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    C O N T E N T S

    2.1 Principal Problem of High Field MRI 212.1.1 Tissue Relaxation Rates 222.1.2 Susceptibility 222.1.3 Contrast Agents 23

    2.1.4 Chemical Shift Artifacts23

    2.1.5 B1-Inhomogeneity Artifacts 242.1.6 SAR Limitations 242.1.7 SSFP Banding Artifacts 25

    2.2 Changes to Accomodate These Problems 282.2.1 Improvement Using Parallel and Fast

    Imaging Techniques 282.2.2 Solving the Banding Artifact at 3 T 30

    References 30

    K E Y P O I N T S

    High-field imaging of the GI tract at 3 T can be veryattractive in terms of signal to noise ratio (SNR)when compared with 1.5 T imaging. However, severalissues need to be addressed and a direct transition ofsequences used at 1.53 T is not possible. The mainconstraints in 3 T imaging are related to changes intissue T1and T2relaxation parameters, susceptibil-ity artifacts, changes in efficiency of contrast agents,chemical shift artifacts, B1-inhomogeneity artifacts,specific absorption rate (SAR) limitations, andsteady-state free precession (SSFP) banding artifacts.In general, a careful adjustment of sequence param-eters is needed to avoid these artifacts. The use ofparallel imaging techniques is more beneficial at 3 T

    when compared with 1.5 T because by reducing scantime SAR problems often present at 3 T can beavoided. For reducing B1-inhomogeneity, hardwareadjustments are needed. Banding artifacts generatedin SSFP-sequences pose a severe limitation to theclinical use of these sequences. Though methodsexist to overcome this problem, a satisfactory solu-tion has not been proposed so far.

    2.1Principal Problem of High-Field MRI

    The use of high-field whole-body MRI systems, i.e. 3Tesla (T) MRI scanners, has become increasingly pop-ular since its introduction in the beginning of the2000s decade. The clinical applications range frombrain to musculoskeletal imaging (Barthet al. 2007).However, several different problems are associatedwith imaging at 3 T and a direct transition from 1.5 to3 T is not possible. In particular, aspects related to

    pulse sequence design, such as timings, radiofrequency

    Aart J. Nederveen, PhDSonia I. Gonalves, PhDDepartment of Radiology, Academic Medical Center, Amsterdam,

    The Netherlands

    MRI of the Gastrointestinal Tract 2

    at High-Field Strength

    Aart J. Nederveen andSonia I. Gonalves

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    22 A. J. Nederveen and S. I. Gonalves

    pulses, and specific absorption rate (SAR) issues haveto be considered when imaging at 3 T. Furthermore,experience gained, for example, in imaging the brainat 3 T does not necessarily apply to improved abdomi-nal imaging (Hussainet al. 2005; Merkleet al. 2006;Merkleand Dale2006; Pataket al. 2007;Schinderaet al. 2006). This is because many of the problems thatare associated with imaging at 3 T are region-specific.The main constraints in 3 T imaging are related tochanges in tissue T1 and T2 relaxation parameters,susceptibility artifacts, changes in efficiency of con-trast agents, chemical shift artifacts, B1-inhomogeneityartifacts, SAR limitations, and steady-state free preces-sion (SSFP) banding artifacts.


    Tissue Relaxation Rates

    Several studies (Bazelaireet al. 2004; Staniszet al.2005) on the variation of T1and T2relaxation param-eters at 3 T showed that the longitudinal relaxationparameterT1, increases with magnetic field strength bysometimes as much as 38% for liver tissue (Bazelaireet al. 2004) or 62% in gray matter. However, the range ofparameter variation among different studies may bequite wide because of the dependence on the T1andT2measuring methods. The increase in Larmor frequency

    from 64 MHz at 1.5 T to 128 MHz at 3 T causes an effi-ciency decrease in the energy transfer between spinsand lattice, thus causing the increase in T1values at 3 T(Bazelaire et al. 2004). In addition to the absolutechanges inT1, the relative differences in T1parametersamong different tissues, change from 1.5 to 3 T. Thiscauses an additional confounding effect when compar-ing T1-weighted images obtained at 1.5 and 3 T usingthe same contrast parameters such as time to repeat(TR). Because of the aforementioned differences, thecontrast in both sets of images will not be comparable.

    The transverse relaxation parameter T2 has beenreported by several studies (Bottomley et al. 1984;Staniszet al. 2005) to be essentially independent ofthe magnetic field strength. However, the study ofBazelaire et al. (2004), where T1 and T2 measure-ments were performed in vivo, reported a decrease inT2values of the order of 8%. This decrease might bepartially explained by the diffusion of water moleculesin the neighborhood of paramagnetic molecules suchas hemoglobin in deoxygenated blood. The spins inwater molecules experience small local field inhomo-geneties in the vicinity of these molecules and as a con-

    sequence, a shortening of T2is observed. These effects,

    though small, tend to increase with field strength andtherefore it is very likely that they become visible at 3 Twhile at 1.5 T they remain essentially unnoticed.

    The change in T1and T2relaxation parameters at3 T implies the decrease of T1contrast inT1-weightedimages or the decrease in the signal-to-noise ratio(SNR) of T2-weigthed images, if time to echo (TE)and TRparameters are used that are identical to theones used at 1.5 T.

    A greater additional source of concern at 3 T is thechange in the T2* relaxation value. At high field, thesphere of homogeneity (Vlaardingerbroek et al.1999) of the magnetic field is smaller. This, together withincreased difficulties in shimming and increased tissuesusceptibility effects cause a significant decrease in T2*,which becomes apparent in gradient-echo sequences, ingeneral, and in SSFP sequences, in particular, where the

    contrast is weighed by T1/T2ratio (see Sect. 2.1.7) andalso by magnetic field inhomogeneities through T2*.



    The magnetic susceptibility of a substance can bedefined as its ability to become magnetized inresponse to an external magnetic field. In the vicinityof interfaces between tissues with different suscepti-

    bilities (e.g., air-soft tissue and bone-soft tissue inter-faces), there are micro-variations in the magneticfield, which cause among others, in-plane image dis-tortion, localized areas of high and low intensity, aswell as localized signal drops due to local shorteningof T2*. This series of image artifacts are referred to astissue susceptibility artifacts and they increase withfield strength (Bernstein et al. 2006; Soher et al.2007). In practice, they are found in gas-filled bowel,since the susceptibility of air is much lower than thatof soft-tissue, or near metallic implants since the sus-ceptibility of metal is much higher than that of tissue.The effects of the field variations introduced by inter-faces between tissues with different susceptibilitiesdepend, however, on the type of sequence. Thus,echo-planar imaging (EPI), which consists of acquir-ing several gradient echoes in each TR, each of theseechoes corresponding to a different phase encodingstep, suffers much more from susceptibility artifactsthan fast gradient-echo imaging sequences. In addi-tion, the effects of susceptibility artifacts on magne-tization preparation prepulses such as inversionrecovery or fat saturation can cause a decrease in the

    efficiency of these prepulses because the spins

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    MRI of the Gastrointestinal Tract at High-Field Strength 23

    affected by the local field inhomogeneities may fail tobe tipped by the prepulse.


    Contrast Agents

    The action of contrast agents also changes with increas-ing field strength because not only the tissue relaxationtimes but also the relaxivity R of the contrast agentchanges. The effect of the contrast agent in theT1relax-ation parameter of a given tissue is given by (Bernsteinet al. 2006; Soheret al. 2007):

    1 1R C,

    T1,C T1= +



    is the longitudinal relaxation in the presenceof a concentration Cof contrast agent;T1is the longitudinal relaxation in the absence of

    contrast agent.For chelated gadolinium contrast agents, there is

    only a decrease of 510% in relaxivity when goingfrom 1.5 to 3 T. When compared with the variationsobserved in T1parameters, the relaxivity variation isvery small. This implies that for the same contrast

    concentration C, the relative change in T1parametersis larger at 3 T because the term 1/T

    1is smaller. When

    comparing images obtained at 1.5 and 3 T using thesame contrast concentration, the image obtained at3 T shows more contrast difference, which means thatless contrast has to be administered at 3 T to obtainimages equivalent to those obtained at 1.5 T. An addi-tional important point is that the increase in contrasteffectiveness at 3 T should be considered when clini-cally evaluating images and increased contrast in cer-tain areas should not be interpreted as pathological.

    Another category of contrast agents are the super-paramagnetic iron oxides (SPIOs) and ultra smallsuperparamagnetic iron oxides (USPIOs). These typesof contrast agents are composed of ferrites consistingof magnemite and magnetite (Fe


    3, Fe


    4), which

    show superparamagnetic behavior (Corotet al. 2006).They have been introduced for hepatic imaging andare nowadays extensively used not only in liver stud-ies but also in molecular imaging (Bulte andKraitchman 2004). The effect of (U)SPIOs on T1image contrast is similar to that described in (2.1) forchelated gadolinium contrast agents with the differ-

    ence that the relaxivity of (U)SPIOs is higher and is

    dependent on the aggregation of the contrast agent.However, the principal effect of (U)SPIOs is T2*effect, i.e. the strong decrease of T2*due to the mag-netization difference between different voxels withinthe image, which results from the inhomogeneous dis-tribution of the contrast agent. This magnetizationdifference becomes important for (U)SPIOs due to thehigh magnetic susceptibility of iron oxide. At higherfield strengths, the T2*effect increases due to higherintra-voxel dephasing, which implies that the use of(U)SPIOs is more efficient at 3 T than at 1.5 T.


    Chemical Shift Artifacts

    The chemical shift artifacts occur due to the differ-

    ence in the resonance frequencies of H1spins in waterand fat molecules and they are divided into chemicalshift artifacts of the first and second kinds. Chemicalshift artifacts of the first kind originate directly fromthe difference in precession frequencies, estimated tobe approximately equal to 3.5 ppm (Soheret al. 2007),and they occur along the frequency encoding and sliceselection directions. The difference in resonance fre-quencies is directly proportional to the field strengthand therefore this type of artifact becomes more prob-lematic at 3 T where the frequency difference is equal

    to 440 Hz. The chemical shift causes a misregistrationof fat, which appears as a dark band toward the lowerpart of the readout gradient field and as a bright bandtoward the higher part of the readout gradient field.Such artifact can be easily seen around the kidneysbut is also present at the boundary of the GI tract andit can occupy from one to several pixels in the image.At 3 T the width of these bands doubles, which meansthat at a constant field of view (FOV) and spatial reso-lution, the receiver bandwidth must be doubled inorder to reduce the artifacts to a size compared withthose obtained at 1.5 T. This bandwidth increase comesat the expense of a decrease in the signal-to-noiseratio (SNR) by approximately 30% (Haacke et al.1999). In clinical practice, however, this type of arti-fact is not very problematic and therefore the noiseconsiderations associated with RF pulse bandwidthbroadening rarely have to be considered.

    The chemical shift artifact of the second typeresults from the intravoxel phase cancellation duringreadout originated by the presence of water and fatmolecules in the same voxel. Contrary to the chemi-cal shift of the first kind, it can be seen along all vox-

    els in a fatwater interface and is therefore not

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    24 A. J. Nederveen and S. I. Gonalves

    restricted to the frequency encoding direction. Thesize of this type of artifact does not change with fieldstrength but the echo times have to be adjusted whenimaging at 3 T in order to account for the higherresonance frequency, which causes the water and fatspins to appear in and out-of-phase at differentmoments when compared with what happens at 1.5 T.In the 3 T case, the spins appear in-phase at 2.2 + i2.2 ms and out-of-phase at 1.1 + i 2.2 ms, where iisan integer larger or equal than 1, whereas at 1.5 T thespins appear out-of-phase at 2.2 + i 4.4 ms. Thus, inorder to capture in-phase or out-of-phase images at3 T, the echo times have to be halved.


    B1-inhomogeneity Artifacts

    Contrary to chemical shift or susceptibility artifacts,which exist both in 1.5 and 3 T and only become moreapparent at 3 T because of the increased SNR, thereare other artifacts such as B1 inhomogeneity andstanding wave artifacts, which are specific of highfield strength MRI. The increase in field strength isassociated with an increase in resonance frequency,which implies a concomitant increase in the RF pulsefrequency. This frequency change caused varioustechnical problems related to coil design (Duensingand Fitzsimmons2006), which had to be optimized

    to achieve a homogeneous B1 field. B1 field homoge-neity is crucial for T2-weigthed pulse sequences suchas turbo spin echo (TSE) images and artifacts due toB1 inhomogeneities often arise in images obtainedwith these protocols (Schicket al. 2005).

    Another problem inherent to the higher frequencyof the RF pulse is the shortening of B1 wavelength. Thehigh dielectric constant of tissue (water) causes the B1wavelength to decrease from 234 cm in free space toapproximately 30 cm in the body. The latter is of theorder of magnitude of the field of view (FOV) of many

    body imaging protocols and therefore artifacts result-ing from the generation of standing waves within theFOV can occur. These artifacts consist of strong signalvariations across the image where areas of high signalintensity occur where there is constructive interferenceand areas of signal drop coincide with areas of destruc-tive interference. These artifacts become more pro-nounced in larger areas of the body and, therefore,imaging of the abdomen can particularly be affectedby this type of artifact. One solution to this problem liesin improving coil design and RF transmission (Alsopet al. 1998; Tomaneket al. 1997; Vaughanet al. 2004).Examples of such solutions are the combination of

    several RF transmit coils whereby the phase and ampli-tude of the signal emitted by each coil are adjusted toobtain an homogeneous B1 field (Thesenet al. 2003)or by parallel RF transmission (Katscheret al. 2006).An additional possibility is to obtain a homogeneousB1 field by means of passive coupling of coils (Schmittet al. 2005). In the latter solution, the current induced ina local coil during excitation is used to improve thehomogeneity of the B1 field. Another type of solution isto place a cushion of dielectric material over the regionto be imaged to avoid the generation of standing waveinterference patterns. The cushion is made of a gel witha high dielectric constant mixed with gadolinium inorder to suppress the signal from the gel itself.

    Conductivity effects also tend to increase B1 fieldinhomogeneity. These effects happen in the neighbor-hood of highly conductive tissue such as ascites where

    current is induced by the rapidly changing RF field.The induced current tends to oppose the RF field,therefore causing local signal drops in the image.

    Figure 2.1 illustrates the effect of B1-inhomogeneityin the abdomen. In healthy volunteer (panel A), onlya small effect can sometimes be visible (see arrow),whereas in a patient with liver cirrhosis, portal hyper-tension, and ascites (panel B), the effect can be muchlarger because of the shielding effect of fluids, mak-ing the entire image useless. In the panels C and D,the result of B1-shimming using multitransmit paral-

    lel RF transmission is shown.


    SAR Limitations

    Specific absorption rate (SAR) is a measure of theenergy deposition in the human body and at 3 T, limita-tions in this parameter play an important role.Quantitatively, the SAR is given by:



    TR 2r




    wheretis the RF pulse duration;TRis the sequence repetition time;N

    Pis the number of pulses;

    NSis the number of slices;

    s and r are tissue conductivity and density,respectively;|E| is the amplitude of the induced electric fieldwithin the tissue.This second equation shows that if the main field

    is increased from 1.5 to 3 T and all other parameters

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    MRI of the Gastrointestinal Tract at High-Field Strength 25

    remain constant, then the SAR increases by a factorof 4. This can potentially cause an increase in tissuetemperature, which has to be taken in considerationin safety issues of pulse sequence design at 3 T tokeep the temperature increase below 1C. Thus, SARaspects frequently introduce limitations in the use of

    sequences such as spin echo (SE) or balanced SSFP(Scheffler and Lehnhardt 2003) that often work closeto SAR limits.


    SSFP Banding Artifacts

    Pulse sequences falling under the classification ofbalanced (b-)SSFP sequences such as b-FFE (balancedfast field echo), True-FISP (fast imaging with steady-state precession), and FIESTA (fast imaging employ-ing steady-state acquisition) have become popular in

    the last years. They belong to the class of rapid gradi-ent-echo sequences where in each TR, the magnetiza-tion reaches the same state thus implying that thesignal is acquired in steady state, this process havingbeen described for the first time by Carr(1958) asSSFP. In a given TR, the RF pulse rotates the magneti-

    zation by a given angle qand in the course of TR, themagnetization will undergo longitudinal and trans-verse relaxation. Furthermore, the transverse magne-tization will experience a certain amount of (spatiallydependent) dephasing bdue to the applied gradientsand magnetic field inhomogeneities. After a numberof TRs typically equal to 5 T1/TR, the magnetiza-tion reaches a steady state.

    In addition to b-SSFP sequences, the most impor-tant steady-state sequences (Chavhanet al. 2008) areT1-FFE (FLASH, SPGR) and T2-FFE (PSIF, SSFP) andthey differ in the type of gradient switching patternthat determines the amount of dephasing in each TR,

    a b

    c d

    Fig. 2.1. Illustration of the effect of B1-inhomogeneity in the

    abdomen (courtesy Bonn University, Germany and PhilipsMedical Systems, Best, The Netherlands). A normal liver is vi-sualized in (a) with B1-inhomogeneity indicated by an arrow.(c) Shows a patient with liver cirrhosis, portal hypertension,

    and ascites. Here, the effect can be much larger due to the

    shielding effect of fluids, making the entire image useless.(band d) Show the effect of B1-shimming using parallel RFtransmission (multitransmit)

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    26 A. J. Nederveen and S. I. Gonalves

    which characterizes the type of steady state that isreached and results in different types of contrast. Inquantitative terms, the amount of dephasing in eachTRis given by:


    ) = gdB


    + gr. G(



    0 (2.3)

    wheregis the gyromagnetic ratio;

    (r)B0 is the magnetic field inhomogeneity atposition (r);G(t)is the applied gradient at time instant t.From all SSFP sequences, b-SSFP has special inter-

    est because of the much higher signal that is obtainedeven at very short TRs (smaller than 10 ms) whencompared with T1-FFE or T2-FFE (Haacke et al.

    1999). The difference of b-SSFP with respect toT1-FFE and T2-FFE is that the gradients are fullybalanced in each TR. This implies that the amount ofdephasing in each TRis only dependent on the mag-netic field inhomogeneities since the second term ofthe right-hand side of (2.2) is zero. A direct conse-quence of this is that b-SSFP is insensitive to (con-stant velocity) flow (Vlaardingerbroekand DenBoer1999).

    Let us consider for simplification that the first RFpulse in the b-SSFP sequence is applied in a situationof thermal equilibrium. This implies that M0=M0z

    and under these circumstances the steady-state mag-netizationM

    SSat the end of each TRis:


    2),M0 (1-E1)E2sinMSS=











    qis the flip angle;bis the dephasing caused by the resonance offsetas given in (2.3);dis defined as

    d= (1-E1cosq) (1-E



    1-cosq) (E



    A detailed mathematical derivation of (2.4) and (2.5)can be found in Haacke et al. (1999) orVlaardingerbroek and Den Boer (1999). From

    (2.4), it is seen thatMSSvaries periodically as a func-

    tion of bas illustrated in Fig. 2.2 where a flip angle of50 was considered and the RF pulse was applied alter-natively along the positive and negative directions ofthe x-axis, i.e. in eachTRthe accumulated phase of thetransverse magnetization is b+ 180.

    Because bis spatially dependent, the appearanceof bands of varying signal intensity in the images isoften a problem in b-SSFP. The banding artifact ischaracterized by large signal drops in locations

    where b is equal to 180 and at 3 T this problembecomes very serious because the B0 inhomogene-ities are larger and shimming is therefore less effi-cient and often unable to prevent these artifacts. As aconsequence, the clinical application of b-SSFP at 3 Tremains limited until now.

    The banding artifacts of b-SSFP at 3 T becomequite prominent. To illustrate this, images obtainedin a phantom are presented in Fig. 2.3 for both at1.5 and 3 T using identical TR, TE, and flip angle.The phantom contains several tubes with differentT1

    -values. The desired contrast is present for bothfield strengths. However, at 3 T the banding arti-fact is much more prominent. At 1.5 T, the areaswithout bands is roughly two t imes larger than at1.5 T. In practice, the banding artifact at 3 Tseverely constrains the clinical use of this sequencein the abdominal region due to the enhanced sus-ceptibility effects, which aggravate magnetic fieldinhomogeneities.

    The signal/contrast characteristics of b-SSFPare, however, quite interesting from a clinical pointof view. In the following simplification, the mag-

    netic field inhomogeneities will be considered to be

    Fig. 2.2. Plot of the transverse magnetization in steady state(M

    SS) as a function of phase offset (b) for the b-SSFP where the

    phase alternation of RF pulses was considered. Flip angle = 50and TR= 10 ms, TE = 5 ms. Signal loss occurs at specificphase offsets

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    MRI of the Gastrointestinal Tract at High-Field Strength 27

    small so that in the case of RF phase alternation,bcan be considered to be equal to 180. In this case,(2.4) simplifies to:





    SS= .


    In the short TR limit, i.e. when TR

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    28 A. J. Nederveen and S. I. Gonalves

    flip angles which, at 3 T, can cause problems related tothe SAR.


    Changes to Accomodate These Problems

    When imaging at 3 T, the changes in relaxation param-eters, increased susceptibility, and chemical shift arti-facts as well as increases in the SAR imply that severalchanges have to be made in pulse sequences to makethem comply with high field strength imaging.

    The increase in theT1relaxation parameter causesa loss of contrast in T1-weigthed sequences. Thus, fora given set of sequence parameters and although at3 T, an increase in the SNR by a factor of 2 is expected,

    the contrast of T1-weighted images obtained at 3 Tcan actually be worse than those obtained at 1.5 T(Soher et al. 2007). A possible change in pulsesequence design to overcome this problem is toincrease the TR (and consequently the total scantime) to account for the increase in T1 relaxationrate. Another possibility is to apply magnetizationinversion prepulses to increase T1contrast withoutincreasing TR. In the same way, the decrease in T2(T2*) relaxation parameter at 3 T may imply a short-ening of the TEto keep the image contrast compara-

    ble with that obtained at 1.5 T. The increase in totalscan time causes additional problems such asincreased motion artifacts or less patient compli-ance. Also, the use of prepulses for magnetizationpreparation can also cause more problems related toSAR because of the additional power deposition. Incertain circumstances, there are additional sequenceparameters, such as number of signal averages, num-ber of acquired k-lines, or number of acquired echoes,which can be adjusted to decrease the total scan time.However, these adjustments are often made at theexpense of a decreased SNR and therefore alternativesolutions that are able to keep SNR levels aredesired.


    Improvement Using Parallel

    and Fast Imaging Techniques

    Additional solutions to decrease scan time withoutcompromising SNR are the use of parallel imagingtechniques (Larkmanand Nunes2007) or 3D fast

    imaging sequences that decrease scan time without

    compromising scan resolution. In parallel imaging,the independent signals recorded by a group ofphased-array coils are combined to generate theimage from the k-space, which is undersampled inthe phase encoding direction. In other words, theundersampling of the k-space by each coil impliesthat the FOV of each coil is reduced. There are twocategories of methods to reconstruct the image fromcoil signals: (1) image domain reconstruction; (2)frequency domain reconstruction. Methods such asSENSE (sensitivity encoding) (Pruessmann et al.1999, Van Den Brink et al., 2003) reconstruct theimage in the image domain by using the sensitivityprofiles of each coil and combining the intermedi-ary images obtained by each coil to obtain the finalimage. On the other hand, methods such as SMASH(simultaneous acquisition of spatial harmonics)

    (Blaimer et al. 2004) reconstruct the image in thek-space domain by computing a weighted average ofthe signals recorded by the different coils in this wayreconstructing the missing k-lines. The weight cor-responding to each coil takes into account the relativeposition of each coil. Finally, GRAPPA (generalizedauto calibrating partially parallel acquisition) algo-rithms (Griswoldet al. 2002) also reconstruct theimage in the k-space domain, but in this case, someof the central k-space lines that would be skippedare acquired and used to calibrate the method in

    order to compute the remaining k-space lines thatare skipped. This procedure lengthens the scan time,which however remains short. Parallel imagingmight be especially useful in EPI sequences wherethe acquisition of multiple echoes make them verysensitive to motion and susceptibility artifacts andwill therefore profit from reduced scan time. EPI,however, is nowadays seldomly applied in the abdo-men except for diffusion weighted imaging wherefast imaging is needed. Here, EPI is usually employedin the form of spin-echo ultra fast EPI. This is ahybrid sequence where after each 90 pulse andbefore each gradient echo readout, an additional180 refocusing pulse is added. As a consequence,several gradient echoes with T2-weigthing, each cor-responding to a different phase encoding step, aresampled in each readout.

    The increased signal strength at 3 T make the useof higher gradient amplitudes and higher gradientswitching speeds more beneficial when comparedwith 1.5 T, namely by allowing the acquisition ofhigh-resolution images within short scan times.Thus, 3D fast imaging sequences take advantage of

    these possibilities to acquire data from an excited

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    MRI of the Gastrointestinal Tract at High-Field Strength 29

    slab where phase encoding is also applied in the slicedirection. In this way, there is an increase in SNR dueto signal averaging during increased total scan timeper slice. This is possible due to increased gradientperformance at 3 T that allows for fast scan timescomparable with those of 2D multislice imaging.

    Both parallel and 3D imaging allow for the acquisi-tion of high-resolution images with fast scan timesbut without having to make a trade-off with the SNR.

    As mentioned before, SAR limitations are animportant problem when imaging at 3 T. From (2.2),it is seen that if all sequence parameters remain thesame, there is an increase by a factor of 4 when goingfrom 1.5 to 3 T. Protocols such as TSE sequences thatoperate near the SAR limits have often to be modi-fied by increasing the TRor decreasing the echo trainlength. However, additional solutions related to RF

    pulse design may be necessary to overcome SARproblems. Methods such as those described in Busseet al. (2004) and Henniget al. (1988, 2001) are able tolower RF energy absorption by factors of 2.56 whilekeeping acceptable levels of SNR and contrast-to-noise ratio (CNR). The method described byHargreaveset al. (2004), i.e. variable-rate selectiveexcitation, is able to lower energy deposition withoutthe need to decrease the flip angle.

    In general, body imaging at 3 T is quite affected bySAR limitations because many imaging protocols

    that are used operate close to allowed SAR limits.Although the above-mentioned methods can be usedto minimize this problem, often the advantages ofimaging at 3 T such as increased SNR, temporal andspatial resolution have to compromise in order toovercome SAR limitations.

    An illustration of a 3D T1-weighted spoiled gradi-ent echo sequence is given for MR enterography inFig. 2.5. Here, wall thickening after gadoliniuminjection at the level of the terminal ileum is clearlyvisible and one can expect that the theoretical signalbenefit of imaging at high field strength can beobtained in clinical practice for these sequences. Fatsuppression at the border of the images may, how-ever, be problematic due the field inhomogeneities atthese locations. Insufficient fat suppression can beobserved in Fig. 2.5 near the borders of the images asindicated by the arrows.

    Figure 2.6 shows a HASTE acquisition in MRcolonography, showing great anatomical detail thatwill be beneficial for colon cancer detection. Bothparallel imaging and decreased refocussing angleswere employed in this acquition in order to limit SAR

    deposition while retaining high image resolution.

    Fig. 2.5. Three-dimensional T1-weighted spoiled gradientecho sequence for MR enterography (TR/TE/FA = 2.1/1.0/10)and parallel imaging factor 2 (SENSE). Artifacts due to in-sufficient fat suppression resulting from magnetic field in-homogeneities are present at the borders and indicated byarrows

    Fig. 2.6. HASTE acquisition in MR colonography (TR/TE =

    500/65) using and parallel imaging factor 2 (SENSE)

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    30 A. J. Nederveen and S. I. Gonalves


    Solving the Banding Artifact at 3 T

    The banding artifact that characterizes b-SSFPsequences also becomes quite prominent at 3 T andparticularly when imaging the abdomen, where sus-ceptibility artifacts are quite intense, it is often verydifficult to use these type of sequences despite its clin-ical interest in imaging the abdomen (Gourtsoyianniset al. 2001). As mentioned before, the effectiveness ofshimming at 3 T is limited and therefore additionalmethods have to be used to minimize banding arti-facts. Banding artifacts appear due to the periodicityof the spectral response of b-SSFP sequences as dis-played in Fig 2.2. If phase alternation is applied to theRF pulse, the passband of the spectral response iscentered on resonance and the stopband coincides

    with a phase offset of 180, as displayed in Fig. 2.2.In steady state, the phase offset is periodic with aperiod equal to TRand therefore it is equivalent tosay that the stopband occurs at frequencies of 1/(2TR) Hz. Thus, a possible solution to avoid bandingartifacts is to decrease the TRso that the nulls in thespectral response fall outside the distribution of fre-quency offsets (Duerk et al. 1998) present in thesample. Figure 2.7 displays the maximum phase off-set as a function of TR for both 1.5 and 3 T fieldstrength when the field inhomogeneities are of the

    order of 3 ppm. As expected, the phase offset shows aperiodic behavior where the period T is such thatgdB

    0Tequals 2p. Note that this is the periodicity of

    themaximumphase offset. This implies that for TR>T, even if spins experiencing the maximum frequency

    offset have completed one period thus having againphase offsets well below 180, other spins experienc-ing a smaller frequency offset may still show phaseoffsets in the neighborhood of the null of the spectralresponse. Thus, to guarantee that the null of the spec-tral response lies completely outside the distributionof phase offsets in the sample,TRshould be kept wellbelow T. From Fig. 2.7, it can be concluded that at 3 T,TRof a b-SSFP sequence should be kept well below1 ms. Methods to solve banding artifacts by design-ing sequences with short TRs have been used at 1.5 T(Luet al. 2004). However, though the hardware at 3 Tallows for faster scan times, it can still be difficult toobtain such small TRvalues due to SAR limitations,gradient heating, high acoustic noise, and reducedcontrast and SNR. Thus, other solutions to the band-ing artifact, which do not require such short scan

    times, may be more desirable. At 1.5 T, the bandingartifact problem has been solved by averaging sets ofcomplex image data sets obtained at different equallyspaced offsets to obtain a uniform spectral response(Zuret al. 1990). More recently, an extension to thismethod has been proposed by Foxall (2002) wherethe RF phase is constantly increased during the entireduration of the experiment. This phase cyclingscheme creates a frequency modulation of the spec-tral response. It is shown that if the images corre-sponding to the different RF phases are averaged

    then good banding artifact cancellation can beobtained without the need to have very short TRspacings. In conclusion, several methods exist toreduce the banding artifact, but a satisfactory solu-tion has not been proposed so far.


    Alsop DC, Connick TJ, Mizsei G, (1998) A spiral volume coil for

    improved RF field homogeneity at higher static magneticfield strength. Magn Reson Med 40(1): 4954Barth MM, Smith MP, Pedrosa I, Lenkinski RE, Rofsky NM

    (2007) Body MR Imaging at 3.0 T: Understanding the oppor-tunities and challenges. Radiographics 27: 14451464

    Bazelaire CMJ, Duhamel GD, Rofsky NM, Alsop DC (2004) MRImaging relaxation times of abdominal and pelvic tissuesmeasured in vivo at 3.0 T: Preliminary results. Radiology230:652659

    Bernstein MA, Huston III J, Ward HA (2006) Imaging artifactsat 3.0 T. J Magn Reson Imaging 24:735746

    Blaimer M, Breuer F, Mueller M, Heidemann RM, GriswoldMA, Jakob PM (2004) SMASH, SENSE, PILS, GRAPPA: howto choose the optimal method. Top Magn Reson Imaging


    Fig. 2.7. Plot of the maximum phase offset as a function ofTRwhen B0 field inhomogeneities are of the order of 3 ppm

    for both 1.5 and 3 T field strengths

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    MRI of the Gastrointestinal Tract at High-Field Strength 31

    Bottomley PA, Foster TH, Argersinger RE, Pfeifer LM (1984) Areview of normal tissue hydrogen NMR relaxation timesand relaxation mechanisms from 1100 MHz: dependenceon tissue type, NMR frequency, temperature, species, exci-sion, and age. Med Phys 11:11841197

    Bulte JWM, Kraitchman DL (2004) Iron oxide MR contrastagents for molecular and cellular imaging. NMR Biomed17: 484499

    Busse RF (2004) Reduced RF power without blurring: correct-ing for modulation of refocusing flip angle in FSE sequences.Magn Res Med 51(5): 10311037

    Carr HY (1958) Steady-state free precession in nuclear mag-netic resonance. Phys Rev 112: 16931701

    Chavhan GB, Babyn PS, Bhavin G, Jankharia BG, Cheng HLM,Shroff MM (2008) Steady-state MR imaging sequences:Physics, classification, and clinical applications. Radio-graphics 28: 11471160

    Corot C, Robert P, Ide JM, Port M (2006) Recent advances iniron oxide nanocrystal technology for medical imaging.Adv Drug Deliv Rev 58: 14711504

    Duensing R, Fitzsimmons J (2006) 3.0 T versus 1.5 T: Coildesign similarities and differences. Neuroimaging Clin NAm 16(2): 249257

    Duerk J, Lewin JS, Wendt M, Petersilge C (1998) Remember trueFISP? A high SNR, near 1-second imaging method forT2-likecontrast in interventional MRI at.2 T. J Mag Res Imaging8:203208

    Foxall DL (2002) Frequency-modulated steady-state free pro-cession imaging. Magn Reson Med 48:502508

    Gourtsoyiannis N, Papanikolaou N, Grammatikakis J, Maris T,Prassopoulos P (2001) MR eneteroclysis between 3DFLASH with fat saturation after intravenous gadoliniuminjection and true FISP sequences. Eur Radiol 11: 908913

    Griswold MA, Jakob PM, Heidemann RM, Nittka M, Jellus V,

    Wang J, Kiefer B, Haase A (2002) Generalized autocalibrat-ing partially parallel acquisitions (GRAPPA). Magn ResonMed 47(6):12021210

    Haacke EM, Brown RW, Thompson MR, Venkatesan R (1999)Magnetic resonance imaging: Physical principles andsequence design. John Wiley & Sons, New York

    Hargreaves BA, Cunningham CH, Nishimura DG, Conolly SM(2004) Variable-rate selective excitation for rapid MRIsequences. Magn Reson Med 52:590597

    Hennig J (1988) Multiecho imaging sequences with low refo-cusing flip angles. J Magn reson 78: 397407

    Hennig J, Scheffler K (2001) Hyperechoes. Magn Reson Med46: 612

    Hussain SM, Wieloplski PA, Martin DR (2005) Abdominalmagnetic resonance imaging at 3.0 T: Problem or a promisefor the future? Top Magn Reson Imaging 16: 325335

    Katscher U, Boernert P, (2006) Parallel RF transmission inMRI. NMR Biomed 19: 393400

    Larkman DJ, Nunes RG (2007) Parallel magnetic resonanceimaging. Phys Med Biol 52(7): R1555

    Lu A, Barger AV, Grist TM, Block WF (2004) Improved spectralselectivity and reduced susceptibility in SSFP using a nearzero TE undersampled three-dimensional PR sequence.J Magn Reson Imaging 19:117123

    Merkle EM, Dale BM, Paulson EK (2006) Abdominal MR imag-ing at 3T. Magn Reson Imaging Clin N Am 1726

    Merkle EM, Dale BM (2006) Abdominal MRI at 3.0 T: Thebasics revisited. AJR 186: 15251532

    Patak MA, von Weymarn C, Froehlich JM (2007) Small bowelMR imaging: 1.5T vs 3T. Magn Reson Imaging Clin N Am15: 383393

    Prassapoulos P, Papanikolaou, Grammatikakis J, Rousomous-takaki M, Maris T, Goutsoyiannis N (2001) MR enteroclysisimaging of crohn disease. Radiographics 21: S161-S172

    Pruessmann KP, Weiger M, Scheidegger MB, Boesiger P (1999)SENSE: Sensitivity encoding for Fast MRI. Magn ResonMed 42: 952962

    Schick F (2005) Whole-body MRI at high field: technical limitsand clinical potential. Eur Radiol 15: 946959

    Stanisz GJ, Odrobina EE, Pun J, Escaravage M, Graham SJ,

    Bronskill MJ, Henkelman RM (2005) T1, T2Relaxation andmagnetization transfer in tissue at 3T. Magn Reson Med 54:507512

    Scheffler K, Lehnhardt S (2003) Principles and applications ofbalanced SSFP techniques. Eur Radiol 13: 24092418

    Schindera ST, Merkle EM, Dale BM, DeLong DM, Nelson RC(2006) Abdominal magnetic resonance imaging at 3.0 T:What is the ultimate gain in signal-to-noise ratio?. AcadRadiol 13: 12361243

    Schmitt M, Feiweier T, Voellmecke E, Lazar R, Krueger G,Reykowski A (2005) B1-homogenization in abdominalimaging at 3T by means of coupling coils. Proc Intl SocMag reson Med 13

    Soher BJ, Dale BM, Merkle EM (2007) A review of MR physics: 3T

    versus 1.5T. Magn Reson Imaging Clin N Am 15: 277290Thensen S, Krueger G, Mueller E (2003) Compensation of

    dielectric resonance effects by means of composite excita-tion pulses. Proc Intl Soc Mag Reson Med 11.

    Tomanek B, Ryner L, Hoult DI, Kozlowski P, Saunders JK (1997)Dual surface coil with high-B1 homogeneity for deep organMR imaging. Magn Res Imaging 15(10): 11991204

    Van den Brink JS, Watanabe Y, Kuhl CK, Chung T, Muthupillai R,Van Cauteren M, Yamada K, Dymarkowski S, Bogaert J, Maki JH,Mato C, Casselman JW, Hoogeveen RM (2003) Implicationsof SENSE MR in routine clinical practice. EJR 46: 327.

    Vlaardingerbroek MT, Den Boer JA (1999) Magnetic resonanceimaging: Theory and practice, 2ndedition. Springer-Verlag,Berlin

    Vaughan JT, Adriany G, Snyder CJ, Tian J, Thiel T, Bolinger L,Liu H, DelaBarre L, Ugurbil K (2004) Efficient high-frequency body coil for high-field MRI. Magn Res Med 52:851859

    Zur Y, Wood M, Neuringer L (1990) Motion-insensitive, steady-state free precession imaging. Mag Res Med 16(3): 444459

  • 8/13/2019 MRI artefact