Practical MR physics: and case file of MR artifacts and pitfalls

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The simplest tagging pulse consists of two rf pulses either side of a single modulating gradient[ 37 ]. The addition of more rf pulses and modulating gradients makes the line pattern sharper[ 38 ]. Immediately after the tagging preparation scheme, a cine image data acquisition is then performed using a fast cine gradient echo pulse sequence to readout the signal at multiple time points throughout the cardiac cycle.

In this example a composite binomial rf pulse is used consisting of three rf pulses with amplitudes in the ratio of Two modulating gradients are applied in the spaces between the rf pulses to de-phase modulate the transverse magnetisation between each rf pulse. The net effect is to cause a variation or modulation of the z-magnetisation, creating a series of parallel lines of tissue with a magnetisation that varies alternately between its equilibrium value untagged and zero tagged.

A spoiler gradient, S, is applied to destroy transverse magnetisation generated by the tagging pulses. T1 relaxation causes the magnetisation of the tagged lines to recover towards equilibrium, while at the same time the magnetisation of the untagged tissue becomes partially saturated by the rf pulses applied as part of the cine imaging sequence.

This causes the contrast between the tagged and untagged lines to reduce as the cardiac cycle progresses. The two short-axis images in b are acquired from separate tagged cine acquisitions with tagging applied in two different directions. The arrows indicate the direction of the modulating gradient in each case. Both image examples correspond to a cardiac phase at around end-systole. For stationary tissue, such as in the chest wall, the tagged pattern has remained fixed and is seen as a series of parallel lines.

:: iMRI :: Investigative Magnetic Resonance Imaging

Within the left ventricle, the line pattern has deformed as it follows the motion of the myocardial muscle. This diagram shows how a SPAMM pulse can produce a magnetisation pattern consisting of a series of parallel lines that are alternately fully magnetised and fully saturated, appearing as bright and dark lines shown on the right. In this example, the composite rf pulse is a binomial pulse, consisting of three rf pulses with relative amplitudes in the ratio of Starting at equilibrium a , the first rf pulse causes all spins to flip through The third On the first image of the cine series immediately after the tagging pulse , the magnetisation pattern appears as a series of low-signal-intensity parallel lines across the image where the magnetisation has been saturated.

As the heart contracts through systole the magnetisation pattern deforms as it follows the contraction of the myocardial muscle Figure 6 b.

As the pattern is generated through saturation of the tissue magnetisation, T1-relaxation causes the magnetisation to return towards its equilibrium value. At the same time the tissue magnetisation at equilibrium becomes partially saturated by the rf pulses applied as part of the cine imaging sequence.

These two effects cause the magnetisation of the tagged and untagged tissue to converge, resulting in a rapid loss of contrast for the tagged lines and fading of the tagging pattern during the cardiac cycle.

The rate at which contrast is lost can be reduced by ensuring a low flip angle is used for the cine gradient echo pulse sequence, and by limiting the number of cine frames. Typically two line patterns are generated at right angles to form a grid pattern[ 38 ]. This can be done by using two tagging preparation pulses within the same acquisition known as grid tagging , or by performing two separate acquisitions with line tagging at right angles, and subsequently combining the two data sets as a post-processing step.

SPAMM is the most well established of the methods used to perform CMR tagging and it is derivatives of this basic method that are implemented by the MR vendors for use in routine clinical practice, with visual assessment being the main method of analysis. There have been many further developments of CMR tagging techniques in the research domain, together with methods used to analyse the tagged images, and many of these are described in the recent review[ 13 ].

MR contrast agents work by modifying the tissue properties that most directly affect image contrast appearances, namely the T1 and T2 relaxation times. The most commonly-used contrast agents exploit a property of the lanthanide ion gadolinium Gd known as paramagnetism. This property exists due to the presence of unpaired electrons in the outer shell of a metal ion, which cause it to become temporarily magnetised when in an externally applied magnetic field creating local magnetic fields over a short range. Gadolinium is particularly strongly paramagnetic as it has seven unpaired electrons in its outer shell, the most of any element.

Local field interactions between the unpaired electrons of the Gd ion and the hydrogen nuclei within adjacent water molecules cause a reduction in both the T1 and T2 of the surrounding tissue. In order for this naturally toxic element to be suitable for use in human subjects, the Gd ion is bound or chelated to a larger electron-donating molecule or ligand.

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This renders the gadolinium safe for in-vivo use in most circumstances although gadolinium-based contrast agents are contraindicated for use in patients with impaired renal function due to their association with nephrogenic systemic fibrosis NSF [ 39 — 41 ]. The ability of a given contrast agent to influence relaxation rates is expressed in terms of its relaxivity which is the change in relaxation rate per unit concentration expressed in mM The higher the value of the relaxivity, the greater is the T1-reducing effect of the contrast medium.

If the concentration in mM of contrast agent is C and the T1 relaxivity is r 1 then the observed relaxation rate of the tissue T 1 observed can be related to its native relaxation rate T 1 native as follows:. There is a corresponding expression for the observed T2 relaxation rate of the substance T 2 observed as follows:. Where r 2 is the T2 relaxivity and T 2 native is the native relaxation rate. Relating T 1 observed to the final image signal intensity SI value is more complicated.

Figure 8 describes a typical plot of SI versus contrast agent concentration. At low concentrations T1 shortening is the dominant effect of the contrast agent so that the SI increases with increasing concentration. However at higher concentrations the T2 shortening effect becomes dominant and SI begins to fall due to the reduction of T2 to very low values. If the purpose of the administered contrast agent is simply to enhance certain structures in the image, as in MRA investigations see later , then the administered dose is designed to maximise the SI and seeks to produce an in-vivo Gd concentration corresponding to the peak in Figure 8.

However if the images are to be used for quantitative analysis then the contrast-induced changes in SI must directly reflect changes in Gd concentration. At low concentrations this assumption holds because the relationship between SI and Gd concentration is approximately linear.

12222 Classic Lectures in Body Imaging with MR

At higher concentrations this relationship becomes non-linear and quantitation of concentration based on uncorrected SI values will yield erroneous results. This is often referred to as signal saturation in quantitative MRI, not to be confused with the saturation of magnetisation caused by the 90 o pulse that generates signal in MRI. It is more helpfully described as a non-linearity effect, generated by a breakdown of the assumption that Gd concentration and SI are linearly related.

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This is discussed in more detail in the next section. The relationship between signal intensity and concentration. Signal intensity values over a range of concentrations for a spoiled gradient echo pulse sequence. The two dashed curves show the separate dependencies of the signal behaviour for T1 or T2 alone. The solid line shows the combined effect of T1 or T2 on signal intensity. At low concentrations the effect of T1 shortening is dominant, while at higher concentrations the T2 shortening effect becomes the dominant factor. A series of samples imaged with increasing percentage concentrations of Gadolinium are shown underneath the plot as a visual demonstration of the effect.

Extra-vascular, extra-cellular contrast agents are most commonly used in clinical practice. These agents are small enough to leak through the capillaries from the vascular space into the extra-vascular, extra-cellular space but not through cell membranes. It is this property of the contrast agent that enables late gadolinium enhancement of myocardial infarcts where the extravascular, extracellular space is enlarged see later. Intravascular contrast agents, which stay within the vascular space, are less commonly used but may be preferable for quantitative perfusion imaging as they allow simpler mathematical models to be used for flow quantification as no account needs to be taken of leakage from the vascular space[ 42 ].

Myocardial perfusion imaging assesses the blood supply to the myocardium and plays an increasing role in the diagnosis of ischemic heart disease[ 43 ]. In this section dynamic contrast enhanced MRI DCE-MRI is introduced and the challenging requirements for performing it in the heart are described with reference to pulse sequences described in previous sections.

Practical MR Physics

A discussion of the necessary trade-offs that should be considered when designing a perfusion imaging protocol is given. In order to assess myocardial perfusion, blood passing into the myocardium needs to alter image signal intensity so that areas of reduced perfusion can be detected. This is typically achieved using a signal enhancing contrast agent. The contrast agent is injected intravenously whilst multiple images of the heart in the same anatomical position and the same point in the cardiac cycle are acquired in successive heart beats.

Typically short-axis images are acquired but a long axis image is also sometimes additionally acquired in order to cover the apex of the heart.

As DCE-MRI generates a movie showing the passage of contrast agent through the body there is a compromise between spatial requirements i. For myocardial perfusion imaging MPI , the need to minimise the effects of cardiac motion adds further constraints, making this application of DCE-MRI one of the most challenging. In the ideal case, the dynamic series of images, when viewed as a movie, should be able to demonstrate a motion-free cross-section of cardiac tissue whose signal intensity rises and falls over time as the bolus of contrast agent passes through the myocardial tissue Figure 9.

Dynamic contrast enhanced cardiac perfusion imaging. Contrast agent is injected intravenously whilst multiple images of the heart are acquired to create a movie showing the contrast agent passing through the heart. Contrast agent can be seen as signal enhancement in the right ventricular cavity RV followed by the left ventricular cavity LV and more gradually in the myocardium, before finally washing out.

The essential requirements of a DCE-MRI cardiac perfusion imaging sequence can therefore be summarised as follows: All data for multiple images must be acquired within a single heart beat and the effects of cardiac and respiratory motion must be minimised. In addition the image contrast must be T1-weighted to maximise the effect of the contrast agent on image signal intensity. In order to fulfil these requirements, the choice of pulse sequence, method of contrast generation and approaches to minimise motion effects must be carefully considered.

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In DCE-MRI the image appearance changes significantly between contiguous frames due to the passage of the contrast agent through the heart so multi-shot k-space imaging strategies that fill k-space over multiple cardiac cycles[ 1 ] are not applicable. All three sequences are fast, having a very short TR partly by virtue of the fact that they avoid the need to wait for remnant transverse magnetization to decay after read-out of the MR signal echo.

The remnant transverse magnetization is then superimposed onto the magnetization generated by subsequent rf pulses, generating high signal from fluid and blood. Conversely the FGE sequence uses spoiler gradients to destroy remnant transverse magnetization after each readout. EPI uses rapidly alternating frequency encoding gradients, interspersed by phase encoding pulses, to refocus multiple gradient echoes following a single rf-pulse.

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