Several studies investigated DCE-MRI approaches for assessing renal filtration in animal models. Annet et al. estimated the GFR in rabbit kidneys by applying a compartment model and compared the MRI-derived estimate with the measured plasma clearance of 51Cr-EDTA [60]. Although MRI-based GFR values were lower than those measured experimentally, a marked correlation was observed between the two techniques. Winter et al. compared renal perfusion values obtained by arterial spin labeling (ASL) or DCE-MRI in healthy rats, with absolute renal cortex perfusion estimates in close agreement with those obtained by ASL [61].

Zimmer et al. performed an inter- and intramethodical comparison by measuring differences in renal blood flow (RBF) in five rats with unilateral ischemic acute kidney injury with both ASL and DCE-MRI. Both, the FAIR-ASL approach and DCE-MRI

Fig. 7 Exemplary illustration of perfusion MRI of a rat with left-side AKI and perfusion maps. (a) True-FISP M0 image of an ASL measurement and the corresponding perfusion map (b). (c) TWIST post–contrast agent injection image and the corresponding RBF map (d). All drawings show the same rat and the same axial slice. Differences between the kidney with AKI and the contralateral kidney are clearly visible on the MRI images as well as on the perfusion maps. (Reproduced with permission from Zimmer et al. [62])

deconvolution technique showed significant differences in RBF between healthy and diseased kidneys as shown in Fig. 7 [62]. The validation of MRI-based GFR estimation by optical imaging was proposed by Sadick et al. [63]. GFR estimation was performed by DCE-MRI on a 3 T whole body scanner using a dedicated animal rat volume coil and by clearance using a fluorescent tracer (FITC-sinistrin) and an optical imaging device. However, the correlation between optical GFR and MRI-based GFR was poor, probably due to the fact that measurements were performed on different days involving two anesthesia thus affecting the physiology of the animals. Zo¨llner et al. recently showed a simultaneous measurement of optical and MR-based GFR in healthy rats and rats with unilateral nephrectomy (UNX) [64]. A two-compartment filtration model was employed to calculate a map of the tubular flow. Subsequent, taking the cortex volume into account single kidney GFR values were calculated with a good correlation observed between the two methods. The reduction of GFR in the UNX rats was about 50% and observed for both techniques.

The role played by the binding to serum proteins in the calculation of DCE-MRI derived estimates was investigated by Notohamiprodjo et al. by administering Gadolinium-based contrast agents with different chemical structure [65]. The two investigated contrast agents, Gd-BOPTA and Gd-DTPA differ in their binding affinity toward serum albumin, hence affecting both the relaxivity and the pharmacokinetic properties [66–68]. In fact, binding to the serum protein results in longer circulation times (hence reduced filtration) and higher contrast capability that are easily observed at low magnetic fields (0.5–1.5 T) [69–73]. Therefore, whereas similar renal perfusion values were obtained by both the two contrast agents, GFR estimates were not accurate when obtained with the albumin-binding one (Gd-BOPTA), thus indicating that even low albumin binding may affect GFR estimates.

### 3.2 DCE-MRI for Assessing Kidney Damages in Animal Models

In the preclinical panorama, DCE-MRI has been widely used to assess kidney function in renovascular and parenchymal diseases that may lead to chronic kidney failure but also in acute events like ischemia–reperfusion, toxicity, infections, or surgery. Thus, different animal models have been explored ranging from classical acute kidney injury (AKI) and chronic kidney disease (CKD) to posttransplant rejection, nephrectomy, and assessment of drug treatment [74–80]. In a murine model of unilateral renal artery stenosis (RAS), measurements of single kidney GFR and perfusion by a modified two-compartment model were compared to those obtained from fluorescein isothiocyanate (FITC)-inulin clearance and ASL, respectively. Both renal GFR and perfusion measurements were in close agreement with the FITC-inulin clearance and ASL methods, with a marked reduction in GFR and perfusion estimates for the stenotic as compared to the control kidneys [81]. Ischemia–reperfusion injury has been exploited to model acute kidney damages and DCE-MRI approaches have been investigated for assessing single-kidney function in both mice and rats [82, 83]. Significant reduction in single kidney GFR (calculated by applying a two compartment renal filtration model) was observed for ischemic kidneys in comparison to contralateral ones in rats [84]. Conversely, for all the other DCE-MRI derived functional parameters (mean transit times and renal blood volume) differences were present, but were not significant between diseased and contralateral kidneys of ischemic rats.

DCE-MRI studies have also been exploited in murine renal transplantations studies for assessing acute rejection and evaluation of immunosuppressive therapy [85]. In this study, a deconvolution approach was applied to the DCE-MRI data and the plasma flow estimate provided restored organ perfusion under Ciclosporin A therapy in syngeneic transplants compared to controls.

### Acknowledgments

The Italian Ministry for Education and Research (MIUR) is gratefully acknowledged for yearly FOE funding to the Euro-BioImaging Multi-Modal Molecular Imaging Italian Node (MMMI).

This publication is based upon work from COST Action PAR-ENCHIMA, supported by European Cooperation in Science and Technology (COST). COST (www.cost.eu) is a funding agency for research and innovation networks. COST Actions help connect research initiatives across Europe and enable scientists to enrich their ideas by sharing them with their peers. This boosts their research, career, and innovation.

PARENCHIMA (renalmri.org) is a community-driven Action in the COST program of the European Union, which unites more than 200 experts in renal MRI from 30 countries with the aim to improve the reproducibility and standardization of renal MRI biomarkers.

#### References

imaging of the kidney: experimental results. Radiology 170(3 Pt 1):713–720. https://doi. org/10.1148/radiology.170.3.2916025

(4):828–836. https://doi.org/10.1111/j. 1365-2133.2011.10465.x

the heptadentate ligand AAZTA. Chemistry 15 (7):1696–1705. https://doi.org/10.1002/ chem.200801803

saturation magnetization transfer. Magn Reson Med 81(2):907–920. https://doi.org/10. 1002/mrm.27442

lung using T1-weighted dynamic contrastenhanced MRI. Magn Reson Med 61 (6):1507–1514. https://doi.org/10.1002/ mrm.21814

Open Access This chapter is licensed under the terms of the Creative Commons Attribution 4.0 International License (http://creativecommons.org/licenses/by/4.0/), which permits use, sharing, adaptation, distribution and reproduction in any medium or format, as long as you give appropriate credit to the original author(s) and the source, provide a link to the Creative Commons license and indicate if changes were made.

The images or other third party material in this chapter are included in the chapter's Creative Commons license, unless indicated otherwise in a credit line to the material. If material is not included in the chapter's Creative Commons license and your intended use is not permitted by statutory regulation or exceeds the permitted use, you will need to obtain permission directly from the copyright holder.

### Noninvasive Renal Perfusion Measurement Using Arterial Spin Labeling (ASL) MRI: Basic Concept

### Min-Chi Ku, Marı´a A. Ferna´ ndez-Seara, Frank Kober, and Thoralf Niendorf

### Abstract

The kidney is a complex organ involved in the excretion of metabolic products as well as the regulation of body fluids, osmolarity, and homeostatic status. These functions are influenced in large part by alterations in the regional distribution of blood flow between the renal cortex and medulla. Renal perfusion is therefore a key determinant of glomerular filtration. Therefore the quantification of regional renal perfusion could provide important insights into renal function and renal (patho)physiology. Arterial spin labeling (ASL) based perfusion MRI techniques, can offer a noninvasive and reproducible way of measuring renal perfusion in animal models. This chapter addresses the basic concept of ASL-MRI.

This chapter is based upon work from the COST Action PARENCHIMA, a community-driven network funded by the European Cooperation in Science and Technology (COST) program of the European Union, which aims to improve the reproducibility and standardization of renal MRI biomarkers. This introduction chapter is complemented by two separate chapters describing the experimental procedure and data analysis.

Key words Magnetic resonance imaging (MRI), Renal, Perfusion, Arterial spin labeling (ASL), Kidney, Mice, Rats

### 1 Introduction

Kidneys are highly perfused organs in the body as they receive 20–25% blood of the resting cardiac output. High blood flow to the kidneys ensures adequate delivery of plasma to be filtered by the glomeruli, a physiological process necessary for the regulation of homeostasis. Therefore, the changes in renal blood flow (RBF) will also change the glomerular filtration rate (GFR) and, consequently, the volume and composition of body fluids. Thus, even a slight change in renal perfusion that exceeds renal autoregulation can have a significant impact on the kidney function. The complex interaction between RBF, renal oxygen delivery, renal oxygen consumption, and GFR made measuring of renal perfusion an attractive target to understand the regulation and homeostasis of renal tissue in health and disease.

Renal perfusion, defined as the blood flow that passes through a unit mass of renal tissue within a given time (mL/min/g) that is also a key determinant of glomerular filtration. A unique aspect of the renal perfusion is that the blood flow is regionally specific. Due to the lack of a reliable technic which allowed to precisely measure the regional RBF, it is roughly estimated that 80–85% of the total renal blood flow supplies renal cortex, 10–15% outer medulla and only 1–5% inner medulla [1].The cortical blood flow is tightly regulated by well-documented mechanisms. Conversely, medullary blood flow regulation is less understood and difficult to measure [2]. The capillary density and surface area is higher in the medulla than in the cortex at least in the mice. It is very challenging to measure the medulla renal perfusion because of the complexity of the microcirculation and its tight connection of arterioles from the cortex. Although inner renal medulla receives less than 1–5% of the total renal blood flow, but changes in this region can still have major effect on electros and water homeostasis and control of arterial blood pressure [3]. Regardless, to understand the physiological regulation of renal perfusion and the role of its dysregulation in separate regions will require methods that allow quantification of renal perfusion both at the whole organ level and at the local tissue level. Most of the imaging techniques for assessing renal perfusion that have been used in humans and animals have significant limitations (Table 1) and by far there is still no gold-standard technique available. For instance, renal perfusion can be measured with dynamic contrast-enhanced MRI (DCE-MRI). However, the major drawback of DCE-MRI is the involvement of contrast media for the measurement of perfusion. The contrast media are easily filtered by renal glomeruli and hence it is difficult to obtain absolute quantification. Other known side effects, such as the development of nephrogenic systemic fibrosis, also limit their application to patients with renal deficits.

A noninvasive, practical, robust, and reproducible method to measure renal perfusion is of great interest but remains elusive, both in the clinic and basic research. The Arterial spin labeling (ASL) MRI technique introduced three decades ago for quantifying tissue perfusion [6, 7], is potentially the only available technique to meet all these needs. ASL is a MRI method generating endogenous contrast specifically related to capillary blood flow. It is therefore free of exogenous contrast media. By magnetically labeling arterial blood proton spins with inversion or saturation as a freely diffusible endogenous tracer, ASL allows for the measurement of RBF. Originally developed for measuring cerebral perfusion, the use of endogenous blood signal to obtain tissue perfusion information is particularly relevant for multiple and longitudinal imaging of patients. This attribute is also important for preclinical

### Table 1 Overview of methods for monitoring renal blood flow [4, 5]

imaging applications, as multiple tail-vein injections and/or catheterizations can cause local inflammation and necrosis resulting in reduced access to veins. The mature and reliable animal models combined with reproducible imaging techniques are an indispensable part of scientifically designed kidney disease studies, and play an important role in resolving the general bottleneck in renal disease treatment that is usually delayed due to difficulties in imaging disease progression.

The ASL approach has been rigorously studied in various preclinical applications including correlating the degree of acute rejection and renal cortical blood flow in rat allograft model [8], comparison of imaging results with histologic examination in ischemia-induced acute kidney injury (AKI) [9], early detection of antiangiogenic treatment responses in a mouse xenograft tumor model [10] and validation of the pathophysiological process of contrast-induced acute kidney injury in rats [11]. In this chapter, we address the basic concept and the developments of renal ASL with the emphasis on kidney disease models in rodents.

This introduction chapter is complemented by two separate chapters describing the experimental procedure and data analysis, which are part of this book.

This chapter is part of the book Pohlmann A, Niendorf T (eds) (2020) Preclinical MRI of the Kidney—Methods and Protocols. Springer, New York.

### 2 Measurement Concept

#### 2.1 Basic Concept of ASL-MRI

ASL provides quantitative parametric images of tissue perfusion. It uses a modified proton magnetization of arterial blood as an endogenous, freely diffusible contrast medium which determines the delivery rate of oxygen and nutrients to the capillary bed and is expressed as the volume of blood traveling through a gram of tissue per minute (mL 100 g1 min1 ). The basic ASL experiment collects two image types, referred to as the "label" (or tag) and "control" (or nonlabeled) image respectively. The label is applied

Fig. 1 Left: All ASL pulse sequences consist of two modules: a preparation module to magnetically label flowing blood, and a readout module to generate paired images. Right: Schematic showing different labeling modules. A single or a train of radiofrequency pulse for labeling followed by a postlabel delay before the image is acquired within one TR.

> using radiofrequency (RF) pulses to alter the longitudinal magnetization of protons in the arterial blood before it enters the desired imaging plane, and an image is collected after a time delay. The control image is acquired at the same time delay without labeling the arterial blood protons. Provided that the inverted magnetization of the inflowing blood is the only difference between the control and label images, a simple subtraction of the label from the control image yields a perfusion-weighted image in which signal intensity is proportional to perfusion. All ASL pulse sequences consist of two components: (1) a preparation module to magnetically label inflowing blood, and (2) an imaging readout scheme to generate paired images of the target tissue under "control" and "tagged" conditions. ASL methods may be classified as to how each of these modules is constructed (Fig. 1).

2.2 Labeling Methods and Strategies for Kidney perfusion Measurement in Rodents

There are two main labeling strategies that have been developed: pulsed ASL (PASL) [12] and continuous ASL (CASL) (Fig. 1, left). Pulsed and continuous ASL labeling methods differ fundamentally in both the spatial extent and the duration of the labeling (Fig. 1, right), and these differences give rise to the strengths and weaknesses of each approach [13]. PASL uses a single short pulse or a limited number of pulses, to invert a thick slab of arterial water spins. It can be further divided into those that label spins asymmetrically with respect to the plane of imaging (Signal Targeting by Alternating Radiofrequency pulses; STAR [14] and Proximal Inversion with Control of Off-Resonance Effects; PICORE [15] or symmetrically (Flow-sensitive Alternating Inversion Recovery; FAIR) [16]. In contrast, CASL techniques use a continuous radiofrequency wave in a plane upstream of the imaging volume for generating a flow-driven inversion of arterial blood flowing through that plane. Later, a modified form called pseudocontinuous arterial spin labeling (pCASL) has been developed, that uses a long series of short RF-pulses together with a strong slice-selection gradient, also resulting in a flow-driven inversion of upstream blood proton spins. pCASL thus combines favorable features of CASL (high signal-to-noise) with those of PASL

#### Table 2 Overview of renal ASL studies in rodents

(lower energy deposition). Today most commercially available ASL products are based on pCASL or a PASL variant. Table 2 shows examples of preclinical studies using ASL for measuring RBF in animal models and list each of the labeling and readout methods.

2.2.1 Pulsed ASL (PASL) The PASL variant (FAIR) is the major method used for measuring renal perfusion in rodents due to its ease of implementation. FAIR uses a single long-duration RF pulse which is applied to invert the arterial magnetization as blood flows through the labeling plane while a gradient is applied in the direction of arterial flow. In FAIR ASL the magnetization is usually prepared using a global inversion which is referred to as the "labeling" and a slice selective inversion which is referred to as "control." The sequence begins with a spatially selective inversion pulse limited to a small region in and around the imaged slice. In the standard labeling scheme the selective inversion slice (control) is in the same plane as the imaging slice with a thickness of 1.5–5 times the thickness of the imaging slice. The inversion pulse is usually slightly wider than the imaging slice to minimize artifacts and ensure uniform inversion. For the "labeling," the same inversion pulse is applied but without the sliceselection gradient. This causes an inversion of spins within the entire sensitive volume of the RF coil (Fig. 2a). By subtracting the control image from the labeling image, a perfusion-weighted image (PWI) is formed, which with the appropriate modeling can be quantified using a perfusion map in units of mL/100 g/min. In various preclinical MRI situations, for instance with dedicated small brain birdcage transmit/receive (TX/RX) resonators, the transmit coil used may not be long enough to provide a sufficiently global inversion for FAIR perfusion measurements.

Fig. 2 Basic principle of arterial spin labeling (ASL) acquisition. It consists of a differential technique during which two acquisitions are carried out: one with the labeling of arterial protons (label image) and a control image. Arterial blood is labelled by inverting its magnetization. After an inflow time (inversion time), which allows the labeled blood to flow into the tissue of interest, an image is acquired with an appropriate readout sequence. (a) FAIR uses a nonselective inversion pulse for the labeled images that becomes selective with the addition of a slice selection gradient for the control. (b) The magnetic labeling of arterial protons is carried out upstream from the volume of interest, by radiofrequency pulses. The labeled protons then migrate via the arterial vessels toward the tissue where they pass from the capillary compartment to the extravascular compartment (example shown: pCASL)

2.2.2 Continuous ASL (CASL) and Pseudo-Continuous ASL (pCASL) CASL uses a continuous flow-driven radiofrequency inversion of blood upstream of the imaging volume. The radiofrequency wave is transmitted either by a dedicated, separate labeling coil or by the volume transmit coil that is also used for imaging. The latter approach has the advantage that no special hardware components are required but the disadvantage of depositing high amounts of RF energy leading to magnetization transfer contrast that has to be properly corrected. One drawback of CASL labeling approach is the consequently high transmitter duty-cycle requirement.

The pCASL technique has evolved from the original CASL technique in both, brain and kidney applications [24]. pCASL uses series of short and high-power RF and gradient pulses instead of the long RF pulses used in CASL to generate a flow-driven inversion. Compared to FAIR, pseudo-continuous ASL (pCASL) provides significant advantages such as increased SNR and better definition of the labeling bolus. FAIR (pulsed ASL) may be attractive for renal imaging because it can label closer to the tissue, but pCASL offers the potential advantage of high SNR of the kidneys and the whole abdomen, which would be desirable in some specific applications (e.g., metastatic). An earlier study compared the performances of FAIR and pCASL in terms of sensitivity, absolute quantification, reproducibility, and flexibility of implementation using healthy mice [18]. It was reported that the application of pCASL at 11.75 T scanner had higher sensitivity compared to FAIR (> þ20%). Finally, since pCASL can be straightforwardly used in multislice mode and in any imaging orientation, it may be considered as a method of choice for mouse kidney perfusion studies.

### 2.3 Imaging Readout A typical ASL MRI acquisition combines an ASL magnetization preparation module (labeling methods) followed by a subsequent rapid imaging module (readouts) to capture the blood flow– weighted contrast. These two aspects of the acquisition are independent of each other. It is important to note that the majority of the ASL imaging developments has focused on the optimization of preparation module of the ASL acquisition [20].

An important prerequisite for successful kidney imaging is the ability to get fast imaging to limit the sensitivity to motion [19]. As such, ASL has been implemented with a wide variety of fast image readout techniques, which in turn can be optimized according to a specific application. The fast image readout techniques can be broadly classified as two-dimensional (2D) or 3D acquisition schemes. Whereas 3D readouts are generally recommended for brain ASL, a consensus of the optimal readout module for renal applications has yet to be established. This consideration has been recently achieved to recommend a 2D readout in humans (unpublished by PARENCHIMA ASL expert panel).

With respect to 2D readouts, to determine the optimal readout for renal ASL, a number of factors should be considered. The optimal readout should have a short echo time (TE) in order to provide the highest image signal-to-noise ratio (SNR), and to reduce the amount of signal dephasing and distortion. At longer TEs, the short tissue intrinsic T2 and T2\* in the abdomen leads to rapid signal drop and the loss of perfusion signal. The ideal readout should be collected in a short shot length to enable multiple slices through the kidney to be acquired prior to the decay of the ASL label, thus enabling whole-kidney perfusion assessment [25]. Many ASL MRI studies have utilized either echo-planar imaging (EPI) or true fast imaging with steady-state free precession (bSSFP; True FISP) readouts [20] though the readout of ASL can be measured with any method. Both EPI and bSSFP readouts are prone to off-resonance artifacts on high field MRI scanners [20] where the majority of animal studies are done by using such high field or ultrahigh field (> ¼ 7 T) scanners. Specifically, B0 inhomogeneity results in increased distortion / ghosting and banding artifacts from EPI and True FISP imaging readouts. These artifacts are particularly problematic for abdominal imaging applications such as kidney where cardiac and respiratory motions well as adipose tissue can make precise shimming difficult. In addition, the increase in T1 relaxation times which is proportional with the field strength can result in spoiled gradient echo images with a lower signal-tonoise ratio (SNR) relative to these other imaging readouts. The low SNR is specifically problematic for ASL MRI techniques as the differential blood flow signal is typically less than 10% of the mean tissue signal.

Though the optimization of EPI acquisitions allow for wholekidney coverage but nevertheless provide sub-optimal and slicedependent perfusion-weighted SNR and is susceptible to off-resonance conditions mentioned above. There is an urgent need to develop a rapid and robust ASL MRI imaging readout that is balanced on blood flow labeling from the ASL labeling module and also avoid B0 inhomogeneities and motion artifacts on high field preclinical MRI scanners.

While several groups independently optimize the bSSFP image readout and shown the advantage over EPI readouts [25], more advanced methods now provide alternative readouts, such as 3D gradient and spin echo (GRASE) [26, 27]. 3D GRASE has superior SNR compared to 2D acquisitions. An alternative to 3D GRASE is 3D rapid acquisition with relaxation enhancement (RARE) combined with a spiral readout trajectory (shorter echo times and superior SNR), but can introduce through-plane blurring due to long readout times.

### 3 Overview of Applications

3.1 ASL for Measuring Kidney Perfusion in Animal Models

The encouraging results using ASL for measuring human renal perfusion mandate further validation of renal ASL measurements using animal models. For example, in vivo interventional swine model has been used with FAIR-ASL technique to track cortical perfusion changes during pharmacologic and physiologic alterations in renal blood flow [28]. For measuring renal hemodynamics and oxygenation in X-ray contrast agent induced acute kidney injury (CIAKI), a rat model was used. Rats were injected with contrast media and followed by ASL measurements [29]. Other models, for instance ischemia induced AKI in mice, was done by transient unilateral clamping of the renal pedicle for either 35 min for moderate AKI or 45 min for severe AKI. Using ASL, renal perfusion impairment was found more pronounced in severe compared to moderate AKI, with severe AKI causing a 33% reduction in renal perfusion [9]. In addition, ASL can be useful for longitudinal assessment of renal perfusion impairment due to acute and chronic renal allograft rejection in translational mouse models [30]. By combining with the histological assessment, perfusion impairment can be validated and revealed its relationship to the severity of

Fig. 3 An example of arterial spin labeling (ASL) acquisition using FAIR-EPI

vascular changes and renal damage. Other applications such as exploring the renoprotective efficacy of drug treatment in ischemia-induced AKI or in renal allograft recipients revealed that the renoprotective effect of the drug was associated with improved renal perfusion [22]. Thus, noninvasive perfusion imaging might provide insights into the underlying disease mechanisms in acute and chronic renal allograft rejection and might facilitate monitoring of novel treatment strategies. Similar approach can be applied to monitor kidney function of most common kidney diseases, such as CIAKI, diabetic nephropathy, and kidney transplant. In addition, ASL techniques can also be used to monitor the therapeutic responses of treatments to common renal hypoxia disease. Figure 3 is an example of ASL measurement using FAIR for labeling and EPI for image readout in a healthy mouse. Table 2 lists all examples of preclinical studies using ASL for measuring RBF in rodent models.

### Acknowledgments

This work was funded in part (Thoralf Niendorf, Sonia Waiczies, Andreas Pohlmann, and Joao Periquito) by the German Research Foundation (Gefoerdert durch die Deutsche Forschungsgemeinschaft (DFG), Projektnummer 394046635, SFB 1365, RENOPROTECTION. Funded by the Deutsche Forschungsgemeinschaft (DFG, German Research Foundation), Project number 394046635, SFB 1365, RENOPROTECTION).

This chapter is based upon work from COST Action PARENCH-IMA, supported by European Cooperation in Science and Technology (COST). COST (www.cost.eu) is a funding agency for research and innovation networks. COST Actions help connect research initiatives across Europe and enable scientists to enrich their ideas by sharing them with their peers. This boosts their research, career, and innovation.

PARENCHIMA (renalmri.org) is a community-driven Action in the COST program of the European Union, which unites more than 200 experts in renal MRI from 30 countries with the aim to improve the reproducibility and standardization of renal MRI biomarkers.

#### References

intravoxel incoherent motion MR imaging to study the renal pathophysiological process of contrast-induced acute kidney injury in rats: comparison with conventional DWI and arterial spin labelling. Eur Radiol 26:1597–1605

rapid and quantitative perfusion technique for high-field MRI. NMR Biomed 27:996–1004

Open Access This chapter is licensed under the terms of the Creative Commons Attribution 4.0 International License (http://creativecommons.org/licenses/by/4.0/), which permits use, sharing, adaptation, distribution and reproduction in any medium or format, as long as you give appropriate credit to the original author(s) and the source, provide a link to the Creative Commons license and indicate if changes were made.

The images or other third party material in this chapter are included in the chapter's Creative Commons license, unless indicated otherwise in a credit line to the material. If material is not included in the chapter's Creative Commons license and your intended use is not permitted by statutory regulation or exceeds the permitted use, you will need to obtain permission directly from the copyright holder.

### Renal pH Imaging Using Chemical Exchange Saturation Transfer (CEST) MRI: Basic Concept

### Dario Livio Longo, Pietro Irrera, Lorena Consolino, Phillip Zhe Sun, and Michael T. McMahon

### Abstract

Magnetic Resonance Imaging (MRI) has been actively explored in the last several decades for assessing renal function by providing several physiological information, including glomerular filtration rate, renal plasma flow, tissue oxygenation and water diffusion. Within MRI, the developing field of chemical exchange saturation transfer (CEST) has potential to provide further functional information for diagnosing kidney diseases. Both endogenous produced molecules as well as exogenously administered CEST agents have been exploited for providing functional information related to kidney diseases in preclinical studies. In particular, CEST MRI has been exploited for assessing the acid-base homeostasis in the kidney and for monitoring pH changes in several disease models. This review summarizes several CEST MRI procedures for assessing kidney functionality and pH, for monitoring renal pH changes in different kidney injury models and for evaluating renal allograft rejection.

This chapter is based upon work from the COST Action PARENCHIMA, a community-driven network funded by the European Cooperation in Science and Technology (COST) program of the European Union, which aims to improve the reproducibility and standardization of renal MRI biomarkers. This introduction chapter is complemented by two separate chapters describing the experimental procedure and data analysis.

Key words Magnetic resonance imaging (MRI), Kidney, Mice, Rats, pH, Iopamidol, Chemical exchange saturation transfer (CEST), Acute kidney injury, Ischemia–reperfusion injury, Renal damage

### 1 Introduction

Renal dysfunction is recognized as a significant health problem originating from a variety of causes leading to acute or chronic kidney diseases. Predisposition to acute kidney injury and premature mortality are frequent outcomes for chronic kidney diseases [1]. Moreover, according to the Global Burden of Disease 2013 study, the age-standardized death rates for chronic kidney diseases showed one of the highest increases in the last two decades [2]. Therefore, reliable and early diagnosis of acute and chronic kidney diseases is needed to preserve renal functionality and

Andreas Pohlmann and Thoralf Niendorf (eds.), Preclinical MRI of the Kidney: Methods and Protocols, Methods in Molecular Biology, vol. 2216, https://doi.org/10.1007/978-1-0716-0978-1\_14, © The Author(s) 2021

improve patients' outcome. Despite these premises, current clinical evaluation of renal function is still based on measurement of serum creatinine, which is well known to have several limitations. In fact, an elevated serum creatinine concentration is only discovered at late stages of the disease when renal functionality is already compromised. Therefore, novel noninvasive imaging approaches are needed for a more accurate and early diagnosis of renal physiology. Magnetic resonance imaging (MRI) has been exploited for tissue anatomic imaging, owing to the high spatial resolution and lack of ionizing radiation. In research, novel MRI techniques have been investigated for assessing multiple functional parameters of the kidneys, including perfusion, filtration, oxygenation and tissue elasticity [3–11]. Several reviews covering the role of these MRI approaches for assessing renal functionality have been published [12–15]. This additional information should enable a better characterization of acute and chronic kidney diseases in comparison to standard urine and serum-based assays.

Recently, Chemical Exchange Saturation Transfer (CEST) MRI has emerged as a novel approach for functional and molecular imaging with great promise for clinical translation [16–18]. In addition, CEST MRI pH imaging has emerged as a valuable approach for assessing extracellular pH values in several tissues, including kidneys and tumors, providing the highest accuracy and spatial resolution achievable so far [19]. Owing to the key role of the kidneys in maintaining the acid-base homeostasis, CEST MRI pH mapping has emerged as a novel and promising approach in monitoring kidney functionality.

In this chapter, we will address the basic concept and the developments of renal CEST pH imaging with the emphasis on kidney disease models in rodents.

This introduction chapter is complemented by two separate chapters describing the experimental procedure and data analysis, which are part of this book.

This chapter is part of the book Pohlmann A, Niendorf T (eds) (2020) Preclinical MRI of the Kidney—Methods and Protocols. Springer, New York.

### 2 Measurement Concept

2.1 Basic Concept of Chemical Exchange Saturation Transfer (CEST) Imaging

CEST is a new technique that enables the indirect detection of molecules possessing mobile protons in exchange with water [20]. Because of this, CEST makes MRI sensitive to endogenous or exogenous molecules that possess suitable protons [21–32]. The generation of contrast is based on a selective irradiation with a radiofrequency pulse at the specific absorption frequency of the exchanging proton, followed by a subsequent transfer, due to chemical exchange with bulk water, of the saturated signal. Upon the application of a long saturation RF (radiofrequency) pulse, the water signal decreases, allowing the amplification of the labeled protons from the molecule to water. As a consequence, low concentration molecules can be specifically and indirectly detected, owing to the frequency selective irradiation of their mobile proton pool(s) [33–38].

Several classes of CEST contrast agents can be used as pH responsive agents and exploited for in vivo experiments.

Several molecules, including natural occurring molecules, possess exchangeable protons that can be indirectly detected within the CEST approach [39]. Most of these labile protons have chemical exchange rates that show a strong pH dependence. Hence differences in CEST contrast can be exploited for assessing the pH of the solution where they are dissolved. Due to the concomitant contribution of concentration and exchange rate to the observed CEST contrast, several approaches have been proposed to rule out the concentration term, based on the ratiometric method, that is, taking the ratio of the observed contrast, providing accurate pH readouts [40, 41]. Most of the pH responsive diamagnetic molecules investigated so far in vivo are iodinated contrast agents used for radiographic procedures, since they have been used in the last 40 years in clinical examinations owing to their high safety profiles [42, 43]. The presence of amide groups in their chemical structure allow them to be exploited as CEST MRI contrast agents, upon selective irradiation of the mobile proton pool [44].

Among the radiographic contrast media, Iopamidol (Isovue®, Bracco Imaging, Italy) (Fig. 1) was the first agent exploited for mapping pH thanks to the presence of two amide groups with different resonance frequencies, at 4.2 ppm and at 5.5 ppm, respectively, that can be selectively irradiated [45–47]. This chemical peculiarity led to the development of a ratiometric approach based on the ratio of the CEST effects at these two frequencies to provide accurate pH measurements. In vitro experiments showed a high pH responsiveness within the physiological range and accurate pH measurements (Fig. 2a, b). Following intravenous administration, it can provide selective contrast at specific frequency offsets and from the ratio of these two parametric maps it is possible to measure renal pH map (Fig. 2c–f). To validate the approach and test the pH-responsiveness of the agent, induced alkalinisation or acidification of urine was obtained in mice upon providing acidic or alkaline drinking water for a week. For control mice, mean pH values calculated for cortex, medulla and calyx regions were: 7.0 0.11; 6.85 0.15 and 6.6 0.20, respectively. In mice drinking acidic water a significant acidification of renal pH values was obtained. Conversely, a marked increase of renal pH values was

### 2.2 CEST Contrast Agents for Imaging Kidneys and Mapping pH

2.2.1 Diamagnetic CEST Agents

Fig. 1 Chemical structures of the reported pH-responsive CEST agents investigated for renal pH mapping

Fig. 2 Iopamidol ratiometric curve obtained from the rations of the CEST contrast upon irradiating at 4.2 and 5.5 ppm, respectively, showing the high pH responsiveness of Iopamidol. Irradiation power levels of 3 μT, Iopamidol 30 mM, 7.05 T, 310 K, irradiation time 5 s (a). Calculated pH values obtained by the ratiometric method (Iopamidol 30 mM, 7.05 T, 310 K, irradiation time 5 s, B1 3 μT) are compared with the values read on the pH-meter (b). Representative images of in vivo renal pH mapping showing the anatomical image (c) and the CEST contrast parametric maps overlaid to the anatomical image at 4.2 ppm (d) and 5.5 ppm (e) and the observed pH map (f) obtained by ratioing maps (d) and (e) and using the calibration curve in (a) for calculating the pH values. (Adapted with permission from Magnetic Resonance in Medicine 2011 (Iopamidol as a responsive MRI-chemical exchange saturation transfer contrast agent for pH mapping of kidneys: in vivo studies in mice at 7 T. Volume: 65, Issue: 1, Pages: 202–211, DOI: https:/doi.org/10.1002/mrm.22608))

observed upon alkaline water administration. Both observations demonstrate the capability to measure in vivo renal pH changes with CEST MRI.

Since the accuracy in measuring pH depends on several factors, including chemical exchange rate, irradiation conditions (saturation power and duration), and main magnetic field, quantification of multisite pH-dependent chemical exchange properties is needed to improve pH accuracy [48]. This characterization of chemical exchange rates and optimal irradiation RF pulses led to the development of an optimized saturation for each single amide proton pools, hence resulting in a higher pH sensitivity. As a result, the capability to measure pH was demonstrated also at magnetic field strengths of 4.7 T, where amide resonances partially overlap. In the study of Wu and coworkers, they demonstrated an improved pH sensitivity, extending the pH detection range from 5.5 to 7.5, with high resolution pH maps of the kidneys in healthy rats (Fig. 3) [49]. More recently, the exploitation of the same ratiometric approach for measuring in vivo pH was also demonstrated at magnetic fields as low as 3 T, which still preserved good pH accuracy [50].

Fig. 3 Extension of pH detection range using the modified ratiometric analysis (red circles) versus the standard ratiometric approach (blue squares) (a). Representative in vivo ratiometric map for the proposed approach (b). pH map shows the renal pH gradually decreasing from the cortex, medulla to calyx (c). Maps are overlaid on the corresponding T2-weighted image. (Adapted with permission from Magnetic Resonance in Medicine 2018 (A generalized ratiometric chemical exchange saturation transfer (CEST) MRI approach for mapping renal pH using Iopamidol, Volume: 79, Issue: 3, Pages: 1553–1558, DOI: https:/doi.org/10.1002/mrm.26817))

Another similar iodinated contrast media, Iopromide (Ultravist®, Bayer Healthcare, Germany), was also demonstrated and compared to Iopamidol for measuring pH, with a pH sensitivity that was not substantially different, although Iopamidol allows for more precise pH measurements [51].

The exploitation of the ratiometric approach requires two distinguishable protons pools on the same molecule, therefore it is limited to a select group of compounds possessing this feature. To overcome this limitation, Longo and Sun proposed a novel approach based on the irradiation of a single pool under different radiofrequency powers [52]. The proposed approach, called ratio of RF power mismatch or RPM, was demonstrated by using Iobitridol (Omnipaque®, GE Healthcare, USA), a radiographic contrast medium possessing only one amide proton pool resonating at 5.5 ppm (Fig. 1). Since the measured CEST contrast is dependent on both pH and irradiation power (B1), a strong pH dependence was demonstrated by calculating the ratio of the CEST contrast at two different B1 power levels. In comparison to the conventional ratiometric approach, a good pH accuracy and an even higher pH sensitivity were demonstrated. When investigated in healthy kidneys, the measured pH values upon iobitridol injection strongly correlated with the pH values obtained following iopamidol injection. Recently, to overcome the specific absorption rate (SAR) limitations when using a continuous wave (CW) irradiation of the mobile proton pools, a pulsed ratiometric approach has been exploited and tested on the Iodixanol (Visipaque®, GE Healthcare, USA) X-ray contrast medium for CEST pH imaging [53].

Diamagnetic contrast agents possess resonances that are close in frequency to the bulk water signal resulting in a reduction of sensitivity (due to an inefficient labeling selectivity) when moving from high to low magnetic fields. Following previous investigations, McMahon and colleagues developed new diamagnetic systems based on the imidazole, salicylate or anthranilate moieties which possess mobile protons shifted very far from the bulk water peak [54–56]. In particular, candidate compounds derived from imidazole-4,5-dicarboxamides (#5 and #8, Fig. 1) show a CEST signal shifted up to 7.8 ppm upfield by exploiting the presence of intramolecular bond shifted hydrogens. Besides the large chemical shift, a good pH sensitivity and a high-water solubility made compound #5 suitable as a pH sensor and so it has been tested in vivo [57]. pH imaging of the kidneys at 11.7 T resulted in average pH values for the whole kidneys of 6.5 0.1, consistent with renal pH values reported by radiographic contrast agents (Fig. 4).

2.2.2 Paramagnetic CEST Agents Paramagnetic chemical exchange saturation transfer (paraCEST) agents typically consist of a paramagnetic metal ion and an organic chelate based on a macrocyclic cage due to the high kinetic stability of these metal chelates to prevent the release of the toxic free lanthanide cation [58]. The CEST signal is therefore generated by selective irradiation of the bound water molecule, or of the slowly exchanging ligand protons, such as hydroxyl, amine, or amide groups [59–62]. The main advantage of paraCEST agents in comparison to diamagnetic ones relies in the exceptionally large chemical shifts (MR frequency relative to the water frequency) due to the hypershift contribution of the lanthanide metal ion. This large range of chemical shifts increases the specificity of the exchanging proton pools, hence reducing the adverse contributions of direct saturation and of the endogenous semisolid macromolecular effects [63]. Since the chemical exchange rate of the paraCEST agent can be altered by environmental factors such as pH, changes in CEST amplitudes can be used as well for deriving pH values.

> The first demonstration of CEST detection in kidneys using a paraCEST agent (TmDOTA-4AmC, Fig. 1) was provided by Vinogradov et al., showing good detectability in the whole kidney regions [64]. Later on, a Europium paraCEST pH responsive agent (europium(III) DO3A-tris(amide) complex, Fig. 1) was exploited to measure pH in mouse kidneys [65]. Owing to its chemical structure, a quite large frequency shift of the ion-bound water molecule due to the delocalization of negative charge coming from deprotonation of phenolic residue was observed. This shift is pH dependent, therefore an alkalinisation from 6.0 to 7.6 at 298 K leads to a 4 ppm shift in frequencies, from 50.5 to

Fig. 4 pH calibration curve of compound #5 (a) and calibration plot using all tubes in the phantom showing experimental versus calculated pH (b). Experimental conditions: CEST data were obtained at 6.25 mM, 25 mM, or 50 mM concentration, saturation time ¼ 3 s, saturation power ¼ 5.9 μT and 37 C. pH measurements were made with a precision of 0.1 unit. T2-weighted image (c) and pH map (d) following administration of compound #5. (Adapted with permission from Contrast Media Molecular Imaging 2016 (Developing imidazoles as CEST MRI pH sensors, Volume: 11, Issue: 4, Pages: 304–312, DOI: https:/doi.org/10.1002/cmmi.1693))

> 54.5 ppm, respectively, that can be exploited for assessing pH. By exploiting a 9.4 T MR scanner and ensuring stable temperature homogeneity, in vivo pH measurements were feasible in kidneys upon the administration of a dose of 0.4 mmol/kg [66].

2.3 Imaging Readout CEST MRI techniques includes continuous wave (CW) or pulsed train RF saturation to prepare the magnetization followed by a fast image readout such as echo planar imaging (EPI), Rapid Imaging with Refocused Echoes (RARE) and/or fast imaging with steadystate precession (FISP) [45, 67, 68]. More advanced methods now provide more sophisticated sampling schemes, such as 3D radial or spiral acquisitions, but those are currently available only on clinical scanners and not on preclinical scanners, thus limiting CEST acquisition in mice and rats to single slice based approaches [69, 70]. However, recent studies have utilized multislice acquisition schemes on preclinical scanners as well [67, 71, 72].

2.3.3 Spin-Echo Echo Planar Imaging (EPI) CEST MRI is often combined with EPI acquisition, which provides fast image readout after a relatively long RF saturation. Whereas both gradient echo and spin echo EPI have been used, SE EPI is often preferred for body application because it is less susceptible to mild magnetic field inhomogeneity distortions that are common in body applications. The use of an EPI readout also enables multislice acquisitions in reasonable times.

### 3 Overview of Applications

CEST imaging has been exploited for assessing renal pH values in healthy and in several models of renal damage, including either bilateral or unilateral acute kidney injury models. Both endogenous CEST approaches and exogenous CEST approaches have been proposed and validated in vivo.

3.1 Endogenous CEST Methods for Assessing Renal Diseases Since diabetic nephropathies (DNs) are associated with changes in renal metabolites, the utility of CEST MRI to detect changes in glucose/glycogen hydroxylic protons was investigated in murine models of diabetic nephropathies. The study was conducted longitudinally from 8 to 24 weeks on two groups of diabetic mice and on nondiabetic mice as control. Based on the variation of glucose/ glycogen composition and the consequent CEST effects measured in kidney regions, a significantly increased CEST of hydroxyl metabolites was observed in diabetic mice during the progression of DN [73].

3.2 CEST pH Imaging for Assessing Renal Diseases Most of the studies reported so far have exploited the pH-responsiveness of Iopamidol for investigating the changes in pH homeostasis following renal injuries. In one of these studies, the pH evolution in an acute kidney injury model induced by intramuscular glycerol injection and consequent rhabdomyolysis was monitored [74]. Renal pH maps acquired at 1, 3, 7, 14, 21 days after the injury reported a marked increase of pH values during the damage evolution up to 3–7 days, followed by recovery of pH toward baseline values at 14 and 21 days. These results were in good agreement with morphological and Blood Urean Nitrogen (BUN) quantification supporting this CEST MRI pH mapping approach for investigating renal function. Furthermore, along with the progression of the damage, a reduction of pixels where Iopamidol was detectable was observed, suggesting that also the percentage of CEST-detected pixels can be used as an imaging biomarker of renal injury.

In another study, Longo and coworkers investigated a unilateral kidney ischemia reperfusion injury (KIRI) model to validate MRI-CEST pH mapping for assessing single kidney functionality [75]. Two different times of ischemia duration, 20 and 40 min, were applied to model moderate or severe KIRI, respectively. Following the damage evolution at days 0, 1, 2, 7, a significant increase in renal pH values was observed even at day 1 in both cases. Furthermore, in the following MRI acquisitions a clear distinction between moderate and severe AKI is possible since a recovery of normal acid-base balance was observed only in the moderate KIRI mice whereas in severe KIRI mice the increased pH values did not restore to baseline values (Fig. 5a, b). In additions, as in the previous study, the percentage of CEST detected pixels, representing a marker of the filtration fraction, showed significant differences between the injured kidneys and the contralateral ones, reflecting the different evolution of moderate-to-severe damage (Fig. 5c, d).

3.3 GlucoCEST Imaging for Assessing Renal Diseases Since the capability of CEST imaging to detect hydroxyl protons, native glucose has also been proposed as a tracer for MRI Gluco-CEST imaging [76–78]. Besides its exploitation in oncological applications, it has been applied for monitoring allograft rejection [79]. In this study, Brown Norway rat kidneys were transplanted into Lewis rats and imaged 4 days following the surgery, before and after glucose administration. By calculating the cortex-to-medulla CEST ratio (CESTR), dependent on the accumulation of the administered glucose, rats that underwent allogeneic transplant showed the highest CESTR values compared to syngeneic transplant group and to control mice.

Fig. 5 MRI-CEST pH mapping detects renal pH changes and regional distribution of damage after moderate and severe unilateral kidney ischemia reperfusion injury (KIRI) showing clamped (right) and contralateral normal kidney (left). Representative MRI-CEST pH maps overimposed onto anatomical images before and after moderate (a) and severe (b) KIRI at different time points (day 1, day 2, and day 7) showing pronounced alkalinization and reduced filtration (noncolored pixels within the renal region, c, d) of the pH-responsive contrast agent in clamped kidney in comparison to contralateral kidney. (Adapted with permission from: NMR in Biomedicine 2017 (Noninvasive evaluation of renal pH homeostasis after ischemia reperfusion injury by CEST-MRI, Volume: 30, Issue: 7, DOI: https:/doi.org/10.1002/nbm.3720))

### Acknowledgments

The Italian Ministry for Education and Research (MIUR) is gratefully acknowledged for yearly FOE funding to the Euro-BioImaging Multi-Modal Molecular Imaging Italian Node (MMMI).

This chapter is based upon work from COST Action PARENCH-IMA, supported by European Cooperation in Science and Technology (COST). COST (www.cost.eu) is a funding agency for research and innovation networks. COST Actions help connect research initiatives across Europe and enable scientists to enrich their ideas by sharing them with their peers. This boosts their research, career, and innovation.

PARENCHIMA (renalmri.org) is a community-driven Action in the COST program of the European Union, which unites more than 200 experts in renal MRI from 30 countries with the aim to improve the reproducibility and standardization of renal MRI biomarkers.

#### References

(1):79–87. https://doi.org/10.1006/jmre. 1999.1956

(2014) Chemical exchange saturation transfer (CEST): an efficient tool for detecting molecular information on proteins' behaviour. Analyst 139(11):2687–2690. https://doi.org/10. 1039/c4an00346b

(6):e3920. https://doi.org/10.1002/nbm. 3920

(2016) In vitro and in vivo assessment of nonionic iodinated radiographic molecules as chemical exchange saturation transfer magnetic resonance imaging tumor perfusion agents. Investig Radiol 51(3):155–162. https://doi. org/10.1097/RLI.0000000000000217

hydrogen. Contrast Media Mol Imaging 10 (1):74–80. https://doi.org/10.1002/cmmi. 1597

saturation transfer MRI. Magn Reson Med 76 (5):1531–1541. https://doi.org/10.1002/ mrm.26045

Open Access This chapter is licensed under the terms of the Creative Commons Attribution 4.0 International License (http://creativecommons.org/licenses/by/4.0/), which permits use, sharing, adaptation, distribution and reproduction in any medium or format, as long as you give appropriate credit to the original author(s) and the source, provide a link to the Creative Commons license and indicate if changes were made.

The images or other third party material in this chapter are included in the chapter's Creative Commons license, unless indicated otherwise in a credit line to the material. If material is not included in the chapter's Creative Commons license and your intended use is not permitted by statutory regulation or exceeds the permitted use, you will need to obtain permission directly from the copyright holder.

# Chapter 15

### Sodium (23Na) MRI of the Kidney: Basic Concept

### James T. Grist, Esben Søvsø Hansen, Frank G. Zo¨ llner, and Christoffer Laustsen

### Abstract

The handling of sodium by the renal system is a key indicator of renal function. Alterations in the corticomedullary distribution of sodium are considered important indicators of pathology in renal diseases. The derangement of sodium handling can be noninvasively imaged using sodium magnetic resonance imaging (23Na MRI), with data analysis allowing for the assessment of the corticomedullary sodium gradient. Here we introduce sodium imaging, describe the existing methods, and give an overview of preclinical sodium imaging applications to illustrate the utility and applicability of this technique for measuring renal sodium handling.

This chapter is based upon work from the COST Action PARENCHIMA, a community-driven network funded by the European Cooperation in Science and Technology (COST) program of the European Union, which aims to improve the reproducibility and standardization of renal MRI biomarkers. This introduction chapter is complemented by two separate chapters describing the experimental procedure and data analysis.

Key words Magnetic resonance imaging (MRI), Kidney, Mice, Rats, 23Na, Sodium

### 1 Introduction

In this chapter, we describe renal sodium imaging, as acquired with magnetic resonance imaging (MRI). An introduction to the many methods that can be used to acquire sodium (23Na) signal is presented, as well as the postprocessing required to quantitatively describe the signal in terms of relaxation or concentration. Example applications of sodium renal imaging are discussed—ranging from the action of diuretic drugs to early sodium handling alterations in acute kidney injury. This introduction is complemented by two separate chapters describing the experimental procedure and data analysis, which are part of this book.

This chapter is part of the book Pohlmann A, Niendorf T (eds) (2020) Preclinical MRI of the Kidney—Methods and Protocols. Springer, New York.

Andreas Pohlmann and Thoralf Niendorf (eds.), Preclinical MRI of the Kidney: Methods and Protocols, Methods in Molecular Biology, vol. 2216, https://doi.org/10.1007/978-1-0716-0978-1\_15, © The Author(s) 2021

### 2 Measurement Concept

birdcage), the dual RF resonator allows for both a homogeneous transmit B1 + field and a highly sensitive receive B1 field [12, 13].

Kalacycian et al. developed a homogeneous transmit-only volume resonator and a highly sensitive RO surface resonator for renal 23Na-MRI for 3 T or 9.4 T MRI scanners [14]. For this setup the sensitivity correction needs to be performed only for the receive profile. Furthermore, the dual resonator allows for increased 23Na-MR signal sensitivity due to the localized signal detection using RO surface coils. For the dedicated task of bilateral kidney imaging, a dual resonator system (TORO) including a two-element phased array at 9.4 T and a receiver saddle shaped RF resonator at 3 T was implemented [13]. These systems were tailored for measuring the absolute renal TSC in rat kidney models with high spatiotemporal resolution, and high concentration measurement accuracy.

2.2.2 Total Sodium Imaging Total sodium imaging is typically performed using a gradient echobased technique. To ensure that the conversion from the imaging signal to the sodium concentration is accurate, a repetition time of approximately 2.5 times the T1 of tissue sodium (approximately 25 ms at 3 T, and longer at higher field strengths) is essential. Further optimization of the sequence readout can be undertaken to capture the fast T ∗ 2 relaxation of the sodium nucleus. For example, an ultrashort echo time (UTE) approach such as radial or twisted projection imaging is preferred over conventional Cartesian strategies. Due to the inherent low SNR of sodium acquisitions, it is common to average a number of sodium scans from the same subject; however, this may lead to long scan times.

An example of total sodium imaging pulse sequence, and renal images, are provided in Figs. 1a and 2.

2.2.3 Triple Quantum Filtering (TQF) Due to the 3 2 spin of the sodium nucleus, a biexponential T ∗ 2 is exhibited by the isotope. A single excitation with a 90-degree hard pulse causes a transition of spins between the outer spin states (3 2 to 1 2 and - 1 2 to - 3 2) as well as the inner states (1 2 to - 1 2 ). The outer spin states exhibit a faster T ∗ 2 decay than the inner. Using phase-cycled radiofrequency pulses it is possible to isolate the signal of slow-moving spins in the intracellular compartment, while removing the extracellular contribution [15].

There are a number of limitations to TQF though. The challenges include long scan times and high energy deposition required due to a large SNR penalty of the TQF technique. Furthermore, inhomogeneity in main magnetic field distribution can lead to spurious results, with corrections in reconstruction required to counter the B0-nonuniformity effects [10]. An example triple quantum filtering sodium imaging pulse sequence is shown in Fig. 1b.

Fig. 1 Sodium imaging pulse sequences. (a) Total sodium imaging sequence using a 3D GRE readout. TR ¼ Repetition time, TE ¼ Echo time. (b) Triple Quantum Filtered sodium imaging pulse sequence using a 3D GRE readout. α ¼ phase of pulse 1, T1 ¼ Mixing time 1, T2 ¼ Mixing time 2. (c) Inversion prepared sodium imaging pulse sequence. TI ¼ Inversion preparation time (determined experimentally)

Fig. 2 Example rodent renal 1 H (left) and sodium (right) imaging acquired at 9.4 T

2.2.4 Inversion Recovery Utilizing the difference in T1 between bound and unbound sodium, it is possible to null the signal contribution fluid compartment using an inversion pulse, whilst retaining the bound signal, as described through Eqs. 1 and 2 [7].

$$S(\text{TI}) = S\_0 \left( 1 - \exp\left(-\frac{\text{TI}}{T\_1}\right) \right) \tag{1}$$

$$\text{TI}\_{\text{null}} = \ln \left( \mathcal{Q} \right) T\_1^{\text{Fluid}} \tag{2}$$

where TI is the inversion time, T1 is the relaxation constant of free sodium, S0 is the initial sodium NMR signal, and S(TI) is the signal at TI. An example of an inversion recovery sodium imaging technique is illustrated in Fig. 1c.

2.2.5 T2\* Mapping T ∗ 2 mapping is performed with either a multiecho based gradient echo sequence, or many individual scans with an incremented echo time (TE). It is currently thought that the shorter T ∗ 2 component (1-5 ms at 3 T) is derived from the intracellular compartment, and the longer (10–30 ms at 3 T) from the extracellular compartment [9, 16]. T2\* mapping therefore could be sensitive to changes in the intracellular–extracellular renal sodium balance.

> A multiecho or incremented echo series can be used to derive both the pool sizes and T ∗ 2 time constant for each compartment through Eq. 3.

$$S(\text{TE}) = \left( a \exp\left( -\frac{\text{TE}}{T\_{2,\text{Short}}^{\ast}} \right) + b \exp\left( -\frac{\text{TE}}{T\_{2,\text{Long}}^{\ast}} \right) \right) \quad (3)$$

where S(TE) is the signal at a given echo time, a and b are the intraand extracellular pool sizes, respectively, and T ∗ 2,Short and T ∗ 2,Long are the biexponential relaxation constants.

ensure short TE in UTE imaging a hard RF pulse is commonly used to excite all spins in the imaging volume. A number of trajectories used for non-Cartesian data readout are available from literature, for example radial imaging, twisted projection imaging (TPI), density adapted radial (DAR) imaging, and 3D CONES, briefly described in the following.

Twisted projection imaging (TPI) was first described by Fernando Boada in 1997, as an extension to 2D TWIRL imaging [18]. The technique utilizes a hybrid radial-spiral sampling of k-space, allowing for acceleration in comparison to radial imaging with an underdamping factor defined during trajectory construction. TPI requires high gradient slew rates for image acquisition, due to the sharp transition between the radial and spiral potions of the trajectory. Further optimizations to the trajectory can be made by optimizing maximum gradient amplitudes throughout scan time [19].

Density adapted radial (DAR) trajectories are described in work by Nagel, whereby concentric rings of k-space are acquired with a radial readout. DAR allows for under sampling of k-space, shortening acquisition times in comparison to radial imaging [20].

The 3D cones approach was initially described for UTE proton imaging and has found applications in sodium acquisitions [21, 22]. Due to the high k-space efficiency of the trajectory, it is possible to acquire a full imaging volume in a very short time, leading to rapid high resolution imaging [23]. However, due to the high k-space efficiency, the SNR of each volume is lower than a radial acquisition; therefore, it is common practice to perform a number of spatial averages in a single acquisition.

2.4 Reconstruction There are three main methods for reconstruction of sodium imaging data, depending on the acquisition scheme used. If Cartesian data is acquired, a Fourier transform is applied in all dimensions (x, y,z,t). If non-Cartesian acquisitions are employed more complex reconstruction strategies such as gridding or nonuniform fast Fourier transforms are used. Compensation for the oversampling of the center of k-space is performed in non-Cartesian imaging using precalculated density compensation functions (see Note 1).

2.5 Sodium Phantoms

External calibration phantoms are used to convert sodium imaging data to quantitative maps, with voxels fitted to a calibration curve formed from measurements of signal from known concentration phantoms as well as from regions of noise in the image.

Phantoms are generally made from agarose doped with sodium chloride, to ensure a T1 and T2 that mimic that of tissue. However, other designs are also possible. Phantoms are placed within the imaging field of view before the start of the experiment to ensure that signal is acquired from both tissue and calibration standards in the same exam.

2.6 Segmentation Segmentation of the kidney is commonly achieved through region of interests initially drawn on proton anatomical imaging and transferred to the sodium images. Other methods have been proposed for the segmentation of the renal system, in particular to measure the corticomedullary sodium gradient, focusing on defining layers of the kidney from an initial whole kidney region [24].

### 3 Overview of Applications

3.1 Total sodium Imaging for Assessing Alterations in the Corticomedullary sodium Gradient

The corticomedullary sodium gradient has been studied with sodium imaging in a number of healthy and pathological scenarios [3–5, 25–29]. Broadly, the current preclinical literature focuses either on the action of diuretic drugs upon the sodium gradient, or upon acute tubular necrosis. The action of furosemide has been estimated by acquiring longitudinal imaging before and after the introduction of the acute diuretic. Furosemide acts as a loop diuretic, halting reabsorption of sodium through the loop of Henley, leading to acute diuresis. Results from studies have shown the capability of sodium MRI to detect the acute renal sodium alterations induced by furosemide, revealing a flattening of the corticomedullary sodium gradient within minutes of administration [3, 4, 26].

Studies focusing upon the early formation of acute kidney injury through tubular necrosis have revealed alterations in the sodium gradient, with a decreased gradient in the damaged kidney [5].

3.2 Potential Future Applications of sodium Imaging Although the literature is limited to the current applications of renal sodium imaging, there is a large scope for further preclinical studies demonstrating the power of this technique to probe disease formation. In particular, the formation of acute kidney disease and chronic kidney disease remain prime targets for further 23Na MRI studies.

3.3 Acute Kidney Disease Acute kidney disease is heralded by a sudden loss of renal function, and impaired reabsorption of electrolytes such as urea and sodium [30]. Indeed, this is a key clinical problem, with a number of patients diagnosed upon admission with AKI and leads to high mortality rates. AKI can occur over a number of hours or days, and is heralded by an decrease of the glomerular filtration rate, leading to retention of nitrogenous waste [31]. Due to the increase in waste retention, common methods for diagnosis are serumbased measures of creatinine clearance and blood nitrogen [31]. Further to the increase in the retention of waste, sodium reabsorption is impaired in AKI. An initial sodium imaging study showed a flattening of the corticomedullary sodium gradient [32]. A particular challenge in diagnosis and management of AKI is the direct estimation of renal health and function pre- and posttherapy, with many imaging methods either requiring the

sodium gradient, to differentiate the earliest changes in the metabolic and functional mismatch that occur in the renal system prior to the development of kidney disease [28, 32, 36]. The use of a combined multimodal anatomical and functional imaging approach with methods such as sodium, arterial spin labeling, BOLD, dynamic contrast enhanced imaging, and diffusion weighted imaging may provide a more complete picture of the alterations in renal function in relation to chronic kidney disease [24, 36–42].

### 4 Notes

1. Example gridding and density compensation calculation functions used for the reconstruction of data acquired with non-Cartesian trajectories can be found at https://www. ismrm.org/mri\_unbound/sequence.htm.

### Acknowledgments

The authors would like to thank Andreas Pohlmann and Thoralf Niendorf for assistance in preparing this chapter (Max-Delbru¨ck Center for Molecular Medicine in the Helmholtz Association, Berlin, Germany.

This chapter is based upon work from COST Action PARENCH-IMA, supported by European Cooperation in Science and Technology (COST). COST (www.cost.eu) is a funding agency for research and innovation networks. COST Actions help connect research initiatives across Europe and enable scientists to enrich their ideas by sharing them with their peers. This boosts their research, career, and innovation.

PARENCHIMA (renalmri.org) is a community-driven Action in the COST program of the European Union, which unites more than 200 experts in renal MRI from 30 countries with the aim to improve the reproducibility and standardization of renal MRI biomarkers.

#### References

MRI of rodent kidney at 3T by using a double-tuned transceiver resonator system. In: International Society for Magnetic Resonance in Medicine. Salt-Lake City, p 4355

kidneys at 7 T: preliminary results. Eur Radiol 24:494–501

Open Access This chapter is licensed under the terms of the Creative Commons Attribution 4.0 International License (http://creativecommons.org/licenses/by/4.0/), which permits use, sharing, adaptation, distribution and reproduction in any medium or format, as long as you give appropriate credit to the original author(s) and the source, provide a link to the Creative Commons license and indicate if changes were made.

The images or other third party material in this chapter are included in the chapter's Creative Commons license, unless indicated otherwise in a credit line to the material. If material is not included in the chapter's Creative Commons license and your intended use is not permitted by statutory regulation or exceeds the permitted use, you will need to obtain permission directly from the copyright holder.

### Hyperpolarized Carbon (13C) MRI of the Kidneys: Basic Concept

### Cornelius von Morze, Galen D. Reed, Zhen J. Wang, Michael A. Ohliger, and Christoffer Laustsen

### Abstract

Existing clinical markers for renal disease are limited. Hyperpolarized (HP) 13C MRI is based on the technology of dissolution dynamic nuclear polarization (DNP) and provides new avenues for imaging kidney structure, function, and most notably, renal metabolism, addressing some of these prior limitations. Changes in kidney structure and function associated with kidney disease can be evaluated using [13C]urea, a metabolically inert tracer. Metabolic changes can be assessed using [1-13C]pyruvate and a range of other rapidly metabolized small molecules, which mainly probe central carbon metabolism. Results from numerous preclinical studies using a variety of these probes demonstrated that this approach holds great potential for monitoring renal disease, although more work is needed to bridge intelligently into clinical studies. Here we introduce the general concept of HP 13C MRI and review the most relevant probes and applications to renal disease, including kidney cancer, diabetic nephropathy and ischemic kidney injury.

This chapter is based upon work from the PARENCHIMA COST Action, a community-driven network funded by the European Cooperation in Science and Technology (COST) program of the European Union, which aims to improve the reproducibility and standardization of renal MRI biomarkers. This introduction chapter is complemented by two separate chapters describing the experimental procedure and data analysis.

Key words Dynamic nuclear polarization, Kidney, Preclinical models, Carbon-13

### 1 Introduction

While renal MRI has traditionally been limited largely to morphologic depiction of vascular disease and neoplasia, hyperpolarized (HP) 13C MRI supplies new metabolic and/or functional insight that could be valuable for a range of kidney diseases. The high renal delivery of intravenously injected small molecules, which provides extraordinary sensitivity for renal imaging of HP 13C MRI probes, suggests great potential for application of this nascent medical imaging modality to kidney disease.

Andreas Pohlmann and Thoralf Niendorf (eds.), Preclinical MRI of the Kidney: Methods and Protocols, Methods in Molecular Biology, vol. 2216, https://doi.org/10.1007/978-1-0716-0978-1\_16, © The Author(s) 2021

Commonly used clinical markers for kidney disease have significant limitations, especially in the context of early disease. Serum creatinine is frequently utilized as a functional marker, but has a wide reference interval, and the mathematical corrections applied for the estimation of glomerular filtration rate (GFR) are inexact. For example, GFR is unpredictably overestimated in glomerulopathic patients as a result of increased creatinine secretion [1]. Even accurately measured GFR, such as by the "gold standard" inulin clearance, is insensitive to early disease [2]. Blood urea nitrogen (BUN) fluctuates due to factors unrelated to kidney function, such as hydration status and diet. Proteinuria at baseline is a significant risk factor for renal disease, but has poor negative predictive value [3]. In general, clinical markers do not exhibit definitive changes until a significant fraction of kidney function is already lost.

Medical imaging has clear potential to address these limitations by providing localized functional data, but this has yet to be translated on a large scale. Nuclear medicine plays a significant role, with (mercaptoacetyltriglycine) MAG3 scintigraphy estimated to account for >400,000 renograms per year [4], commonly for evaluation of renal function (often pre- or post-transplant), split function, collecting duct/urinary tract obstruction, renovascular hypertension, and renal artery stenosis. However, nuclear scans typically have relatively poor spatial resolution and lack the rich tissue contrast of MRI. Furthermore, these studies carry the risks of exposure to ionizing radiation. Ultrasound is also commonly used for assessing kidney disease [5], but is generally limited to evaluating advanced damage reflected in alterations in kidney size and shape, and to some extent blood flow. For patients with impaired renal function, iodinated Computer Tomography (CT) contrast media carry an increased risks of acute kidney injury [6]. Gadolinium based MRI contrast media are associated with nephrogenic systemic fibrosis [7].

HP 13C MRI has unique potential for improved, safer clinical imaging studies of renal structure and function, and perhaps most notably offers an unprecedented avenue toward assessing renal metabolism in vivo. Indeed, results of numerous preclinical studies conducted to date, as summarized below, show the clear potential of this new approach to fill pressing unmet clinical needs for improved markers of various diseases affecting the kidneys, both malignant and nonmalignant. More preclinical work with realistic renal disease models is, however, needed in order to bridge effectively into clinical studies.

This introduction chapter is complemented by two separate chapters describing the experimental procedure and data analysis, which are part of this book.

This chapter is part of the book Pohlmann A, Niendorf T (eds) (2020) Preclinical MRI of the Kidney—Methods and Protocols. Springer, New York.

### 2 Measurement Concept

#### 2.1 Basic Concept of HP 13C MRI

HP 13C MRI is based on a transient enhancement of the nuclear magnetism of 13C-enriched liquids, by up to five orders of magnitude over states readily attainable in a clinical MRI magnet, thereby offsetting the limitation of poor sensitivity that previously hampered in vivo 13C MRI. Although multiple alternate approaches exist for producing HP material, the method of dissolution dynamic nuclear polarization (DNP) [8] provides an especially robust avenue for hyperpolarizing a wide variety of 13C-labeled substrates, using commercially available instruments. In this approach, the 13C substrate is first prepolarized in a separate magnet at low temperature (see Note 1), primarily through microwave irradiation applied near the electron paramagnetic resonance (EPR) frequency in the presence of a stable organic free radical (aka electron paramagnetic agent or EPA, usually a trityl radical) (see Note 2). Following a period of polarization buildup, the solid sample is rapidly dissolved into a liquid state where it temporarily maintains its hyperpolarization, and is quickly transferred to the MRI scanner for intravenous injection and imaging (see Note 3). The overall process is illustrated in Fig. 1. The transient hyperpolarized state persists for a period of time determined by the substrate's T1 relaxation time, an exponential time constant governing the rate of decay of polarization (see Note 4). Inverse scaling of T1 with molecular size limits the scope of applicable probes to small molecules, with labels preferably located distal to bonded hydrogen atoms, which can destroy the 13C magnetization via 13C-1 H dipolar coupling (see Note 5). During the decay period, 13C MRI data is collected, capitalizing on the sensitivity enhancement afforded by hyperpolarization. Specific methods for data acquisition and processing of HP 13C MRI data are described in the relevant sections to follow.

2.2 Renal Functional and Metabolic Investigations Using HP 13C MRI

HP 13C MRI offers many new possibilities for renal functional and metabolic investigations. Dynamic images of the real-time delivery of metabolically inert tracers such as 13C urea from arterial vasculature to the kidneys can be used to quantify renal function. Realtime in vivo processing of metabolically active HP probes by the kidneys can be tracked by spectroscopic imaging, which yields separate images of injected tracer(s) and their downstream metabolic product(s) based on chemical shift differences. Notably, multiple probes can also be simultaneously copolarized to yield multiparametric data sets [9], for example spanning both renal functional and metabolic parameters. In this section, we review the most promising HP 13C MRI probes of renal function/metabolism proposed to date, and the relevant associated principles. These probes and their metabolic pathways are illustrated in Fig. 2.

Fig. 1 Illustration of the process of HP 13C MRI via dissolution dynamic nuclear polarization. The agent (e.g., [1-13C]pyruvate) is mixed with the EPA and cooled to 0.8 K in a magnetic field of 5 T, then irradiated at 140GHz to transfer polarization from unpaired electrons in the EPA to the 13C nuclei. The sample is then rapidly dissolved and transferred to the MRI scanner for intravenous injection and MR imaging

2.2.1 [13C]Urea The key osmolyte and metabolic end product urea was one of the first 13C-labeled molecules to be hyperpolarized (T1 ¼ 45 s at 3 T) [8] and imaged in vivo [10], and high quality dynamic images of the kidneys and the feeding arterial vasculature can readily be generated [11–13]. Quantitative estimates of renal function require dynamic measurements from both the renal parenchyma and arterial vasculature [11, 14]. Urea is largely reabsorbed by urea transporters (UTs) expressed at the inner medullary collecting ducts (IMCD), and accordingly renal imaging shows distinct modulation by hydration state [13, 15], based on differential action of vasopressin. Despite this significant reabsorption, it has been suggested that urea can also be used to estimate glomerular filtration rate (GFR) [14], based on the first-order transfer coefficient between arterial vasculature and renal parenchyma. Secondary labeling of [ 13C]urea with 15N2 is helpful in terms of prolonging both T1 (at low magnetic fields only) [16] and T2 relaxation times [12] of urea, with the latter factor allowing imaging rat kidneys at spatial resolutions approaching 1 mm isotropic (using refocused image acquisition) [13]. Furthermore, relaxation mapping shows great promise for improved contrast between the individual kidney compartments and various functional states [12, 13, 17]. Urea is distinguished by an exceptionally good safety profile among medical imaging contrast agents, even in patients with reduced kidney function, and therefore has potential for clinical translation for imaging patients with kidney diseases.

Fig. 2 HP 13C MRI probes of interest for monitoring renal disease, and the associated metabolic conversions (if any). Positions of the 13C labels are indicated by stars

2.2.4 [1-13C] Dehydroascorbate HP [1-13C]dehydroascorbate (T1 ¼ 56 s at 3 T) [27, 28] can be used to probe oxidative stress in a direct manner. Vitamin C, a key physiologic antioxidant, exists in a NADP(H)-mediated equilibrium with its oxidized form, dehydroascorbate (DHA). Conversion of injected HP [1-13C]DHA, which is rapidly taken up by glucose transporters, to [1-13C]vitamin C is detectable in vivo. The observed extent of conversion of HP DHA to vitamin C, a reaction whose reducing power is likely mostly derived from reduced glutathione (GSH), is attenuated by oxidative stress.

2.2.5 [13C]Acetoacetate The ketone body acetoacetate, a universal oxidative fuel, rapidly interconverts with its reduced form, beta-hydroxybutyrate, in mitochondria via beta-hydroxybutyrate dehydrogenase. In analogy with lactate/pyruvate, the BOHB–AcAc ratio has frequently been interpreted as a marker of the free mitochondrial NADH–NAD+ ratio, suggesting an analogous interpretation of conversion of HP acetoacetate [29, 30] as being driven by the mitochondrial NAD (H) redox state. Metabolic conversion of injected [1,3-13C2]acetoacetate (58 s at 3 T, C1 position) can be detected in rat kidney in vivo (via the C1 label) [29], suggesting applicability of this new probe to interrogate mitochondrial-driven kidney disease.

### 3 Overview of Applications

3.1 Kidney Cancer Multiple studies have reported elevated conversion of HP [1-13C] pyruvate to [1-13C]lactate and rapid export of [1-13C]lactate in renal cell carcinoma (RCC), in both RCC cells ex vivo [31–33] and in vivo in mice orthotopically implanted with human RCC cells [34]. These findings suggest that HP [1-13C]pyruvate MRI could have clinical value for improved characterization of kidney cancer in patients. Moreover, a recent study showed that HP [1-13C]pyruvate MRI could be used to predict RCC treatment response to mTOR inhibition, which varies greatly among individual patients, thus potentially informing treatment decisions [35].

3.2 Diabetic Nephropathy Diabetes results in profound shifts in central carbon metabolism, directly shifting several pathways accessible to HP 13C MRI. Although diabetes is detectable using relatively simple testing, sustained hyperglycemia over time leads to serious end-organ complications including diabetic nephropathy, a significant source of morbidity, the detection/prediction of which is a difficult clinical problem where HP 13C MRI could be valuable. Though cancer has been the main focus of the HP 13C MRI community, several recent studies have reported differences in renal HP 13C signals detected in type 1 and type 2 diabetes models as compared to normal controls, based on multiple HP probes including [1-13C]pyruvate [23, 36, 37], [13C]urea [38, 39], and [1-13C]DHA [40]. Not surprisingly, standard antidiabetic agents including insulin and metformin have been found to induce large renal HP signal changes as well [29, 41, 42]. Renal HP lactate levels are attenuated with hyperbaric [43] and antioxidant treatment [44] in diabetic rats. HP [13C]acetate has so far failed to show any significant metabolic change associated with the diabetic kidney [45]. Further work is needed in animal models of frank diabetic nephropathy, which may display changes additive to the fundamental metabolic shifts observed in diabetes, to address this important clinical problem.

#### 3.3 Ischemic Kidney Injury Effective detection of acute kidney injury (AKI) is an unresolved clinical problem especially affecting hospitalized patients. Rapid treatment of AKI is critical but is impeded by existing clinical indicators which are poor in diagnosing early AKI. Several studies of models of ischemia reperfusion injury (i.e., unilateral renal artery clamp) have shown clear effects of AKI on renal HP 13C MRI using [ 13C]urea [46, 47], pyruvate [48], and [1-13C]pyruvate–[1-13C] DHA combination [49], and fumarate [50]. However, these relatively simple models likely do not reflect the etiology of ischemic injuries more frequently encountered in patients, such as those resulting from atherosclerotic disease. Further work using more realistic models of ischemic kidney injury is needed to answer, for example, the important question of predicting treatment response in renal artery stenosis. Finally, HP [1,4-13C2]fumarate has also been investigated as a probe of tissue necrosis in a toxic folic acid model of acute tubular necrosis (ATN) [26]. A nonzero urinary fumarase activity could potentially be used as a clinical indicator for a hyperpolarized fumarate examination to investigate the extent and the origin of the renal damage [50].

### 4 Notes

### Acknowledgments

Grant support from National Institutes of Health (NIH) K01DK099451 is gratefully acknowledged.

This chapter is based upon work from COST Action PARENCH-IMA, supported by European Cooperation in Science and Technology (COST). COST (www.cost.eu) is a funding agency for research and innovation networks. COST Actions help connect research initiatives across Europe and enable scientists to enrich their ideas by sharing them with their peers. This boosts their research, career, and innovation.

PARENCHIMA (renalmri.org) is a community-driven Action in the COST program of the European Union, which unites more than 200 experts in renal MRI from 30 countries with the aim to improve the reproducibility and standardization of renal MRI biomarkers.

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treatment response to mTOR inhibitors in patient-derived ccRCC xenograft models. Cancer Res 79(1):242–250. https://doi.org/10. 1158/0008-5472.CAN-18-2231

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N2 -urea T2 relaxation changes in acute kidney injury. Magn Reson Med 80:696–702. https://doi.org/10.1002/mrm.27050

Open Access This chapter is licensed under the terms of the Creative Commons Attribution 4.0 International License (http://creativecommons.org/licenses/by/4.0/), which permits use, sharing, adaptation, distribution and reproduction in any medium or format, as long as you give appropriate credit to the original author(s) and the source, provide a link to the Creative Commons license and indicate if changes were made.

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### Functional Imaging Using Fluorine (19F) MR Methods: Basic Concepts

### Sonia Waiczies, Christian Prinz, Ludger Starke, Jason M. Millward, Paula Ramos Delgado, Jens Rosenberg, Marc Nazare´, Helmar Waiczies, Andreas Pohlmann, and Thoralf Niendorf

### Abstract

Kidney-associated pathologies would greatly benefit from noninvasive and robust methods that can objectively quantify changes in renal function. In the past years there has been a growing incentive to develop new applications for fluorine (19F) MRI in biomedical research to study functional changes during disease states. 19F MRI represents an instrumental tool for the quantification of exogenous 19F substances in vivo. One of the major benefits of 19F MRI is that fluorine in its organic form is absent in eukaryotic cells. Therefore, the introduction of exogenous 19F signals in vivo will yield background-free images, thus providing highly selective detection with absolute specificity in vivo. Here we introduce the concept of 19F MRI, describe existing challenges, especially those pertaining to signal sensitivity, and give an overview of preclinical applications to illustrate the utility and applicability of this technique for measuring renal function in animal models.

This chapter is based upon work from the COST Action PARENCHIMA, a community-driven network funded by the European Cooperation in Science and Technology (COST) program of the European Union, which aims to improve the reproducibility and standardization of renal MRI biomarkers. This introduction chapter is complemented by two separate chapters describing the experimental procedure and data analysis.

Key words Magnetic resonance imaging (MRI), Fluorine (19F), Kidney, Inflammation, Tissue oxygenation

### 1 Introduction

The need for methods to assess renal function that are both robust and noninvasive is increasingly becoming recognized. One approach is to use MRI in combination with contrast agents that are responsive to physiological changes [1]. The most common type of MR contrast agents are paramagnetic agents that enhance contrast on MR images. These contrast agents act by modulating proton (1 H) relaxation, either during recovery (T1 component) or

Andreas Pohlmann and Thoralf Niendorf (eds.), Preclinical MRI of the Kidney: Methods and Protocols, Methods in Molecular Biology, vol. 2216, https://doi.org/10.1007/978-1-0716-0978-1\_17, © The Author(s) 2021

decay (T2 component) of the MR signal. In order for a contrast agent to be useful to identify changes in a specific physiological phenomenon, its sensitivity to detect and quantify physiological changes on MRI and its pharmacokinetic properties must be balanced [2].

From a technical point of view, paramagnetic contrast agents have a concomitant effect on all components of MR relaxation; the interplay between T1 and T2 relaxation depends on the dose and location of the contrast agent as well as the MR imaging sequence used [3]. This complexity is exacerbated by intrinsic factors such as deoxygenated blood or air cavities that may confound the signal being investigated [4]. A case in point is when imaging cells labeled with contrast agent: changes in contrast created by such intrinsic tissue factors could render the labeled cells indistinguishable from the surrounding tissue. This drawback is heightened by the signal sensitivity constraints due to the low concentrations of contrast agent (<μM range). As a result, any quantification of the physiological observation is hampered, and in some cases, no definitive conclusions can be made.

The use of fluorinated agents in combination with fluorine (19F) MRI overcomes the hurdle of signal ambiguity because the 19F MR signal is specific to the administered 19F agent. While the 19F nucleus has similar MR sensitivity and resonance frequency as 1 H, unlike 1 H, it is practically absent in vivo, that is, there are no endogenous 19F organic forms in vivo. Therefore, any exogenous fluorinated compounds introduced in vivo can be detected with high selectivity and absolute specificity. This has been exploited to noninvasively track cells in vivo using 19F MRI after labeling them with 19F compounds, typically in the form of nanoparticles (NP). Interestingly, the chemical shift range for 19F is much larger than that of 1 H. Organofluorine compounds span a chemical shift range of over 350 ppm [5]. Thus multiple 19F nuclei resonances for organofluorine compounds can be easily separated, even under rather inhomogeneous conditions [6]. The unique properties of 19F nuclei as well as 19F compounds have been the driving force behind the increasing interest in 19F MRI. However, the application of 19F MRI to a wide range of biomedical research fields has been restricted by the low abundance of 19F nuclei in vivo following exogenous administration. This constraint is compounded by the fact that the signal sensitivity of current MR hardware remains limited, making the detection of 19F compounds present at low concentrations extremely challenging. Therefore, one strategy in 19F MRI to circumvent this is to boost the signal-to-noise ratio (SNR) as a function of time, that is, the SNR efficiency to enhance signal sensitivity and improve detection limits.

From a biological perspective, several key components must be considered when using contrast agents to study renal function. Among other physicochemical properties, size will determine the pharmacokinetic properties of the contrast agent as well as its uptake by cells in vivo. For example, low-molecular-weight-gadoliniumbased contrast agents (GBCAs) are filtered through the glomerulus and are not reabsorbed by the renal tubule. Conversely, larger contrast agents such as paramagnetic iron oxide particles or 19F nanoparticles commonly follow a one-compartment pharmacokinetic model within the blood [7]. These nanoparticles are too large to be filtered by the glomerulus and are eventually taken up by phagocytic blood cells. Such nanoparticles including the iron replacement therapy ferumoxytol have a long intravascular half-life (>14 h) and are therefore useful for vascular and perfusion-weighted MRI [8]. Unless inflammatory processes are ongoing, these particles will localize primarily in the liver and will eventually undergo elimination via hepatic clearance. The clearance of nanoparticles via liver or kidney is dependent on their size and surface properties [9] but also on the health status of these organs, such that clearance should be critically studied [10] especially prior to patient studies.

This introduction chapter is complemented by two separate chapters describing the experimental procedure and data analysis, which are part of this book.

This chapter is part of the book Pohlmann A, Niendorf T (eds) (2020) Preclinical MRI of the Kidney—Methods and Protocols. Springer, New York.

### 2 Overview of Applications

hypoxic injury [11]. While O2 consumption determines perfusion in most nonrenal tissue such as the brain, perfusion determines O2 consumption in the kidney [12]. Therefore, a sudden temporary impairment of the blood flow to the kidney results in acute kidney injury (AKI) as a result of impaired O2 delivery. Approaches for identifying and treating AKI as early as possible should aim at detecting and restoring imbalances the homeostasis between renal oxygenation and factors such as nitric oxide that control the renal microcirculation [13].

Noninvasive methods to measure kidney function in vivo, specifically tissue pO2 and blood perfusion, are crucial. Arterial spin labeling MRI determines intrarenal perfusion without the need of exogenous contrast agents, and will be reviewed in a separate chapter of this book. MRI techniques that utilize the native 1 H contrast from blood—blood-oxygenation-level-dependent (BOLD) MRI—have been proposed [14]. The level of renal oxygenation within each voxel is quantified by mapping the shortening of 1 H T2 \* induced by deoxyhemoglobin—deoxyHb [15]. However, the relation of T2 \* to tissue pO2 may be confounded by factors such as hematocrit, oxyHb dissociation curve, and the vascular volume fraction. This underscores the need for careful scrutiny of this relation in order to make quantitative interpretations [16].

Perfluorocarbon (PFC) emulsions offer an alternative to quantify changes in tissue pO2 during vascular pathology [17, 18]. More recently PFC NPs were also employed as a theranostic strategy for studying perfusion and treating AKI [19]. Primarily PFC NPs have been used to assess perfusion in the renal microvasculature following AKI [20]. 19F MRI sensitively detected decreased RBV and pO2 in the cortical–medullary (CM) junction 24 h following unilateral (left) renal ischemia–reperfusion [20], where increased T2 \* was also observed by 1 H BOLD MRI (Fig. 1a). The T2 \* , 19F signal, and pO2 in the renal cortex of injured kidneys were all comparable to the contralateral control kidney, suggesting recovery of perfusion and oxygenation in this region. In the inner medulla, however, vascular leakage and hemorrhage-induced extravascular retention of PFC NPs (Fig. 1b, c) resulted in a false readout of reduced 1 H T2 \* , increased 19F signal intensity, and unchanged pO2.

PFCs and oxygen exhibit similarly low cohesive energy densities, such that both require little energy for mutual solubility [21]. The high oxygen-dissolving capacity of PFC emulsions makes them ideal as O2 carriers. PFCs were, in fact, considered as artificial blood substitutes [22]. Fluosol-DA—a mixture of perfluorodecalin (PFD) and perfluorotributylamine—was the first PFC emulsion studied in humans in a trial of Japanese patients [23]. It was soon recalled due to its rather low oxygen transport capacity compared to that of red blood cells [24]. PFCs ultimately failed as blood substitutes (even as next generation products

Fig. 1 Assessing perfusion in the renal microvasculature following acute kidney injury using PFCs. (a) Composite Representative 1 H T2\*-weighted image (top row), 19F spin density weighted image (middle row), and pO2 map (bottom row) in the left control and right injured kidneys of the same mouse. (b) Fluorescence images showing FITC-lectin labeled perfused blood vessels in control and injured kidneys (top row) and zoomin view of FITC-lectin (green) labeled perfused vessels in the CM junction (bottom row). (c) Fluorescence images showing rhodamine labeled PFC NPs in injured and control kidneys (top row) and zoom-in view showing PFC NPs (red) in the renal medulla (bottom row). Adapted from Hu L. et al. (2013) [20] with permission from John Wiley & Sons Inc.

appeared in the clinic). Although O2-deprived tissues easily extract O2 in PFCs, the oxygen dissociation curve of PFC emulsions is linear and monotonic, in contrast to the sigmoidal hemoglobin dissociation curve. Therefore, most of the O2 dissolved in PFCs is released at high O2 partial pressure in the arteries, with little O2 being available for the capillaries where O2 partial pressure is low but needed the most [22]. Interestingly Fluosol-DA (the first PFC that was used in patients [23]) was also used to acquire the first 19F MR images in vivo, showing its accumulation in the rat liver [25]. The PFC in this emulsion gives complex MR spectra and has since been replaced by other PFCs with better imaging properties [26].

The longitudinal relaxation rate (1/T1) of 19F in PFCs was shown to depend linearly on O2 partial pressure [27]. This led to the first noninvasive in vivo tissue O2 assessments in tumor tissue and liver [28] and later in myocardial tissue [29]. The latter study provided the first proof of concept study for rapid, noninvasive measurements of O2 oxygen tension changes in response to ischemia and reperfusion [29]. It is claimed that the size of PFC nanoparticles (>100 nm) confers upon them a good safety profile without renal toxicity, since they are not expected to be cleared through glomerular filtration. Therefore, unlike other imaging agents that undergo renal clearance (e.g., iodinated CT compounds, gadolinium MRI chelates), PFC emulsions do not increase the workload of the kidney and are not immediately considered as nephrotoxic [10].

2.1.2 Acid–Base Homeostasis in the Kidney and pH Sensors

The kidney plays a major role in acid–base homeostasis; this balance depends on its capacity to reabsorb filtered bicarbonate, mainly from the proximal tubule, and to excrete an equivalent amount of protons at the distal tubule [30]. CEST MRI was originally developed to measure pH with exogenous pH-responsive MRI contrast agents [31]. The pH in the kidney following AKI gradually increased from 6.7 to 7.2 as measured with CEST-MRI pH-responsive probes [32]. A major drawback of paramagnetic probes is that signal intensity is not only proportional to the pH but also to contrast agent concentration. Particularly in tumors, in vivo pH mapping assumes a constant contrast agent concentration across tissue notwithstanding poor perfusion where acidolysis occurs [33].

Approaches such as ratiometric 19F/1 H probes have been developed to overcome the issue of quantification; the signal is normalized to the concentration of the contrast agent by means of a concentration reporting (19F) moiety [34]. Another advantage of 19F MR probes is the larger range in chemical shifts in 19F (compared to 1 H). These chemical shifts are highly sensitive to the immediate surroundings, such that small chemical changes can lead to dramatic changes in chemical shift [6]. The first 19F MR probe identified as a candidate to probe cellular pH was the vitamin B6 analog 6-fluoropyridoxol (6-FPOL) reported to have a pKa in the physiological range and large chemical shift response [35]. Subsequently the feasibility of using 6-FPOL to probe cellular pH was demonstrated in whole blood [36]. The concept of employing 19F MR probes for in vivo imaging was introduced when a series of fluorinated vitamin B6 analogs with a chemical shift sensitivity in the range 7.4–12 ppm were synthesized as pH indicators [37]. The sensitivity of 19F MRS probes can be further improved by incorporating paramagnetic metal ions that further amplify the chemical shift between protonated and deprotonated forms by the presence of covalent and dipolar contributions [38, 39]. Iron (II) complexes were also explored to exploit the pH-dependent spin-state population change; the 19F NMR chemical shift was in a sensitivity range of 13.9 ppm per pH unit at 37 C [40]. More recently fluorinated nickel (II) complexes with diamagnetic and paramagnetic properties were designed for ratiometric pH mapping by 19F MRI [41]. Supramolecular nanostructures formed by fluorinated peptide amphiphiles transitioning from cylindrical to ribbon-like shape as pH increases from 4.5 to 8.0 resulting in enhanced MR signals. These nanostructures represent a good strategy to design stimulus-responsive 19F MRI contrast agents for responses to pH changes in vivo [42].

2.2 19F MRI to Study Renal Inflammation In Vivo

Together with renal tissue hypoxia and tubular injury, inflammation is one underlying component in the pathology of AKI [43]. Studies suggest that systemic inflammation causes ischemic injury in one organ and that repercussions follow in distant organs downstream

Fig. 2 Inflammation as one underlying factor in the pathology of AKI. AKI can be triggered by a variety of clinical disorders ranging from renal ischemic events to sepsis and exposure to nephrotoxic substances. The kidney is tightly connected with the interior milieu of the body such that even the slightest of changes will be rapidly detected and communicated between the kidney and the rest of the body

of the ischemic event [44–46] (Fig. 2). Therefore, methods that can non-invasively quantify the level of inflammation in different organs would be crucial to identify the causality and significance of inflammation during both systemic and kidney-specific disease.

19F MRI is commonly used in association with an intravenous administration of PFC emulsions to study the distribution of inflammatory cells in vivo [26, 47–51]. 19F MRI is performed after an in vivo labeling of inflammatory cells with a 19F compound—typically in nanoparticle (NP) form. The 19F-NPs are administered intravenously in order for them to be taken up by inflammatory cells that are migrating through the systemic circulation into inflamed organs. Typically, PFCs such as perfluoro-15 crown-5 ether (PFCE) are used to prepare NPs for cell labeling. PFCE is biologically stable and provides a high payload of 19F nuclei since it contains 20 chemically equivalent 19F spins.

In a rat model of orthotopic kidney transplantation, 19F MRI was used to study inflammation in the kidney undergoing rejection [48]. Three days following transplantation, PFCE NPs were administered intravenously and 1 H/19F images acquired 2 and 24 h following NP infusion (Fig. 3a). At the earlier time point, the 19F signal was mainly found in the circulation but also in small but roughly equal amounts in both the native and transplanted kidneys (Fig. 3a). By day 4 posttransplantation (i.e., 24 h following PFCE NP), the transplanted kidney underwent moderate (Grade 2) rejection and showed an intense signal in the allograft kidney and in liver and spleen, but the 19F signal was virtually undetectable above the level of image noise in the native kidney (Fig. 3a). T2\*-weighted

Fig. 3 Inflammation in the kidney in a rat model of orthotopic kidney transplantation experiencing rejection. (a) Composite 1 H/19F images of an orthotopic kidney transplant recipient 2 h post–PCE injection on POD 3 with 19F signal found in the vasculature and equally in both the native (upper) and graft (lower) kidneys (left) and ~24 h post–PCE injection with 19F signal in the allograft kidney, as well as other organs (right). (b) Allograft kidney on POD 4 following an injection (left) of 6 mg/kg BW USPIO (upper row) or following an injection of saline (lower row) and 24 h later (right). Adapted from Hitchens T.K. et al. (2011) [48] with permission from John Wiley & Sons Inc.

> imaging reveals hemorrhaging in the medulla on day 4 post-kidney transplantation, which can be detected both in animals administered ultrasmall superparamagnetic iron oxide (USPIO) NPs but also in transplants without USPIO (saline control) (Fig. 3b). This can confound the detection of USPIO-labeled macrophages and exhibits one of the advantages of 19F MRI over iron oxide-based NPs.

> Apart from labeling inflammatory cells in vivo, NPs have been used to label specific immune cell types such as dendritic cells (DCs) ex vivo, prior to their administration in vivo [52–54]. This kind of labeling is particularly useful to label cellular therapies that will be administered in patients as part of a treatment strategy [55]. In addition to cancer patients, kidney transplantation patients would also benefit from cell-based therapies, specifically the types of cells inducing tolerance [56]. Monocyte-derived suppressor cell products are currently being investigated alongside immunosuppressive therapies in kidney transplant recipients with promising results [57]. These candidate therapies would benefit from a labeling strategy that will enable them to be tracked in vivo over time with 19F MRI. The possibility of quantifying the 19F MR signal would allow an objective assessment of the spatial and temporal cellular distribution within the same individual. This kind of labeling and 19F MRI would result in absolute cell quantification, since the 19F MR signal is generated from the cells alone, by direct detection of the 19F spins in the cells labeled with 19F NPs, without any confounding background MR signal. However, several biological and technical considerations need to be made, in order to achieve a truly absolute MR signal quantification: not all cell types can be labeled with the 19F NPs currently available (Subheading 3.1.1).

### 3 Basic Concepts of 19F MRI

3.1.1 19F Probes for Cellular Uptake

3.1 19F imaging Probes The design and synthesis of novel fluorine-rich molecular labels and imaging probes is an important domain for 19F MRI. For several years, a wide range of fluorine probes has been available for experimental and even clinical applications [58, 59].

> Current procedures used to label cells with nanoparticles take advantage of the intrinsic nature of cells to engulf foreign bodies, a cellular property known as phagocytosis. Dendritic cells are intrinsically phagocytic and therefore ideal candidates for tracking with 19F MRI [52, 53, 55, 60–62]. Cells labeled with clinical grade 19F nanoparticles have been administered as immunotherapies in cancer patients [55]. This tracking method could support several other cell therapy studies, including those dealing with kidney transplantation [63].

> Enhancing cellular uptake is desirable, even for cells that are intrinsically phagocytic. Several studies have investigated the incorporation of phosphatidylserine (PS) in nanoparticle-based therapeutics to enhance their uptake [64]. After renal ischemia–reperfusion, the signal from PS-containing microbubbles was two-fold higher than that for standard lipid microbubbles [65]. Phospholipids such as PS and phosphatidylethanolamine (PE) are important components of biological cell membranes and are important components for endocytosis in phagocytic cells such as DCs [66]. We previously reported that enrichment of 19F nanoparticles with 1,2-dipalmitoylsn-glycero-3-PE (DPPE) resulted in a striking increase (at least one order of magnitude) in cytoplasmic uptake in DCs [62]. Interestingly, membrane lipid-raft processes have been proposed as trafficking mechanisms for transporting therapeutic cargo such as siRNA [67] or targeting ligands [68] into the cytoplasm, circumventing the usual endosomal uptake pathway.

> Labeling cells with low intrinsic phagocytic properties (such as T cells) will prove more difficult using the above strategy, even if the nanoparticle size is lower than 180 nm [69]. Some research groups claim to have labeled T cells using the same above principles [70]. T cells are not able to internalize nanoparticles, even small iron oxide nanoparticles (core diameter ¼ 12.5 nm), as has been clearly shown by transmission electron microscopy (TEM), considered to be a definitive method, and also by confocal microscopy [71]. For the latter method nanoparticles with fluorescent tags were used to corroborate the in vivo 19F signals with ex vivo fluorescence microscopy signals, but the authors also included a cell membrane stain [71]. This is a necessary control to distinguish membrane association from intracellular distribution, in light of the point spread function contribution [72]. Another key point is that the fluorescent tag dissociates from the fluorine label over time, both in vitro and in vivo [73]. This has a critical impact on the interpretation of long-term experiments validated by standard fluorescent techniques.

3.1.2 19F Probes

Sensitivity

Before taking the next steps toward administering cells in vivo, it is essential to understand how nanoparticles are recognized, internalized, trafficked, and distributed within the specific host cell [74]. This is critical when contemplating labeling of cells that may ultimately be applied in vivo to patients in a therapeutic context. Cells that are not intrinsically phagocytic will not readily internalize NPs but might still exhibit a 19F MR signal ex vivo, due to reasons such as transitional binding of NPs to the cell membrane or incomplete clearing of noninternalized NPs from the cell culture medium. The 19F MR signals following administration of ineffectively labeled cells with contaminating NPs would be misrepresentative of the actual distribution of the cells in vivo. Therefore, we need to ask the right questions regarding the specific cell type to be labeled, and the methods used prior to transfer of the cells into living organisms [75].

with Increased Signal Fluorinated probes are commonly designed with the goal of improving signal strength by increasing the number of fluorine atoms per molecule [59]. Linear 19F polymers such as perfluoropolyethers (PFPEs) are the simplest choice to achieve increased fluorine content but the NPs produced with these polymers are polydisperse and the presence of multiple chemical shifts presents several peaks in the 19F spectra [76]. Superfluorinated branched 19F probes such as 19FIT [77] and PERFECTA [78] promise several advantages, specifically a very high fluorine nuclei density with a single 19F spectral peak. These probes therefore afford enhanced sensitivity and better imaging performance than linear perfluoropolymers. Unlike PERFECTA [78], the bispherical fluorocarbon 19FIT is water soluble [77] and therefore possesses low organ retention, but it cannot be emulsified into nanoparticles for cell labeling strategies. Increasing the density of 19F nuclei can only go so far in terms of increased sensitivity of 19F probes. Ultimately, a limit to this density will be reached.

Another strategy to promote 19F MR signal is to manipulate the MR properties of the 19F probes in order to favor ideal conditions for MR signal acquisition. One approach is to decrease the T1 of the 19F nuclei, which is typically in the range of 0.5–3 s for diamagnetic compounds. Paramagnetic metal ions such as lanthanides(III) or iron(II) reduce the T1 of 19F compounds [38]. Paramagnetic centers affect the relaxation of 19F nuclei (paramagnetic relaxation enhancement or PRE effect) via a number of mechanisms, most notably electron–nucleus dipole–dipole interaction and Curie relaxation [39], both of which are influenced by the effective magnetic moment (μeff) of the paramagnetic metal [79]. Other than the μeff of the metal ion, the PRE effect is also influenced by the distance separating the 19F nuclei and the paramagnetic metal ion [39]. While long distances will severely diminish PRE, short distances will substantially shorten the spin–spin relaxation time T2. Thus T1 reduction by molecular design should be dealt with carefully since it could also lead to a reduction in the T2 and thereby line broadening of the resonance, which ultimately affects sensitivity [80]. The most sensitive metal-based probes are, in fact, not those whose metal ions have the highest μeff, but rather those who have the ideal distance separating 19F from the paramagnetic metal ion (4.5–7.5 A˚) thus resulting in higher T2/T1 ratios [81]. Nonetheless even at larger 19F–M distances (>10 A˚), the PRE effect can still be notable as is the case for a DOTA-based 19F chelator [82]. In paramagnetic PFC nanoparticles, the 19F MR relaxation enhancement has also been explained by diffusion mechanisms attributed to the changes in Brownian motion of PFC molecules inside the NP in conjunction with their ability to permeate into the lipid surfactant coating [83].

Introduction of paramagnetic metals in the design of 19F probes not only results in a decrease T1 of 19F nuclei but also leads to paramagnetic induced chemical shifts. The CEST signal generated by paramagnetic involvement (PARACEST) will also be modulated by pH (Subheading 2.1.2). A macrocyclic Fe (II) complex bearing both exchangeable 1 H and 19F was shown to behave as a ratiometric pH-responsive MRI contrast agent, whereby the CEST, but not the 19F MR signal was pH-dependent [84]. This enables accurate pH mapping without the use of multiple probes. More recently a DOTA macrocyclic structure complexed with Ni(II) metal was shown to exhibit larger separations between the 1 H CEST signal and water [85]. Though the in vivo application of 19F PARACEST probes remains to be investigated, the available complexes will likely provide valuable information in the near future when probing pH changes in vivo.

3.2 MR Data Acquisition

3.2.1 Pulse Sequence Optimization

One way of improving signal sensitivity is to avoid signal losses such as those occurring during spin relaxation. Most fluorinated compounds generally have long T1 relaxation times, which effectively increases the acquisition time. Together with the fact that there is a large range of relaxation times for 19F MR, this motivates more reliable optimization of pulse sequences that is informed by the MR properties of the 19F compounds under investigation. Sequence performance is largely dependent on the relaxation times of the addressed spin system but also on the conditions the nuclei are in, such as tissue oxygenation level [17].

The choice of the best pulse sequence and acquisition method is dictated by the MR properties, particularly the relaxation times, of the nuclei being measured. The acquisition protocol for each sequence should be optimized by tailoring detection schemes to the properties of the nucleus—also in relation to its environment using relationships empirically drawn from simulations [86]. Relaxation times determine the time frames of signal losses, so the time orchestration for data acquisition should be assigned correctly. Other parameters, such as acquisition bandwidths, should also be properly planned to avoid recording unnecessary noise and to achieve the highest possible SNR [87].

A number of 19F MR studies using different pulse sequences have aimed to optimize the acquisition method to promote SNR efficiency and signal sensitivity [17, 87–90].  This method reduces acquisition time by accumulating multiple echoes (dictated by the turbo factor) within a single repetition time [92]. Optimal settings for echo train lengths and repetition time (TR) according to measured relaxation times (T1/T2 combinations) were calculated for RARE to improve sensitivity thresholds [90].

Although RARE is commonly been for studying PFCs in vivo, other sequences will be better suited for 19F compounds that unlike typical PFCs have short T2 relaxation times. In such cases signal detection might be technically challenging and will demand acquisition strategies properly adapted for the compound and its environment. Gradient echo (GE) sequences were used to maximize SNR efficiency for imaging paramagnetic macrocyclic 19F molecules complexed to lanthanides [87]. These molecules exhibit T1 values in the order of 1–15 ms and T2\* values in the order of 0.4–12 ms [88]. In the latter study a radial zero echo time (ZTE) sequence yielded a sensitivity gain of 27-fold when compared to nonparamagnetic molecules, while the GE sequence yielded a gain of 11-fold [88]. Additionally, the highest SNR efficiency gain was achieved when T2/T1 were close to 1, particularly for the ZTE pulse sequence, and when T1 was in the range of 1–5 ms [88]. Even 19F molecules with unique spectral properties will require a more individualized adjustment of the MR parameters; such molecules might benefit from sequences such as ultrashort echo time (UTE) that acquire signal before the spins dephase, taking into consideration destructive phase interference [93]. An efficient k-space sampling strategy will be necessary to achieve maximal SNR and artifact-free MR images [94]. To image perfluorooctyl bromide (PFOB) that has a rich spectrum with large chemical shifts, a radial UTE with a balanced steady-state free precession (bSSFP) pulse sequence showed superiority above other sequences, such as SE, GE, or bSSFP with Cartesian sampling [89]. However, radial sampling might not always be viable since k space sampling pattern effects need to be considered, which might altogether result in reduced SNR efficiency [86]. Another way to speed up data acquisition is synchronous multinuclear imaging. For this, the MR system needs to have the architecture suited to acquire images from more than one nucleus in a single pulse sequence [95]. Simultaneous proton and 19F MRI was performed in a preclinical model using a clinical MR scanner at 3 T [95].

3.3 Hardware Considerations The similarity between 19F and 1 H MR properties with respect to sensitivity and resonance frequency is an advantage, from an engineering perspective since similar pulse power settings can be used, and dual-tuned radio frequency (RF) coils for 19F and 1 H MRI are more convenient to design [50, 96]. One way of improving the 19F MR signal in vivo is to increase SNR by improving the sensitivity of the measuring instrument. Two hardware strategies to improve sensitivity for 19F MR are to increase the strength of the static magnetic field (B0), and to increase the sensitivity of the RF coil by cryogenic cooling.

3.3.1 Cryogenic Cooling The sensitivity of the RF coil is an important factor limiting SNR and thereby the level of spatial and temporal resolution that is targeted for 19F MRI [97]. Even when employing an RF coil with a homogeneous transmit field (B1 + ) distribution and dedicated for brain 19F and 1 H MRI, the spatial resolution of in vivo 19F MRI was limited to approximately 600 μm [50]. This level of spatial resolution prohibits a detailed understanding of inflammation dynamics. To overcome the sensitivity constraints in 19F MR and improve detail of inflammatory cell location, we applied the concept of cryogenically cooling RF coil hardware to improve SNR by reducing thermal noise [98–100]. The cryogenically cooled RF coil (CryoProbe™) technology is commercially available for small animal MRI of 1 H and X-nuclei such as carbon (13C) and phosphorous ( 31P). For these X-nuclei this technology is expected to be particularly useful for increasing SNR. A gain in SNR of 3.0- to 3.5-fold was reported for the 13C CryoProbe [101]. We recently compared a 19F CryoProbe with a room temperature 19F RF coil of similar size and reported on a maximal SNR gain of 15-fold and were able to study inflammation at an isotropic resolution of 150 μm [100].

> Surface RF coils provide a very high SNR close to the region where they are applied. However, a decrease in signal amplitude with increasing distance from their surface is inherent to the geometry of their design. Inhomogeneities in the transmit (B1 + ) field cause a spatially varying flip angle (FA) and as a result signal quantification is hampered. Likewise, the receive (B1 ) field of surface RF coils is position-dependent [102]. In the particular case of the 19F CryoProbe, we observed a strong FA decrease with increasing distance from the surface of the coil [100]. Using a receive-only surface RF coil in combination with a volume resonator used for transmission would overcome B1 + inhomogeneities. With this configuration the transmitted FAs will be homogeneous, thus enabling quantification [103]. This option is not available when using CryoProbes, since they are typically built in the form of transceive (transmit–receive) RF coils with no active decoupling.

Particularly for quantitative 19F MRI, a correction of the B1 + inhomogeneity will be necessary when comparing concentrations of PFC in vivo with an external standard of known 19F concentration. Differences in the B1 + field between the region of interest in vivo and the external standard will result in errors when quantifying in vivo 19F signals. Several approaches have been implemented to correct B1 + inhomogeneities retrospectively, even for 19F MRI [104, 105]. These approaches typically include RF pulse sequences for which an analytical knowledge of the signal intensity dependency on the FA exists.

3.3.2 Magnetic Field Strengths Increasing the strength of the static magnetic field (B0) will increase SNR of the measured signal [106], a strategy which is actively being pursued for clinical application [107]. Intrinsic SNR is expected to grow at least linearly with increasing B0 strengths [97, 108, 109]. The principle of reciprocity can be used at low field strengths to calculate the receive field (B1) sensitivity of a single channel RF coil in terms of the transmit field (B1 + ) that can be easily measured [110]. In studies performed at frequency ( f ) <1 MHz the maximum sensitivity was calculated to be proportional to the f: sensitivity / f1.75 [97, 108]. When going to higher magnetic field strengths, the increasing f and the influence of wave propagation need to be considered [106]. The homogeneity of the B1 + field is expected to decline with increasing B0, thereby influencing the overall SNR gain. An experimental study investigating SNR dependency on B0 in the human brain revealed that SNR / B01.65 in the range of 3.0–9.4 T [109]. A recent simulation study also showed that SNR grows superlinearly with frequency, particularly in deeper regions of the sample; in less deep regions the SNR versus B0 trend approached linearity [111].

> An increase in magnetic field strength is particularly beneficial for X-nuclei imaging. An increase in B0 from 9.4 to 21.1 T results in an SNR gain of ~3 compared to a gain of ~2 for 1 H MR [112]. We recently studied the influence of increasing B0 on improving 19F MR signal sensitivity [113]. Prior to the practical experiments, we estimated the expected SNR gain when moving from 9.4 to 21.1 T to be 2.7 [113]. We estimated SNR gain using the principle of reciprocity [110] and ignoring all flip angle, sequence and relaxation-dependent effects. The estimated SNR gain represents a baseline expectation, based on the information available for the specific hardware of each system. When introducing the relationship of the reference power for both MR systems, we estimated an SNR gain of ~3.8. In our practical experiments the ratio in SNR efficiency (SNR/√t ratio) was estimated to be 7.29, when comparing the best possible conditions for both 21.1 and 9.4 T, and when including T1 relaxation effects [113]. Differences between the actual SNR gains determined experimentally and those expected

from simulations are conceivable, due to minor inaccuracies in the assumptions made for the EMF simulations.