Contrast-Enhanced Magnetic Resonance Imaging

Introduction

Contrast agents used in magnetic resonance imaging (MRI) can be either endogenous or exogenous. The most common type of endogenous contrast-enhanced MRI is arterial spin labeling (ASL) techniques, which use a dynamic labeling of internal water to measure brain perfusion. ASL is noninvasive, and the sequences needed to acquire this type of imaging are included on most major vendors' platforms. However, ASL still suffers from poor signal-to-noise ratios (SNRs) and limited spatial coverage. Exogenous contrast agents for MRI, first introduced in the early 1980s, can substantially increase image contrast to facilitate diagnosis of lesions in both brain and body imaging. Beyond simply increasing the conspicuity of lesions, exogenous contrast agents are also commonly used to measure tissue properties such as blood perfusion and permeability. Approximately 30 million MRI examinations are performed in the United States annually, and exogenous contrast agents are used to improve image quality in about 30% of cases. Exogenous contrast agents for MRI are administered either by venous injection or orally and usually include gadolinium (Gd)-, iron-, or manganese-based compounds, with Gd-based contrast agents being the most common. The applications of iron- and manganese-based agents are still in early-stage development. In this chapter, we focus on Gd-based exogenous contrast agents and their applications.

Gadolinium (Gd, atomic number 64) is a paramagnetic heavy metal. Although the Gd(III) ion is quite toxic to humans, chelated-Gd(III) compounds are much less toxic and are referred to as Gd-based contrast agents (GBCA). GBCAs are manufactured by a chelating process in which large organic molecules encapsulate the gadolinium. This procedure reduces the chances of toxicity due to free Gd ions because these stable chelated compounds are predominantly eliminated via the kidneys before the compounds degrade and release free Gd ions. The U.S. Food and Drug Administration (FDA) first approved a GBCA for use in MRI in 1988. Since then, eight additional GBCAs have been approved for use in clinical studies. FDA-approved GBCAs are as follows: Ablavar (gadofosveset trisodium), Dotarem (gadoterate meglumine, Gd-DTOA), Eovist (gadoxetate disodium, Gd-EOB-DTPA), Gadavist (gadobutrol, Gd-DO3A-Butriol), Magnevist (gadopentetate dimeglumine, Gd-DTPA), Multihance (gadobenate dimeglumine, Gd-BOPTA), Omniscan (gadodiamide, Gd-DTPA-BMA), Optimark (gadoversetamide, Gd-DTPA-BMEA), and Prohance (gadoteridol, Gd-DO3A-HP). GBCAs can be divided into two groups: ionic and nonionic agents. Ionic contrast agents typically, but not always, have higher osmolality and more side effects. Nonionic contrast agents have lower osmolality and tend to have fewer side effects. The properties of GBCAs are summarized in Table 5-1 .

TABLE 5-1

FDA-Approved Gadolinium-Based Contrast Agents (GBCAs)

Drug Name	Active Ingredients	Nonionic/Ionic	Recommended Dose (mL/kg)	Indication	Children/Adults	Elimination Half-Life (hours)	Elimination Percentage (%)	Year Approved
Ablavar	Gadofosveset trisodium	Ionic	0.12	MRA	A	0.48/16.3 ^a	78.7 ^c	2008
Dotarem	Gadoterate meglumine (Gd-DTOA)	Ionic	0.2	CNS	C/A	1.4 in females, 2.0 in males ^b	72.9 in females, 85.4 in males ^d	2013
Eovist	Gadoxetate disodium (Gd-EOB-DTPA)	Ionic	0.1	Liver	A	0.9 ^b	None	2008
Gadavist	Gadobutrol (Gd-DO3A-butriol)	Nonionic	0.1	CNS	C/A	1.81 ^b	>90 ^e	2011
Magnevist	Gadopentetate dimeglumine (Gd-DTPA)	Ionic	0.2	CNS/ body	C/A	0.2/1.6 ^a	91 ^e	1988
MultiHance	Gadobenate dimeglumine (Gd-BOPTA)	Ionic	0.2	CNS/ MRA	C/A	2.02 ^b	>80 ^e	2004
Omniscan	Gadodiamide (Gd-DTPA-BMA)	Nonionic	0.2	CNS/ body	C/A	0.062/1.30 ^a	95 ^e	1993
Optimark	Gadoversetamide (Gd-DTPA-BMEA)	Nonionic	0.2	CNS/ Liver	C/A	0.23/1.73 ^a	95 ^e	1999
Prohance	Gadoteridol (Gd-DO3A-HP)	Nonionic	0.2	CNS	C/A	0.2/1.57 ^a	94 ^e	1992

CNS, central nervous system; MRA, magnetic resonance angiography.

a Using the two-compartment model;

b using the one-compartment model;

c elimination percentage within 72 hours;

d elimination percentage within 48 hours;

e elimination percentage within 24 hours.

Although GBCAs are considered to be relatively safe in clinical studies, they are not completely without risks. Recently, associations between nephrogenic systemic fibrosis (NSF) and GBCAs have been reported in a small subset of patients. NSF may result in fatal or debilitating fibrosis. Patients with acute, chronic, and severe kidney disease (glomerular filtration rate [GFR] < 30 mL/min/1.73 m ² ) have very poor elimination of GBCAs and could have an increased risk of developing NSF due to GBCAs. Because GBCAs increase the risk of NSF, patients with poor elimination should not be administered such contrast agents. Patients in MRI examination with GBCAs should be properly screened and be closely observed during the procedure. Upon any reaction to the drugs, the appropriate therapy must be used, and any case of NSF should be reported to the FDA (1-800-FDA-1088 or www.fda.gov/medwatch ).

Because GBCAs are paramagnetic, they generate a magnetic moment when placed within a static magnetic field, influencing water protons within their local magnetic field and shortening T1, T2, and T2* relaxation times to enhance image contrast. T1 is the spin-lattice or longitudinal relaxation time, which measures how fast the proton magnetization recovers to its equilibrium position. Shortening T1 increases the recovery speed and further increases MR signal in the tissue. T2 is the spin-spin relaxation time, which represents how fast MR transverse magnetization disappears. T2 shortening increases the rate at which the transverse magnetization disappears and further decreases MR signal in the tissue. T2* is the transverse relaxation time including T2 and magnetic field inhomogeneity effects. Although T2* has an effect on MR signal that is similar to that of T2, T2*'s effect is usually much smaller than that of T2. In MR imaging, spin echo sequences generate T2-weighted images and gradient echo sequences usually generate T2*-weighted images.

GBCAs do not cross an intact blood-brain barrier (BBB) and therefore do not accumulate in normal brain. However, disruption of the BBB allows accumulation of GBCAs in lesions such as neoplasms. Within the body, there is no barrier to hinder extravasation of the smaller GBCAs into the surrounding tissue. The rate and degree to which this extravasation occurs are highly elevated in many neoplasms and can be used to accentuate the conspicuity of the lesions. In this chapter, we will describe the measurement of perfusion and permeability in tissues performed by using dynamic contrast-enhanced MRI (DCE-MRI) and dynamic susceptibility contrast MRI (DSC-MRI).

DCE-MRI and DSC-MRI

Dynamic MRI with GBCA can be used to measure two basic physiologic properties: perfusion and capillary permeability. Perfusion is a faster process and requires a faster acquisition with high temporal resolution to capture the first pass of the bolus of contrast agent while it is still intravascular. DSC-MRI is mostly used to measure perfusion with high temporal and low spatial resolution. Permeability is usually a relatively slower process and requires a longer acquisition time to characterize the slow transfer of GBCA from capillaries to the interstitial space. DCE-MRI is generally used to measure permeability using higher spatial resolution and a longer acquisition time.

DSC-MRI acquires a series of T2*-weighted or T2-weighted dynamic MR images before, during, and immediately after a bolus injection of a GBCA. DSC-MRI is widely used for perfusion measurements in brain imaging, which generates a larger signal change with the same concentration of GBCA than does DCE-MRI; this is because of the higher T2* relaxivity of GBCA. DSC-MRI can be routinely acquired as a relatively low spatial resolution volume data set with a temporal resolution of approximately 1 second and an acquisition time of approximately 60 seconds, making DSC-MRI ideal for the study of first-pass kinetics of the GBCA bolus. However, it is difficult for DSC-MRI to measure permeability because the decrease of signal due to T2* effects counteracts the increase in signal caused by T1 effects during the leakage and thus confounds the signal change.

DCE-MRI acquires a series of T1-weighted dynamic MR images before, during, and after a bolus injection of a GBCA. T2* effects can be minimized by using a short echo time. The T1 relaxation rate has a linear relationship with drug concentration, and the T1 relaxivity is largely independent of tissue type. When first introduced, DCE-MRI had a relatively low temporal resolution (>1 minute per image) because of acquisition sequence design, hardware limitations, and the requirement for higher spatial resolution. However, because of improvements in the acquisition sequences and hardware improvements, DCE-MRI can now be acquired as a volume data set with a temporal resolution of approximately 3 seconds. DCE-MRI can be used to measure perfusion and permeability simultaneously but has to have both a high temporal resolution and a sufficiently long acquisition time of approximately 8 to 10 minutes. DCE-MRI measures of perfusion have a lower contrast-to-noise ratio (CNR) than DSC-MRI does because of its lower T1 relaxivity in comparison with the high T2* relaxivity in DSC-MRI. This drawback can be largely mitigated when imaging highly vascularized cerebral tumors, where high drug concentrations can be reached.

DCE-MRI

A complete DCE-MRI study usually includes baseline T1 mapping, dynamic data acquisition, arterial input function measurement, and dynamic data analysis. In the following sections, we will discuss each part of the whole process. DCE-MRI images can be analyzed by using a heuristic method or a suitable tracer-kinetic model to get the proper physiologic metrics related to perfusion and permeability. For convenience, all physiologic parameters in this chapter are summarized in Table 5-2 for later reference.

TABLE 5-2

Summary of Physiologic Parameters Modeled in DCE- and DSC-MRI

Category	Parameters	Normal Range	Unit	Interpretation
Perfusion	F _p	20-400	mL/min/mL	Plasma flow
Perfusion	v _p	1-30	%	Plasma volume <5% for brain
Mixed	K ^trans	E × F _p	1/min	Volume transfer constant
Mixed	k _ep	K ^trans /v _e	1/min	Interstitium-to-plasma rate constant
Permeability	PS	0.01-0.1	mL/min/100 mL	Permeability–surface area product for normal muscle
	v _e	20-30	%	Interstitial volume fraction (extracellular extravascular space [EES] fraction)
	E	∝PS < 10-20	%	Extraction fraction
	τ	<5	Second	Mean intracellular water molecule lifetime

DCE, dynamic contrast-enhanced; DSC, dynamic susceptibility contrast.

Baseline T10 Measurement

The baseline longitudinal relaxation time T10, before drug injection, is required for pharmacokinetic modeling analysis. Several methods can be used to measure T1. The variable flip angle (VFA) method is one of most widely used methods, calculating a T1 map by using multiple spoiled gradient echo acquisitions with different flip angles. The VFA method allows rapid high-resolution three-dimensional (3D) acquisitions because of the modern implementation with very short repetition time. Radio frequency (RF) field (B1) inhomogeneity leads to variable actual flip angles in different locations even when the same nominal flip angle is prescribed. To ensure more accurate T1 mapping, B1 mapping has to be assessed for correction of errors caused by B1 inhomogeneity. B1 inhomogeneity correction increases the complexity of the data processing and requires additional acquisition time, which diminishes the advantage of the time-efficient VFA method. A potential alternative method for T1 mapping, T1 inversion recovery, uses a single-shot fast-spin echo sequence with multiple inversion of times (TI). These sequences are less sensitive to B1 inhomogeneity and require similar acquisition times. An example of T1 protocols for the VFA method and the inversion recovery method are given in Table 5-3 for readers' reference.

TABLE 5-3

Summary of Major Parameters for T1 Mapping Protocols

Parameters	Variable Flip Angle			Inversion Recovery
B ₀	1.5 T			1.5 T
Sequence	3D spoiled gradient echo			3D single-shot fast spin echo
Vendors	GE	Phillips	Siemens	GE	Phillips	Siemens
Sequence name	SPGR	T1-FFE	FLASH	Single-shot FSE	Single-shot TSE	HASTE
Time of echo (TE)	1.0 ms			78 ms
Time of repetition (TR)	5.0 ms			4000 ms
Flip angles	2°, 5°, 10°, 15°, 20°, 25°, 30°			90°-180°
Time of inversion (TI)	None			100, 300, 900, 1500, 2200, 3300 ms
Slice number	16			16
Slice thickness	5 mm			5 mm
Acquisition matrix	256 × 160 × 16			256 × 128 × 16
Average number	4			1
Partial Fourier	None			6/8 in slice; 5/8 in phase
Time/measures	0.85 minutes			0.83 minutes
Total time	5.97 minutes			5 minutes

DCE-MRI Data Acquisition

DCE-MRI data should be acquired by using a 3D T1-weighted spoiled gradient echo sequence, which is widely available on 1.5 and 3 T scanners from major vendors. For better B1 homogeneity, a body coil should be used as the transmit coil; the receiver coil can be a body or phased array coil, depending on the anatomy being studied. The acquisition volume should cover the whole lesion (or the major lesions) and a feeding vessel with in-plane flow for measuring arterial input function (AIF) if feasible. The ideal echo time (TE) should be less than 1.5 milliseconds to improve SNR and to diminish the T2* effects on signal change. The repetition time (TR) should be approximately 3 to 5 milliseconds to achieve higher temporal resolution and optimal T1 quantification. The temporal resolution should be less than 10 seconds for quantifying perfusion. A small flip angle could cause saturation effects when used with higher tracer concentrations. Therefore 25- to 30-degree flip angles are recommended to minimize saturation effects and balance specific absorption rate (SAR), which measures the amount of energy deposited into the human body by radiofrequency pulses. Faster acquisition techniques, including partial Fourier and parallel imaging, can be used with additional attention to ensure consistency on all scanners. The total acquisition time should be no less than 5 minutes; the optimal is about 8 to 10 minutes to ensure the accuracy of estimating kinetic parameters related to the relatively slow process of permeability. An example of a DCE-MRI protocol is shown in Table 5-4 and can be modified for specific applications. Additional details and sample protocols can be found in a workshop report by the Imaging Committee of the Experimental Cancer Medicine Centres (ECMC) and in the Quantitative Imaging Biomarkers Alliance (QIBA) protocols and profiles.

TABLE 5-4

An Example DCE-MRI Protocol

Parameters	DCE-MRI
B ₀	1.5 T
Sequence	3D spoiled gradient echo
Vendors	GE	Phillips		Siemens
Sequence name	SPGR	T1-FFE		FLASH
Time of echo (TE)	1.2 ms
Time of repetition (TR)	3.5 ms
Flip angles	30°
Parallel imaging	GRAPPA R=2, Nref=24
Slice number	16
Slice thickness	5 mm
Acquisition matrix	256 × 192 × 16
Bandwidth	360 Hz/pixel
Partial Fourier	6/8 in slice and phase
Time/measures	7 seconds
Measurement number	50 (minimum)		80 (optimal)
Total time	5:50 min		9:20 min

Contrast injection should start after at least five measurements have been acquired to allow for a robust nonenhanced baseline reference. The preferred method for contrast agent delivery is a bolus-like injection consisting of a fast constant rate injection of generally 2 to 4 mL/s to ensure that the most meaningful perfusion information is obtained in the early enhancement phase. Because of the relatively small volumes of GBCA being administered, the injection should be immediately followed by a saline flush of 20 to 40 mL administered at the same rate as the bolus to ensure that all contrast agent is cleared from the line. Although a manual bolus injection can be used, a power injector synchronized with the MRI scanner should be used for the most consistent results.

MRI Signal vs. Tracer Concentration

MRI signal is measured in arbitrary units and varies substantially with many factors, including the scanners, coils, field inhomogeneity, and imaging location. Therefore MRI signal is not a suitable measure with which to directly quantify physiologic parameters in large clinical studies. However, MRI signal can be converted to tracer concentration for further modeling analyses. The first step in this conversion is to convert the gradient echo signal into T1 relaxation rate (R ₁ = 1/T1). The second step is to calculate the tracer concentration by using an assumed linear relationship: R ₁ = R ₁₀ + r ₁ ⋅ C, where the baseline R ₁₀ = 1/T10 is measured in a precontrast scan as described previously, r ₁ is the T1 relaxivity of the specific GBCA administered, and C is the GBCA concentration. It has been demonstrated that the difference in T1 relaxation rates (ΔR ₁ = R ₁ − R ₁₀ ) is approximately linearly proportional to the GBCA concentration. Most in vivo studies use an r ₁ of approximately 4.0 kg/s/mmol, which is a value originally measured in cartilage. However, r ₁ has also been reported to vary in different anatomic locations in animal models for different agents. For example, r ₁ for Gd-EOB-DTPA changes from 2.0 in kidney to 10.7 in liver, and r ₁ for Gd-DTPA changes from 1.2 in kidney to 4.8 in liver. Additional care should be taken when using a constant r ₁ value in studies that may involve various types of tissue.

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