all schedule, without increasing the burden ontechnologists or taking up too much time. Theperiodic collection of data alone is not enough:Only the results of regular analyses of the data canpoint toward a trend in the performance of theMR imaging system or its subsystems and allowthe measurement of performance against thresh-
old specifications. Trend analysis enables preven-tive maintenance of important components thatmay fail, thereby preventing substantial systemdowntime (54).
Most people agree that a quality-control pro-gram is beneficial for the smooth operation of anMR imaging system. However, very few sitesimplement rigorous quality control measures. Toensure a uniform quality of diagnosis across allMR imaging centers and all MR imaging systemvendors, the ACR instituted an MR system ac-creditation program in 1996. Furthermore, it isthe goal of the ACR to ensure that sites have along-standing commitment to quality control.Participation in the ACR accreditation program isvoluntary. The program is designed to be educa-tional while providing a framework for qualitycontrol. It also includes an evaluation of thequalifications of technologists, radiologists, andother personnel who will read images from thesystem that is to be accredited, as well as evalua-tion of the equipment performance and the qual-ity of clinical images. The accreditation processinvolves the purchase of a standardized phantom,and discounts for the accreditation process areavailable if an imaging center has more than oneMR imaging system. There are seven quantitativemeasures that the ACR is looking for that can bedetermined from the ACR phantom images. Theseinclude geometric accuracy, spatial resolutionof high-contrast objects, section thickness accu-racy, section position accuracy, image intensityuniformity, percentage of signal ghosting, andlow-contrast object detectability (55) (Fig 26).Accreditation is valid for 3 years and is renewable.A newly instituted regulation requires that newapplicants and previously accredited sites submitan annual MR imaging system performanceevaluation form, to be completed by a medicalphysicist or MR scientist. In addition, sites mustsubmit copies of their weekly data from onsitequality-control testing for the most recent quarter(55,56).
It is highly recommended that all sites applyfor ACR accreditation, for the simple reason thatthe process of accreditation provides a means ofestablishing a quality-assurance or quality-controlprogram that is recognized by leading authorities.An even more important and more practical rea-son to seek accreditation is that more and moreinsurance companies (starting with Aetna in2001) require that sites be accredited beforeclaims are submitted for insurance payment. Theinitial application process may seem daunting,but accreditation is well worth the effort. For
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Teaching PointIt is highly recommended that all sites apply for ACR accreditation, for the simple reason that the process of accreditation provides a means of establishing a quality-assurance or quality-control program that is recognized by leading authorities. An even more important and more practical reason to seek accreditation is that more and more insurance companies (starting with Aetna in 2001) require that sites be accredited before claims are submitted for insurance payment.
those who find the application process too timeconsuming, MR imaging system vendors and anumber of consultants will provide the service fora fee. Although the quality-assurance process in-stituted by the ACR may enhance the smoothrunning of an MR imaging center, it might nothelp identify the cause of increased ghosting onecho-planar images or help resolve patient moni-toring unit issues. However, the ACR quality-assurance guidelines can serve as the skeleton fora quality-control task list, to which individualsites may add items specific to their needs. Over-all, following the ACR guidelines and conductingregular internal analyses of the data collected willenable imaging sites to achieve and maintain highstandards of MR imaging quality.
ConclusionsMR imaging is a versatile and important modalityfor the diagnosis of various pathologic entities. Aquality-control program that incorporates safepractices in the MR imaging environment andthat reflects a good understanding of methods foravoiding artifacts will ultimately enable improve-ments in the efficiency of the use of this technol-ogy and, more important, in patient care.
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Teaching Points for AAPM/RSNA Physics Tutorial for Residents: MR Artifacts, Safety, and Quality Control
Jiachen Zhuo, MS and Rao P. Gullapalli, PhD
Page 276 We no longer look to MR imaging to provide only structural information, but also functional information of various kinds. Information about blood glow, cardiac function, biochemical processes, tumor kinetics, and blood oxygen levels (for mapping of brain function) are just a few examples of the data that can be obtained with MR imaging today (2–6). Page 277 System-related artifacts might be due to transient effects generated within one or more of the subsystems or could be a sign of degradation of some of the electronic components in the subsystem. System-related artifacts can be minimized through the adoption of a good quality assurance or quality control program, such as the one recommended by the ACR. Artifacts related to physiology are dependent on complex interactions between the subject and the MR imaging system. Page 278 Fourier transform of k-space is then applied to convert the data into an image. Each pixal in the resultant image is the weighted sum of all the individual points in k-space. Therefore, the information in each pixel is derived from a fraction of every point in k-space. On the basis of these facts, we can conclude that nay disruption of k-space, whether by motion, extraneous frequencies, or frequency spikes, has the potential to corrupt the entire image. Page 291 The magnetic field strengths of clinical MR imaging systems vary from 0.2 to 3.0 T. Static magnetic field strengths grater than 1.5 T may induce mild sensory effects when a person moves within the field. Time-varied gradient field strength (ie, the rate of change in the magnetic field, rendered in mathematic terms as dB/dt), which is measured in gauss per centimeter per second, may cause peripheral nerve stimulation. In addition, the RF field has the potential to heat tissue because the conductivity of the tissues supports the transfer of RF energy into conductive electricity within the tissue. Page 294 It is highly recommended that all sites apply for ACR accreditation, for the simple reason that the process of accreditation provides a means of establishing a quality-assurance or quality-control program that is recognized by leading authorities. An even more important and more practical reason to seek accreditation is that more and more insurance companies (starting with Aetna in 2001) require that sites be accredited before claims are submitted for insurance payment.
RadioGraphics 2006; 26:275–297 ● Published online 10.1148/rg.261055134 ● Content Codes:
tp last page template zhuo.pdfJiachen Zhuo, MS and Rao P. Gullapalli, PhD