**2. Basics of MRI**

Before reviewing the work that has been done with MR, a brief overview of the theory be‐ hind MR should be covered. MR physics can essentially be understood using principles of classical physics, however for a more comprehensive understanding the reader is directed to an excellent review by Plewes and Kucharzck [4,10]. Briefly, MR takes advantage of the be‐ havior of a system of protons in the presence of a magnetic field and how this behavior changes based on the micro- and macroscopic environment. Magnetic strength is generally reported in units of Tesla (T), and MRI scanners have very high field strengths. In human research, 1.5T, 3T, and 7T scanners are commonly used, though 1.5T and 3T scanners pre‐ dominate the clinical setting. The magnet's strength and direction is represented by the vec‐ tor B0 (see Figure 1), and lies along the Z-axis (generally from foot to head).

Due to the large amount of water that constitutes tissue (~80-99% depending on tissue type), most MR is specifically focused on the protons on water molecules. Protons have an intrinsic spin that in nature is oriented randomly. In the presence of a magnetic field however, these spins align themselves on average parallel or antiparallel to the axis of the field (Figure 1a). The number of protons aligned parallel to the field is very slightly larger than the number of protons aligned antiparallel, and it is this difference that produces the net magnetization vector in a voxel. When a radiofrequency (RF) pulse is applied at the proper frequency (Lar‐ mor frequency), the longitudinal (z) component of the magnetization vector is tipped away from the axis of the main magnetic field, but continues to spin around the longitudinal axis or "precess" (Figure 1b). When the pulse is removed, the longitudinal component of the magnetization vector will realign itself with the field with a unique time constant that varies with the local environment. commonly used, though 1.5T and 3T scanners predominate the clinical setting. The magnet's strength and direction is represented by the vector B0 (see Figure 1), and lies along the Z-axis (generally from foot to head). Due to the large amount of water that constitutes tissue (~80-99% depending on tissue type), most MR is specifically focused on the protons on water molecules. Protons have an intrinsic spin that in nature is oriented randomly. In the presence of a magnetic field

ing a difficult task, performing MR work in advanced AD cases is quite challenging and as such, most studies are limited to mild and very mild cases. While these studies are typically performed at a time when pathology is irreversible, the results of this work point to changes that may be apparent before cognitive decline has become clinically ap‐ parent. For this reason, studies that examine differences between people who will even‐ tually develop AD and people who will not develop AD provide insight into both the

It is impossible to predict with certainty who will develop AD, but there are several fac‐ tors that increase the risk. These at-risk populations include individuals in the prodromal stage of AD, termed amnestic mild cognitive impairment (aMCI), and people at a genetic risk for developing AD. A diagnosis of aMCI indicates that there is more memory de‐ cline than would be expected based on the person's age and education level, however memory impairment is not interfering with daily activities. It is estimated that 10-20% of people 65 and older have aMCI, and out of those 10-15% will progress to develop AD in 3-4 years. [8,9] Because approximately 30% of people diagnosed with aMCI will remain stable or improve over time, it is important to find biomarkers that will identify those

This chapter will focus on the use of MR in the early detection of AD. Major advances have been made in structural imaging of both gray and white matter using proton density, T1 and T2- weighted imaging, and Diffusion Tensor Imaging (DTI). Functional imaging in AD will also be reviewed, and Blood Oxygenation Level-Dependent (BOLD) functional Magnet‐ ic Resonance Imaging (fMRI) will be broken down into its primary contributors: Cerebral Blood Flow (CBF), Cerebral Blood Volume (CBV), and the Cerebral Metabolic Rate of Oxy‐ gen (CMRO2). Finally, hemodynamic fMRI contrast can be complemented using measures of neurochemistry, including measuring the balance between excitatory (glutamatergic) and inhibitory (γ-aminobutyric acid; GABAergic) neurotransmission. This can be achieved with new single-voxel chemical imaging techniques such as Magnetic Resonance Spectroscopy

Before reviewing the work that has been done with MR, a brief overview of the theory be‐ hind MR should be covered. MR physics can essentially be understood using principles of classical physics, however for a more comprehensive understanding the reader is directed to an excellent review by Plewes and Kucharzck [4,10]. Briefly, MR takes advantage of the be‐ havior of a system of protons in the presence of a magnetic field and how this behavior changes based on the micro- and macroscopic environment. Magnetic strength is generally reported in units of Tesla (T), and MRI scanners have very high field strengths. In human research, 1.5T, 3T, and 7T scanners are commonly used, though 1.5T and 3T scanners pre‐ dominate the clinical setting. The magnet's strength and direction is represented by the vec‐

tor B0 (see Figure 1), and lies along the Z-axis (generally from foot to head).

(MRS), or more recently using multi-voxel MRS imaging (MRSi)

cause and the physiology of the disease.

226 Understanding Alzheimer's Disease

most likely to progress to AD.

**2. Basics of MRI**

Manipulating the timing of the RF pulses controls the magnetization and creates the desired contrast. The most fundamental timing parameters of relevance are repetition time (TR), echo time (TE), and in some cases inversion time (TI). TR is the time between consecutive acquisitions, and TE is the time from the onset of the excitation pulse that is used for prepar‐ ing the signal for detection to the signal refocusing and in most cases acquisition. In an in‐ version recovery pulse sequence, TI refers to the time between the inversion pulse and the excitation pulse. Importantly, simply by manipulating the timing of the above parameters a range of MR contrasts can be obtained with varying sensitivity to different tissue types. A simple pulse sequence indicating RF and gradient timing is illustrated in Figure 1c. however, these spins align themselves on average parallel or antiparallel to the axis of the field (Figure 1a). The number of protons aligned parallel to the field is very slightly larger than the number of protons aligned antiparallel, and it is this difference that produces the net magnetization vector in a voxel. When a radiofrequency (RF) pulse is applied at the proper frequency (Larmor frequency), the longitudinal (z) component of the magnetization vector is tipped away from the axis of the main magnetic field, but continues to spin around the longitudinal axis or "precess" (Figure 1b). When the pulse is removed, the longitudinal component of the magnetization vector will realign itself with the field with a unique time constant that varies with the local environment. Manipulating the timing of the RF pulses controls the magnetization and creates the desired contrast. The most fundamental timing parameters of relevance are repetition time (TR), echo time (TE), and in some cases inversion time (TI). TR is the time between consecutive acquisitions, and TE is the time from the onset of the excitation pulse that is used for preparing the signal for detection to the signal refocusing and in most cases acquisition. In an inversion recovery pulse sequence, TI refers to the time between the inversion pulse and the excitation pulse. Importantly, simply by manipulating the timing of the above parameters a range of MR contrasts can be obtained with varying sensitivity to different tissue types. A simple pulse sequence indicating RF and

gradient timing is illustrated in Figure 1c.

(1) ܯ ή ܯήܥןܵ

Figure 1. Physics underlying magnetic resonance. Hydrogen atoms align parallel and antiparallel to a strong magnetic field, producing a net magnification vector (a). When a radio frequency pulse is applied at the Larmor frequency, the net magnetization vector is tipped away from the main magnetic field (b). Example of a simple pulse sequence showing timing parameters of the application of radio frequency pulse (RF), the onset of gradients in the Z direction (Gz), and the timing of signal acquisition (Signal) (c). In most cases, the detectible signal (S) that is measured in MRI is a combination of three primary factors: water proton density (C; **Figure 1.** Physics underlying magnetic resonance. Hydrogen atoms align parallel and antiparallel to a strong magnetic field, producing a net magnification vector (a). When a radio frequency pulse is applied at the Larmor frequency, the net magnetization vector is tipped away from the main magnetic field (b). Example of a simple pulse sequence show‐ ing timing parameters of the application of radio frequency pulse (RF), the onset of gradients in the Z direction (Gz), and the timing of signal acquisition (Signal) (c).

ml tissue /100 ml water ), magnetization in the longitudinal plane (MZ) and magnetization in the transverse plane (MXY):

The two major contrasts or "weighting" that are used for generating contrast are T1 and T2. T1 and T2 are independent measures and reflect different properties of the tissue of interest, with T1 governing the Mz term and T2 the Mxy term in Eq. 1 above. The time it takes for the magnetization to realign itself longitudinally is measured using T1 weighting (Figure 2a), and is achieved with a short TR and a short TE sequence. T1 is a constant that is unique for each tissue type and is equal to the point when 63% of longitudinal magnetization is recovered (Figure 2b). At the times selected for T1 imaging, there is a high amount of contrast between gray and white matter and therefore T1 weighted imaging is useful for viewing structural changes in the brain (Figure 2c).

In most cases, the detectible signal (S) that is measured in MRI is a combination of three pri‐ mary factors: water proton density (C; ml tissue /100 ml water ), magnetization in the longi‐ tudinal plane (MZ) and magnetization in the transverse plane (MXY):

$$S \propto \mathbb{C} \bullet M\_Z \bullet M\_{XY} \tag{1}$$

Note that in the absence of a prepulse (α=0), the TR determines the T1-weighting. When the RF pulse is applied, individual protons will also precess in synchrony in the transverse plane. When the pulse is removed, the protons will lose that synchrony or dephase, which results in a reduced MXY (Figure 3a). This is referred to as T2 decay. Like T1, the T2 time constant is unique for each tissue (Figure 3b). Unlike T1, T2 weighting is achieved with a

(Figure 3b). Unlike T1, T2 weighting is achieved with a long TE and long TR.

is a constant that is unique to each tissue type (b). Example of T2 weighted images (c).

takes into account both the T1 and the T2 properties of the tissue:

*S* ∝*C* ∙(1 - *αe*

**Figure 3.** T2 weighted imaging. Protons lose synchrony after removal of an RF pulse (a). The amount of time it takes for protons to lose synchrony is a constant that is unique to each tissue type (b). Example of T2 weighted images (c).

The MXY component of MRI is based on pulse timing as well as the T2 constant of the tissue

The three equations can be combined to form one overall equation for the MR signal that


T1 and T2 components are each present whenever a proton is flipped out of alignment, but by manipulating the pulse sequences one can contribute to the signal more than the other. This is referred to as weighting. If neither the T1 nor the T2 signal contributes strongly to the signal, only the C component is left. These images are referred to as proton density images.


following a pre-pulse with flip angle, , is given by:

భ ሻ (2)

భ ݁షೃ

for each tissue type based on the magnetic field strength that is applied (b). Example T1 weighted images (c).

Figure 2. T1 weighted imaging. After removal of an RF pulse the magnetization vector recovers longitudinally (a). The recovery time is a constant

http://dx.doi.org/10.5772/54445

229

Using Magnetic Resonance Imaging in the Early Detection of Alzheimer's Disease

The MZ component of the magnetization vector is based on pulse timing as well as the T1 of tissue, and for magnetization

Note that in the absence of a prepulse (=0), the TR determines the T1-weighting. When the RF pulse is applied, individual protons will also precess in synchrony in the transverse plane. When the pulse is removed, the protons will lose that synchrony or dephase, which results in a reduced MXY (Figure 3a). This is referred to as T2 decay. Like T1, the T2 time constant is unique for each tissue

Figure 3. T2 weighted imaging. Protons lose synchrony after removal of an RF pulse (a). The amount of time it takes for protons to lose synchrony

The three equations can be combined to form one overall equation for the MR signal that takes into account both the T1 and the T2

*MXY* <sup>=</sup>*e*-*TE*/*<sup>T</sup>* <sup>2</sup> (3)

*<sup>T</sup>* <sup>1</sup> ) ∙*e*-*TE*/*<sup>T</sup>* <sup>2</sup> (4)

The MXY component of MRI is based on pulse timing as well as the T2 constant of the tissue area, and can be written:

long TE and long TR.

<sup>ܯ</sup> ൌ ሺͳ െ ߙ݁ష

ܯ ൌ ݁ି்ாȀ்ଶ (3)

area, and can be written:

properties of the tissue:

The two major methods or "weighting" that are used for generating contrast are T1 and T2. T1 and T2 are independent measures and reflect different properties of the tissue of interest, with T1 governing the Mz term and T2 the Mxy term in Eq. 1 above. The time it takes for the magnetization to realign itself longitudinally is measured using T1 weighting (Figure 2a), and is achieved with a short TR and a short TE sequence. T1 is a constant that is unique for each tissue type and is equal to the point when 63% of longitudinal magnetization is recov‐ ered (Figure 2b). At the times selected for T1 imaging, there is a high amount of contrast be‐ tween gray and white matter and therefore T1 weighted imaging is useful for viewing structural changes in the brain (Figure 2c).

Figure 2. T1 weighted imaging. After removal of an RF pulse the magnetization vector recovers longitudinally (a). The recovery time is a constant for each tissue type based on the magnetic field strength that is applied (b). Example T1 weighted images (c). The MZ component of the magnetization vector is based on pulse timing as well as the T1 of tissue, and for magnetization **Figure 2.** T1 weighted imaging. After removal of an RF pulse the magnetization vector recovers longitudinally (a). The recovery time is a constant for each tissue type based on the magnetic field strength that is applied (b). Example T1 weighted images (c).

following a pre-pulse with flip angle, , is given by: <sup>ܯ</sup> ൌ ሺͳ െ ߙ݁ష భ ݁షೃ భ ሻ (2) The MZ component of the magnetization vector is based on pulse timing as well as the T1 of tissue, and for magnetization following a pre-pulse with flip angle, α, is given by:

(Figure 3b). Unlike T1, T2 weighting is achieved with a long TE and long TR.

is a constant that is unique to each tissue type (b). Example of T2 weighted images (c).

ܯ ൌ ݁ି்ாȀ்ଶ (3)

properties of the tissue:

$$M\_Z = \begin{pmatrix} 1 \ -ae \ \stackrel{\cdot n}{\dashrightarrow} + e^{\frac{\cdot n}{\dashrightarrow}} \end{pmatrix} \tag{2}$$

which results in a reduced MXY (Figure 3a). This is referred to as T2 decay. Like T1, the T2 time constant is unique for each tissue

Figure 3. T2 weighted imaging. Protons lose synchrony after removal of an RF pulse (a). The amount of time it takes for protons to lose synchrony

The three equations can be combined to form one overall equation for the MR signal that takes into account both the T1 and the T2

The MXY component of MRI is based on pulse timing as well as the T2 constant of the tissue area, and can be written:

Note that in the absence of a prepulse (α=0), the TR determines the T1-weighting. When the RF pulse is applied, individual protons will also precess in synchrony in the transverse plane. When the pulse is removed, the protons will lose that synchrony or dephase, which results in a reduced MXY (Figure 3a). This is referred to as T2 decay. Like T1, the T2 time constant is unique for each tissue (Figure 3b). Unlike T1, T2 weighting is achieved with a long TE and long TR. The MZ component of the magnetization vector is based on pulse timing as well as the T1 of tissue, and for magnetization following a pre-pulse with flip angle, , is given by: <sup>ܯ</sup> ൌ ሺͳ െ ߙ݁ష భ ݁షೃ భ ሻ (2) Note that in the absence of a prepulse (=0), the TR determines the T1-weighting. When the RF pulse is applied, individual protons will also precess in synchrony in the transverse plane. When the pulse is removed, the protons will lose that synchrony or dephase, which results in a reduced MXY (Figure 3a). This is referred to as T2 decay. Like T1, the T2 time constant is unique for each tissue

(Figure 3b). Unlike T1, T2 weighting is achieved with a long TE and long TR.

In most cases, the detectible signal (S) that is measured in MRI is a combination of three pri‐ mary factors: water proton density (C; ml tissue /100 ml water ), magnetization in the longi‐

The two major methods or "weighting" that are used for generating contrast are T1 and T2. T1 and T2 are independent measures and reflect different properties of the tissue of interest, with T1 governing the Mz term and T2 the Mxy term in Eq. 1 above. The time it takes for the magnetization to realign itself longitudinally is measured using T1 weighting (Figure 2a), and is achieved with a short TR and a short TE sequence. T1 is a constant that is unique for each tissue type and is equal to the point when 63% of longitudinal magnetization is recov‐ ered (Figure 2b). At the times selected for T1 imaging, there is a high amount of contrast be‐ tween gray and white matter and therefore T1 weighted imaging is useful for viewing

for each tissue type based on the magnetic field strength that is applied (b). Example T1 weighted images (c).

(Figure 3b). Unlike T1, T2 weighting is achieved with a long TE and long TR.


**Figure 2.** T1 weighted imaging. After removal of an RF pulse the magnetization vector recovers longitudinally (a). The recovery time is a constant for each tissue type based on the magnetic field strength that is applied (b). Example T1

The MZ component of the magnetization vector is based on pulse timing as well as the T1 of

is a constant that is unique to each tissue type (b). Example of T2 weighted images (c).

following a pre-pulse with flip angle, , is given by:

*MZ* =(1 - *<sup>α</sup><sup>e</sup>*

tissue, and for magnetization following a pre-pulse with flip angle, α, is given by:

భ ݁షೃ భ ሻ (2)

The MZ component of the magnetization vector is based on pulse timing as well as the T1 of tissue, and for magnetization

which results in a reduced MXY (Figure 3a). This is referred to as T2 decay. Like T1, the T2 time constant is unique for each tissue

*<sup>T</sup>* <sup>1</sup> ) (2)

Figure 3. T2 weighted imaging. Protons lose synchrony after removal of an RF pulse (a). The amount of time it takes for protons to lose synchrony

The three equations can be combined to form one overall equation for the MR signal that takes into account both the T1 and the T2

The MXY component of MRI is based on pulse timing as well as the T2 constant of the tissue area, and can be written:

*S* ∝*C* ∙*MZ* ∙*MXY* (1)

tudinal plane (MZ) and magnetization in the transverse plane (MXY):

structural changes in the brain (Figure 2c).

228 Understanding Alzheimer's Disease

<sup>ܯ</sup> ൌ ሺͳ െ ߙ݁ష

weighted images (c).

ܯ ൌ ݁ି்ாȀ்ଶ (3)

properties of the tissue:

Figure 3. T2 weighted imaging. Protons lose synchrony after removal of an RF pulse (a). The amount of time it takes for protons to lose synchrony is a constant that is unique to each tissue type (b). Example of T2 weighted images (c). **Figure 3.** T2 weighted imaging. Protons lose synchrony after removal of an RF pulse (a). The amount of time it takes for protons to lose synchrony is a constant that is unique to each tissue type (b). Example of T2 weighted images (c).

The MXY component of MRI is based on pulse timing as well as the T2 constant of the tissue area, and can be written: ܯ ൌ ݁ି்ாȀ்ଶ (3) The MXY component of MRI is based on pulse timing as well as the T2 constant of the tissue area, and can be written:

$$M\_{XY} \triangleq e^{-TE/T\,2} \tag{3}$$

Figure 2. T1 weighted imaging. After removal of an RF pulse the magnetization vector recovers longitudinally (a). The recovery time is a constant The three equations can be combined to form one overall equation for the MR signal that takes into account both the T1 and the T2 properties of the tissue:

properties of the tissue:

$$S \propto C \bullet \begin{pmatrix} 1 \ -ae^{\frac{\cdot \cdot \cdot}{T \cdot 1}} + e^{\frac{\cdot \cdot \cdot \cdot}{T \cdot 1}} \end{pmatrix} \bullet e^{-TE/T2} \tag{4}$$

Note that in the absence of a prepulse (=0), the TR determines the T1-weighting. When the RF pulse is applied, individual protons will also precess in synchrony in the transverse plane. When the pulse is removed, the protons will lose that synchrony or dephase, T1 and T2 components are each present whenever a proton is flipped out of alignment, but by manipulating the pulse sequences one can contribute to the signal more than the other. This is referred to as weighting. If neither the T1 nor the T2 signal contributes strongly to the signal, only the C component is left. These images are referred to as proton density images.
