MRI Measures of Aging
Methodological Issues
MRI 測量老化的方法論問題

Magnetic resonance imaging (MRI) is one of the most widely used imaging techniques in studies of cognitive neuroscience, including in brain aging. Compared to other medical imaging modalities, MRI has three important features that have contributed to this broad acceptance. First, MRI does not involve ionizing radiation and, as a matter of fact, most MRI scans do not require any exogenous contrast agents. Therefore, it is particularly suitable for studies on healthy participants in whom injection of exogenous tracer or agent is not desirable. For similar reasons, one can repeat the MRI scan as frequently as needed in a follow-up or longitudinal setting. Second, MRI provides an excellent spatial resolution, allowing the visualization of the brain with exceptional details. It also shows clear image contrast between soft tissues, unlike several other imaging technologies such as X-ray and computertomography (CT), thus can easily distinguish the gray and white matter based on their characteristic MR properties. Finally, MRI is also a versatile imaging modality. Within the same imaging session, one can perform several MRI pulse sequences and therefore obtain multiple domains of structural, functional, and physiological information from the participant. This advantage reduces subject burden and allows the integration of multi-parametric datasets in understanding cerebral aging. This chapter will provide a review of the basic principles of MRI, describe several major MRI techniques that are commonly used in cerebral aging, and introduce new, emerging techniques that are on the horizon.
磁共振成像（MRI）是認知神經科學研究中最廣泛使用的成像技術之一，包括大腦老化的研究。與其他醫學成像方式相比，MRI 有三個重要特徵促成了其廣泛的接受度。首先，MRI 不涉及電離輻射，事實上，大多數 MRI 掃描不需要任何外源性對比劑。因此，它特別適合於對健康參與者的研究，因為不希望注射外源性示蹤劑或藥劑。出於類似的原因，在後續或縱向研究中，可以根據需要頻繁重複 MRI 掃描。第二，MRI 提供了優秀的空間解析度，能夠以卓越的細節可視化大腦。它還顯示出軟組織之間的清晰影像對比，與其他幾種成像技術（如 X 光和電腦斷層掃描（CT））不同，因此可以輕鬆區分灰質和白質，基於其特徵性的 MRI 屬性。最後，MRI 也是一種多功能的成像方式。 在同一次影像檢查中，可以執行多個 MRI 脈衝序列，因此可以從參與者那裡獲得多個結構、功能和生理信息的領域。這一優勢減少了受試者的負擔，並允許在理解大腦老化時整合多參數數據集。本章將回顧 MRI 的基本原理，描述幾種在大腦老化中常用的主要 MRI 技術，並介紹一些即將出現的新技術。

The source of MRI signal is from the endogenous nuclei (e.g., ${}^1\mathrm{H},{}^{23}\mathrm{Na},{}^{31}\mathrm{P},{}^{17}\mathrm{O}$${}^1\mathrm{H},{}^{23}\mathrm{Na},{}^{31}\mathrm{P},{}^{17}\mathrm{O}$^(1)H,^(23)Na,^(31)P,^(17)O{ }^{1} \mathrm{H},{ }^{23} \mathrm{Na},{ }^{31} \mathrm{P},{ }^{17} \mathrm{O}, and ${}^{19}\mathrm{F}$${}^{19}\mathrm{F}$^(19)F{ }^{19} \mathrm{~F} ) that form the human body. The vast majority of the MRI studies used in cerebral aging are based on the detection of the protons in ${}^1\mathrm{H}$${}^1\mathrm{H}$^(1)H{ }^{1} \mathrm{H} nucleus (more specifically water protons); thus we will focus our discussion on this nuclear species.
MRI 信號的來源是來自於形成人體的內源性核（例如， ${}^1\mathrm{H},{}^{23}\mathrm{Na},{}^{31}\mathrm{P},{}^{17}\mathrm{O}$${}^1\mathrm{H},{}^{23}\mathrm{Na},{}^{31}\mathrm{P},{}^{17}\mathrm{O}$^(1)H,^(23)Na,^(31)P,^(17)O{ }^{1} \mathrm{H},{ }^{23} \mathrm{Na},{ }^{31} \mathrm{P},{ }^{17} \mathrm{O} 和 ${}^{19}\mathrm{F}$${}^{19}\mathrm{F}$^(19)F{ }^{19} \mathrm{~F} ）。大多數用於腦部老化的 MRI 研究都是基於對 ${}^1\mathrm{H}$${}^1\mathrm{H}$^(1)H{ }^{1} \mathrm{H} 核中的質子的檢測（更具體地說是水質子）；因此我們將重點討論這種核物種。

Each proton can be viewed as a small magnet. When these small magnets are placed in an environment without any significant external magnetic field (e.g., outside the MRI scanner), they will be oriented randomly (Figure 1.1A) and, as a result, their net strength, referred to as magnetization, is 0 . On the other hand, when these magnets are placed inside a strong external magnetic field (e.g., inside the MRI scanner), their orientation is no longer random. Slightly more than half of them will orient themselves to be parallel to the external field, while slightly less than half will be anti-parallel to the external field (Figure 1.1B). Thus, the net strength of the proton magnets is no longer 0 but is parallel to the external field (see $M$$M$MM vector in Figure 1.1B), since slightly more magnets are pointing along that direction.
每個質子可以被視為一個小磁鐵。當這些小磁鐵置於沒有任何顯著外部磁場的環境中（例如，在 MRI 掃描儀外），它們將隨機排列（圖 1.1A），因此它們的淨強度，稱為磁化，為 0。另一方面，當這些磁鐵置於強外部磁場中（例如，在 MRI 掃描儀內），它們的排列不再是隨機的。稍微多於一半的磁鐵將會與外部磁場平行，而稍微少於一半的磁鐵將會與外部磁場反平行（圖 1.1B）。因此，質子磁鐵的淨強度不再是 0，而是與外部磁場平行（見圖 1.1B 中的 $M$$M$MM 向量），因為稍微多一些的磁鐵指向該方向。

The magnitude of the net strength, which ultimately determines the intensity of our MRI signal, is given by:
淨強度的大小，最終決定我們 MRI 信號的強度，表達為：

Figure 1.1 Illustration of the formation of MRI signal. (A) When the water protons are placed in an environment without a significant magnetic field, they are oriented randomly, thus the net signal (referred to as magnetization, $M$$M$MM ) is zero. (B) When the water protons are placed in a strong magnetic field, they are aligned along the axis of the external field, ${\mathrm{B}}_0$${\mathrm{B}}_0$B_(0)\mathrm{B}_{0}, with a slightly larger fraction parallel to the field (relative to anti-parallel). Thus, the net magnetization is along the $B_0$$B_0$B_(0)B_{0} direction. This net magnetization is what an MRI system measures when performing a scan.
圖 1.1 MRI 信號形成的示意圖。 (A) 當水中的質子置於沒有顯著磁場的環境中時，它們隨機取向，因此淨信號（稱為磁化， $M$$M$MM ）為零。 (B) 當水中的質子置於強磁場中時，它們沿著外部磁場的軸對齊， ${\mathrm{B}}_0$${\mathrm{B}}_0$B_(0)\mathrm{B}_{0} ，其中有稍微更大的一部分平行於磁場（相對於反平行）。因此，淨磁化沿著 $B_0$$B_0$B_(0)B_{0} 方向。這個淨磁化是 MRI 系統在進行掃描時所測量的。
where $\rho$$\rho$rho\rho is the proton density (PD) of the issue, ${\mathrm{B}}_0$${\mathrm{B}}_0$B_(0)\mathrm{B}_{0} is the strength of the external field, T is the temperature in Kelvin, and C is a parameter related to several physics constants. While mathematically simple, several observations can be made from Equation [1]. First, one can appreciate that the MRI signal is greater at a higher field strength, ${\mathrm{B}}_0$${\mathrm{B}}_0$B_(0)\mathrm{B}_{0}. Second, proton density, i.e., the amount of water in the tissue, plays a major role in MRI signal intensity. In the brain, CSF contains the highest water density (about 1 g water $/\mathrm{ml}$$/\mathrm{ml}$//ml/ \mathrm{ml} ) while the gray ( 0.89 g water $/\mathrm{ml}$$/\mathrm{ml}$//ml/ \mathrm{ml} ) and white matter ( 0.73 g water $/\mathrm{ml}$$/\mathrm{ml}$//ml/ \mathrm{ml} ) contain less (Herscovitch and Raichle, 1985). Therefore, in a proton-density MRI, CSF is expected to be brighter than the gray matter, which is in turn brighter than the white matter. Third, the lower the temperature, the greater the MRI signal. Unfortunately, this is not something that one can easily exploit under in vivo conditions.
其中 $\rho$$\rho$rho\rho 是組織的質子密度 (PD)， ${\mathrm{B}}_0$${\mathrm{B}}_0$B_(0)\mathrm{B}_{0} 是外部場的強度，T 是以開爾文為單位的溫度，C 是與幾個物理常數相關的參數。雖然在數學上簡單，但從方程式 [1] 中可以得出幾個觀察結果。首先，可以欣賞到在較高的場強下，MRI 信號更強， ${\mathrm{B}}_0$${\mathrm{B}}_0$B_(0)\mathrm{B}_{0} 。其次，質子密度，即組織中水的含量，在 MRI 信號強度中扮演著重要角色。在大腦中，腦脊髓液 (CSF) 含有最高的水密度（約 1 克水 $/\mathrm{ml}$$/\mathrm{ml}$//ml/ \mathrm{ml} ），而灰質（0.89 克水 $/\mathrm{ml}$$/\mathrm{ml}$//ml/ \mathrm{ml} ）和白質（0.73 克水 $/\mathrm{ml}$$/\mathrm{ml}$//ml/ \mathrm{ml} ）則含有較少的水（Herscovitch 和 Raichle，1985）。因此，在質子密度 MRI 中，腦脊髓液預期會比灰質更亮，而灰質又會比白質更亮。第三，溫度越低，MRI 信號越強。不幸的是，這在體內條件下並不是容易利用的。

While the application of the external ${\mathrm{B}}_0$${\mathrm{B}}_0$B_(0)\mathrm{B}_{0} field can induce a net magnetic field in the tissue, this field along the $B_0$$B_0$B_(0)B_{0} direction cannot be detected. Therefore, a second external field, referred to as radiofrequency (RF) magnetic field (also known as the $B_1$$B_1$B_(1)B_{1} field) is needed to “excite” the tissue spins. With these steps, one can detect a signal in the receiving coil. However, no spatial information is present yet. The spatial information is encoded by applying a magnetic field gradient such that water protons in different spatial locations experience different magnetic fields and therefore have different oscillation frequency. Once these signals are received and recorded by the coil, an image reconstruction algorithm can then be used to obtain the final MRI images.
雖然外部 ${\mathrm{B}}_0$${\mathrm{B}}_0$B_(0)\mathrm{B}_{0} 場的應用可以在組織中誘導出淨磁場，但沿 $B_0$$B_0$B_(0)B_{0} 方向的磁場無法被檢測到。因此，需要第二個外部場，稱為射頻（RF）磁場（也稱為 $B_1$$B_1$B_(1)B_{1} 場），來“激發”組織自旋。通過這些步驟，可以在接收線圈中檢測到信號。然而，尚未存在空間信息。空間信息是通過施加磁場梯度來編碼的，使得不同空間位置的水質子經歷不同的磁場，因此具有不同的振盪頻率。一旦這些信號被線圈接收和記錄，就可以使用圖像重建算法來獲得最終的 MRI 圖像。

The signal one obtains from the final MR image is further modulated by MR properties of the tissue, in particular T1 and T2. Specifically, the MR signal received is given by:
從最終的 MR 影像獲得的信號進一步受到組織的 MR 特性的調制，特別是 T1 和 T2。具體而言，接收到的 MR 信號為：

where FA is flip angle, TR is repetition time, and TE is echo time. In Equation [2], the MR properties T1 and T2 should be differentiated from imaging parameters, e.g., repetition time (TR) and echo time (TE). The difference is that imaging parameters can be chosen by the researcher, whereas T1 and T2 are intrinsic properties of the tissue that the researcher has no control over. However, by properly choosing TR and TE (and other imaging parameters for more advanced sequences), one can make the overall image intensity weighted in a pre-defined fashion, resulting in different contrasts such as T1 weighted or T2 weighted MRI.
其中 FA 是翻轉角度，TR 是重複時間，TE 是回波時間。在方程式 [2] 中，MR 性質 T1 和 T2 應與成像參數區分開來，例如重複時間 (TR) 和回波時間 (TE)。區別在於，成像參數可以由研究者選擇，而 T1 和 T2 是組織的內在特性，研究者無法控制。然而，通過適當選擇 TR 和 TE（以及其他更高級序列的成像參數），可以使整體圖像強度以預定的方式加權，從而產生不同的對比，例如 T1 加權或 T2 加權 MRI。

When evaluating the quality of a set of MRI data, three criteria can be considered. One is to identify whether the image contains any artifact, which is represented by the appearance of signals in unexpected areas. Figure 1.2 illustrates a brain image in which the fat signal was insufficiently suppressed and appeared inside the brain due to chemical shift between fat and water protons. Many reasons can cause image artifacts, and it is preferable to have the images reviewed by an experienced imaging scientist, especially at the beginning of a project, before proceeding with a large sample
在評估一組 MRI 數據的質量時，可以考慮三個標準。一個是識別圖像是否包含任何伪影，這由意外區域出現信號來表示。圖 1.2 顯示了一個腦部圖像，其中脂肪信號未被充分抑制，並因脂肪和水質子之間的化學位移而出現在腦內。許多原因可以導致圖像伪影，最好由經驗豐富的影像科學家對圖像進行審查，特別是在項目開始時，在進行大樣本之前。

Figure 1.2 An example of brain images with and without artifacts induced by undesired fat signal (arrows).
圖 1.2 一個有和沒有由不必要的脂肪信號（箭頭）引起的伪影的腦部影像示例。
scanning. Another criterion is the spatial signal-to-noise ratio (SNR) of the image, which is a useful index in evaluating the quality of structural MR images. The signal in the spatial SNR calculation can be obtained by drawing a region-of-interest (ROI) in a representative brain region and calculating the mean of all voxels, while the noise can be estimated by defining an ROI outside the brain and calculating the mean or standard deviation of the voxels. For MRI data that have multiple time points, such as functional MRI (fMRI) or arterial-spin-labeling (ASL) MRI, a third criterion, temporal SNR, is often used to assess data quality. In the calculation of temporal SNR, the signal term is defined similar to that for spatial SNR. For the noise term, however, it is defined as the standard deviation across the time points. Temporal SNR can be defined on a voxel-by-voxel basis or on an ROI. The estimation of temporal SNR is preferred in fMRI data assessment because it considers contributions from both thermal and physiological noise, whereas spatial SNR only considers thermal noise. Since physiological noise is known to be a major, if not the predominant, component of the noise source in fMRI, the examination of temporal SNR is more appropriate for determining the reliability of a functional dataset.
掃描。另一個標準是影像的空間信號噪聲比（SNR），這是一個評估結構性磁共振影像質量的有用指標。在空間 SNR 計算中，信號可以通過在代表性腦區劃定感興趣區域（ROI）並計算所有體素的平均值來獲得，而噪聲則可以通過在腦外定義 ROI 並計算體素的平均值或標準差來估算。對於具有多個時間點的 MRI 數據，例如功能性磁共振成像（fMRI）或動脈自旋標記（ASL）MRI，通常使用第三個標準，即時間 SNR，來評估數據質量。在時間 SNR 的計算中，信號項的定義類似於空間 SNR。然而，噪聲項的定義是跨時間點的標準差。時間 SNR 可以在每個體素的基礎上或在 ROI 上定義。在 fMRI 數據評估中，時間 SNR 的估算更受青睞，因為它考慮了來自熱噪聲和生理噪聲的貢獻，而空間 SNR 僅考慮熱噪聲。 由於生理噪聲被認為是功能性磁共振成像中噪聲來源的主要成分（如果不是最主要的話），因此檢查時間信噪比更適合用來確定功能數據集的可靠性。

As mentioned above in Equation [1], higher magnetic field strength usually corresponds to a greater sensitivity. 3 T is therefore preferred over 1.5 T for cognitive aging studies. This advantage has been experimentally demonstrated for virtually all brain MRI pulse sequences (although for cardiac and body MRI, 1.5T is sometimes preferred). As far as the magnitude of the enhancement effect is concerned, it depends on specific pulse sequence and spatial resolution. Going from 1.5T to 3T, a typical SNR increase of $50\mathrm{\%}-100\mathrm{\%}$$50\mathrm{\%}-100\mathrm{\%}$50%-100%50 \%-100 \% is often reported (Willinek and Kuhl, 2006; Bradley, 2008). High-resolution scans tend to manifest a greater gain because in those scans thermal noise is usually the dominant source of noise relative to physiological noise, which scales with signal. Some sequences such as BOLD fMRI and ASL perfusion benefit from additional factors related to
如上所述，在方程[1]中，較高的磁場強度通常對應於更大的敏感度。因此，在認知老化研究中，3 T 比 1.5 T 更受青睞。這一優勢在幾乎所有的腦部 MRI 脈衝序列中都得到了實驗證明（儘管在心臟和全身 MRI 中，有時會偏好 1.5T）。至於增強效應的大小，則取決於特定的脈衝序列和空間解析度。從 1.5T 到 3T，通常報告的典型信噪比增益為 $50\mathrm{\%}-100\mathrm{\%}$$50\mathrm{\%}-100\mathrm{\%}$50%-100%50 \%-100 \% （Willinek 和 Kuhl，2006；Bradley，2008）。高解析度掃描往往表現出更大的增益，因為在這些掃描中，熱噪聲通常是相對於生理噪聲的主要噪聲來源，而生理噪聲則隨信號而變化。一些序列，如 BOLD fMRI 和 ASL 灌注，受益於與之相關的額外因素。
enhanced magnetic susceptibility effects and longer T1 at 3T. Other sequences such as FLAIR benefit less because slower signal recovery at higher field offsets some of the advantages.
增強的磁敏感性效應和在 3T 下更長的 T1。其他序列如 FLAIR 的好處較少，因為在較高的場強下信號恢復較慢，抵消了一些優勢。

7T MRI is also becoming increasingly available in some research institutions. There are about 50 to 60 human 7T systems around the world. However, the use of 7T in cognitive aging studies is still in an early stage. Despite the promise of an increased sensitivity, 7T MRI still suffers from several technical limitations at present. These limitations are due to magnetic field inhomogeneity (thereby resulting in image inhomogeneity), higher power deposition (thus tissue heating becomes an issue, which is usually not a problem when going from 1.5 T to 3 T ), and physiological side effects (i.e., participant may feel dizzy when entering the scanner), aside from its high costs. However, these limitations may be resolved with technical efforts and advances, as shown by highly promising results from the Human Connectome Project.
7T 磁共振成像在一些研究機構中也變得越來越普及。全球大約有 50 到 60 台人類 7T 系統。然而，7T 在認知老化研究中的使用仍處於早期階段。儘管提高靈敏度的前景令人期待，但目前 7T 磁共振成像仍然面臨幾個技術限制。這些限制是由於磁場不均勻性（因此導致影像不均勻性）、更高的功率沉積（因此組織加熱成為一個問題，這在從 1.5 T 到 3 T 的過程中通常不是問題）以及生理副作用（即參與者在進入掃描儀時可能會感到頭暈），此外還有其高昂的成本。然而，這些限制可能會通過技術努力和進步得到解決，正如人類連接組計劃所顯示的非常有前景的結果。

It is recommended that each scan session be less than 60 minutes, as excessive motion is often observed when the subject has been inside the scanner for a long period of time. When that happens, the data collected are of low quality and usually need to be excluded. This is especially likely for elderly participants. For studies that require more than 60 minutes, one should consider allowing the subject to come out of the scanner to take a break before entering again for the remainder of the scans.
建議每次掃描會話少於 60 分鐘，因為當受試者在掃描儀內待的時間過長時，通常會觀察到過度運動。當這種情況發生時，收集到的數據質量較低，通常需要排除。這對於老年參與者尤其可能發生。對於需要超過 60 分鐘的研究，應考慮允許受試者在再次進入掃描儀進行剩餘掃描之前先出來休息。

It is also useful to prioritize the scans such that the pulse sequences that are most important for the study hypothesis are performed first. It is not uncommon that a subject would decide to abort the scan session or show excessive motion after staying in the scanner for a while. Thus, arranging the scan order based on priority can ensure the successful data collection of the most relevant sequences. Some investigators also found it helpful to place the functional and physiological scans at the beginning of the session when the subject is most alert and attentive. The structural scans can usually be performed even when the subject is asleep.
優先安排掃描是有用的，這樣對於研究假設最重要的脈衝序列可以優先進行。受試者在掃描儀中待了一段時間後，決定中止掃描會話或出現過度運動的情況並不罕見。因此，根據優先順序安排掃描順序可以確保成功收集最相關序列的數據。一些研究者還發現，將功能性和生理性掃描放在會話開始時進行是有幫助的，因為此時受試者最為清醒和專注。結構性掃描通常即使在受試者睡著的情況下也可以進行。

When calculating the total scan session duration, it should be remembered that the scan duration displayed on the scanner console often underestimates the real-life scan time, as the preparation time necessary at the beginning of every scan is usually not included in the displayed time. The preparation time is inherent to every MRI pulse sequence and usually includes $B_0$$B_0$B_(0)B_{0} shimming, RF center frequency determination, RF power optimization, and dummy scans to allow the magnetization to approach a steady state. Collectively, the actual data collection time for a 60-minute session may be around 45 minutes. It is useful to keep this in mind during planning. A small number of pilot scans should be performed before finalizing the protocol and obtaining a more accurate estimation of the session duration.
在計算總掃描會話持續時間時，應該記住掃描儀控制台上顯示的掃描持續時間通常低估了實際的掃描時間，因為每次掃描開始時所需的準備時間通常不包括在顯示的時間內。準備時間是每個 MRI 脈衝序列固有的，通常包括 $B_0$$B_0$B_(0)B_{0} 的調整、射頻中心頻率確定、射頻功率優化和虛擬掃描，以使磁化接近穩定狀態。總體而言，60 分鐘會話的實際數據收集時間可能約為 45 分鐘。在計劃過程中記住這一點是有用的。在最終確定協議並獲得更準確的會話持續時間估算之前，應進行少量的試點掃描。

Table 1.1 provides a list of MRI techniques commonly used in cerebral aging studies. They measure various aspects of the brain structure and function. Later chapters will
表 1.1 提供了在腦部老化研究中常用的 MRI 技術列表。它們測量大腦結構和功能的各個方面。後面的章節將

Table 1.1 List of MRI techniques and associated imaging parameters at 3T.
表 1.1 3T MRI 技術及相關影像參數列表。

cover the analyses and interpretation of some of these techniques. The present chapter will primarily focus on the image acquisition strategies.
涵蓋這些技術的一些分析和解釋。本章將主要集中於影像獲取策略。

Cerebral aging is associated with pronounced brain volumetric changes (Pfefferbaum et al., 1994; Good et al., 2001; Ge et al., 2002; Sowell et al., 2003; Raz et al., 2005). Thus, the ability to assess brain volume and cortical thickness is important in cerebral aging studies. Brain volumetric parameters are usually assessed with a high-resolution anatomical MRI scan that provides sufficient contrast between gray matter, white matter, and CSF. In principle, this can be achieved with any one of the three main image contrasts in MRI, i.e., T1-weighted, T 2 -weighted, and proton-density weighted. In practice, however, the vast majority of such studies have used the T1-weighted pulse sequence (Pfefferbaum et al., 1994; Good et al., 2001; Ge et al., 2002; Sowell et al., 2003; Raz et al., 2005). This is due to scan time considerations. A proton-density weighted pulse sequence usually utilizes a long TR, thus is less time-efficient. A T2weighted pulse sequence usually uses a long TR, again lengthening the scan duration. Therefore, a T1-weighted pulse sequence, which uses a short TR and a short TE, is considered an optimal approach for rapid, strong-contrast, and high-resolution acquisition of brain structure information.
腦部老化與顯著的腦部體積變化有關（Pfefferbaum et al., 1994; Good et al., 2001; Ge et al., 2002; Sowell et al., 2003; Raz et al., 2005）。因此，評估腦部體積和皮質厚度的能力在腦部老化研究中是重要的。腦部體積參數通常通過高解析度的解剖 MRI 掃描來評估，該掃描提供了灰質、白質和腦脊液之間的足夠對比。原則上，這可以通過 MRI 中的三種主要影像對比之一來實現，即 T1 加權、T2 加權和質子密度加權。然而，在實踐中，絕大多數此類研究使用了 T1 加權脈衝序列（Pfefferbaum et al., 1994; Good et al., 2001; Ge et al., 2002; Sowell et al., 2003; Raz et al., 2005）。這是由於掃描時間的考量。質子密度加權脈衝序列通常使用較長的 TR，因此效率較低。T2 加權脈衝序列通常也使用較長的 TR，進一步延長掃描時間。 因此，T1 加權脈衝序列使用短的 TR 和短的 TE，被認為是快速、強對比和高解析度獲取腦結構信息的最佳方法。

T 1 of CSF is greater than gray matter T 1 , which is in turn greater than white matter T1. At 3 Tesla, these values are $4,163\mathrm{ms}$$4,163\mathrm{ms}$4,163ms4,163 \mathrm{~ms} (Lin et al., 2001), $1,135\mathrm{ms}$$1,135\mathrm{ms}$1,135ms1,135 \mathrm{~ms} (Lu et al., 2005), and 732 ms (Lu et al., 2005), respectively. Thus, in a T1-weighted image, white matter is expected to have a higher signal intensity than gray matter, and CSF will have the least signal intensity. Figure 1.3A shows an example of T1-weighted image at a resolution of $1\times 1\times 1{\mathrm{mm}}^3$$1\times 1\times 1{\mathrm{mm}}^3$1xx1xx1mm^(3)1 \times 1 \times 1 \mathrm{~mm}^{3}. Using image segmentation functions that are readily available in standard software packages, probability maps of gray matter
CSF 的 T1 大於灰質的 T1，而灰質的 T1 又大於白質的 T1。在 3 特斯拉下，這些值分別為 $4,163\mathrm{ms}$$4,163\mathrm{ms}$4,163ms4,163 \mathrm{~ms} （Lin et al., 2001）、 $1,135\mathrm{ms}$$1,135\mathrm{ms}$1,135ms1,135 \mathrm{~ms} （Lu et al., 2005）和 732 毫秒（Lu et al., 2005）。因此，在 T1 加權影像中，白質的信號強度預期會高於灰質，而 CSF 的信號強度則最低。圖 1.3A 顯示了一個解析度為 $1\times 1\times 1{\mathrm{mm}}^3$$1\times 1\times 1{\mathrm{mm}}^3$1xx1xx1mm^(3)1 \times 1 \times 1 \mathrm{~mm}^{3} 的 T1 加權影像範例。使用標準軟體包中 readily available 的影像分割功能，可以得到灰質的概率圖。

Figure 1.3 Example of T1-weighted image and the corresponding probability maps of gray matter, white matter and CSF. (A) T1-MPRAGE image obtained from a MRI scan. (B) Gray matter probability map. © White matter probability map. (D) CSF probability map. (b-d) are obtained by segmenting T1-MPRAGE image shown in (a).
圖 1.3 T1 加權影像及其對應的灰質、白質和腦脊髓液的概率圖示例。(A) 從 MRI 掃描獲得的 T1-MPRAGE 影像。(B) 灰質概率圖。(C) 白質概率圖。(D) 腦脊髓液概率圖。(b-d) 是通過對(a)中顯示的 T1-MPRAGE 影像進行分割獲得的。
(Figure 1.3B), white matter (Figure 1.3C), and CSF (Figure 1.3D) can be obtained from the T 1 -weighted image.
（圖 1.3B）、白質（圖 1.3C）和腦脊髓液（圖 1.3D）可以從 T 1 加權影像中獲得。

In the early years, T1-weighted high-resolution images were acquired using a plain short-TR, short-TE gradient-echo sequence. Different vendors have different names for this sequence. On MRI scanners made by General Electric, it is called Spoiled Gradient Echo (SPGR). On Siemens systems, it is called Fast Low Angle Shot (FLASH). On Philips scanner, it is called Fast Field Echo (FFE). More recently, an improved version of the T1-weighted pulse sequence has gained popularity. This sequence is called Magnetization Prepared Rapid Acquisition of Gradient Echo (MPRAGE) (Mugler and Brookeman, 1991), which uses a preparation pulse to enhance the tissue contrast between gray matter, white matter, and CSF, thereby allowing better tissue segmentation and image registration. The MRPAGE sequence is the most widely used technique for brain volumetric studies at present.
在早期，T1 加權高解析度影像是使用簡單的短 TR、短 TE 梯度回波序列獲得的。不同的廠商對這個序列有不同的名稱。在通用電氣製造的 MRI 掃描儀上，它被稱為 Spoiled Gradient Echo (SPGR)。在西門子系統上，它被稱為 Fast Low Angle Shot (FLASH)。在飛利浦掃描儀上，它被稱為 Fast Field Echo (FFE)。最近，一種改進版的 T1 加權脈衝序列變得越來越受歡迎。這個序列被稱為磁化準備快速獲取梯度回波（Magnetization Prepared Rapid Acquisition of Gradient Echo，MPRAGE）（Mugler 和 Brookeman，1991），它使用準備脈衝來增強灰質、白質和腦脊髓液之間的組織對比，從而允許更好的組織分割和影像配準。目前，MRPAGE 序列是腦體積研究中最廣泛使用的技術。

Brain volumetric images are usually collected at a high resolution. The typical voxel size of a brain volumetric image is around $1\times 1\times 1{\mathrm{mm}}^3$$1\times 1\times 1{\mathrm{mm}}^3$1xx1xx1mm^(3)1 \times 1 \times 1 \mathrm{~mm}^{3}, although a slightly nonisotropic voxel size (e.g., $1\times 1\times 1.2{\mathrm{mm}}^3$$1\times 1\times 1.2{\mathrm{mm}}^3$1xx1xx1.2mm^(3)1 \times 1 \times 1.2 \mathrm{~mm}^{3} ) is sometimes also used. This level of spatial resolution is needed in order to provide sufficient number of voxels across the thickness of cortical ribbon, which is around $2-4\mathrm{mm}$$2-4\mathrm{mm}$2-4mm2-4 \mathrm{~mm}. A resolution higher than 1 mm is of course desirable, but often at a cost of scan duration and/or SNR. Thus, sub-millimeter resolution (e.g., 0.7 mm ) is only used when one is specifically interested in a small structure of the brain, for example identifying subfields of hippocampus. Under these circumstances, usually the image will focus on the specific structure but not cover the entire brain, i.e., only partial brain coverage.
腦部體積影像通常以高解析度收集。腦部體積影像的典型體素大小約為 $1\times 1\times 1{\mathrm{mm}}^3$$1\times 1\times 1{\mathrm{mm}}^3$1xx1xx1mm^(3)1 \times 1 \times 1 \mathrm{~mm}^{3} ，雖然有時也會使用稍微非各向同性的體素大小（例如， $1\times 1\times 1.2{\mathrm{mm}}^3$$1\times 1\times 1.2{\mathrm{mm}}^3$1xx1xx1.2mm^(3)1 \times 1 \times 1.2 \mathrm{~mm}^{3} ）。這種空間解析度的水平是必要的，以便在皮質帶的厚度上提供足夠的體素數量，該厚度約為 $2-4\mathrm{mm}$$2-4\mathrm{mm}$2-4mm2-4 \mathrm{~mm} 。當然，超過 1 毫米的解析度是理想的，但通常會以掃描時間和/或信噪比為代價。因此，只有在特別關注腦部的小結構時，才會使用亞毫米解析度（例如，0.7 毫米），例如識別海馬體的亞區。在這種情況下，影像通常會專注於特定結構，而不覆蓋整個腦部，即僅部分腦部覆蓋。

Brain volumetric images are acquired in 3D. In MRI acquisition, there is a clear distinction between “3D” and “multislice” acquisitions. In multislice acquisition, although the stack of 2D slices can cover the same brain volume as a 3D scan, the through-plane resolution is never as good as that of a 3D acquisition. That is, even if one prescribes the slice thickness in a multislice acquisition to be 1 mm , the actual resolution cannot reach 1 mm due to imperfection in slice selection profile. This is because the RF pulse used to excite the spins has a frequency-selection band that is not perfect. As a result, the spins adjacent to the intended slice are also partially excited, and this effect decays with distance. Fortunately, brain volumetric images are always collected in a 3D mode, thus the images will look equally sharp when viewed in axial, sagittal, or coronal planes. A corollary of this notion is that, from the image quality point of view, it does not matter if the original scan is performed in axial, sagittal, or coronal orientation, because the data can be reformatted into any orientation without loss of spatial resolution. From the acquisition efficiency (thereby scan duration) point-of-view, sagittal orientation is often used. The main reason for this choice is that the human brain is usually the shortest along the left-right direction, typically about $160-180\mathrm{mm}$$160-180\mathrm{mm}$160-180mm160-180 \mathrm{~mm}. Thus, it requires the least number of slices to cover the brain from the left to right side. A second reason is that there is no other source of MRI signal to the left or right of the brain (unless the cushion or headset stabilizing the head contains fluid). Thus, even if the slice selection profile is not perfect, no spurious signal will be excited or detected outside the intended excitation volume. Such signals would have manifested themselves as foldover artifacts in which a spin to the left of the intended volume will overlap with the right end of the image, and vice versa.
腦部體積影像以 3D 方式獲取。在 MRI 獲取中，“3D”和“多切片”獲取之間有明顯的區別。在多切片獲取中，儘管 2D 切片的堆疊可以覆蓋與 3D 掃描相同的腦部體積，但其平面內的解析度永遠不如 3D 獲取的好。也就是說，即使在多切片獲取中將切片厚度設置為 1 毫米，實際解析度也無法達到 1 毫米，因為切片選擇輪廓存在不完美。這是因為用於激發自旋的射頻脈衝的頻率選擇帶並不完美。因此，鄰近預定切片的自旋也會部分被激發，這種效應隨距離而衰減。幸運的是，腦部體積影像總是以 3D 模式收集，因此在軸向、矢狀或冠狀平面查看時，影像看起來都同樣清晰。這一觀念的推論是，從影像質量的角度來看，原始掃描是以軸向、矢狀或冠狀方向進行並不重要，因為數據可以在不損失空間解析度的情況下重新格式化為任何方向。 從獲取效率（因此掃描持續時間）的角度來看，通常使用矢狀面方向。這種選擇的主要原因是人腦在左右方向上通常是最短的，通常約為 $160-180\mathrm{mm}$$160-180\mathrm{mm}$160-180mm160-180 \mathrm{~mm} 。因此，從左到右側覆蓋大腦所需的切片數量最少。第二個原因是大腦的左側或右側沒有其他 MRI 信號源（除非穩定頭部的墊子或耳機內含有液體）。因此，即使切片選擇輪廓不完美，也不會在預期的激發體積之外激發或檢測到虛假信號。這些信號會表現為折疊伪影，其中位於預期體積左側的自旋將與圖像的右端重疊，反之亦然。

Without any acquisition acceleration schemes (e.g., parallel imaging), typical scan duration for an MRPAGE sequence is between $8-10$$8-10$8-108-10 minutes. With a parallel imaging acceleration factor of 2 , this becomes $4-5$$4-5$4-54-5 minutes. When describing the MRPAGE sequence in a report, aside from the typical parameters such as voxel size, field-ofview, and echo time (TE), it is useful to specify an imaging parameter referred to as the inversion time (TI). In the MPRAGE sequence, the TI is crucial in determining the contrast of the image. In general, it is useful to describe the imaging parameters as thoroughly as possible when reading or writing an article.
在沒有任何獲取加速方案（例如，平行成像）的情況下，MRPAGE 序列的典型掃描時間為 $8-10$$8-10$8-108-10 分鐘。使用平行成像加速因子 2，這變為 $4-5$$4-5$4-54-5 分鐘。在報告中描述 MRPAGE 序列時，除了典型參數如體素大小、視野和回波時間（TE）外，指定一個稱為反轉時間（TI）的成像參數也是有用的。在 MPRAGE 序列中，TI 對於確定圖像的對比度至關重要。一般來說，在閱讀或撰寫文章時，盡可能詳細地描述成像參數是有用的。

Functional MRI (fMRI) provides an approach to probe the brain function noninvasively (Bandettini et al., 1992; Kwong et al., 1992; Ogawa et al., 1992). The most commonly measured fMRI signal is based on an indirect effect of neural activity on blood flow and blood oxygenation. This signal is referred to as blood-oxygenation-level-dependent (BOLD) signal (Ogawa et al., 1990), since it is sensitive to blood oxygenation in the venous vessels. There are two forms of fMRI, task-related fMRI and resting-state fMRI. In task-related fMRI, one aims to measure BOLD signal changes in response to a time-controlled task. Thus, the researcher usually presents a predefined stimulus time series to the participant while the MRI scanner is continuously acquiring images of the brain. Then, in data analysis, mathematical algorithms such as linear regression are used to search for the expected temporal signal pattern throughout the brain. If any voxel manifests the expected pattern, one usually calls this voxel “activated” and labels it with a pseudocolor. In resting-state fMRI, no explicit task is applied during the experiment while images are continuously collected for a period of a few minutes. Then, in analysis, one either selects a seed voxel and looks throughout the brain for voxels that have a signal pattern similar to that of the seed, or one uses more advanced algorithms such as independent component analysis to identify clusters of voxels that have similar temporal fluctuation, presumably due to direct or indirect synaptic connections. Thus, resting-state fMRI is also referred to as functional connectivity MRI (fcMRI).
功能性磁共振成像（fMRI）提供了一種非侵入性探測大腦功能的方法（Bandettini et al., 1992; Kwong et al., 1992; Ogawa et al., 1992）。最常測量的 fMRI 信號是基於神經活動對血流和血氧飽和度的間接影響。這種信號被稱為血氧水平依賴（BOLD）信號（Ogawa et al., 1990），因為它對靜脈血管中的血氧飽和度敏感。fMRI 有兩種形式，任務相關 fMRI 和靜息態 fMRI。在任務相關 fMRI 中，目的是測量對時間控制任務的 BOLD 信號變化。因此，研究者通常會在 MRI 掃描儀持續獲取大腦圖像的同時，向參與者呈現預定義的刺激時間序列。然後，在數據分析中，使用數學算法如線性回歸來搜索整個大腦中預期的時間信號模式。如果任何體素顯示出預期的模式，通常會將該體素稱為“激活”，並用偽顏色標記。在靜息態 fMRI 中，實驗期間不施加明確的任務，同時持續收集幾分鐘的圖像。 然後，在分析中，選擇一個種子體素並在整個大腦中尋找與該種子具有相似信號模式的體素，或者使用更先進的算法，如獨立成分分析，來識別具有相似時間波動的體素集群，這可能是由於直接或間接的突觸連接。因此，靜息態功能性磁共振成像也被稱為功能連接性磁共振成像（fcMRI）。

Both task-related and resting-state fMRIs use a T2*-weighted echo-planarimaging (EPI) technique. The fMRI pulse sequence uses a multislice acquisition scheme to cover the whole brain, in which the images are taken on a slice-byslice basis. For each slice, a fast acquisition technique, EPI, is used to collect the data. For a typical brain volume containing 40 slices, it takes up to $3,000\mathrm{ms}$$3,000\mathrm{ms}$3,000ms3,000 \mathrm{~ms} to complete the data collection. Since neural activity and the associated hemodynamic responses are transient, it is desirable to expedite the rate of data collection. This can be achieved using a recently developed multiband acquisition technique in which two or more slices are acquired simultaneously (Feinberg et al., 2010; Mueller et al., 2010; Setsompop et al., 2012). As such, the time it takes to complete the data collection of one volume can be shortened substantially. At present, the TR in an fMRI scan can be as short as 600 ms or less. These recent technologies have been utilized in major brain imaging initiatives such as the Human Connectome Project (Van Essen et al., 2012) and are gradually becoming the standard procedure for fMRI studies.
任務相關和靜息狀態的功能性磁共振成像（fMRI）均使用 T2*-加權回聲平面成像（EPI）技術。fMRI 脈衝序列使用多切片獲取方案來覆蓋整個大腦，其中圖像是逐片獲取的。對於每個切片，使用快速獲取技術 EPI 來收集數據。對於包含 40 個切片的典型大腦體積，完成數據收集最多需要 $3,000\mathrm{ms}$$3,000\mathrm{ms}$3,000ms3,000 \mathrm{~ms} 。由於神經活動及其相關的血流動力學反應是瞬時的，因此希望加快數據收集的速度。這可以通過最近開發的多帶獲取技術來實現，其中兩個或更多切片同時獲取（Feinberg 等，2010；Mueller 等，2010；Setsompop 等，2012）。因此，完成一個體積的數據收集所需的時間可以大大縮短。目前，fMRI 掃描中的 TR 可以短至 600 毫秒或更少。這些最近的技術已被應用於主要的大腦成像計劃，如人類連接組計劃（Van Essen 等，2012），並逐漸成為 fMRI 研究的標準程序。

The typical spatial resolution of fMRI is between $3-3.5\mathrm{mm}$$3-3.5\mathrm{mm}$3-3.5mm3-3.5 \mathrm{~mm}, but there is a trend toward a higher spatial resolution of $2-2.5\mathrm{mm}$$2-2.5\mathrm{mm}$2-2.5mm2-2.5 \mathrm{~mm} with recent advances in fast imaging technologies, including multiband, parallel imaging, and high-field MRI.
功能性磁共振成像（fMRI）的典型空間解析度介於 $3-3.5\mathrm{mm}$$3-3.5\mathrm{mm}$3-3.5mm3-3.5 \mathrm{~mm} 之間，但隨著快速成像技術的最新進展，包括多頻帶、平行成像和高場磁共振成像，對於更高空間解析度的趨勢正在增強，達到 $2-2.5\mathrm{mm}$$2-2.5\mathrm{mm}$2-2.5mm2-2.5 \mathrm{~mm} 。

FMRI signal value at any particular time point, often written in arbitrary units with a typical range of several hundred, has no physiological meaning. It is the temporal fluctuation or change of the signal that contains information about neural activity. Taking task-related fMRI, for example, a task stimulus may cause the fMRI signal to increase by a fraction of a percent. A signal change of $1\mathrm{\%}$$1\mathrm{\%}$1%1 \% is considered a large signal in fMRI data, in particular for cognitive tasks. Does a $1\mathrm{\%}$$1\mathrm{\%}$1%1 \% signal change mean that neural activity has increased by $1\mathrm{\%}$$1\mathrm{\%}$1%1 \% ? Not really. FMRI signal is a complex function of several physiological parameters including cerebral blood flow (CBF), cerebral blood volume (CBV), and cerebral metabolic rate of oxygen (CMRO2) (Davis et al., 1998; Hoge et al., 1999; Arthurs and Boniface, 2002). Figure 1.4A illustrates the pathway that leads to the observation of fMRI signal change due to a stimulus. Neural activation induced by the stimulus results in an elevation in metabolism and release of neurotransmitters. These factors could either directly dilate blood vessels or do so via activation of glial cells such as astrocytes, which have end feet attached to vessels. These effects cause an increase in CBV and CBF, which are referred to as the hemodynamic responses of the brain. It is important to note that the increase in oxygen supply (i.e., CBF) during brain activation is more prominent than the increase in oxygen demand (i.e., CMRO2), thus the oxygenation status of the venous blood is increased. Therefore, given a task stimulus such as shown in Figure 1.4B, BOLD fMRI signal change could be detected (Figure 1.4C).
在任何特定時間點的 FMRI 信號值，通常以任意單位表示，典型範圍為幾百，並沒有生理意義。信號的時間波動或變化包含了有關神經活動的信息。以任務相關的 fMRI 為例，任務刺激可能會使 fMRI 信號增加百分之一的某個小數。信號變化 $1\mathrm{\%}$$1\mathrm{\%}$1%1 \% 在 fMRI 數據中被視為一個大的信號，特別是在認知任務中。信號變化 $1\mathrm{\%}$$1\mathrm{\%}$1%1 \% 是否意味著神經活動增加了 $1\mathrm{\%}$$1\mathrm{\%}$1%1 \% ？並不一定。FMRI 信號是幾個生理參數的複雜函數，包括腦血流量（CBF）、腦血容量（CBV）和腦氧代謝率（CMRO2）（Davis 等，1998；Hoge 等，1999；Arthurs 和 Boniface，2002）。圖 1.4A 說明了由於刺激而導致 fMRI 信號變化的觀察路徑。刺激引起的神經激活導致代謝升高和神經遞質的釋放。這些因素可以直接擴張血管，或通過激活如星形膠質細胞等膠質細胞來實現，這些細胞的末端附著在血管上。 這些效應導致 CBV 和 CBF 的增加，這被稱為大腦的血流動力學反應。重要的是要注意，在大腦激活期間，氧氣供應的增加（即 CBF）比氧氣需求的增加（即 CMRO2）更為明顯，因此靜脈血的氧合狀態增加。因此，給定如圖 1.4B 所示的任務刺激，可以檢測到 BOLD fMRI 信號的變化（圖 1.4C）。

Figure 1.4 Illustration of BOLD fMRI. (A) Illustration of the neurovascular coupling pathway that leads to the observation of fMRI signal change due to a stimulus. (B) Example of taskfMRI stimulus paradigm. © Corresponding BOLD signal in visual cortex from an fMRI scan using the paradigm in (b). Red bars indicate stimulus periods. (See color plate also)
圖 1.4 BOLD fMRI 的插圖。(A) 神經血管耦合通路的插圖，該通路導致由於刺激而觀察到的 fMRI 信號變化。(B) 任務 fMRI 刺激範式的示例。© 使用(b)中的範式進行的 fMRI 掃描中視覺皮層的相應 BOLD 信號。紅色條形表示刺激期間。(另見彩色圖版)

How is venous blood oxygenation level related to fMRI signal? It is fortuitous that hemoglobin in the blood has a differential magnetic property between oxygenated and deoxygenated states, thus can serve as an endogenous contrast agent for MRI. When oxygenated, the hemoglobin is in a so-called diamagnetic state, which does not distinguish itself from the surrounding water in terms of magnetic property. In the deoxygenated state, on the other hand, it is in a paramagnetic state, which generates a small magnetic field in its surrounding, causing a disturbance to the homogeneity of the magnetic field. Since T2* of a voxel is closely associated with the homogeneity of the field, this MR parameter is dependent on the abundance of deoxyhemoglobin in the voxel and thereby the oxygenation status of the blood.
靜脈血氧合水平與功能性磁共振成像信號有何關聯？血液中的血紅蛋白在氧合和去氧狀態之間具有不同的磁性特性，因此可以作為 MRI 的內源性對比劑。當血紅蛋白氧合時，它處於所謂的抗磁性狀態，在磁性特性上與周圍的水無法區分。另一方面，在去氧狀態下，它處於順磁性狀態，這會在其周圍產生一個小的磁場，造成磁場均勻性的擾動。由於體素的 T2*與磁場的均勻性密切相關，因此這個 MR 參數依賴於體素中去氧血紅蛋白的豐富程度，從而影響血液的氧合狀態。

When integrating these cascades in a quantitative manner, the fMRI signal change due to task stimulation can be written as (Davis et al., 1998; Hoge et al., 1999):
當以定量方式整合這些級聯時，因任務刺激而引起的 fMRI 信號變化可以寫成（Davis 等，1998；Hoge 等，1999）：

in which $\mathrm{\Delta }\mathrm{CMRO}/\mathrm{CMRO}2$$\mathrm{\Delta }\mathrm{CMRO}/\mathrm{CMRO}2$DeltaCMRO//CMRO2\Delta \mathrm{CMRO} / \mathrm{CMRO} 2 indicates task-induced change in brain metabolic rate; $\mathrm{\Delta }\mathrm{CBF}/\mathrm{CBF}$$\mathrm{\Delta }\mathrm{CBF}/\mathrm{CBF}$DeltaCBF//CBF\Delta \mathrm{CBF} / \mathrm{CBF} is the change in blood flow; $M,\alpha$$M,\alpha$M,alphaM, \alpha, and $\beta$$\beta$beta\beta are variables related to scanner field strength and imaging parameters, blood flow-volume relationship, and vessel geometry, respectively. To provide the readers with a realistic example of these parameters in the primary visual cortex, Hutchison et al. (2013) found that flashing grating caused a $\mathrm{\Delta }\mathrm{CMRO}2/\mathrm{CMRO}2$$\mathrm{\Delta }\mathrm{CMRO}2/\mathrm{CMRO}2$DeltaCMRO2//CMRO2\Delta \mathrm{CMRO} 2 / \mathrm{CMRO} 2 of $15\mathrm{\%}$$15\mathrm{\%}$15%15 \% and a $\mathrm{\Delta }\mathrm{CBF}/\mathrm{CBF}$$\mathrm{\Delta }\mathrm{CBF}/\mathrm{CBF}$DeltaCBF//CBF\Delta \mathrm{CBF} / \mathrm{CBF} of $35\mathrm{\%}$$35\mathrm{\%}$35%35 \%, which resulted in a fMRI signal increase by $0.7\mathrm{\%}$$0.7\mathrm{\%}$0.7%0.7 \%.
在其中 $\mathrm{\Delta }\mathrm{CMRO}/\mathrm{CMRO}2$$\mathrm{\Delta }\mathrm{CMRO}/\mathrm{CMRO}2$DeltaCMRO//CMRO2\Delta \mathrm{CMRO} / \mathrm{CMRO} 2 表示任務引起的大腦代謝率變化； $\mathrm{\Delta }\mathrm{CBF}/\mathrm{CBF}$$\mathrm{\Delta }\mathrm{CBF}/\mathrm{CBF}$DeltaCBF//CBF\Delta \mathrm{CBF} / \mathrm{CBF} 是血流的變化； $M,\alpha$$M,\alpha$M,alphaM, \alpha 和 $\beta$$\beta$beta\beta 分別是與掃描儀場強度和成像參數、血流-體積關係以及血管幾何形狀相關的變量。為了向讀者提供這些參數在初級視覺皮層中的現實例子，Hutchison 等人（2013）發現閃爍的格子圖案引起了 $\mathrm{\Delta }\mathrm{CMRO}2/\mathrm{CMRO}2$$\mathrm{\Delta }\mathrm{CMRO}2/\mathrm{CMRO}2$DeltaCMRO2//CMRO2\Delta \mathrm{CMRO} 2 / \mathrm{CMRO} 2 的 $15\mathrm{\%}$$15\mathrm{\%}$15%15 \% 和 $\mathrm{\Delta }\mathrm{CBF}/\mathrm{CBF}$$\mathrm{\Delta }\mathrm{CBF}/\mathrm{CBF}$DeltaCBF//CBF\Delta \mathrm{CBF} / \mathrm{CBF} 的 $35\mathrm{\%}$$35\mathrm{\%}$35%35 \% ，這導致 fMRI 信號增加了 $0.7\mathrm{\%}$$0.7\mathrm{\%}$0.7%0.7 \% 。

More discussions on fMRI experimental design and analyses in aging studies can be found in chapters 4 and 6.
有關老年研究中 fMRI 實驗設計和分析的更多討論可以在第 4 章和第 6 章中找到。

DTI provides a powerful tool to evaluate microstructural tissue integrity and whitematter connectivity, beyond those of conventional structural imaging (Basser et al., 1994). Diffusion MRI can reflect structural properties of the brain because tissue structures such as cell membrane act as barriers that prevent water molecules from free, random motion. Furthermore, some structures such as axonal membrane and myelin sheath disrupt free diffusion in an orientation-dependent manner. That is, along the white-matter fiber’s principal direction, the water diffusion is minimally affected, whereas across the fiber direction the diffusion is heavily influenced. As such, DTI can also provide an anisotropy measure of water diffusion, which can be used to determine white-matter fiber orientation (Mori et al., 1999; Le Bihan et al., 2001).
DTI 提供了一個強大的工具來評估微觀組織的完整性和白質連接性，超越了傳統結構成像的範疇（Basser et al., 1994）。擴散 MRI 可以反映大腦的結構特性，因為組織結構如細胞膜充當障礙，阻止水分子自由隨機運動。此外，一些結構如軸突膜和髓鞘以方向依賴的方式干擾自由擴散。也就是說，沿著白質纖維的主要方向，水的擴散受到的影響最小，而在纖維方向的交叉處，擴散則受到很大影響。因此，DTI 也可以提供水擴散的各向異性測量，這可以用來確定白質纖維的方向（Mori et al., 1999; Le Bihan et al., 2001）。

DTI measures the diffusion pattern of water molecules by applying two magnetic field gradients in the pulse sequence (Figure 1.5A), which can encode the distance of water diffusion during this period. The strength of the magnetic field gradients can be represented by an imaging parameter called “b value,” which is associated with both magnitude and duration of the gradients as well as their time gap:
DTI 測量水分子擴散模式，通過在脈衝序列中應用兩個磁場梯度（圖 1.5A），這可以編碼在此期間水擴散的距離。磁場梯度的強度可以用一個稱為“b 值”的成像參數來表示，該參數與梯度的大小和持續時間以及它們的時間間隔有關：
(A)

(B)

$\mathrm{b}=0\mathrm{s}/{\mathrm{mm}}^2\mathrm{b}=1000\mathrm{s}/{\mathrm{mm}}^2$$\mathrm{b}=0\mathrm{s}/{\mathrm{mm}}^2\mathrm{b}=1000\mathrm{s}/{\mathrm{mm}}^2$b=0s//mm^(2)quadb=1000s//mm^(2)quad\mathrm{b}=0 \mathrm{~s} / \mathrm{mm}^{2} \quad \mathrm{~b}=1000 \mathrm{~s} / \mathrm{mm}^{2} \quad ADC map   $\mathrm{b}=0\mathrm{s}/{\mathrm{mm}}^2\mathrm{b}=1000\mathrm{s}/{\mathrm{mm}}^2$$\mathrm{b}=0\mathrm{s}/{\mathrm{mm}}^2\mathrm{b}=1000\mathrm{s}/{\mathrm{mm}}^2$b=0s//mm^(2)quadb=1000s//mm^(2)quad\mathrm{b}=0 \mathrm{~s} / \mathrm{mm}^{2} \quad \mathrm{~b}=1000 \mathrm{~s} / \mathrm{mm}^{2} \quad ADC 地圖
©

(D)

Figure 1.5 Illustration of diffusion tensor imaging (DTI). (A) A simplified DTI pulse sequence. (B) Diffusion images with two different $b$$b$bb values and the calculated ADC map. © Components of the diffusion tensor matrix. (D) ADC can be determined by the tensor matrix.
圖 1.5 擴散張量成像 (DTI) 的示意圖。 (A) 簡化的 DTI 脈衝序列。 (B) 具有兩個不同 $b$$b$bb 值的擴散影像及計算出的 ADC 地圖。 (C) 擴散張量矩陣的組成部分。 (D) ADC 可以通過張量矩陣來確定。