introduction to mr imaging - university of california, davis...the t1 time (longitudinal recovery)...

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Introduction to MR Imaging

IntroductionThe development of magnetic resonance (MR) imaging hasrevolutionized diagnostic imaging. Prior to the introduction ofMR imaging in the late 1970s, medical imaging relied primarilyon radiography, which is limited by poor soft tissue contrastand superimposition of multiple structures. The advent ofdiagnostic ultrasound and x-ray computed tomography (CT)allowed visualization of "slices" of a patient’s anatomy, butultrasound is unable to penetrate gas or mineral, and CTachieves modest improvements in soft tissue contrastcompared with radiographs. MR imaging creates imagesbased on how protons (i.e., hydrogen atoms) behave in amagnetic field. Like ultrasound, MR imaging does not requireionizing radiation to generate an image. The ubiquity of waterand other molecules containing protons throughout the bodymakes MR imaging a powerful tool for evaluating parts of apatient’s anatomy that are inaccessible or poorly visible byother imaging methods. The sensitivity of MR imaging tochanges in tissue water content makes it particularly wellsuited to detecting pathology, such as edema, infection, andneoplasia. Knowledge of the basic principles of MR imagingwill help the dentist or physician to interpret images obtainedby commonly used MR imaging protocols, and to know whenspecial protocols may be beneficial for diagnosis.

MR Imaging System HardwareMR imaging requires sophisticated hardware and software togenerate an image, including

• Strong primary magnetic field; most clinical MR imagingsystems use superconducting magnet to generatemagnetic field with strength of 1.5 or 3.0 Tesla (T)

• Series of 3 gradient coils; gradient coils add or subtractfrom primary magnetic field along X-, Y-, or Z-axis of MRimaging system

• Radiofrequency (RF) coil; RF coil transmits pulses that"excite" protons within patient and then receives signalemanating from excited protons

Nuclear Magnetic ResonanceMR imaging is based on the phenomenon of nuclear magneticresonance, in which certain atoms act like tiny bar magnetswith positive and negative ends. Protons are the mostabundant magnetically active atoms in biological tissues andare the source of signal in conventional MR imaging. Whenplaced in a magnetic field, such as the bore of an MR imagingsystem, protons will align parallel (a.k.a., "spin up") orantiparallel (a.k.a., "spin down") with the magnetic field. Moreprotons will align in the lower energy, spin-up state, resultingin net magnetization parallel to the external magnetic field,referred to as longitudinal magnetization. Quantummechanics predicts that protons will not perfectly align withthe axis of the external magnetic field but at an angle to themagnetic field. The protons then wobble, or "precess," aroundthe axis of the magnetic field with a frequency, known as theLarmor frequency, that depends on the external magneticfield strength. At equilibrium, a group of protons will berandomly distributed around their respective precessionalorbits, causing no net magnetization perpendicular to the axisof the magnetic field; i.e., there is no net transversemagnetization. The equilibrium condition, in which there ispositive longitudinal and zero transverse magnetization, canbe perturbed by briefly applying a secondary magnetic field(known as the RF pulse) that oscillates at the Larmorfrequency. Application of an RF pulse flips some protons from

the spin-up state to the spin-down state and aligns theprotons' orbits, decreasing the net longitudinal magnetizationand inducing transverse magnetization. Transversemagnetization is the source of signal detected by the MRimaging system for constructing an image. A 90° RF pulseresults in zero longitudinal and positive transversemagnetization; this effect can be visualized as "tipping" thenet magnetization 90° from the longitudinal to the transverseplane. Following excitation by an RF pulse, a group of protons"relaxes" toward the equilibrium state by 2 independentprocesses: (1) recovery of longitudinal magnetization and (2)decay of transverse magnetization. The rates of relaxation bya tissue depend on inherent properties of the tissue known asthe T1 time (longitudinal recovery) and T2 time (transversedecay).

Anatomy of MR Imaging Pulse SequenceAn MR imaging pulse sequence consists of a coordinatedseries of RF pulses, gradient coil application, and dataacquisition to create an image with a specific type of contrast.An MR imaging operator can choose from many possible pulsesequences and refine some pulse sequence parametersdepending on the prescribed application. The events in a basicpulse sequence serve 1 of the following purposes: (1)excitation, (2) spatial encoding, or (3) refocusing.

ExcitationAll pulse sequences employ an excitation RF pulse. Many pulsesequences begin with a 90° excitation RF pulse, but otherexcitation tip angles are also used. The amount of timebetween consecutive excitation RF pulses is known as therepetition time (TR). A long TR will allow tissues to recovertheir longitudinal magnetization before the next excitation,whereas a short TR will only allow tissues with rapidlongitudinal recovery to fully recover their longitudinalmagnetization.

Spatial EncodingSpatial encoding involves applying gradient coils so that signalarising from a group of excited protons contains informationregarding the 3D location of the protons within the patient. Agradient coil works by adding to or subtracting from theprimary magnetic field to create a magnetic field gradient thatvaries with location along 1 axis of the MR imaging system.Because the Larmor frequency of a proton depends on thestrength of the surrounding magnetic field, a magnetic fieldgradient will result in protons precessing at differentfrequencies depending on their location in space. Spatialencoding typically consists of 3 steps: slice selection, phaseencoding, and frequency encoding, which respectively definethe z-, y-, and x-axes of the MR images. Slice selection occursby turning on a gradient coil(s) to vary the magnetic fieldstrength along the z-axis during the excitation RF pulse. Theexcitation RF pulse is then applied at a predetermined rangeof frequencies. Only protons with Larmor frequenciesmatching the RF pulse will become excited, creating a "slice"of excited protons within the patient. The slice-select gradientstrength and frequencies of the RF pulse determine thethickness and location of the slice along the z-axis.

Immediately following slice selection, all protons in theexcited slice will be aligned in their precessional orbits and are"in phase." Phase encoding works by applying a gradient coil(known as the phase encode gradient) along the y-axis.Protons located at points along the y-axis with a strongermagnetic field will speed up their precession and advance

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their phase. Protons experiencing a weaker magnetic field willprecess more slowly, retarding their phase. Once the phase-encode gradient is turned off, the excited protons return totheir original Larmor frequency but retain their new phases.Importantly, constructing a typical MR image requires manyphase-encoding "steps," but only a single or few phase-encoding steps can be performed per TR interval using basicpulse sequences. The need for phase encoding is 1 of themajor rate-limiting steps of MR imaging.

Frequency encoding occurs during data acquisition. Thefrequency encode gradient is applied during data acquisitionto vary the precessional frequencies of protons along the x-axis. Protons at different positions along the x-axis will thuscontribute waves with varying frequencies to the signalcollected during data acquisition. Many different frequenciescan be encoded simultaneously, and frequency encoding doesnot typically affect image acquisition time. At the time of dataacquisition, signal from a tissue will be encoded with spatialinformation pertaining to all 3 axes. The image position alongthe z-axis is defined by slice selection, and spatial informationin the y- and x-axes is encoded by the phases and frequenciesof the signal collected during data acquisition.

RefocusingImmediately following the excitation RF pulse, excitedprotons in a tissue generate a signal; however, without spatialencoding, the signal generated by the excitation RF pulsecannot create an image. During basic MR imaging, this initialsignal decays prior to and during spatial encoding. Therefore,it is necessary to refocus the decayed signal to form an image.

Transverse magnetization decays much more rapidly thanlongitudinal magnetization recovers. Imperfections in theprimary magnetic field, gradient coil application, and tissueinhomogeneity further hasten the loss of transversemagnetization. By the time image data are acquired, theoriginal transverse magnetization is undetectable. Thisproblem is overcome by inducing an "echo" of the originaltransverse magnetization using either a 180° refocusing RFpulse for spin-echo imaging or by reversing the polarity of aspecific gradient coil for gradient-echo imaging. Both spin-echo and gradient-echo imaging "rephase" the decaying signalto form a signal echo. The time between the excitation RFpulse and the echo formation is known as the echo time (TE).

With the use of multiple 180° refocusing pulses after anexcitation pulse, it is possible to generate and detect multiplesignal echoes during a single TR interval. This technique,known as fast spin-echo (FSE) imaging, is often used for clinicalpulse sequences to reduce image acquisition time. FSE pulsesequences slightly alter the expected contrast of an imagecompared with spin-echo pulse sequences and are primarilyused for creating images with T2W contrast. Multiple echoescan also be generated per TR interval during gradient-echoimaging, but the rapid decay of transverse magnetizationduring gradient-echo pulse sequences limits the number ofuseful echoes that can be obtained.

Mechanisms of Contrast in MR ImagingContrast between different tissues in MR images depends oninherent tissue properties and user-defined pulse sequenceparameters. Inherent properties that influence theappearance of a tissue in an MR image include the tissueproton density, T1 time, and T2 time. The major pulsesequence parameters that affect tissue appearance are the TRand TE. The TR determines how much time is available to

recover longitudinal magnetization between excitation RFpulses. Tissues that incompletely recover their longitudinalmagnetization will have less magnetization available to betipped into the transverse plane by the next excitation pulse,resulting in decreased signal in the final image. A relativelyshort TR will thus create contrast between tissues withdifferent T1 times, yielding a T1W image. The pulse sequenceTE determines how sensitive the image will be to differencesin transverse decay among tissues. Setting a relatively long TEwill allow more time for transverse magnetization to decay,creating contrast between tissues with different T2 times toproduce a T2W image. A pulse sequence with a long TR andshort TE will maximize longitudinal recovery while minimizingtransverse decay, resulting in minimal dependence on tissueT1 or T2 times. Contrast in such an image will depend primarilyon the density of protons in tissues, creating a protondensity-weighted (PDW or PD) image. It is important to notethat 1 characteristic of a tissue may exert a large influence onthe appearance of the tissue in an image, regardless of thecontrast intended by the pulse sequence. For example, atissue with very scarce protons will appear dark on T1W, T2W,and PDW images regardless of the T1 or T2 times of thetissue. Also, gradient-echo pulse sequences are moresusceptible to magnetic field heterogeneity and rapidly losetransverse magnetization by a mechanism known as T2*decay. Gradient-echo images with relatively long TE are said tohave T2* weighting. Many other mechanisms of generatingcontrast in MR imaging are possible but beyond the scope ofthis chapter.

Inversion Recovery and Tissue NullingInversion recovery is a technique that employs an extra RFpulse with a tip angle of 180° to begin the pulse sequence.Immediately after this initial 180° RF pulse, known as theinversion pulse, all excited tissues will have negativelongitudinal magnetization. The inversion pulse is followed bya delay known as the inversion time (TI), then a 90° excitationRF pulse. During the TI, different tissues experiencelongitudinal recovery according to their respective T1 times,and each tissue will pass through a zero point when the tissuehas no net longitudinal magnetization. Tissues with short T1times (e.g., fat) will pass through the zero point at relativelyshort TI and tissues with long T1 times (e.g., water orcerebrospinal fluid) will pass through the zero point atrelatively long TI. If a tissue has zero net longitudinalmagnetization at the time of the excitation pulse, that tissuewill not contribute signal and will appear black in the finalimage. Commonly used inversion pulse sequences includeshort tau inversion recovery (STIR), which nulls the signal fromfat, and fluid-attenuating inversion recovery (FLAIR), whichnulls the signal from low protein fluids.

Contrast EnhancementContrast enhancement typically refers to using exogenouscontrast media to aid in the detection of pathology or increasevisibility of normal anatomic structures. The most commonclinically used MR imaging contrast media are based ongadolinium. Gadolinium is a metal ion with magneticproperties that shorten the T1 time of a tissue. Othermechanisms of positive and negative contrast enhancementare used in research and may have specific clinical applicationsbut are not routinely used. Gadolinium contrast medium istypically injected intravenously and taken up by highly vasculartissues and tissues with increased vascular permeability. In thepresence of gadolinium, a tissue will have a shortened T1 timeand appear brighter in a T1W image compared with a

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precontrast image. Gadolinium contrast enhancement can behelpful for identifying inflammation, blood vessels, and sometumors.

Spatial ResolutionImage spatial resolution determines the ability to resolve 2adjacent structures as separate entities in an image. Greaterspatial resolution will allow closer and smaller individualstructures to be resolved. In MR imaging, spatial resolution isdefined by the image slice thickness, field of view (FOV), andimage matrix dimensions. Increasing slice thickness will resultin averaging of signal from adjacent tissues in the z-dimensionof the patient, blurring the edges of some structures anddegrading the image spatial resolution. FOV defines theboundaries of the image to be acquired. If all otherparameters are maintained, increasing the FOV will causepoorer spatial resolution; however, FOV must be large enoughto contain all of the patient anatomy within the RF coil in theplane of the image. Selecting an FOV smaller than thepatient’s anatomy can lead to an effect known as "wraparound" in which signal from body parts outside the FOV willbe superimposed on other structures in the image. The matrixdimensions of an image define the number of rows andcolumns of pixels in the image. Together, the slice thickness,FOV, and matrix dimensions determine the size of each pieceof tissue that will be represented by individual pixels in animage. Thin slices, small FOV, or a large matrix size will allowresolution of fine details in an image but are all associatedwith long acquisition time &/or poorer signal:noise ratio (SNR).

MR imaging can obtain images in 2D or 3D modes ofacquisition. Both 2D and 3D modes of acquisition produce 2Dimages representing slices of the patient’s anatomy. 2Dimaging operates by exciting a relatively thin slice of tissue,then encoding the remaining spatial information as describedabove. 2D imaging excites and acquires image data from 1slice of tissue at a time, and typical 2D MR imaging pulsesequences require a gap between adjacent slices to preventthe occurrence of certain artifacts. 3D imaging excites a thick"slab" of tissue at the beginning of the pulse sequence, thenuses 2 phase-encoding steps plus 1 frequency-encoding stepto encode spatial information within the slab. Slices are thenreconstructed from the spatially encoded slab of image data.A 3D approach allows creation of images with very thin,contiguous slices and fine spatial resolution. The disadvantageof 3D MR imaging is that many phase-encode steps arenecessary, and image acquisition can be time consuming.

Positioning of the patient and accurately prescribing thedesired image orientation are essential for accurate diagnosis.Sagittal, axial, and coronal image planes are all used forvisualizing specific anatomic features and abnormalities of thehead. In general, an image best portrays anatomic featuresand abnormalities in the plane of the image and poorlyrepresents features orthogonal to the image plane. Two ormore imaging planes are often necessary to completelyvisualize all surfaces of an anatomical region. Compleximaging planes can be prescribed to capture curved or obliquesurfaces. It may also be necessary to acquire images withchanges in patient position to characterize a pathologicprocess. For example, MR imaging of the TMJ is oftenperformed in closed-mouth and open-mouth positions toevaluate for potential changes in position of the TMJ disc.

Signal:Noise Ratio and Acquisition TimeAs with any diagnostic imaging modality, MR image qualitydepends on spatial resolution, contrast, and signal:noise ratio(SNR). A major disadvantage of MR imaging compared withother diagnostic imaging modalities is that an MR imagingstudy typically requires much longer acquisition time thanradiographs or CT. Improving MR imaging spatial resolution&/or SNR generally requires more time for image acquisition.For clinical MR imaging studies, it is necessary to balanceimage quality with acquisition time.

All digital imaging modalities are subject to electronic noisethat contributes randomly to the final image and degradesthe image quality. For radiography and CT, the number of x-ray photons contributing to image formation (i.e., the signal)far outweighs any electronic noise, and the noise is oftenimperceptible. Conversely, the signal generated during MRimaging is relatively small, and MR imaging SNR is often thefactor that limits image quality. SNR can be improved by usingdifferent hardware as well as changing certain pulse sequenceparameters. Using an MR imaging system with greater fieldstrength increases the SNR for other factors being equal but isnot practical in clinical settings. One of the most effectiveways to improve SNR is to use an RF coil that is specially suitedto the region of interest because proximity of the RF coil tothe region of interest has a major influence on signal. Surfacecoils are RF coils intended to be placed directly on the surfaceof the region to be imaged. Surface coils sacrifice the ability toimage deep into the body but achieve excellent SNR fortissues close to the surface. Surface coils are frequently usedto image the TMJs and spine.

Nearly all pulse sequence parameters have the potential toinfluence the image SNR. For example, changing the imagespatial resolution will also influence the SNR. Decreasing theslice thickness, decreasing FOV, &/or increasing the matrix sizewill cause a smaller sample of tissue to generate the signal foreach pixel in the image. Thus, creating fine spatial resolutionwill sacrifice SNR. Increasing TR and decreasing TE will tend toincrease the amount of signal for image formation but mayalter the image contrast. When preservation of contrast isessential, the number of excitations (NEX) can be increased.Using a NEX > 1 simply repeats the pulse sequence 2 or moretimes and averages the signal from each excitation to createthe final image. Importantly, doubling the NEX will onlyincrease the SNR by 41% but double the acquisition time.Other pulse sequence parameters can also influence imageSNR but are beyond the scope of this chapter.

SummaryMR imaging system hardware includes a primary magneticfield, gradient coils, and RF coil. A pulse sequence generatesan MR image via coordinated gradient coil application, RFpulses, and data acquisition. Inherent tissue properties (e.g.,T1 time, T2 time, and proton density) and pulse sequenceparameters (e.g., TR and TE) determine the appearance of atissue in an image. MR imaging can create images withdifferent tissue appearances (e.g., T1W, T2W, PDW, and otherimages).

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(Top) (A) In a magnetic field at equilibrium, a small excess of protons will align in the "spin up" orientation, creating a net longitudinalmagnetization parallel to the magnetic field. At equilibrium, protons will be randomly distributed around their precessional orbits suchthat there is no net transverse magnetization. (B) Immediately following a 90° radiofrequency (RF) pulse, equal numbers of protons willbe in the spin-up and spin-down orientations, and protons will be aligned in their precessional orbits to create transverse magnetization.(Bottom) A basic spin-echo pulse sequence consists of coordinated RF pulses, gradient coil application, and data acquisition to achieveproton excitation, spatial encoding, signal refocusing, and image generation. The repetition time (TR) is defined as the time intervalbetween consecutive 90° excitation RF pulses. The echo time (TE) is the time interval between excitation and formation of a signalecho. Data acquisition for image generation occurs at the TE.

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Fatty marrow of trabecularbone (high signal)

Cortical bone (low signal)

Synovial fluid (intermediate tohigh signal)

Disc (intermediate to lowsignal) Retrodiscal tissues

(intermediate to high signal)

Air in external auditory canal(low signal)

Posterior attachment(intermediate to low signal)

(Top) T1W images are obtained using pulse sequences with relatively short TR and short TE to create contrast between tissues withdifferent T1 times. Tissues with short T1 times (e.g., fat) will rapidly recover longitudinal magnetization and appear bright in an image.Tissues with long T1 times (e.g., collagenous tissues and tissues with high water content) will appear relatively dark in the image.(Bottom) T1W image of the TMJ in closed-mouth position shows the signal intensities of normal structures primarily due to differencesin T1 relaxation. The cortical bone is low signal (black), and the fatty marrow of the trabecular bone is relatively high signal (bright).The collagenous TMJ disc and attachments are intermediate to low signal (dark gray), and the synovial fluid and retrodiscal tissues areintermediate to high signal (light gray).

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Synovial fluid (high signal)

Disc (intermediate to lowsignal)

Posterior attachment(intermediate to low signal)

Cortical bone (low signal)

Fatty marrow in trabecularbone (intermediate to high

signal)

Air in the external auditorycanal (low signal)

Retrodiscal tissues (highsignal)

(Top) T2W images are obtained using pulse sequences with relatively long TR and long TE to create contrast between tissues withdifferent T2 times. Tissues with short T2 times (e.g., chronic hemorrhage and cartilage) will rapidly lose transverse magnetization andappear dark in the image. Tissues with long T2 times (e.g., fluid) will appear bright in the image. Note that, in clinical T2W fast spin-echo images, fat will appear bright although fat has a relatively short T2 time. (Bottom) T2W fast spin-echo image of a TMJ in theopen-mouth position shows the signal intensities of the normal structures due to T2 decay; cortical bone is low signal (black) and thefatty marrow of the trabecular bone is intermediate to high signal (light gray). The collagenous TMJ disc and the posterior attachmentare intermediate to low signal (dark gray/black), and the synovial fluid and the retrodiscal tissues filling with blood in the open positionare high signal (bright).

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(Left) Short tau inversionrecovery (STIR) pulse sequenceis for nulling signal from fat.STIR pulse sequences beginwith an inversion RF pulse,causing negative longitudinalmagnetization, followed by adelay during whichmagnetization recoversaccording to T1 times. A tissuewith zero longitudinalmagnetization at theexcitation pulse will appearblack in the image. (Right)Sagittal STIR image shows theTMJ with decreased signalfrom fatty tissues, e.g., fattymarrow st, compared with aT1W or T2W image.

(Left) Sagittal T1W image is atthe level of the TMJ . Notethe subtle, hypointense masscompressing the brain superiorto the TMJ st. (Right)Sagittal, fat-saturated (FS),postcontrast T1W image is atthe level of the TMJ . Thereis marked contrastenhancement of the mass stin the brain superior to theTMJ. The mass was diagnosedas a meningioma. This was anincidental finding on a TMDpatient's CBCT scan. Thelesion presented as faintcalcification within the middlecranial fossa on CBCT, whichappears dark on this MR.

(Left) MRs can be acquiredusing 2D or 3D modes ofacquisition. During 2D MR,thin slices of anatomy areexcited and generate images 1at a time. During 3D MR, athick slab of tissue is excitedto generate a volume of imagedata. The 3D volume of data isthen used to create 2D imagesrepresenting very thin (often<1 mm), contiguous slices ofanatomy. (Right) TMJ RFsurface coils are typicallyplaced directly against theskin centered on the patient'sTMJ to acquire images withhigh signal:noise ratio.

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(Left) Six axial images (samepatient) show the appearanceof structures at the level ofthe condyles with differentMR sequences: On T1WI MR,fat is bright, and CSF staround the brainstem is low tointermediate signal. Air and cortical bone st areblack. (Right) Contrast-enhanced (C+) T1WI MR is atthe level of the condyles .Gadolinium-based contrastcauses increased signalintensity of highly vascularstructures, such as the nasalmucosa , and of somepathologic processes, such asmalignancies.

(Left) Axial T2WI fast spin-echo MR is at the level of themandibular condyles . Fat and CSF st are bright. Air and cortical bone st arelow signal intensity. (Right)Axial FLAIR MR at the level ofthe mandibular condyles shows that, on this sequence,normal CSF st and other low-protein fluid will appear oflow signal intensity, butedema and otherproteinaceous fluids willappear bright (notrepresented on this image).

(Left) Axial T2*WI gradient-echo MR is at the level of themandibular condyles . Fluid,such as CSF st, appearsbright, similar to a T2W image.T2*W images are sensitive totissue inhomogeneity, causingdecreased signal or artifact atair-tissue boundaries .T2*W images are useful fordetecting hemorrhage (notpresent in this case). (Right)Proton density-weighted (PD)MR at the level of the condyles shows that PD imagesprovide excellent signal:noiseratio but may be less sensitiveto pathology compared withT2W images.

introduction to mr imaging - university of california, davis...the t1 time (longitudinal recovery)...

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