Pre-clinical functional Magnetic Resonance Imaging Part I: The kidney.

ZEMEDI-10534; No. of Pages 21

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ÜBERSICHTSARBEIT

Pre-clinical Functional Magnetic Resonance Imaging Part I: The kidney Frank G. Zöllner ∗,1 , Rafﬁ Kalayciyan 1 , Jorge Chacón-Caldera, Fabian Zimmer, Lothar R. Schad Computer Assisted Clinical Medicine, Medical Faculty Mannheim, Heidelberg University, Mannheim, Germany Received 19 November 2013; accepted 19 May 2014

Abstract The prevalence of chronic kidney disease (CKD) is increasing worldwide. In Europe alone, at least 8% of the population currently has some degree of CKD. CKD is associated with serious comorbidity, reduced life expectancy, and high economic costs; hence, the early detection and adequate treatment of kidney disease is important. Pre-clinical research can not only give insights into the mechanisms of the various kidney diseases but it also allows for investigating the outcome of new drugs developed to treat kidney disease. Functional magnetic resonance imaging provides non-invasive access to tissue and organ function in animal models. Advantages over classical animal research approaches are numerous: the same animal might be repeatedly imaged to investigate a progress or a treatment of disease over time. This has also a direct impact on animal welfare and the reﬁnement of classical animal experiments as the number of animals in the studies might be reduced. In this paper, we review current state of the art in functional magnetic resonance imaging with a focus on pre-clinical kidney imaging.

Keywords: Pre-clinical research, functional MRI, kidney disease

Präklinische Funktionelle Magnetresonanztomographie Teil 1: Niere Zusammenfassung Fallzahlen chronischer Nierenerkrankungen (CKD) steigen weltweit. In Europa sind z.B. mindestens 8% der Bevölkerung betroffen. CKD ist der Ausgangpunkt für eine reduzierte Lebenserwartung und einer hohen Komorbidität, welches sozio-ökonomischen Kosten nach sich zieht. Daher sind eine frühe Erkennung und eine adäquate Therapie von Nierenerkrankungen wichtig. Präklinische Forschung kann dabei helfen, die Grundmechanismen dieser Krankheitsbilder zu verstehen, aber auch neue Medikamente für eine mögliche Therapie zu testen. Funktionale Magnetresonanztomographie (MRT) bietet dabei die Möglichkeit nicht-invasiv Zugang zu Gewebe- und Organfunktion zu erhalten. Vorteil gegenüber klassischen Tierversuchen ist, dass die Tiere hierbei nicht getötet werden müssen und mehrmals untersucht werden können, z.B. in Verlaufsuntersuchungen. Neben dem wissenschaftlichen Gewinn trägt dies auch direkt zum Tierschutz bei, da mit MRT-Techniken klassische Tierversuche überarbeitet werden können und so Tierzahlen reduziert werden können. In diesem Beitrag stellen wir den aktuellen Stand der präklinischen funktionalen MRT mit einem Fokus auf Nierenbildgebung vor. Schlüsselwörter: Präklinische Forschung, funktionelle MRT, Nierenerkrankung

∗ Corresponding author. Frank G. Zöllner, Computer Assisted Clinical Medicine, Medical Faculty Mannheim, Heidelberg University, Theodor-Kutzer-Ufer 1-3, 68167 Mannheim, Germany. Tel.: +49(0)6213835117; fax: +49(0)6213835124. E-mail: [email protected] (F.G. Zöllner). 1 both authors contributed equally.

Z. Med. Phys. xxx (2014) xxx–xxx http://dx.doi.org/10.1016/j.zemedi.2014.05.002 http://journals.elsevier.de/zemedi

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F.G. Zöllner et al. / Z. Med. Phys. xxx (2014) xxx–xxx

Introduction

Functional Renal Imaging Techniques

Magnetic resonance imaging (MRI) is recognized as a valuable imaging tool that allows excellent soft tissue contrast and therefore, non-invasive morphological imaging without radiation exposure to the patient. However, magnetic resonance imaging techniques have been developed for functional imaging like perfusion and diffusion [1–4], blood oxygenation level dependent (BOLD) [5–10] and also X-nuclei MRI [11–15] allowing for assessing the organ status or tissue viability and aiming at to establish imaging biomarkers for various diseases [16,17]. Moreover, functional MRI is nowadays not only used for diagnostics but also emerging in therapy and treatment control [18–20]. To establish such biomarkers a detailed understanding of the targeted disease is needed which could only be derived via e.g. ex-vivo and in-vivo tissue or animal studies [21,22]. Thereby, preclinical MRI can play an important role. With the availability of dedicated hardware, MRI as a non-invasive imaging technology is emerging into pre-clinical biomedical research. Advantages over classical approaches are numerous: the same animal might be repeatedly imaged to investigate a progress or a treatment of disease over time [23–25]. This has also a direct impact on animal welfare and the refinement of classical animal experiments [26,27] as the number of animals in the studies might be reduced. Of course, this can also be realized using other imaging techniques like x-ray, positron emission tomography, or computed tomography [28,29], however, the animals are not exposed to ionizing radiation using MRI. For rodents, extensive genetic information and technology to create e.g. knockout lines is available. Combined with MRI, effects of knock out genes might be investigated on the phenotype level [26,30]. Furthermore, new contrast agents and nanoparticles have been developed providing new insights but can also be tracked and observed by MRI. Since these are usually not approved for use in humans in the first place they are tested within animals. The combination of new methods for physiological MR imaging based on proton techniques (perfusion, diffusion, oxygenation, etc.), especially in the abdomen (kidney, lung, heart) [31] as well as functional sodium imaging for noninvasive measurement of tissue viability could potentially improve the overall diagnostic role of multimodal MRI techniques. Therefore, it is anticipated to translate in-vivo preclinical studies using state-of-the-art MRI techniques (1 H and 23 Na) from pre-clinical experiments at ultra-high field to clinical MRI applications, and thus increasing the accuracy of diagnostic methods which are pre-clinically validated. In this first part, we review functional imaging of the kidney of small animals implemented at dedicated ultra-high field small animal scanner and their translation to clinical whole body scanners. In the second part of this review, the focus was set on cardiac magnetic resonance imaging (CMRI) [32].

The prevalence of chronic kidney disease (CKD) is increasing worldwide [33]. In Europe alone, at least 8% of the population currently has some degree of CKD. CKD is associated with serious comorbidity, reduced life expectancy, and high economic costs [34]; hence, early detection and adequate treatment of kidney disease are important. Magnetic resonance imaging has an increasingly role in the workup of chronic kidney diseases since it allows for minimal invasive measurements of structure and function. A recent review covers techniques and data analysis methods for assessing kidney volume and therefore, morphological changes related to kidney disease [35]. In this review we will focus on functional imaging techniques in small animal models. Animal handling For pre-clinical kidney imaging, either at dedicated small animal scanners or at whole body human scanners, certain animal handling has to be performed. This comprises anaesthesia, monitoring of the physiological conditions (heart rate, breathing) and temperature monitoring, especially in small rodents. Dedicated small animal scanners can be equipped with respective animal beds/ holders providing means for heating, i.e. to allow for temperature regulation of the animal, gas anaesthesia, and animal positioning and fixation (e.g. Bruker Biospin, Ettlingen, Germany). For whole body scanners, small animal holders with air heating systems can be purchased (e.g. Rapid, Rimpar, Germany). For both systems, external triggering and monitoring systems for physiological signals (heartbeat, breathing, temperature) as well as gating of the image acquisition are commercially available (e.g. Biopac Systems Inc., Essen, Germany, SAI Inc, Stony Broke, NY, USA) but low cost in-house built systems also exist [36]. These system are specific for the high heart and breathing rates of small rodents (heart rate of 300-800 bpm in mice and 300-500 bpm in rats; respiration rate of 100-200/min in mice and 70-110/min in rats, [31]). Besides the monitoring of the physiological condition of the animal, these systems are also used for cardiac and or respiratory gating of the image acquisition to enhance image quality [31,36]. For complete control of the respiration of the animal, intubation or tracheotomy is required. With respect to the papers on kidney functional imaging reviewed in this work, only one study reported a gating strategy [37]. Few studies relied on image registration [38–40] as post processing step while all others omitted motion correction. However, triggering and respiratory gating is more commonly applied in cardiac functional imaging (see second part of this review [32]). Eventually, in all studies performed at MRI systems, the animals have to be sedated using an appropriate anaesthesia. The

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selection of the anaesthesia thereby heavily depends on the study aims and the preparation of the animals beforehand the MRI examination. Therefore, it has to be adjusted accordingly. Briefly, there are two main application schemes in small animal imaging; gas anaesthesia using inhalant anaesthetics like isoflorane or halothane or the injection of the anaesthetics via subcutaneous, intraperitoneal, or intervenous administration. While the first is commonly employed in studies using dedicated small animal scanners (cf. [23,38]), the latter is more often used at whole body scanners that are not commonly equipped with a dedicated small animal gas inhalation unit [41–43]. Moreover, some drugs alter the physiological parameters like blood pressure, breathing rate, or temperature; others have side effects during the awaking of the animal. All this might influence the study outcome and should be taken into account when planning the experiment. An overview of different anaesthesia regimes, pro and cons of the used substances, etc., is given by Hanusch et al. [44]. Glomerular Number Glomeruli are nodes of capillaries in the kidney that filtrate and purify plasma, thus maintaining the salt and water and the pH balance in the blood. Waste and excess from the cells are released from the body after the filtration via the urine [45]. The number of glomeruli in the kidney, officially known as glomerular number (Nglom), has been proposed as a direct measurement of the condition of the kidneys by itself [46] and also multiplied by the glomerular size to find the glomerular volume and the filtration surface area [47]. Several publications have established a relation between glomerular numbers and various factors like age [48], weight at birth [49], and body surface area [47], as well as illnesses and diseases including diabetes mellitus [50], hypertension [51], and chronic kidney disease (CKD) [52]. Diagnosing anomalies in the number and/ or size of the glomeruli could prevent or slow down the effect of degenerative diseases via the application of early treatments to avoid the need for dialysis or a kidney transplant. Additionally, better planning and early arrangements could be made for those procedures. Then, even after transplantation, the rejection of the transplanted kidney can produce focal segmental glomerulosclerosis that could lead to renal failure [53]. A mean to obtain glomerular numbers and sizes could help evaluating the degree of rejection of a transplanted kidney, avoiding or minimizing damage to the organ and the patient. There are five methods to estimate glomerular numbers. Four of them involve the partial or total destruction of the kidney to allow for manual glomerular identification and counting from small samples of the kidney e.g. some physical slices to calculate the whole glomerular number based on those randomly chosen samples [54]. The fifth method, based on MRI [55,56] has been recently introduced to image, segment, and count individual glomeruli in whole kidneys. It was initially presented for in vitro rat kidneys labeled with cationized

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ferritin. Given the drawbacks of the gold-standard method (i.e. it destroys the kidney, is time consuming, and is highly operator dependent), it is evident that this method is neither efficient nor useful for animal models nor assessment of patients’ kidneys. The MRI approach to provide Ngloms (e.g. of rat kidneys) can be divided into labeling, imaging and image post-processing. The labeling is achieved by synthesizing cationized ferritin (CF) from horse spleen [55,57] and injecting it into the rodent as described in [56,58]. The kidney is then perfused, extracted from the animal and embedded in agarose gel. The imaging has been so far done in 7, 9.4, and 18.8T animal/ research scanners using gradient echo (GRE) sequences to obtain images with effective transverse relaxation time (T2 * ). This accentuates the main effects of the CF: it creates local inhomogeneities that extend beyond the glomeruli and reduces the T2 * of the CF-labeled glomeruli. The details of the effects of the CF as contrast medium are beyond the scope of this review but can be consulted in the literature [59,60]. The images need to be taken at a very high resolution, e.g. up to 35x35x35 ␮m3 [56]. This is a challenge in MRI because the signal-to-noise ratio (SNR) decreases with the decrease in size of the imaged voxels, which requires the use of long scanning times. The total acquisition time for a 3D dataset of a full rat kidney has reportedly taken no less than 4 hours [56]. After the imaging, the full 3D datasets are postprocessed and the glomeruli are segmented, quantified and sized using custom algorithms. The results were validated by comparison to the gold-standard method. Filtration Currently, the most important and widely employed assessment of the kidney function is the glomerular filtration rate (GFR) [61]. GFR is commonly analyzed from blood samples and provides information about the filtration efficiency of the kidneys but it does not provide anatomical information. Several tracers are available for GFR evaluation, including endogenous markers such as creatinine and cystatin C, or exogenous markers, including iothalamate [62] or inulin, which is accepted as gold standard for determination of renal function. Current techniques for the measurement of GFR, such as clearance of inulin or sinistrin, scintigraphy with radio-labeled markers, and creatinine clearance are limited as they are invasive, expensive, result in a radiation exposure, or are inaccurate, e.g. because of serum creatinine dependency on muscular tissue mass and nutritional factors [63,64]. Thus, GFR is calculated only approximately using approaches such as the Chronic Kidney Disease Epidemiology Collaboration (CKD-EPI) or Modification of Diet in Renal Disease (MDRD) [65,66]. Alternative GFR measurement techniques are dynamic contrast enhanced magnetic resonance imaging (DCE-MRI) [4,67,68] or transcutaneous approaches based on fluorescent tracer molecules [69–72]. Both approaches allow for a non-invasive and radiation free way to measure GFR.

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However, the transcutaneous methods reported only provide a whole kidney GFR while DCE-MRI also allows for single kidney GFR and even spatially resolved analysis of the GFR through calculation of parametric maps by applying different physiological compartment models [73–75]. Imaging is performed in 2D, i.e. either capturing a single slice [73] or performing a multi-slice acquisition [43] and in 3D [42,76,77]. In all examinations, contrast agents are used to allow for bolus tracking experiments. Since all of these techniques can similarly be used to assess kidney perfusion, a more detailed description of the employed sequences is given in the next section. Achieved spatial resolution range from 0.3 x 0.3 x 1.0 mm3 (7T, dedicated animal system) to 1.25 x 1.25 x 4.00 mm3 (1.5T, whole body system). Slices or volumes were recorded at low temporal resolution (16s) but also high temporal resolutions up to 0.5 s were reported. A detailed overview of the temporal-spatial resolutions and the employed scanners is given in Table 1. Perfusion Another important parameter for kidney function assessment is the measurement of tissue perfusion. The perfusion describes the amount of blood flowing through the capillary bed supplying the tissue with oxygen, nutrients, and electrolytes. There are basically two ways of measuring the perfusion using MRI, arterial spin labeling (ASL) and DCEMRI. Briefly, DCE-MRI involves the injection of a contrast agent, usually gadolinium (Gd) based, as a tracer. By repeatedly (fast) imaging the same field of view (FOV), the passage of tracer can be recorded (see Fig. 1). To assess renal blood flow (RBF), different pharmacokinetic models [73–75] are applied to the signal intensity/ concentration time curves. In pre-clinical MRI, model free approaches like deconvolution techniques [41] as well as (multi-) compartment models were employed [23,76–78]. Details on general aspects of perfusion weighted imaging as well as the data analysis and modeling are omitted here. The reader is referred to recent reviews [2,3,72,79]. Regarding the choice of a suitable model for quantification of the renal perfusion in small animals, Zöllner et al. compared RBF estimated using a model free analysis (deconvolution) and a 2-compartment filtration model in healthy rats and rats subjected to AKI [78]. Significant differences in RBF between models could not be observed. Moreover, both models could differentiate healthy from diseased kidney at similar accuracy. In (small) animal perfusion MRI, 2D and 3D acquisitions are performed. The latter require fast image acquisition strategies to capture the bolus and to sufficiently sample the tissue signal curve but also the arterial input function (AIF) since the blood circulation times are high, especially in rodents. Therefore, recent studies performed on clinical whole body scanners employed view sharing techniques like Time-resolved angiography WIth Stochastic Trajectories (TWIST) [41,76,78] or

Time Resolved Imaging of Contrast KineticS (TRICKS) [80] to achieve temporal resolutions of up to 0.5 s and in plane resolutions between 0.5 mm and 1.25 mm (see Table 2). Furthermore, Winter et al. [77] used a dual bolus approach to allow for a very high sampling rate of the AIF which is important for quantification of the perfusion. A pre bolus, i.e. a small fraction of the total amount of contrast agent was used for sampling the AIF at high temporal but lower spatial resolution. Thereafter, the remaining amount of contrast agent was injected and a classical bolus tracking experiment was performed with a lower temporal resolution but therefore, increased spatial resolution. Other studies employed 2D image acquisition techniques to aim at higher temporal resolutions. The study of Pedersen et al. [81] achieved a temporal resolution of 0.6s and half millimeter in plane resolution acquiring a single slice while Sadick et al. [43] performed a multi slice approach, however, the temporal resolution was in the range of the aforementioned 3D approaches. Regarding dedicated ultra-high field animal scanner, most employed acquisition scheme were parameterized towards higher spatial resolutions rather than temporal resolution (see Table 2) most likely due to the small size of rodents. Only one study could achieve a high spatial and also a 1s temporal resolution [82]. In the study of Yang et al. [83] a temporal resolution of 0.3 s was reached, here the spatial resolution lay within the range of whole body clinical scanners. In recent years, cases of nephrogenic systemic fibrosis (NSF) associated with gadolinium based contrast agents have been reported [84]. ASL does not have this disadvantage. It is a completely non-invasive method that uses magnetically labeled water protons in arterial blood as an endogenous tracer [85,86]. Briefly, two images are acquired, one with magnetical spin preparation via inversion recovery (also called tag image) and one without labeling (control image). By subtracting these two images a perfusion weighted image is created. Technically, by combining one of the three main labeling schemes (pulsed labeling (PASL) [86–89], continuous labeling (CASL) [90] or pseudo-continuous labeling (pCASL) [91]) with an adequate imaging sequence, a perfusion sensitive experiment can be designed. As no exogenous tracer is used, it is a technique that can be repeatedly employed on the same patient within short time intervals. To quantify perfusion from the perfusion weighted image, either the Bloch equation approaches by Detre et al. [85] and Williams et al. [90] or the general kinetic model of Buxton et al. [92] can be used. In the case of using a single inversion time, also the QUIPSS/ QUIPSS II (quantitative imaging of perfusion using a single subtraction) [93] or Q2TIPs (QUIPSS II with thin-slice TI 1 periodic saturation) model [94] are utilized to compensate for varying bolus arrival times within the tissue. In ASL imaging an intrinsic problem is that a sufficient signal to noise ratio is needed for proper quantification as the obtained perfusion signal is only about 1%. To overcome this problem, lower spatial resolutions including thick read out


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Table 1 Overview of GFR estimation in pre-clinical MRI studies. Besides the used MRI system also the temporal/spatial resolution achieved is given. Note, that Winter et al. used a pre-bolus approach for higher temporal sampling of the arterial input function. MRI system

Spatial resolution (mm)

Temporal resolution (sec)

Animal model

Reference

3T, wbs

1.0 x 1.0 x 2.5 0.6 x 0.6 x1.2 0.9 x 0.9 x 4.0 1.3 x 1.3 x 4.0 0.3 x 0.3 x 1.0

0.8 0.9 1.13 0.5 / 2.8 16

Rat Rat Canine Canine Rat

Sadick et al. [43] Zöllner et al. [42,76] Annet et al. [73] Winter et al. [77] Oostendorp et al. [23]

1.5 T, wbs 7T, dsas

Abbreviations: wbs = whole body scanner, dsas = dedicated small animal scanner.

slices (> 2 mm) and a high number of averages are usually used to increase the SNR. Furthermore, high field or even ultra-high field scanners are preferred. While in plane spatial resolutions on clinical whole body scanner can be improved utilizing a field strength of 3.0 T (see Table 2), a gain in spatial resolution and also a smaller slice thickness are reached on dedicated animal systems [38,80]. All reviewed studies used a imaging sequence acquiring single 2D slice and the fluid attenuated inversion

recovery (FAIR) labeling scheme [87] (see Fig. 2). Only Liu et al. [95] and Winter et al. [77] used a multi-slice approach covering larger parts of the kidneys rather than a central slice. Oxygenation The presence of paramagnetic deoxyhemoglobin in venous blood generates variations in magnetic susceptibility within the vessel and its surroundings and is well known as blood

Figure 1. Illustration of dynamic contrast enhanced MRI of the kidney in a rat model of AKI. Image acquisition was performed using a 3D T1-weighted TWIST sequence with spatial resolution of (0.6 x 0.6 x 1.2) mm3 and a temporal resolution of 0.9 s (see [76] for details). Here, a central slice of the 3D volume was selected for illustration purposes. AKI was performed on the left kidney. Contrast uptake in the kidney: A) before injection (no uptake), B) first pass, and C) washout phase. D) illustration of ROI placement for deriving the signal intensity curves displayed in E). These signals are used to calculate functional parameters like RBF and GFR using dedicated pharmacokinetic models.

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Table 2 Temporal and spatial resolution for perfusion MRI sequences comprising ASL and DCE-MRI used in pre-clinical studies at different field strengths and scanner hardware. Technique

MRI system

Spatial resolution (mm)

Temporal resolution (sec)

Animal model

Reference

ASL

3T, wbs

0.5 x 0.5 x 4.0 0.8 x 0.8 x 2.0 2.8 x 2.8 x 8.0 2.5 x 2.5 x 5.0 2.7 x 2.7 x 8.0 0.5 x 1.0 x 2.0 0.3 x 0.3 x 2.0 0.5 x 0.5 x 2.0 1.3 x 1.3 x 4.0 1.0 x 1.0 x 2.5 0.6 x 0.6 x1.2 0.5 x 0.5 0.4 x 0.8 x 1.2 0.2 x 0.4 x 1.0 0.9 x 0.9 x 2.0 0.5 x0.5 x 1.5 0.3 x 0.3 x 1.0

N/A

Rat

Zimmer et al. [41] Liu et al. [95] Artz et al. [159] Winter et al. [77] Wentland et al. [160] Wang et al. [80] Hüper et al. [38] Pedersen et al. [167] Winter et al. [77] Sadick et al. [43] Zöllner et al. [41,76,78] Vexler et al. [157] Sari-Sarraf et al. [164] Beckmann et al. [82] Yang et al. [83] Gaschen et al. [155] Oostendorp et al. [23]

1.5 T, wbs

DCE-MRI

4.7 T, dsas 7T, dsas 1.5 T, wbs 3T, wbs 2T, dsas 4.7 T, dsas

7T, dsas

Swine Canine Swine

0.6 0.5 / 2.8 0.8 0.9 1.5 6.0 1.0 0.3 0.9 16

Mice Rat Canine Rat Rat Rat Rat Rat

Rat


oxygen level-dependent (BOLD) MRI [5]. This technique was invented by Ogawa in 1990 [96] and thereafter was mainly employed in fMRI. However, when the tissue oxygenation state is altered, the ratio between oxyhemoglobin and deoxyhemoglobin concentrations will change accordingly, resulting in a local change in magnetic susceptibility. These differences can be detected via changes in the apparent transverse relaxation rate R2 * , where an increase in R2 * indicates a decreased oxygenation, and vice versa. This makes this technique also a valuable tool for functional imaging of the kidney [97,98]. The oxygenation can be measured indirectly by MRI via the transverse relaxation time T2 * using multi-echo GRE sequences [99–101]. These sequences are either performed as 2D acquisition [8,81,102–104], e.g. one slice, or as multi slice acquisition covering the whole kidney [23,39,40]. The number of employed echo times range from two to 12 echoes. Image in plane resolution including interpolation was between 0.16 mm2 (at 9.4 T) and 0.5 mm2 at whole body clinical scanners (see Table 3). To enhance the recorded signal, up to 16 averages per image were performed. Diffusion Diffusion weighted imaging (DWI) assesses the microscopic mobility of water in tissues, with diffusion describing the random thermal motion of gaseous or liquid molecules. MRI can detect signal changes due to microscopic positional changes of water molecules. In a homogenous medium the diffusion can be described by a Gaussian distribution. However, in tissue, the diffusion of (water) molecules is restricted. This restriction can be resolved by MRI, too, and thus, provides insight into structural changes on a microscopic level. Therefore, DWI is nowadays applied for diagnosing diseases

like stroke, cancer, or functional status of organs [2,105,106]. Thereby, the apparent diffusion coefficient (ADC) could be extracted from a diffusion weighted MR examination but also more detailed DWI models are emerging, namely intra-voxel incoherent motion (IVIM) diffusion weighted imaging [107] or kurtosis diffusion weighted imaging (DKI) [108]. Compared to clinical whole body scanners with a limited gradient system, dedicated ultra-high field scanner provides stronger and fast switching gradients that allow for large b-values and more demanding DWI measurements like Q-ball imaging at reasonable image acquisition times. For DTI, in the pre-clinical setting, the reported image acquisition approaches comprise of 6 directions with two b-values [109]. For diffusion weighted MRI, studies employed several b-values up to b = 1000 s2 /mm to allow for bi-exponential fitting, i.e. to employ the IVIM model [37,109–112]. In all studies, a spin echo echo planar imaging sequence (SE-EPI) was used. To reduced bulk motion a gated spin echo DWI sequence was employed by Yang et al. [37]. In plane resolution, achieved on the different MR systems, range from 0.5 – 0.9 mm2 for rodents and up to 1.7 mm2 for minipigs, respectively (see also Table 4). Sodium MRI Sodium ions are the most abundant cations in the extracellular space in the body regulating the cell membrane resting potential. The amount of Na+ is strongly linked to the vital pump mechanism Na+ -K+ -ATPase. This energy dependent antiporter enzyme located in the plasma membrane maintains the cell membrane potential by removing three Na+ from and moving two K+ into the cell at the cost of 1 ATP molecule.


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Figure 2. Illustration of FAIR ASL labeling scheme in a healthy rat. On the right-hand side, exemplary tag, control and M0 images as recorded during the 2D ASL measurements of renal perfusion in rats are shown. Tag and control image are scaled identically. On the left, the positioning of imaging (red) and inversion slice (green) is illustrated. The underlying image depicts a T2-weighted coronal slice through the rat abdomen that allows a good identification of the kidneys. The image shows a rat with two healthy kidneys.

Hence, the extracellular Na+ concentration is about 5-10 times higher than the concentration inside the cell, which results in a mean tissue sodium concentration (TSC) of about 45 mM in healthy tissue. This concentration gradient can be altered under pathological conditions as associated with ischemia, stroke, or cancer. Therefore, the non-invasive 23 Na MRI represents a potential tool for monitoring these changes. In early ex-vivo experiments [113–115] and in recent in vivo renal 23 Na MRI studies [11,116–119], the concentration of Na+ ions was measured to have a linear corticomedullary profile which is increasing from the cortex (60 – 96 mM) to the inner medulla (180-350 mM) depending

on the dehydration state of the healthy kidney. The renal concentration gradient can be altered due to renal diseases, drugs, and diuretic agents. Regarding the nuclear characteristic properties and the natural abundance in tissue, sodium (23 Na) represents the most convenient X-nucleus for MRI after proton with nonzero nuclear spin. Non-invasive 23 Na MRI does not rely on administration of contrast agents or ionizing irradiation, but instead, it makes use of naturally abundant Na+ ions in biological tissue. Furthermore, 23 Na MRI allows for the absolute quantification of the tissue sodium concentration (TSC) using the 23 Na MR signal. This is an essential tool for diagnostic

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Table 3 Spatial resolution achieved for BOLD imaging in pre-clinical studies at different field strengths and scanner hardware. MRI system

Resolution (mm)

Animal model

Reference

1.5 T / wbs

0.5 x 0.5 x 3.0

Rat

2.3 x 2.3 x 5.0 0.6 x 0.6 x 3.0 1.3 x 1.5 x 5.0 1.7 x 1.7 x 4.0 0.2 x 0.2 x 0.5 0.5 x 0.5 0.2 x 0.2 x 1.0 0.3 x 0.3 x 2.0 0.2 x 0.4 x 1.4

Swine Mice Swine Minipigs Mice Rat Rat Rat Rat

Prasad et al. [8] Pedersen et al. [81] Pedersen et al. [167] Ries et al. [112] Wentland et al. [160] Haneder et al. [111] Prasad et al. [103] Vexler et al. [165] Oostendorp et al. [23] Rognant et al. [104] Pohlmann et al. [39,40]

3T / wbs 2T / dsas 7T / dsas 9.4 T / dsas


imaging due to the direct link between Tissue Sodium Concentration (TSC) and tissue integrity and viability. Therefore, the ability of monitoring the concentration changes using functional 23 Na MRI could provide special insights into the pathophysiological processes on the cellular level. However, in vivo (renal) 23 Na MRI is challenging since the signal sensitivity is c.a. four magnitude lower compared to 1 H MRI. This is caused by a lower in vivo concentration in tissue, a lower gyromagnetic ratio, fast sodium signal decay, and the required small voxel sizes (< 4 ␮l) in the rodent kidney. Therefore, several SNR optimization steps are required to establish 23 Na MRI as a standardized functional imaging modality. These can be summarized into three main topics: the measurement of TSC, the development of particular RF resonator systems for 23 Na MRI and the development of appropriate MR measurement sequences. The absolute quantification of the tissue sodium concentration (TSC) values using 23 Na MRI is based on a linear relationship between the 23 Na MR signal and the sodium concentration. This linear relationship requires that 23 Na MR images are recorded using neither with T1 , T2 * relaxation time weighting nor with variations in flip angles across the sample. In the literature, there are several methods for the quantification of the sodium concentration. Ouwerkerk et al. [120] introduced a method based on a sample and reference scan to quantify TSC in breast tissue. This method included a loading correction factor. It is the most common and established method applied for the TSC quantification. The quantification

method based on the reference scan has been previously applied to quantify 23 Na concentration in rodent brain tissue [121], in the human breast [120], and in the rodent kidney [11]. Further improvements concerning RF coil development and MR imaging techniques allowed an accurate TSC quantification method independent of the relaxation times, the coil loadings, and the B1 profile effects using separated transmit and receive elements (referred to as the dual RF resonator system) [122]. The accuracy of this quantification method was assessed as follows: First, the homogeneous phantom scan is performed prior to the sample scan including the target region. Second, the sample scan is co-registered with the reference scan due to the positioning vials fixed on top of the coil. Third, signal sensitivity variations caused by different coil loadings are corrected, and finally the quotient image of both co-registered images is weighted by the known concentration generated the concentration map. The known concentration can be extracted from external reference vials or internal references, e.g. spinal fluid. Furthermore, errors due to B1 homogeneity of the RF resonator can be corrected by means of B1 mapping methods already suggested for 1 H MRI. For the acquisition of 23 Na MRI signal without any T1 or T2 * weighting, the homogeneity of the transmit B1 -field is of essential importance. For RF excitation with uniform flip angles, the most appropriate coil design is a transceiver volume resonator [119]. However, volume resonators suffer from low SNR due to their large diameter and consequently due to their bad filling factor. The sodium signal sensitivity can be

Table 4 Pre-clinical diffusion MRI studies. Besides achieved image resolutions and employed scanner hardware also the diffusion weightings and animal model are given. MRI system

Resolution (mm)

b- values (s/mm2 )

Animal model

Reference

1.5T / wbs

0.6 x 0.6 x 3.0 1.7 x 1.7 x 4.0 0.8 x 0.8 x 2.0

Mice Minipigs Rat

Ries et al. [112] Haneder et al. [111] Cheung et al. [109]

7T / dsas

0.5 x 0.5 x 2.0

b = 0, 130, 260, 390 b = 0, 50, 100,400, 800 b = 0, 300 (DTI, x directions) 8 b-values range 0 – 1000 (DWI) b = 5, 20, 42, 72, 142, 260

Rat

Yang et al. [37]



9

Figure 3. Dual RF resonator system for bilateral 23 Na MRI of rat kidneys at 3T using the whole body 23 Na RF resonator (34 cm height and 47.5 cm width) combined with a receive-only saddle surface coil.

further improved using transceiver surface coils [116,117], which in turn suffer from insufficient B1 -field homogeneity. For the quantification of the TSC, besides a dual resonator system, also a transmit-only receive-only (TORO) system [122,123] can be used, combining the advantages of both coil designs, i.e. the B1 transmit field homogeneity of a volume resonator with the high sensitivity of a receive-only surface coil. For the first time, Barberi et al. introduced the transmit-only receive-only (TORO) system, so-called dual RF resonator, which bases on the idea of separated transmit and receive RF elements [123]. In contrast to the standard RF resonators in transceiver (TXRX) mode, e.g., surface coil [124] or birdcage [125], the dual RF resonator allows for both a homogeneous transmit B1 field and a highly sensitive receive B1 field. This technique is well established in human MRI due to an in-built 1 H volume resonator combined with multi-channel receive arrays, which also allows for taking the advantage of the parallel imaging technique [126]. However, for 23 Na MRI there is no standard RF hardware in clinical scanners. An overview of the newly developed dual RF resonator systems applied to kidney 23 Na MRI in 3T or 9.4T MRI scanners is presented in Figures 3 and 4. It is composed of a homogeneous transmit-only volume resonator and a highly sensitive RO surface resonator, so that the sensitivity correction needs to be performed only for the receive profile. Furthermore, the dual resonator allows for increased 23 NaMR 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 9.4T and a receiver saddle shaped RF resonator at 3T. These systems are developed for measuring the absolute renal sodium concentration (RSC) in in vivo rat kidney models with high spatiotemporal resolution, and high concentration measurement accuracy. A comparison of 23 Na MRI experiments using a dual resonator system at 9.4T and 3T for in vivo rat kidney experiment in the furosemide model was presented in Figure 5. For 23 Na MRI, the pulse sequences are required to be designed based on the special relaxation properties of the 23 Na nuclei [127], i.e. sequences with short TE [128–131]. The recently published 23 Na rodent or murine kidney studies [117,119] used Gradient Echo (3D-GRE) or Chemical Shift Imaging (3D-CSI) sequences. In a recent work, an alternative 3D MR pulse sequence with anisotropic radial acquisition technique, so called Ultra-Short Time-to-Echo (3D-UTE) sequence, was integrated for bilateral renal 23 Na MRI [11]. Furthermore, anisotropic cones [132], or anisotropic TPI [133] sequences already applied for human 23 Na MRI could be integrated into the small animal imaging at high fields.

Applications After outlining functional techniques for renal imaging, in this section dedicated applications of these techniques in

Figure 4. The dual RF resonator system for bilateral 23 Na MRI of rat kidneys at 9.4T using a home-built transmit volume resonator (inner bore diameter of 7.2 cm) combined with a two-element 23 Na receive loop array.

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Figure 5. Transversal 23 Na MR images of the rat kidney at 9.4T with a spatial resolution of (1x1x4) mm3 acquired before and after furosemide application compared to the 23 Na MR images acquired at 3T with a spatial resolution of (1.5x1.5x1.5) mm3 before and after furosemide application. For both experiments at 3T and 9.4T an acquisition time of 10 min was needed by using the 3D radial acquisition technique and a zero-filling by factor 2.

pre-clinical studies are reviewed. Table 5 gives an overview about animal models and techniques employed to measure renal function discussed in the following. Counting Glomeruli using MRI MRI starts appearing as a very promising method to provide accurate glomerular numbers and a remarkable alternative to all other four methods used for this purpose. Nevertheless, the gold-standard method to estimate glomerular numbers and measure glomerular sizes remains the design-based stereology [134,135] even though the dissector method [136] destroys the kidney, and relies on the Cavalieri or fractionator [137–139] principles to help an operator to do a manual count. A good

overview of all the methods to obtain Ngloms was presented by Bertram et al. in 2013 [54]. In vivo images of glomeruli labeled with CF were shown in 2008; the labeling method was presented then [57]. After that, the first glomerular numbers were obtained using MRI in 2011 [55] and 2012 [56]. In 2013, efforts to quantify glomeruli in human transplanted kidneys [140] and mice [140] were presented. In vivo images of glomeruli have been shown in rats with an alternative paramagnetic T1 contrast agent [141]. Development of new coils was also performed to compare cryogenic probes (Bruker cryoprobe, Ettlingen, Germany) and volumetric Alderman-Grant resonators to visualize glomeruli [142] (see Fig. 6). These recent results are very promising but more work needs to be done in order to achieve the first in vivo glomerular quantifications. Even after the labeling, the small size of the glomeruli in comparison to the FOV necessary to scan the whole kidney (e.g. in rats, diameter of CF labeled glomeruli = 122 ␮m, FOV = 26.88x17.92x8.96 mm3 [55]) forces the datasets to be acquired at very high resolutions (e.g. mice = 34x34x54 ␮m [140], rats = 62x62x78 ␮m [141], humans = 117x117x117 ␮m [58]). These resolutions require strong gradient moments and high magnetic fields to obtain enough SNR and contrast to noise ratio (CNR) to segment the glomeruli. Although ultra-high fields have been used (7, 9.4T, and 18.T), the number of repetitions required force the scanning times to be much longer than any type of in-vivo experiment could last for. An overview of the employed scanner hardware and achieved image resolution is given in Table 6. Scanning patients and turning glomerular quantification into a clinical routine diagnostic procedure would be the final purpose. It has been already suggested that in vivo human trials could be possible in the future [140] but some challenges must be addressed first. The current MRI based methods will take very long to be adapted to measure humans, whereby also lower magnetic fields are used (1.5, 3, and less commonly 7 T) which decreases the effect of the CF [57]. Additionally, restrictions in the use of power are given via the specific absorption rate (SAR) regulations. Weaker gradient moments need to be used to protect patients. This introduces additional challenges in minimum achievable FOV, maximum resolution, SNR, and CNR. Ranging from ∼4 (in rats) to 10 hours (in human transplanted

Table 5 Overview of renal disease model and functional MRI techniques employed to study the models. Kidney disease model Kidney transplantation Acute kidney injury Renal artery stenosis Unilateral nephrectomy Diabetes Hypo / hyperperfusion Acute tubular necrosis Diuresis

BOLD

Perfusion

GFR

Diffusion

Sodium

x x

x x x

x

x x

x

x

x x x x

x x

x x

x

NGlom

x x x


11

Figure 6. Kidney images visualizing glomeruli. Improved homogeneity shown in the central slices in the slab for each axis. A to C were acquired with a cryogenic surface probe. D to F were acquired with a dedicated Alderman-Grant probe.

kidneys) per sample, the main obstacle at the time this review is written is the acquisition time. A new approach needs to be taken in order to reduce this time before proceeding to in vivo trials, i.e. a total acquisition time < 2 hours for in vivo studies). Perfusion In animals, DSC/DCE-MRI and ASL have been used to assess perfusion in various organs and disease models. This comprises peripheral tissue perfusion in diabetic rats [143], myocardial blood flow [144] in infarcted rat hearts [145], fMRI in rodent brain [146,147], cerebral blood flow [148], e.g. after traumatic brain injury [149] or stroke [150,151], breast cancer bone metastasis [152] and even retinal and choroidal blood flow responses to acute hypertension [153]. Due to its potential as an indicator for renal impairment and the high renal blood flow, the measurement of kidney

perfusion [154] is of great interest. Earlier studies have already used dedicated animal models, primarily rat. The models were chosen with regard to a later application of the findings in clinical routine. This way, for example the perfusion in transplanted [80,82,83,155,156] or injured kidneys [38,41,81,157] was investigated. At dedicated high field animal scanners only few studies investigated renal perfusion in small animals so far. Hüper et al. compared renal perfusion in AKI of mice measured by ASL with renal histological examination and inulin and para-aminohippuric acid (PAH) clearance [38]. Their results showed that the degree of perfusion impairment measured by means of MR imaging was related to kidney volume loss, the severity of histopathologic alterations of renal tissue, and impairment of renal function. Oostendorp et al. also evaluated renal perfusion in AKI but in a rat model [23]. They performed a DCE-MRI scan 24 hours post the induction of the AKI. The

Table 6 Estimation of glomerular number by MRI. Here reported system and achieved image resolution along with the used models is given. MRI System

Resolution (␮m)

Animal model

Reference

19T / spectrometer

62 x 62 x 78 34 x 34 x 34 35 x 35 x 35 25 x 25 x 25 117 x 117 x 400 117 x 117 x117

Rat (ex-vivo) Mice (ex-vivo) Rat (ex –vivo) Mice (ex–vivo) Rat Human transplant (ex-vivo)

Beeman et al. [55] Beeman et al. [140] Heilmann et al. [56] Chacon et al. [142] Beeman et al. [141] Beeman et al. [140]

9.4 / dsas 7T / dsas


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Figure 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 from [41].

renal blood flow was significantly different between AKI and healthy kidney. While the experiments on small animals have been mainly conducted on dedicated animal scanners, recently, clinical whole body scanners were also used for small animal studies. Zimmer et al. [41] 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 (see Fig. 7). Both, the FAIR-ASL approach and DCE-MRI deconvolution technique showed significant differences in RBF between healthy and diseased kidneys. The study was conducted on a 3 Tesla whole-body scanner and used imaging protocols that can easily be adapted to measure RBF in humans. Alternations in RBF induced by the injection of contrast agent have been investigated by Liu et al. [95]. Like Zimmer and co-workers a dedicated ASL sequence with a FAIR labeling scheme was used and the measurements were conducted on a clinical 3 Tesla system. RBF was assessed at four different points in time after the injection of a contrast agent. The authors showed that an iodinated contrast agent significantly decreased the RBF in outer medulla and cortex of spontaneously hypertensive rats compared to normointensive rats which only showed a decrease in the outer medulla. While MRI of small animals can always be conducted on both dedicated animal scanners and clinical whole-body systems the (functional) imaging of larger animals is restricted

to the latter. Due to their physiological similarity to humans, swine have been used for interventional studies such as renal artery stent placement [158], pharmacologic and physiologic alterations in renal blood flow [159,160] and kidney transplantation [161]. Lüdemann et al. investigated the possibility of absolute regional quantification of renal medullary and cortical perfusion in a pig model using DCE-MRI and a 1.5 Tesla scanner [162]. An inflatable cuff around the renal artery allowed for four different states of perfusion in overall 11 pigs. The values were highly correlated with the absolute renal blood flow measured by an ultrasound transit time flow probe around the renal vein and the regional blood flow in medulla and cortex were correlated with the degree of flow reduction. A similar study was conducted by Artz and co-workers who measured the RBF of 11 swine under four different pharmacologic and physiologic conditions using a FAIR-ASL approach on a 1.5 Tesla clinical scanner [159]. There, the RBF measurement with fluorescent microspheres served as a gold standard. Both techniques were well correlated and tracked the expected RBF changes. Regarding absolute perfusion, ASL values were systematically lower compared with microsphere perfusion. In comparison, the goal of the study of Notohamiprodjo et al. was to measure errors in perfusion and filtration parameters derived from DCE-MRI and caused by protein binding of contrast agents [163]. Therefore, the RBF of eight healthy swine was assessed with DCE-MRI using


13

Figure 8. Image examples obtained by the MR measurements. A) depicts one slice of a T2 weighted morphological MR image used for cortex delineation, B) shows a parametric map of the tubular flow (filtration) calculated from the DCE- MRI perfusion data in units of ml/min/100 ml tissue. The map is superimposed onto the corresponding slice of the T2 weighted image. Reproduced from [42].

two different contrast agents and evaluating the data with a two-compartment filtration model. The authors showed that perfusion values of both agents are comparable. Changes in renal function by pharmacologically induced hyper- and hypoperfusion have been assessed by Wentland et al. in 11 swine [160]. Besides BOLD and phase contrast flow measurements, renal regional perfusion was measured with ASL. The cortical perfusion demonstrated the expected changes given the induced changes in renal function. A study that investigated the renal perfusion in transplanted porcine kidneys was conducted by Soendergaard et al. [161]. In a model of donor recipient size mismatch the effect of remote ischaemic conditioning was investigated for which the perfusion was significantly higher compared to a control group. Cortical and medullary perfusion was assessed with DCE-MRI using a 1.5 Tesla machine. GFR Glomerular filtration rate as renal perfusion can be calculated from DCE-MRI. Since DCE-MRI studies for investigation renal perfusion were performed seldom at high field animal scanners, similar, the number of papers reporting studies of GFR is low. Change of GFR with respect to injection of nitric oxide synthase inhibitor NG-nitro-l-arginine methylester (l-NAME) was investigated by Sari-Sarraf et al. [164]. To measure the GFR, the authors performed DCE-MRI in healthy rats at 4.7 T. A direct calculation of GFR by pharmacokinetic modeling was not performed but only semi-quantitative parameters like time to peak were reported. Results showed that with increasing dose of L-NAME, the uptake of contrast agent in the kidney is reduced. The authors concluded that the GFR is altered in the presence of L-NAME. Oostendorp et al. employed for their study of AKI a two compartment filtration model that allows for calculating renal blood flow and GFR [23]. Similar to their findings regarding renal blood flow, also the GFR is significantly reduced in the AKI kidney.

At whole body MR scanners several studies investigated renal filtration in animal models. Annet et al. estimated the GFR in the rabbit kidney [73]. Winter et al. compared renal perfusion estimated by ASL and DCE-MRI in healthy rats [77]. As byproduct they also calculated the GFR using a separable two compartment filtration model [75]. The validation of MR-based GFR estimation by optical imaging was proposed by Sadick et al. [43]. GFR estimation was performed by DCE-MRI on a 3T 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 MR-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. Zöllner et al. recently showed a simultaneous measurements of optical and MR-based GFR [42]. In this study healthy rats and rats with unilateral nephrectomy (UNX) were used. The two compartment filtration model [75] employed produces a map of the tubular flow (see Fig. 8). Based on this parameter, the single kidney GFR could be calculated taking the cortex volume into account. GFR values between methods correlated well. The reduction of GFR in the UNX rats was about 50% as expected. Also this reduction was the same for both techniques. Zöllner et al. [76] recently investigated if the GFR and additional parameters like renal blood volume or mean transit times, that are calculated by the employed 2-compartment model, can discriminate between healthy kidneys and kidneys subjected to AKI. While the GFR showed significant differences between healthy and diseased kidney, no significant differences between healthy and diseased kidneys could be obtained for the other parameters. Oxygenation Renal oxygenation is known to reflect renal diseases like acute kidney injury or renal artery stenosis and has been

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Figure 9. Change of renal T2 * and T2 during hypoxia, hyperoxia, ischemia and reperfusion. Shown are T2 * and T2 difference maps of the kidney (color-coded, overlay on anatomical MR image) between the last time point in each experiment phase and baseline. The three interventions (hypoxia, hyperoxia, ischemia) led to a reasonably homogeneous decrease/increase in both parameters. During ischemia the magnitude of the T2 * / T2 reduction exceeded that during hypoxia. Parameter changes in the reperfusion phase clearly differentiated cortex and medulla: while medullary T2 * / T2 remained reduced throughout reperfusion, in the cortex T2 * returned to baseline and T2 rose above baseline. Reproduced from [40].

identified as important marker [5,100]. Renal oxygenation studies in small animals were already conducted in 1993 using a 2T small animal device [165]. However, the results to that time, due to limited hardware allowed only for qualitative results. With new technologies, especially dedicated high field scanners, animal models of different kidney disease were analyzed with respect to oxygenation. Oostendrop et al. investigated the change in oxygenation in an unilateral rat model of acute kidney injury [23]. Directly after IRI, the R2* increases, i.e. nearly doubles, but already after one hour after reperfusion it reaches nearly the baseline. However, compare to the contralateral kidney, the oxygenation is significantly reduced. Similar results were reported by Pohlmann et al. [40]. In their study, in addition to the IRI study also the change in oxygenation of the kidney during hypoxia / hyperoxia was investigated (see Fig. 9). Changing the oxygen concentration in the inhaled gas is reflected in the oxygenation of the kidney (change about 30%). In another study by Pohlmann et al. a combination of noninvasive MRI and invasive perfusion and oxygenation measurements using MRI compatible pO2 and flow probes was proposed [39].Using such setup, the correlation between pO2 and R2 * in living tissue can be measured. Ragnant et al. investigated oxygenation in a model of chronic renal artery stenosis (RAS) [104]. They measured BOLD MRI of the stenosed kidney before and several times after surgery to follow the change of R2 * over time. Due to the stenosis at post 4 days a steep drop of R2 * could be observed, thereafter, the R2 * slowly normalized to base line again.

BOLD MRI was also performed on whole body clinical scanners. Compared to the ultra-high field, today’s 1.5T scanners provide a homogenous B0 field that allows for quantitative oxygenation measurements. Several animal models and renal diseases were subject to research. Prasad et al. investigated the change in renal oxygenation under a drug challenge [8]. Three different drugs including one radiocontrast agent were injected to rats in different combinations. All three drugs and their combination let increase the R2 * of the renal medulla while R2 * of the renal cortex showed no or only small changes. The effect of diuretics to the renal oxygenation and its dynamic over time was reported by Kusakabe et al. [166] while Pedersen et al. investigated change in oxygenation in a rat RAS model [81]. Edlund et al. and Ries et al. investigated the relation of R2 * to renal function in diabetic rats [102,112]. Prasad et al. recently showed similar results for mice at a 3 T scanner [103]. Besides small animals, BOLD MRI was performed in larger animals like swine [111,158,167]. Diffusion MRI Diffusion weighted imaging of small rodents is mainly performed in the animal brain [168,169] but recently, a few studies applied DWI to the kidney in small animal models. Cheung et al. applied diffusion tensor (DTI) and diffusion weighted imaging to a rat model of renal ischemia reperfusion injury [109]. ADC and fractional anisotropy (FA) showed reduced values in the injured kidney compared to the control animals and also the IVIM parameters D* (ADC corrected for


perfusion effects) and perfusion fraction f showed a reduction compared to a healthy kidney. Yang et al. investigated changes in ADC in transplanted and normal kidneys [37] using a similar set up as described above. No changes between left and right kidney in ADC were found. Allografts exhibited decreased ADC values and isografts exhibited similar ADC values compared with native kidneys. Reducing the blood flow injecting angiotensin II, a significant reduction in ADC for renal cortex and medulla was observed. Similar to the other functional imaging techniques like BOLD or perfusion MRI, DWI was applied to renal animal models also at clinical whole body systems [110]. However, recent studies mainly focused on in small rodent models with either cerebral ischemia or subcutaneous and intrahepatic tumors [110]. Ries et al. used a 1.5 T whole body scanner to evaluate DWI in a rat model of diabetic nephropathy [112]. Their results show that in the diabetes rat the ADC is reduced compared to the control while a furosemide challenge did not show any changes of the ADC. Sodium imaging The sodium ion is a natural biomarker giving functional information about the renal tissue viability in renal diseases, or about the effect of drugs, e.g. diuresis. The very first functional parameter investigated in renal 23 Na MRI was suggested to be the corticomedullary gradient which has a linear profile in intact renal tissue. Maril et al. investigated this gradient in healthy, obstructed, and diuretic rat kidneys by means of the TSC as a further functional parameter [118]. The MR measurements were performed using a 3D gradient-echo sequence at a 4.7T small animal scanner and distinct changes were observed in the corticomedullary sodium gradient showing the feasibility and utility of 23 Na MRI under various physiologic and pathologic conditions. In a further study, Maril et al. investigated the corticomedullary sodium gradient before and after the administration of diuretic agents, e.g. furosemide and mannitol [117]. The relative sodium signal intensity change was monitored and 23 Na MRI was suggested as a tool for quantifying renal physiology despite methodological limitations. In a last rodent kidney study of the same research group 23 Na MRI was applied for the detection of acute tubular necrosis (ATN), which resulted in sodium signal reduction in cortico-outer medullary and the inner medulla to cortex under conditions of ATN compared to the control kidney [116]. In a later work about the same rat kidney model at 9.4T, Atthe et al. observed the alterations in renal sodium distribution in the rat kidney during ischemia and reperfusion injury using 23 Na MRI and a 3D gradient-echo sequence [119]. The timeresolved monitoring revealed that 23 Na MRI can distinguish between reversible and irreversible ischemic tissue damage inducing ATN by means of the TSC investigation in tissue. In another study, the temporal high resolution was used for monitoring renal sodium changes after furosemide application [170]. For the first time, Neuberger et al. performed 23 Na MRI

15

on a mouse kidney model at 17.6T – the highest field strength used for rodent renal 23 Na MRI to date [171]. By means of the density-weighted 3D CSI acquisition technique the TSC was monitored in diuretic mouse kidney and the results agreed well with rodent kidney studies published before. A number of 23 Na MRI studies in animals were performed with particular focus on the rodent brain models such as stroke 23 Na MRI where increased TSC in affected tissue was observed [121,172–175]. Others studied the corticomedullary RSC change in the rodent kidney after application of diuretic agents [117,118], or in the acute tubular necrosis (ATN) model for early monitoring [116,119]. For the future vision of translational MRI, it is essential to realize the pre-clinical experiments in clinical MRI systems for easier method transfer between small animal models and human pathologies. Given that the animal models are also needed in the development of pharmaceuticals, 23 Na MRI represents an essential functional MRI tool for pre-clinical applications, e.g., in investigating the effects of drugs. However, pre-clinical MR experiments have overall higher demand in spatial resolution due to the smaller size of the target object. Furthermore, the translational 23Na MRI for the small animal experiments on a clinical MRI system at 3T is a challenging task firstly, due to the field strength penalty, secondly, due to limitations about animal handling instruments, e.g., narcosis, respiration and temperature control units, and lastly, due to the lack of mechanical stability in the magnet bore, e.g., bed positioning system. Therefore, further improvements are necessary in imaging techniques and in hardware development especially regarding RF resonators. Although differences in hardware exist, image resolutions achieved so far were rather similar (see Table 7). The benefit of the high field strength and therefore, increase in SNR was mainly invested into shorter acquisition times (about 5 minutes, see Neuberger et al. [171]). Sodium MRI of the rodent kidney was realized for the first time on a 3T whole body MRI scanner by means of a dual RF resonator system in order to investigate the ATN model in the rodent kidney [176]. Such a modular RF system allows imaging of different organs by simple modification of the surface coil geometry, which may also benefit the quantification of Na+ concentration in human [177]. Furthermore, the dual resonator allows for increased 23 Na-MR signal sensitivity due to the localized signal detection using RO surface coils. A further initial pre-clinical experiment about the diuresis model of the rodent kidney was performed at 3T using a double-tuned TXRX coil [178]. In order to implement the in-house developed multi-tuned TXRX RF resonator, the 3T clinical MRI system’s standard RF hardware had to be equipped with an additional coil interface including a transceiver switch. The ultra-short TE and the sensitivityoptimized surface resonator enabled fast acquisition of 23 Na kidney images within 10 min of scanning time, which allowed for studying the fast 23 Na change after furosemide injection. The achieved (isotropic) image resolution of 1.5 mm3 is

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Table 7 Sodium MRI of small animals at different field strengths. MRI system

Resolution (mm)

Animal model

Reference

17.6T / dsas 9.4T / dsas

1.0 x 1.0 x 2.5 1.0 x 1.0 x 4.0 0.9 x 0.9 x 3.8 0.9 x 0.9 x 3.0 1.5 x 1.5 x 1.5

Mice Rat Rat Rat Rat

Neuberger et al. [171] Kalayciyan et al. [11,170] Atthe et al. [119] Maril et al, [116–118] Kalayciyan et al. [176,178]

4.7T / dsas 3T / wbs


comparable to similar studies performed at ultra-high field systems (see Table 7). The 23 Na signal decrease in the medulla and the increase in the cortex were measured and the herein reported values matched those in the studies performed at the 9.4T pre-clinical MRI system very well [179]. In conclusion, the pre-clinical renal MR investigation in a whole-body human scanner was demonstrated to be feasible despite the field strength penalty. Further measurements are needed to determine the accuracy of 23 Na-MR measurements in the renal tissue, and to establish this powerful imaging technique for absolute quantification of the sodium concentration.

Discussion and conclusion Pre-clinical functional MRI of the kidney is an emerging field of research. Various kidney diseases are targeted by MRI methods using dedicated animal models (see Table 5). All state-of-the-art functional imaging techniques reported for kidney imaging (perfusion, diffusion, and oxygenation) are nowadays implemented for small animals on dedicated ultrahigh field scanners and also on whole body clinical systems. Furthermore, new emerging X-nuclei imaging techniques like sodium MRI have progressed in functional imaging. Considering the close link of the tissue sodium concentration to the tissue viability [14], the 23 Na MRI technique represents a powerful tool to understand the renal physiology and pathologies, as well as the pharmacological effect of drugs. The recently introduced quantification techniques allow for measuring the RSC in in-vivo rodent kidneys with high spatiotemporal resolution, and accuracy at 9.4T. The dual RF resonator system with an optimized receiver coil and the shortTE sequences are main factors in maximizing SNR in 23 Na MRI. Eventually, the novel RF resonator techniques at 9.4T were transferred to a 3T clinical MRI system allowing for experiments in the rodent kidney and for quantitative measurement of the sodium concentration in humans including spine [178], breast [180], and abdominal [178] imaging. At ultra-high field animal scanners, the visualization of the individual glomeruli became also possible. Though, this is yet mostly performed ex-vivo current research directed towards in vivo glomeruli counting. Magnetic resonance imaging of small animals is a challenging task and poses several limitations. Compared to the human, smaller voxel sizes are needed for the imaging of the

organs like the rodent kidney and thus, a high image resolution for morphological and functional imaging is required. A higher image resolution is usually paid by either a loss in SNR or temporal resolution, e.g. in perfusion imaging (see Table 2). Furthermore, the animal’s physiology like heart and breathing rate is much faster and needs gating or triggering strategies. MRI at dedicated ultra-high field scanners, however, provides intrinsically higher SNR due to the higher B0 . MRI techniques like ASL, sodium imaging, or glomeruli counting directly benefit of higher SNR. In addition to their field strength, the small animal scanners are equipped by about 20 times stronger and about 30 times faster gradient systems, which further improves the imaging quality compared to a whole body human scanner. Further improvement of the SNR was achieved by the implementation of cryogenically cooled RF resonators, which is a worthwhile technique at lower frequencies, especially for imaging of small regions of interest, e.g. the rodent kidney or brain. Nevertheless, a number of studies reporting on the research of kidney diseases utilize whole body scanners with dedicated RF coils. Also, on these systems, adequate image resolution could be reached [181] and the full spectrum of functional imaging techniques were reported. Certainly, translating some techniques from the ultra-high field system to the clinical whole body scanner like counting glomeruli is challenging. A possible benefit of pre-clinical small rodent MRI at whole body scanners is that the employed sequences are the same as for human applications. In principle, if a certain technique shows promising results in small animals, it should be applicable also in humans. Certainly, one has to keep in mind that there are physiological differences between rodents and humans and that experiments like drug testing underlie certain ethical restrictions. In conclusion, pre-clinical functional imaging of the kidney offers a large variety of imaging techniques that allow access to essential functional parameters non-invasively. Therefore, a multi-parametric assessment of tissue and organ function is possible which might provide even more comprehensive insights into kidney diseases. Nearly all of these imaging approaches are available on today’s imaging hardware, either dedicated small animal devices or whole body clinical scanners. This allows for performing studies using the optimal hardware demands but also a translation of knowledge from basic research towards a clinical application.


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