Fig. 4 Changes of invasively measured parameters of renal hemodynamics and oxygenation during ramp-wise reduction in renal perfusion pressure followed by ramp-wise pressure restoration in anesthetized rats. Conductance values (the reciprocal of vascular resistance) were calculated by dividing the respective perfusion values by renal perfusion pressure, in order to distinguish flow changes that result from passive circular distension/compression of vessels from those actively exerted by vascular smooth muscles. Three groups were studied: a healthy control group, a group in which a diabetes mellitus type 1-like disorder (DM) was induced by streptozotocin 4 weeks before obtaining the data, and a third group in which DM was induced and the antidiabetic liraglutide administered for 3 weeks (DM + LIRA). Values (mean SEM) are given as relative changes from baseline [32]

[22]. While the amount of deoxyHb in arterial blood is barely changed, the increase in blood pO2 in intrarenal veins that results from the higher arteriovenous pO2 difference decreases venous deoxyHb. While primarily altering blood oxygenation, the hyperoxic stimulus has also secondary effects: it results in vasoconstriction, preferentially in nonrenal vascular beds, which leads to an increase in arterial pressure [22, 31, 39].

The primary effect of reducing the FiO2 (typically to either 8% or 10%, with durations of 3–12 min, in rat studies) is a decrease in oxygenation of arterial blood (hypoxemia) with the consequent reduction in renal O2 supply. With ongoing O2 consumption, this does per se result in a decrease in renal tissue pO2. Yet renal O2 supply is further diminished by hypoxia-induced extrarenal vasodilation that results in a drop in arterial pressure with ensuing decrease in RBF [22, 31, 39, 59]. Whether this is aggravated of alleviated by constriction or dilation, respectively, of the renal vasculature depends on the degree of hypoxia: in anesthetized rats, FiO2 of 8% results in renal vasoconstriction while 10% results in vasodilation [22, 31, 39]. The combined effect of hypoxemia and reduced RBF on renal O2 supply leads to a major mismatch with O2 consumption, that massively reduces tissue pO2 as well as T2\* (see Fig. 1) [22, 39]. A further secondary effect of arterial hypoxemia is increased ventilation triggered by arterial chemoreceptors. The ensuing decrease in arterial pCO2 shifts the oxyHb dissociation curve to the left, that is, O2 is hindered from being released by Hb, which further aggravates the tissue hypoxia [61, 62].

With the hypercapnic stimulus (increasing FiCO2 to 5%) the opposite effect is achieved, namely a rightward shift of the oxyHb dissociation curve. This would per se result in a decrease in StO2 and an increase in blood and tissue pO2. However, while the increase in tissue pO2 is substantial, the StO2 decrease is meagre [39]. The major reason that StO2 does not decrease much is that increased pCO2 of arterial blood is a very strong stimulus for ventilation, again mediated by arterial chemoreceptors [61, 62].

Hyperoxic, hypoxic, and hypercapnic tests have been used in a multitude of preclinical in vivo studies. With regard to the kidney this includes but is not limited to studies on the control of renal hemodynamics and oxygenation in healthy animals and models of kidney diseases [26, 27, 31, 53–55, 59], experiments that aimed at calibration of T2\* by means of the MR-PHYSIOL setup (see Fig. 5), [22] studies on the T2\* effect of an X-ray contrast medium, [36] assessment of possible unwarranted effects of the USPIO ferumoxytol, [38] and experiments en route to the PHYSIOL-NIRS setup [39].

Fig. 5 Time courses of selected invasively measured data and MR parameters acquired simultaneously throughout baseline, a period of hypoxia (FiO2 ¼ 8%), and recovery in anesthetized rats by means of a dedicated MR-PHYSIOL hybrid setup (for details see text). Data (mean SEM) are relative changes from baseline, redrawn from Ref. 22

Given the broad therapeutic use of pure oxygen, short-term tests with 100% FiO2 should pose no problem for studies in humans, whereas hypoxic challenges are precluded in humans, for obvious ethical reasons. Hypercapnia has been used for decades in humans, in particular, for the study of cerebrovascular reactivity, and should thus be employed in preclinical and clinical studies on renal hemodynamics and oxygenation as well [63, 64].

3.4 Administration of Drugs and Endogenous Vasoactive Substances Furosemide is the "classic" loop diuretic: its major action is the inhibition of the sodium–potassium–two-chloride cotransporter in the apical membrane of tubular epithelial cells of the thick ascending limb of Henle's loop. The primary effect is an increase in urine flow rate and in urinary sodium and potassium excretion. As less tubular resorption necessitates less renal O2 consumption, administration of furosemide leads to an increase in renal tissue pO2 [65, 66]. In accordance, increases in renal T2\* (or decreases in its reciprocal value R2\*) have been observed in a multitude of preclinical as well as clinical studies (see Fig. 6) [37, 66–69]. It must be noted, however, that the increase in T2\* upon furosemide does not solely rely on improved oxygenation. First, the increase in tubular fluid downstream of the thick ascending limb will increase the transmural pressure gradient, thereby compressing intrarenal

Fig. 6 Effect of injections of furosemide (5 mg/kg body mass), hydralazine (5 mg/kg), angiotensin II (0.5 μg/ min/kg), and saline (repeatability) on renal BOLD as recorded by means of a 1.5 T clinical MR scanner (Magnetom Avanto, Siemens Healthcare), using a multiple gradient echo sequence (TR ¼ 300 ms, TE ¼ 5, 10, 20, 30, and 40 ms, voxel size 0.6 0.6 mm in-plane and 5 mm slice thickness) in rats in vivo. Data are mean SEM of median ROI values [37]

vessels with the ensuing decrease in the amount of deoxyHb per tissue volume. Second, furosemide inhibits the TGF, thereby compromising renal autoregulation with the possible consequence of an increase in RBF [34, 35, 50].

The use of the furosemide test in both preclinical and clinical MR studies is nowadays as widespread that it is almost regarded as a gold standard. However, whether it fulfils all expectations regarding its use as a diagnostic tool in patients suffering from various kidney diseases, remains to be seen [68].

Administration of furosemide is a reversible intervention insofar as its direct effects vanish with the excretion of the drug. However, it leaves the organism with deficits in water, sodium, and potassium. These should ideally be replenished—be it per os or by means of infusions of a balanced electrolyte solution.

Bolus injections of adenosine cause a rapid drop in arterial pressure due to its vasodilatory effect on nonrenal resistance vessels. In the renal cortical vascular bed, it exerts vasoconstriction [39, 49]. The consequence of these two effects is a substantial decrease in RBF followed by a smaller decrease in cortical tissue pO2. All these effects vanish rapidly, lasting less than a minute for the hemodynamics and less than 2 min for the cortical pO2 in rats [39].

Whereas the role of adenosine in various renal control mechanisms including the TGF as well as the potentially beneficial effect of adenosine receptor antagonists for prevention of X-ray contrast media-induced AKI have been intensively studied, [49, 70] the adenosine test has seldom been used to study renal hemodynamics and oxygenation. This may appear surprising, as injections of adenosine—be it intravenously or into coronary arteries—in patients suffering from coronary disease is quite established [71, 72]. While the risk for a decrease in renal cortical pO2 in patients must not be ignored—even if it is lasting less than 2 min, the test should at least find wider use in preclinical studies.

### Acknowledgments

This work was funded, in part (Kathleen Cantow and Erdmann Seeliger) by the German Research Foundation (Gefoerdert durch die Deutsche Forschungsgemeinschaft (DFG), Project number/ Projektnummer 394046635, SFB 1365, RENOPROTECTION). The authors wish to thank Ariane Anger and Andrea Gerhardt for expert technical assistance.

This chapter 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.

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clinical 1.5T platform. BMC Nephrol 17 (1):142. https://doi.org/10.1186/s12882- 016-0356-x

fraction variations En route to renal magnetic resonance oximetry. Tomography 3 (4):188–200. https://doi.org/10.18383/j. tom.2017.00012

(12):3271–3275. https://doi.org/10.1161/ 01.cir.94.12.3271

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.

### Preparation of Ex Vivo Rodent Phantoms for Developing, Testing, and Training MR Imaging of the Kidney and Other Organs

### Jason M. Millward, Joa˜ o S. Periquito, Paula Ramos Delgado, Christian Prinz, Thoralf Niendorf, and Sonia Waiczies

### Abstract

Here we describe a simple and inexpensive protocol for preparing ex vivo rodent phantoms for use in MR imaging studies. The experimental animals are perfused and fixed with formaldehyde, and then wrapped with gauze and sealed with liquid latex. This yields a phantom that preserves all organs in situ, and which avoids the need to keep fixed animals and organs in containers that have dimensions very different from living animals. This is especially important for loading in MR detectors, and specifically the RF coils, they are usually used with. The phantom can be safely stored and conveniently reused, and can provide MR scientists with a realistic phantom with which to establish protocols in preparation for preclinical in vivo studies—for renal, brain, and body imaging. The phantom also serves as an ideal teaching tool, for trainees learning how to perform preclinical MRI investigations of the kidney and other target organs, while avoiding the need for handling living animals, and reducing the total number of animals required.

This protocol chapter is part of the PARENCHIMA initiative "MRI Biomarkers for CKD" (CA16103), a community-driven Action of 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.

Key words Ex vivo methods, Rat, Mouse, Phantom, Magnetic resonance imaging (MRI), Kidney

### 1 Introduction

Prior to starting any preclinical in vivo MRI study, it is essential to invest time to test novel imaging techniques, optimize MR pulse sequence parameters and develop a workflow tailored to address the specific scientific questions at hand [1]. Much of this work can be done without using living animals, by constructing synthetic phantoms using materials that have similar MR properties as the tissue of interest (e.g., agarose). Ultimately, a point is reached when the limitations of synthetic phantoms impede progress, and phantoms that are more biologically realistic are needed. In some cases this can be accomplished using ex vivo tissue samples removed from experimental animals postmortem. The tissues are chemically fixed (usually with formaldehyde), and samples embedded in tubes or containers for imaging. However, this removes the organs from their surrounding environment. In the case of renal MR, it is advantageous to image the kidney in situ, within its biological context, including the surrounding organs and perirenal fat. The typical solution for this problem is to use ex vivo phantoms, which consist of the intact experimental animal (usually rats or mice) placed inside a sealed container along with the fixation solution. This solution is suboptimal. The containers used generally have a cylindrical or rectangular shape, which is substantially different from the shape of an animal. This can be problematic given the space constraints of the MR hardware (e.g., RF coil diameter, size of the animal holding bed). In some cases the container size may mean that a larger field of view needs to be acquired than would otherwise be necessary for a living animal. It may also be difficult, or impossible, to use surface RF coils to scan samples embedded in a container. For example, when scanning the brain or the kidney, positioning the sample such that the target organ is as close as possible to the RF coil could be problematic if the container is in the way. Furthermore, the additional mass of the fixation liquid in a container with a rat or mouse body can substantially change the loading in the coil, which could cause problems with matching and tuning. The shape of the container can also cause other problems; the curved and conical shapes of typical laboratory plastic tubes can interfere with shimming the main magnetic field B0.

Here we describe a simple and inexpensive method to prepare ex vivo phantoms using mice or rats that are perfused and chemically fixed, covered in gauze and sealed with latex. This solution results in a phantom that preserves the biological context of the organs of interest, and which is much more similar to the original animal in terms of dimensions and loading. It avoids the use of containers of toxic formaldehyde solution. The phantoms can be readily used by imaging scientists and scientists from related fields who are not trained in the handling of living experimental animals. This makes the phantom an ideal teaching tool for educating trainees learning how to perform preclinical studies, while also avoiding the confounding effects of motion artifacts. Finally, the phantoms can be easily stored and reused for extended periods of time. This will substantially reduce the total number of animals required for developing new MRI methods, and will facilitate more efficient and effective in vivo studies, thus consistent with the 3R principles of animal research [2, 3].

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 Materials

2.2 Perfusion and Dissection All steps of the protocol should be conducted in accordance with procedures approved by the local Animal Welfare Department and authorities and conform to guidelines to minimize discomfort to animals (86/609/EEC).

> The perfusion should be done in a lab environment equipped for dealing with hazardous fumes, using a chemical fume hood (alternatively fume extraction dome hoods) due to the respiratory hazards of formaldehyde. The lab should also have appropriate mechanisms in place for disposing of formaldehyde liquid waste, solid materials contaminated with formaldehyde (e.g., paper tissues), as well as animal carcasses (e.g., freezer for temporary storage; capacity for incineration).

### 2.3 Latex Covering 1. Liquid latex (e.g., 0.3% ammonia prevulcanized, low-viscosity—various suppliers).

### 3 Methods

3.1 Perfusion and Fixation

used by the previous needle (look for a small dark spot). \*RAT—attach the 60 ml syringe with PFA directly to the tube connected to the butterfly needle. At this point the exsanguination and perfusion is complete.

	- 1. Remove the skin around the joint of the forelimb and the shoulder (scapular joint) with surgical scissors. Using a scalpel, dissect the musculature around the scapula (Fig. 1a). Using the fingers, feel for the location of the joint (Fig. 1b). When the joint is exposed, use rongeurs to separate the scapula, and remove the entire forelimb (Fig. 1c). Repeat the process on the other side.
	- 2. The procedure for the hind limbs is similar. Remove the skin around the joint of the femur and hip. Using a scalpel, dissect the musculature around the femur, and feel for the location of the ball and socket hip joint (where the femur fits in the acetabulum of the pelvis). When the joint is exposed, cut the tendons with a scalpel, or cut the bone directly using rongeurs (Fig. 1a). Repeat on the other side.
	- 3. Use a scalpel to dissect the musculature of the tail at the proximal end. Expose the coccyx, and separate the tail using rongeurs. The disarticulation is now complete.
	- 4. Proceed with removing the skin of the remaining body. Using scissors, cut through the skin at the level of the neck—the skin of the head can remain in place. Remove the skin and any extraneous tissue from the exposed thoracic cavity and the dorsal thoracic region. Use the fingers to separate the skin and expand the subcutaneous space in the dorsal lumbar region. Make a midline cut of the skin with scissors, and use a scalpel to separate the rest of the skin from the underlying

Fig. 1 Schematic for disarticulation. The location of the cut-points for removing the forelimb and hind limb is depicted by the gray lines (a). After cutting the overlying musculature, the joint of the scapula can be visualized (b). The forelimb is reflected back, and the joint cut using rongeurs (c)

tissue. Continue removing skin from the ventral lumbar region, taking care not to pierce the peritoneum.

5. The skin of the head can remain in place, since the head will not be covered with gauze. Remove the ears (exposing the auditory cavity), trim the whiskers and remove extraneous tissue from the lower mandible region.

#### 3.3 Coating with Gauze and Latex In the next steps the fixed and disarticulated body will be wrapped in gauze and covered in liquid latex.

1. Starting from the neck, tightly wrap the thoracic region with gauze. It may be necessary to fill part of the thoracic cavity that was opened during the perfusion step with gauze, in order to eliminate empty space and have a securely wrapped phantom. Continue wrapping the lumbar/sacral region with gauze; a

Fig. 2 Workflow for preparation of the phantom. The rat body is wrapped tightly with gauze, which is then fixed securely to the underlying musculature using surgical sutures (a). The entire body is covered with gauze, except for the head, which remains unwrapped (scale bar ¼ 22 cm) (b). Liquid latex is applied directly to the surface of the phantom using a paintbrush (c, d). The phantom is place on an improvised holder, here a laboratory tube rack with pipettes. A hair dryer is used to accelerate the drying process (e). An example of a completed, labeled mouse phantom (f)

> small extension of gauze can remain at the posterior end to serve as a grip when applying the latex.

covered, although avoid excessive latex. It is preferable to apply the latex with a brush, rather than simply submerging the phantom in latex.

### 4 Example Images

In Fig. 3 we show representative images of MRI scans of the ex vivo phantoms. These scans were performed on a Bruker Biospec 9.4 Tesla USR94/20 animal MRI scanner (Bruker Biospin, Ettlingen, Germany).

4.1 Rat MRI scans targeting the rat kidney were acquired in the coronal (Fig. 3a, b) and sagittal (Fig. 3d, e) planes. The identical MRI hardware setup (including the RF coils) and pulse sequence parameters were used to acquire the images of the ex vivo phantom (Fig. 3a, d) and an in vivo rat of the same strain and approximate mass as the animal used to prepare the phantom (Fig. 3b, e). The in vivo scans were done as described in the chapter by Pohlmann A et al. "Experimental Protocol for MRI Mapping of the Blood Oxygenation-Sensitive Parameters T2 \* and T2 in the Kidney." The ex vivo and in vivo scans show a comparable level of detail, and the kidneys (denoted by "K") can be seen in relation to the surrounding tissues. MRI sequence parameters are listed in Note 4.

Fig. 3 Representative ex vivo and in vivo MR images. MRI scans targeting the rat kidney were acquired in the coronal (a, b) and sagittal (d, e) planes. Identical MRI hardware and pulse sequence parameters were used to acquire the images of the ex vivo phantom (a, d) and in vivo images of a rat of the same strain and approximate mass as the animal used to prepare the phantom (b, e). The kidneys are denoted by "K." MRI scans targeting the mouse brain were acquired in the coronal plane in the ex vivo phantom (c, f). Images show the brain in relation to surrounding structures such as the skull, and the eyes (denoted by "e")

4.2 Mouse MRI scans targeting the mouse brain were acquired in the coronal plane in the ex vivo phantom (Fig. 3c, f). A mid-brain slice shows good contrast between gray matter and white matter structures, and the skull can be seen on the exterior of the brain (Fig. 3c). A lower brain slice (from a more inferior level) shows the brain in relation to surrounding tissue structures, for example the eyes (denoted by "E") (Fig. 3f). MRI sequence parameters are listed in Note 5.

### 5 Notes

### Acknowledgments

This work was funded in part (Thoralf Niendorf, Sonia Waiczies, 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

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.

# Part III

### Basic Concepts of Measurement Techniques

# Chapter 6

### Quantitative Assessment of Renal Perfusion and Oxygenation by Invasive Probes: Basic Concepts

### Kathleen Cantow, Roger G. Evans, Dirk Grosenick, Thomas Gladytz, Thoralf Niendorf, Bert Flemming, and Erdmann Seeliger

### Abstract

Renal tissue hypoperfusion and hypoxia are early key elements in the pathophysiology of acute kidney injury of various origins, and may also promote progression from acute injury to chronic kidney disease. Here we describe basic principles of methodology to quantify renal hemodynamics and tissue oxygenation by means of invasive probes in experimental animals. Advantages and disadvantages of the various methods are discussed in the context of the heterogeneity of renal tissue perfusion and oxygenation.

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 a separate chapter describing the experimental procedure and data analysis.

Key words Renal hemodynamics and oxygenation, In vivo methods, Quantitative invasive probes

### 1 Introduction

Kidney diseases are a global health burden with steadily increasing prevalence and incidence [1–5]. Animal studies have provided evidence that acute kidney injury (AKI) of various origins shares a common link in the pathophysiological chain of events, ultimately leading to AKI, as well as to progression from AKI to chronic kidney disease (CKD): imbalance between renal oxygen delivery and oxygen demand [3, 6–13]. Renal tissue hypoperfusion and hypoxia have also been suggested to play a pivotal role in the pathophysiology of other kidney diseases including diabetic kidney disease [14–16].

The majority of the preclinical studies that generated this concept utilized a set of invasive probes to measure renal hemodynamics and oxygenation in anesthetized animals [12, 14, 17–20]. These methods have also been used (1) to study mechanisms of

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\_6, © The Author(s) 2021

physiological control of renal hemodynamics and oxygenation and (2) the effects of various substances on this control in healthy animals, (3) to test several putative preventive or therapeutic approaches for AKI and CKD in animal models, and (4) for the purpose of calibrating functional magnetic resonance imaging (MRI) including renal blood oxygenation level-dependent MRI (BOLD-MRI) that makes use of the quantitative MR parameter T2\*, which is a surrogate for blood oxygenation [18–28].

The invasive probes typically include (1) a perivascular flow probe for measurement of total renal blood flow, (2) laser-Doppleroptodes for assessment of local tissue perfusion, (3) Clark-type electrodes or fluorescence-quenching optodes for measurements of local tissue partial pressure of oxygen (pO2), and (4) devices for invasive measurement of arterial blood pressure. These methods are considered the gold standard for the study of renal hemodynamics and oxygenation, because the methods, with the exception of the laser-Doppler fluxmetry, provide calibrated quantitative data [25, 29, 30]. Step-by-step protocols for application of these techniques in two setups, one for stand-alone experiments in laboratories for integrative physiology, and the other for experiments within small-animal MR scanners by use of a hybrid setup, are provided in the chapter by Cantow K et al. "Monitoring Renal Hemodynamics and Oxygenation by Invasive Probes: Experimental Protocol".

As with all established modalities available in today's experimental and translational research, these techniques have shortcomings and methodological limitations. In particular, their invasiveness, for obvious ethical reasons, precludes their use in humans. The invasiveness is, furthermore, a considerable hurdle for implementation of these methods in long-term and longitudinal studies: although invasive probes have been used in conscious, chronically instrumented animals [22, 24, 31–37], the vast majority of studies has been performed in "terminal experiments" in anesthetized animals. This is a major limitation of all invasive techniques because the measurements are by necessity potentially confounded by the effects of anesthesia (see the chapter by Kaucsar T et al. "Preparation and Monitoring of Small Animals in Renal MRI").

The measurement of intrarenal tissue perfusion and tissue oxygenation is quite challenging, because of their vast spatial heterogeneity both among the renal layers (cortex, outer medulla, and inner medulla) and within the layers. In this chapter, the causes and the degree for this spatial variability are outlined first, before the basic principles of methodology to quantify renal hemodynamics and oxygenation by means of invasive probes in experimental animals are described together with a discussion of the pros and cons of the various methods.

This introduction chapter is complemented by a separate chapter describing the experimental procedure, which is 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 Heterogeneity of Intrarenal Perfusion and Oxygenation—Causes and Degree of Variability

Intrarenal local tissue pO2 is determined by the balance between local O2 supply and local O2 consumption. Additionally, determinants of local O2 diffusion play a role. Local O2 supply depends on local blood perfusion, local vessel density, and local blood pO2 and blood O2 content. Local O2 consumption is foremost determined by local metabolic demand. Blood perfusion within the kidney is much more heterogeneous than within most other organs, due to its particular vascular architecture [8, 38–40].

From the branches of the renal artery that enter the kidney at its hilus, interlobar arteries arise, which align with the medullary–cortical axis. From interlobar arteries emerge arcuate arteries that run along the boundary between the cortex and the medulla, that is, perpendicular to the medullary–cortical axis. From the latter, interlobular arteries branch off at regular intervals at angles of about 90 degrees, so that they all enter the renal cortex. For this reason, virtually all the blood flowing into the kidney perfuses the cortex first [8, 25, 38, 41–43]. The interlobular arteries are in close proximity to interlobular veins, together forming what might be called "vascular bundles" that run in parallel at regular distances from the corticomedullary boundary toward the capsule. Within the cortex, interlobular arteries branch off into afferent (preglomerular) arterioles, from which glomerular capillaries arise that, in turn, rejoin into efferent (postglomerular) arterioles. The efferent arterioles from the subcapsular and mid-cortical glomeruli branch off into networks of peritubular capillaries. The efferent arterioles of the deep cortical (juxtamedullary) glomeruli, however, branch off at about the corticomedullary boundary to form capillary-like vessels, the descending vasa recta [38–40]. These very straight and exceptionally long vessels descend toward the papilla. The descending vasa recta provide virtually all blood that perfuses the medulla. As a consequence, only 10–15% of blood that previously passes through the cortex, will reach the medulla, and the vascular density is much lower in the medulla than the cortex [8, 25, 38, 41–43]. The descending vasa recta are in close proximity with the ascending vasa recta, thus forming distinct "vascular bundles" that run in parallel at regular distances through the medulla [38–40].

These particularities of the intrarenal vascular architecture have a number of salient consequences with regard to local blood

Fig. 1 Scheme depicting the range of intrarenal tissue partial pressure of oxygen (pO2). The red curve depicts values at the greatest distance from two neighboring "vascular bundles," the blue one that directly adjacent to "vascular bundles" (for details see text); OS (outer stripe), IS (inner stripe). Original data were obtained by measurements in in anesthetized dogs and rats [43–45]

perfusion and thus the local O2 supply. First, the striking difference in the quantity of blood perfusion between the cortex and the medulla, together with the difference in their vascular densities, is one of the major reasons for the low tissue pO2 in the medulla. Second, the Fa˚hræus–Lindqvist effect that lowers the hematocrit and thereby the O2 content of blood in the vasa recta that supply the medulla, contributes to low medullary O2 delivery [39, 40]. Third, diffusive shunting of O2 between the arcuate and interlobular arteries and their respective veins in the cortex as well as between the descending and ascending vasa recta in the medulla, owing to the close proximities of the respective vessels, are a further major reason behind the "physiological hypoxia" of the medulla. Thus, blood pO2 is higher in the venous vessels than in the capillaries these venous vessels drain. Fourth, due to the particular vascular architecture with distinct "vascular bundles" that comprise arterial vessels with high pO2 and venous vessels with relatively high pO2 (due mainly to the low fractional O2 extraction of the renal tissue), there are huge differences in local pO2 even within a given renal layer under physiological conditions [25, 41, 43–47]. Fifth, plasma skimming at intrarenal vessel branches results in different hematocrit and therefore O2 content of blood perfusing the daughter vessels, which also contributes to the local pO2 differences [39, 48].

The schematic shown in Fig. 1 depicts the range of local pO2 values as measured by invasive Clark-type electrodes in the kidneys of rats and dogs under physiological conditions in vivo [43– 45]. These measurements reveal that local tissue pO2 levels in the cortex vary dramatically, foremost according to the distance to the "vascular bundles" of interlobular arteries and veins: pO2 values range from 10 mmHg at the greatest distance from two neighboring "bundles," to up to 70 mmHg around the "bundles." In addition, local cortical tissue pO2 varies with the distance to the arcuate arteries located at the corticomedullary boundary. Also, the range of tissue pO2 variability in the outer medulla is huge (see Fig. 1): whereas pO2 is less than 10 mmHg in locations at maximum distance from the "vascular bundles" of descending and ascending vasa recta, pO2 is up to 75 mmHg in close proximity to these "bundles." In the inner medulla, tissue pO2 is rather homogenous: in general, it usually amounts to less than 15 mmHg.

The huge differences in tissue pO2 among the renal layers not only reflect the differences in O2 supply but also the different O2 consumption of the distinct sections of the tubules located in the respective renal layers. Tubular resorption relies on active transport processes, which account for about 85% of the kidney's energy expenditure and, therefore, its O2 consumption [8, 49]. The proximal tubule and the distal convoluted tubule, which constitute the majority of tubular segments within the cortex, consume much O2 because a huge amount of osmolytes is resorbed by active transport in these segments. On the other hand, tubular sections such as the thin descending and the thin ascending limbs of Henle's loop and the medullary collecting duct, which are the most abundant tubular sections within the medulla, feature very low to nonexisting active transport. Yet the thick ascending limbs of Henle's loop, which are located in the outer medulla, consume much O2, because of their active resorption of osmolytes [8, 25, 41, 43–47, 50]. The high O2 demands of the thick ascending limbs in relation to the rather low O2 supply of the outer medulla, consequent to its particular vascular architecture, exposes this renal layer to the greatest risk for hypoxic tissue damage [6–14, 43, 51, 52].

### 3 Methods: Methodologic Principles, Advantages, and Disadvantages

Techniques for invasive measurements of renal hemodynamics and oxygenation comprise readily (commercially) available methods and devices that are employed in many laboratories of integrative kidney physiology worldwide, but also techniques that are established only in one or a few laboratories. The widely used methods are described first. They include (1) perivascular flow probes for measurement of total renal blood flow, (2) laser-Doppler-optodes for assessment of local tissue perfusion, (3) Clark-type electrodes for measurements of local tissue pO2, (4) fluorescence-quenching optodes for the same purpose, (5) devices for invasive measurement of arterial blood pressure, and (6) pimonidazole adduct immunohistochemistry for assessment of renal tissue hypoxia.

The pimonidazole technique is included here because of its widespread usage, although it is an exceptional case for two reasons. First, in contrast to the other techniques, this approach does not allow investigators to monitor a given parameter over time, but only provides a snapshot: Pimonidazole is injected at a single time point in vivo, and pimonidazole adducts formed in severely hypoxic tissue are visualized in ex vivo samples [53–55]. Second, although the technique has been used for two decades, it was subjected to close scrutiny only recently [55]. Upon this careful investigation major pitfalls were identified and detailed—which, at least in part, cast doubts on the reliability of previously published results—and methods to address and overcome these issues have been proposed [55].

In the final part of this chapter, a few selected examples of methods are discussed that are not (yet) generally available but presently established in only one or a few laboratories.

3.1 Transonic Flow Probes Today's gold standard for whole kidney blood flow measurements is a flow probe attached to the renal artery or to the renal vein that makes use of ultrasound transit time differences. This method was established in the late 1970s and provides continuous absolute measurements of blood flow (in mL per min) independent of the actual flow velocity profile, the actual vessel diameter, and the actual hematocrit [56]. The commercially available device (Transonic™, Transonic Systems Inc.) comprises an electronic flow detection unit (flowmeter) and perivascular probes [57, 58]. The flow probe consists of two ultrasonic transducers and a fixed acoustic reflector (see Fig. 2a). The flow probe is positioned on the blood vessel so that the vessel lies within the acoustic sensing window defined by

Fig. 2 Photographs of typical probes for invasive measurements. (a) Perivascular Transonic™ flow probe for measurement of total renal blood flow. (b) Combined fluorescence-quenching-laser-Doppler probe (OxyLite/OxyFlo™) for simultaneous measurements of local tissue pO2 and assessment of local tissue perfusion in the same circumscribed area. (c) Clark-type electrode for measurement of local tissue pO2

the transducers and the metallic reflector. Probes are available in various sizes to fit the dimensions of the blood vessel of interest.

The flowmeter directs the flow probe through measurement cycles [57, 58]. First, a wide beam of ultrasound is emitted by the downstream transducer. It intersects the vessel in the upstream direction, is reflected by the acoustic reflector, then intersects the vessel again before being received by the upstream transducer. The flowmeter records the transit time. Then, the transit–receive sequence is repeated, but with the transmitting and receiving functions of the transducers reversed, that is, the vessel is bisected by the beam in the downstream direction. The transit time during the upstream cycle is higher than during the downstream cycle, because the ultrasound beam travels against the blood flow in the first, and with the blood flow in the latter cycle. The difference in the transit times is directly proportional to blood volume flow. Due to the wide ultrasonic beams, all flowing blood is accounted for; thus, the transit time is sampled at all points across the vessel diameter, and the volume flow measurement is independent of the flow velocity profile [57, 58].

An accurate measure of volumetric flow is obtained by the perivascular probes even when the vessel is smaller than the acoustic window, because the parts of the ultrasonic beam that do not intersect the vessel do not contribute to the signal [57, 58]. However, air effectively blocks ultrasound transmission, and even small air bubbles compromise measurement accuracy. Therefore, all spaces between the vessel and probe must be filled with a suitable acoustic coupling agent. In case of "terminal experiments" in anesthetized animals, where the abdomen is opened, this can be achieved by simply filling the abdominal cavity with isotonic saline (at 37 C) [27]. Acoustic coupling gels can also be used and are available from the supplier of the probes. Fatty tissue also affects the ultrasonic beam; thus, the vessel of interest must be thoroughly prepared by gently removing fatty tissue [27].

Measurements of total renal blood flow by means of Transonic™ devices have also been performed in chronically instrumented animals [22, 31, 33, 58, 59]. Here, the perivascular probe becomes encapsulated by tissue, which stabilizes its position. En route to calibration of functional MRI for the study of renal hemodynamics and oxygenation, an integrated multimodality approach was developed that enables measurements of invasive gold standard parameters and functional MR parameters simultaneously for the same kidney in an ultrahigh field small animal scanner (MR-PHYSIOL) [25–28]. For such a setup, a special Transonic™ flow probe is used: its acoustic reflector is made of Macor ceramics instead of stainless steel or brass to meet the safety and compatibility requirements of MRI (see the chapter by Cantow K et al. "Monitoring Renal Hemodynamics and Oxygenation by Invasive Probes: Experimental Protocol").

With the introduction of the Transonic™ technique, all other methods used for quantitative assessment of total renal blood flow in experimental animals including but not limited to electromagnetic probes and microsphere techniques were rendered obsolete.

3.2 Laser-Doppler Optodes In the 1970s, laser-Doppler probes were developed that estimate local blood perfusion within circumscribed areas of tissue [60, 61]. For this purpose, pulsed light of a distinct wavelength (e.g., 785 nm) is guided via an optical fiber into the tissue subjacent to the tip of the fiber, where it is scattered within the tissue. A fraction of the light encounters erythrocytes, is reflected, and is fed back via the fiber to the photodetecting device of the apparatus. From the amount of reflected light the concentration of erythrocytes per tissue volume is estimated. Because erythrocytes move with the bloodstream, the reflected light becomes frequency shifted due to the Doppler effect. The photodetected signal comprises a broad spectrum of Doppler frequency shifted signals from which the average velocity of erythrocytes is derived. An estimate of microvascular blood perfusion is than calculated by the apparatus as the product of mean erythrocyte velocity and mean erythrocyte concentration in the volume of tissue under illumination from the probe [61].

> The most important limitation of the laser-Doppler technique is that it can only provide arbitrary, nonabsolute, units for blood perfusion rather than an absolute value of blood flow, which is why the method is often termed laser-Doppler-fluxmetry [25]. This is due to the fact that the actual sampling volume could only be determined if it were possible to identify which erythrocytes have interacted with the light reflected from the tissue. The latter depends on the optical scattering and absorption coefficients of the tissue. Since these coefficients vary considerably with regard to the local microvasculature and its perfusion at the time of measurement, it is impossible to determine the actual sampling volume and thus the blood flow in absolute terms [61].

> Estimated mean sampling depths for mammal tissues range from 0.5–1.5 mm, corresponding to "captured" volumes of roughly 1 mm3 of tissue. However, these values must be lower in regions with large light absorption due to large amounts of erythrocytes per volume such as the renal cortex, and lower in the inner medulla with its low amount of erythrocytes per volume [38– 40, 61].

> Considering the rather small volume of tissue for which the laser-Doppler data are obtained versus the heterogeneity of intrarenal perfusion, extrapolation of the results of one area to other areas, even of the same renal layer, is unfeasible. Because of the impossibility to provide absolute data, laser-Doppler-fluxmetry allows investigators only to report relative changes occurring while the optode is fixed in place. Therefore, only continuous and

Fig. 3 Original tracings of invasively measured parameters of renal hemodynamic and oxygenation in an anesthetized rat (left panel) and sketch of the left kidney with the locations of the used probes (right panel). The rat was exposed to staircasewise reductions in renal perfusion pressure followed by staircasewise pressure restorations performed by means of a servo-controlled inflatable occluder placed around the suprarenal aorta (pressure servo) [22]. This was done under normoxic (inspiratory fraction of O2, FiO2 ¼ 0.2), hyperoxic (FiO2 ¼ 1.0), and hypoxic (FiO2 ¼ 0.1) conditions

Fig. 4 Time courses of invasively measured parameters of renal hemodynamics and oxygenation in two groups of anesthetized rats that both received a 1.5 mL bolus injection of an X-ray contrast medium (iodixanol 320 mg iodine per mL) into the thoracic aorta at time 0. In the blue group, a low dose of nitrite was administered by continuous infusion, while in the red group saline was infused (control). Conductance values (the reciprocal of vascular resistance) were calculated by dividing the respective perfusion values by arterial blood pressure, in order to distinguish flow changes that result from passive circular distension/compression of vessels from those actively exerted by vascular smooth muscles. Values (mean SEM) are given as relative changes from baseline. Redrawn from Ref. 20

quantifiable changes relative to baseline (control) measurements are recorded, typically induced by acute interventions such as changes in renal arterial pressure or the inspiratory O2 fraction, or by the administration of drugs (see Figs. 3 and 4) [18–20, 22, 23, 25, 28, 62]. Yet even interpretation of changes in laser-Doppler data recorded during such maneuvers must be done with due caution: they could rely on changes in local hematocrit, for example, due to changes in plasma skimming, rather than changes in blood flow [25, 39, 48].

In many studies on renal hemodynamics, data from two probes are recorded, one placed within the cortex and the other one in the (outer) medulla. This approach has been also employed in chronically instrumented rats [22]. It must be noted, that the potential tissue injury inflicted by the insertion of the optodes (usual tip diameters range from 50 to 300 μm) may confound the results. Laser-Doppler fluxmetry devices are supplied by various manufacturers. Oxford Optronics, Ltd., provides combined fluorescencequenching and laser-Doppler optodes (OxyLite/OxyFlo™) that enable simultaneous recordings of local tissue pO2 (vide infra) and of relative changes in local tissue perfusion in the same circumscribed area of renal tissue (see Fig. 2b).

3.3 Clark-Type Electrodes These polarographic electrodes were developed in the 1950s by Clark and colleagues [63]. The electrodes consist of a noble metal (e.g., silver, gold, platinum) which reduces O2 due to a negative polarizing voltage. The difference in voltage between the reference electrode (anode) and the measuring electrode (cathode) is proportional to the amount of O2 molecules being reduced on the cathode. This enables quantification of tissue pO2 (absolute values in mmHg). The critical technical advance was electrical insulation of the anode and cathode with a gas-permeable but liquidimpermeable membrane. This approach provided the foundation for the development of miniaturized versions (see Fig. 2c) with tips as small as 1–5 μm, that enable quantification of pO2 even in specific vascular and tubular elements of the kidney [64, 65]. Using such small electrodes, Lu¨bbers, Baumg€artl, and colleagues were the first to document the vast heterogeneity of renal tissue pO2 among as well as within the renal layers that is depicted by Fig. 1 [43–45]. These authors generated "spatial maps" with the frequency distribution of pO2 values ("pO2 histograms") from the cortex to the inner medulla, by moving the electrodes through the renal tissues using a nanostepper technique

(step width 50 μm).

While "pO2 histograms" have also been obtained in a few recent studies [66, 67], the majority of studies uses fixed electrodes, making sure that the tips remain at the same location within the kidney throughout the experiment. Often, one electrode is placed within the cortex and the other one in the (outer) medulla. The tissue volume that is "captured" by the electrodes has probably about the same diameter as the tip of the electrode; thus, standard electrodes of 10 μm diameter obtain pO2 values that are representative for a relatively small tissue volume only [67]. In the face of the spatial heterogeneity of intrarenal pO2, this constitutes a major limitation of this technique. On the other hand, the potential tissue injury inflicted by the insertion of these small electrodes is less than with bigger electrodes or fluorescence quenching optodes that enable better "spatial averaging" (vide infra) [24, 35, 67]. Another drawback of Clark-type electrodes is O2 consumption due to their electrochemical reduction, which is particularly unfavorable in areas with a fragile balance between O2 delivery and demand such as the renal medulla [67–70].

Furthermore, it is not technically feasible to make repeated measurements in the same anesthetized animal over time-frames longer than a few hours. Yet, recently, a telemetric method using carbon paste electrodes has been developed that allows assessment of renal tissue pO2 in conscious rats [35]. This method provides exquisite temporal resolution [24]. The major limitations of the method are (1) that it requires considerable skill and expertise to construct and implant the devices, (2) that tissue pO2 can only be monitored at a single location in each animal, and (3) that only relative changes in tissue pO2, rather than absolute levels are measured.

3.4 Fluorescence-Quenching Optodes In the 1990s, a fiber-optic oxygen-sensing device was developed that measures pO2 from the lifetime of fluorescence of a (ruthenium- or platinum-based) luminophore at the tip of an optode [70, 71]. The probe is connected to a pO2 meter developed by Oxford Optronics, Ltd., that generates short pulses of blue LED light, which are transmitted via the fiber to the tip of the optode [70]. The resulting emission of fluorescent light is quenched by the presence of O2 molecules. The OxyLite™ instrument measures the lifetime of each pulse of fluorescence, which is approximately inversely proportional to the pO2 according to the Stern-Volmer equation at pO2 values between 0 and 100 mmHg [68, 71].

Direct comparisons between Clark-type electrodes and quenching optodes reveal their respective advantages and disadvantages [67, 68, 70]. Both techniques share the disadvantage of invasiveness, and the advantage of enabling absolute quantification of pO2 (in mmHg). Advantages of the optode include the absence of O2 consumption and a greater accuracy at low pO2, which is of particular importance when it comes to measurements in the renal medulla. In contrast to the electrodes that "capture" a tissue volume that is approximately equal to their tip diameter, optodes do not "capture" a proper volume. The pO2 measured by the optodes is that within the probe-head (at the tip) of the optode itself, which covers a circular area of 230 μm diameter [67, 68, 70]. The larger diameter of the optodes as compared with that of standard pO2 electrodes increases the risk for tissue damage. On the other hand, it facilitates a favorable "spatial averaging" of renal tissue pO2 values in the face of their huge spatial heterogeneity. In fact, "pO2histograms" of renal cortical pO2 obtained in rabbits by a fluorescence optode were more tightly distributed than those provided by a Clark-type electrode, which may aid detection of (patho) physiologically significant changes in intrarenal oxygenation [67].

As mentioned earlier, combined fluorescence-quenching-laser-Doppler probes (OxyLite/OxyFlo™) enable simultaneous measurements of local tissue pO2 and assessment of local tissue perfusion in the same circumscribed area (see Fig. 2b).

3.5 Monitoring of Arterial Blood Pressure It is mandatory for in vivo studies in anesthetized experimental animals, be it studies on control of renal hemodynamics and oxygenation or on every other functional parameter, to monitor cardiovascular variables such as arterial blood pressure and heart rate throughout the experiment in order to ascertain that physiological cardiovascular control is maintained. In anesthetized animals this is best achieved by use of invasive techniques: a fluid-filled catheter inserted into a conduit artery (e.g., femoral or carotid arteries) connected to one of the various commercially available pressure transducers and quantitatively calibrated via a recording device (see the chapter by Cantow K et al. "Monitoring Renal Hemodynamics and Oxygenation by Invasive Probes: Experimental Protocol").

Direct intra-arterial measurements in unanesthetized animals is the most physiologically relevant means of arterial pressure determination. Telemetric pressure monitoring is the gold standard in conscious animals, and various telemetric devices are commercially available [72, 73]. A limitation of this technique is that implantation surgery and maintenance of catheters require considerable skill [72]. The tail-cuff method is a simple noninvasive plethysmographic technique to assess arterial pressure in conscious rats and mice [74, 75]. Since the conditions for tail-cuff measurements necessarily involve heating and restraint of the animal, which almost inevitably alter the measured parameter (arterial pressure), the physiological meaningfulness of the results is questionable [73].

3.6 Pimonidazole Adduct Immunohistochemistry This method was originally developed for detection of hypoxia in tumors, but subsequently adapted for use in the kidney [53]. For this technique, pimonidazole (usually 60 mg/kg of body mass) is administered to the experimental animal in vivo [54]. Pimonidazole forms adducts within hypoxic (<approximately 10 mmHg) cells. These adducts are immunohistochemically visualized in ex vivo samples of the kidney. The advantage of this technique is that it provides good spatial resolution of renal tissue hypoxia at the cellular level. Its major limitations include (1) that it detects severe hypoxia only, with a cutoff of about 10 mmHg, (2) that hypoxia can only be assessed postmortem, and (3) only at a single time point. Those using this technique must also be aware of two major sources of false-positive detection of hypoxia, as recently revealed [55]. Firstly, pimonidazole adducts can form postmortem in cells that were not hypoxic in the living animal. For this reason, it is recommended that kidney tissue be perfusion fixed, rather than immersion fixed. Secondly, when the standard (mouse) monoclonal antibody is used, false positive staining has been observed in damaged rat kidney tissue, presumably as a result of the anti-mouse secondary antibody binding to mouse-like antigens in the damaged tissue. This problem can be avoided by use of a newly commercially available polyclonal anti-pimonidazole antibody raised in rabbits [55].

### 3.7 Examples of Methods That Are Not Generally Available

3.7.1 Bladder pO2

It is not technically feasible to measure renal tissue pO2 in humans. It can also be technically difficult to measure renal tissue pO2 in experimental animals unless they are deeply anesthetized. However, it is technically feasible to measure the pO2 of urine in both experimental animals and man. There has been a long history of interest in this approach for monitoring renal medullary tissue oxygenation [76]. However, with the recent development of fiber-optic probes that can be inserted into a standard Foley bladder catheter, it has been possible to subject this method to close scrutiny [77]. Available evidence now indicates that urinary pO2 can provide a good estimate of renal medullary pO2 [77]. Because this technique can be deployed in humans equipped with a bladder catheter, it provides a translational pathway for development of approaches to improve renal medullary oxygenation in human patients at risk of AKI. It is also noteworthy that low urinary pO2 during cardiac surgery requiring cardiopulmonary bypass predicts development of postoperative AKI. In light of this, urinary pO2 measurements may provide a real-time measure of medullary oxygenation and predict the risk of AKI in critically ill patients. The major limitation of this approach is that the combination of low urine flow and high arterial pO2 can confound the relationship between urinary pO2 and medullary pO2 [77].

3.7.2 Near Infrared Spectroscopy Near infrared spectroscopy (NIRS), together with models for diffuse optical imaging affords the measurement of (1) tissue concentrations of oxygenated hemoglobin (oxyHb) and deoxygenated hemoglobin (deoxyHb), (2) oxygen saturation of Hb (StO2) within tissues, and (3) tissue blood volume fraction in vivo. In particular, spatially resolved reflectance NIRS (SRR-NIRS), together with Monte-Carlo simulations of diffuse light propagation, has been applied to monitor these parameters within renal cortical tissue [78]. Quantification of the underlying absorption and reduced scattering coefficients of the tissue is based on analyzing the decrease of the intensity of diffusely remitted light with increasing distance from the light injection point. Recording techniques for the renal cortex of anesthetized rats have been established, with a linear fiber array placed on the ventral surface of the exposed kidney. Measurements have been done at selected nearinfrared wavelengths between about 670 and 850 nm by multiplexing millisecond pulses from the different laser sources in time. In this way, quantitative measurements of renal cortex parameters have been achieved at a repetition rate of multiple Hz [78].

SRR-NIRS has been successfully combined with "classical" invasive techniques to study renal hemodynamics and oxygenation in rats [78]. Thus, SRR-NIRS-derived parameters such as the amount of Hb per tissue volume and StO2 were obtained simultaneously with parameters including total renal blood flow, local cortical tissue perfusion and tissue pO2 in the same kidney. Control of renal hemodynamics and oxygenation was studied by dedicated reversible test interventions including brief occlusions of the renal artery or the renal vein, short-term hypoxia, hyperoxia, and hypercapnia, and administration of adenosine. The results demonstrate that the combined approach, by providing different but complementary information, enables a more comprehensive characterization of renal hemodynamics and oxygenation [78].

The SRR-NIRS technique could also serve to help calibrate renal T2\* MRI. This MR surrogate parameter for blood oxygenation is related to the amount of deoxyHb per tissue volume. Confounders of the T2\* to tissue pO2 relationship include the vascular volume fraction, the local hematocrit, and shifts of the oxyHb dissociation curve [25–28]. Since SRR-NIRS enables quantitative monitoring of the amount of Hb per tissue volume and of StO2 it is an ideal candidate to unravel and quantify the contributions of confounding factors to renal T2\* MRI.

By extending the wavelength range beyond 900 nm, the tissue water content can additionally be assessed, which, together with the reduced scattering coefficients at the different wavelengths, can serve as a surrogate for other (patho)physiologically relevant parameters such as the tubular volume fraction.

Application of SRR-NIRS in reflection mode constrains measurements to the renal cortex of rats. Recently it was demonstrated that changes of rat medullary Hb parameters can also be assessed by SRR-NIRS in transmission mode using an additional fiber placed on the dorsal surface of the kidney [79].

Perhaps the "final frontier" in methods for assessing renal oxygenation are techniques that provide cellular resolution, in real time, in living animals. Hirakawa and colleagues have recently reported use

3.7.3 Intravital Microscopy

of a technique they call "intravital phosphorescence lifetime imaging microscopy" [47]. This technique allows cellular pO2 in the renal cortex of live animals to be resolved to the level of individual tubular cross-sections. It requires injection of a molecule whose phosphorescence is quenched by O2. Thus, pO2 can be measured by assessing the decay curve of the phosphorescence. Such probes are available that accumulate intracellularly or are retained within the vasculature. Thus, it is possible to use this approach to assess the heterogeneity of renal blood and tissue pO2 in real time with unprecedented cellular resolution. Currently, the technique is limited for use in the superficial cortex.

### Acknowledgments

This work was funded, in part (Kathleen Cantow, Thoralf Niendorf, and Erdmann Seeliger), by the German Research Foundation (Gefoerdert durch die Deutsche Forschungsgemeinschaft (DFG), Projektnummer/Project number 394046635, SFB 1365, RENOPROTECTION).

The authors wish to thank Ariane Anger and Andrea Gerhardt for expert technical assistance and Dr. Connie Ow for the image shown in Fig. 2c.

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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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 7

### Ultrasound and Photoacoustic Imaging of the Kidney: Basic Concepts and Protocols

### Sandra Meyer, Dieter Fuchs, and Martin Meier

### Abstract

Noninvasive, robust, and reproducible methods to image kidneys are provided by different imaging modalities. A combination of modalities (multimodality) can give better insight into structure and function and to understand the physiology of the kidney. Magnetic resonance imaging can be complemented by a multimodal imaging approach to obtain additional information or include interventional procedures. In the clinic, renal ultrasound has been essential for the diagnosis and management of kidney disease and for the guidance of invasive procedures for a long time. Adapting ultrasound to preclinical requirements and for translational research, the combination with photoacoustic imaging expands the capabilities to obtain anatomical, functional, and molecular information from animal models. This chapter describes the basic concepts of how to image kidneys using different and most appropriate modalities.

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, Ultrasound imaging (US), Photoacoustic imaging (PA), Multimodality, Doppler, Contrast, Animals, Rats, Mice

### 1 Introduction

Magnetic resonance imaging, while very powerful by itself, can be complemented by a multimodal imaging approach to streamline workflow processes, obtain additional information, or include interventional procedures. Ultrasound and computed tomography are modalities of primary choice in renal imaging on patients, with the disadvantage of computed tomography exposing the body to harmful irradiation. Due to its fast, cost-effective, versatile, and nonionizing characteristics, ultrasound imaging is well-established in the clinical practice, often used for first-line diagnostics and follow-ups and as a complementary method to MRI.

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\_7, © The Author(s) 2021

In the clinic, renal ultrasound has been essential for the diagnosis and management of kidney disease and for the guidance of invasive procedure for decades. Nevertheless, new developments to advance clinical diagnostics and medical treatments are required and strongly rely on appropriate preclinical models. Adapting the ultrasound technology to match preclinical requirements for translational studies and combining it with photoacoustic imaging, expands the researchers' capabilities to obtain anatomical, functional, and molecular information from their animal models.

MRI has the advantage of superior soft-tissue contrast, which facilitates and enables the detection and characterization of renal lesions. This chapter will give an introduction to how detailed kidney images can be created, blood flow visualized and quantified, and capillary as well as renal function assessed in vivo, noninvasively, and in real time using high-frequency ultrasound.

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 Concepts of Ultrasound and Photoacoustics

The basic principle of ultrasound imaging relies on detecting differences in tissue density by using sound waves with frequencies above the range of human hearing—typically in the MHz range. Using piezoelectric elements short, harmless ultrasound pulses are sent into the body. These sound waves are partially reflected back to the transducer by different layers of tissue and the return signals (often referred to as "echo"—hence the term "echography" which is used as a synonym to ultrasound) are recorded. Based on the time it takes for the echo to travel into the tissue and back and the strength of the signal, an image can be computed and displayed in real time. The resulting two-dimensional, grayscale image of an area of interest is referred to as Brightness Mode or B-Mode. Tissue closest to the transducer is displayed at the top and structures further away are shown below; a scale provides depth and size information in real time. With increasing density more echo is recorded resulting in a brighter signal on the screen; therefore, dense structures (e.g., skin) are represented in white and light gray tones, whereas ultrasound gel used for coupling and fluid-filled structures inside the body appear black. Most tissues and organs are displayed in a range of grayscales in between and have a characteristic appearance ("speckle pattern") that aids in their identification. B-Mode is the standard imaging mode on ultrasound systems and worked with the most because it allows for effective visualization of anatomical structures. It can be complemented by additional imaging modes to obtain functional information which are outlined in more detail in Subheading 3.

In ultrasound, the frequency of the sound waves sent into the tissue is of importance as it determines image resolution and penetration depth. In a clinical setting, renal ultrasound scans on adult patients are normally performed with a 3–6 MHz curved array transducer where the term "curved" describes the shape of the transducer surface. Routine measurements include determining the length of the kidney in a sagittal view, the measurement of cortex thickness and of blood flow in the renal and interlobar arteries [1].

By using an appropriate setup and ultrasound imaging system, these parameters and a variety of others can be obtained in preclinical animal models. Dedicated small-animal micro-ultrasound systems provide high spatial resolution down to 30 μm by employing higher frequencies of up to 70 MHz. While early systems for small animal imaging were equipped with mechanical probes, state-ofthe-art micro-ultrasound systems are working with digital linear array (flat transducer surface) probes and are suitable for a broad range of applications [2]. In addition to providing higher resolution than clinical machines, ultrasound systems dedicated to preclinical research also:

Since ultrasound relies on the detection of soundwaves, no contrast agent is required to visualize anatomical structures, 3D volumes, or to assess blood flow velocities. It is noteworthy that anatomical structures and volumes measured with dedicated small animal ultrasound match results acquired with dedicated small animal MRI or μCT systems [3].

Above and beyond the capabilities of micro-ultrasound to gather anatomical and functional information, photoacoustic imaging can also be applied in a preclinical setting to add molecular information. Photoacoustic imaging is a hybrid imaging modality that combines the strengths of optical and ultrasound imaging and can be used to monitor endogenous parameters such as blood content or oxygen saturation or various exogenous contrast agents.

In this imaging approach (Fig. 2), short (a few nanoseconds) pulses of near-infrared laser light are transmitted into the tissue and

Fig. 1 Standard renal exam: animal and transducer positioning; (a) transverse position for left kidney imaging, (b) transverse US B-Mode image of the left kidney with length measurement, (c) sagittal position for left kidney imaging, (d) sagittal US B-Mode image of the left kidney with length measurement. Note the millimeter scale on the right and the physiological monitoring (ECG in green, respiration in yellow) below the US images

are selectively absorbed according to the optical properties of the tissue components. Light absorption results in a thermoelastic expansion creating a sound wave that can be localized by the ultrasound transducer and at the same resolution as the ultrasound image [4].

Fig. 2 Photoacoustic imaging; (a) the photoacoustic effect, (b) coregistered B-Mode image providing anatomical information (left) and photoacoustic image displaying oxygen saturation (right) in the spleen and left kidney

### 3 Overview of Applications

This section provides an overview of applications of high-frequency ultrasound and photoacoustic imaging in preclinical renal research. Examples of specific disease models and how multimodal imaging is applied are summarized in a brief review of selected publications in Table 1.

### Table 1

### Examples of specific disease models and how multimodal imaging is applied

(continued)

3.1 Evaluating Kidney Size and Anatomy

B-Mode, as introduced in Measurement concepts (2.), is used to locate the kidneys in the abdomen, observe the surrounding tissue and anatomical structure and to obtain organ size.

When viewed using ultrasound, the kidneys are located lateral of the abdominal aorta. The right kidney is positioned slightly more cranial in the thoracic cavity then the left kidney, and both can be identified by their characteristic appearance in ultrasound: the medulla in the medial portion of the kidney is darker than the cortex and surrounding the cortex, the kidney capsule appears as a bright thin line. Once the appearance of kidneys and surrounding in ultrasound are familiar, it is easy to identify any changes: Cysts are liquid-filled structures and therefore are visible as black, round structures with in the tissue. Hydronephrosis is qualitatively recognized as distension and dilatation of the renal pelvis and a method to quantify hydronephrosis based on the proportion of renal parenchyma in a longitudinal ultrasound image as been established [5] and underlying etiologies have been studied [6]. Tumor masses will appear in a different grayscale than the surrounding tissue and when becoming larger will distort the regular bean shaped form of the kidney. The observation and sizing of the kidney in 2D images can be performed in the transverse (Fig. 1a, b) and sagittal imaging (Fig. 1c, d) plane as required, based on preference. Moving through the whole organ in in both imaging planes in 2D allows to systematically screen for changes inside and also in the surroundings (e.g., ascites in the peritoneum). In addition, ultrasound imaging also may be used to study cardiovascular disease linked to chronic kidney disease [19].

Most preclinical ultrasound systems also allow to gather threedimensional information. The principle for 3D image acquisition is similar between ultrasound and MRI where both acquire sequential 2D images with a known distance between them, determining the resolution along the z-axis (slice thickness). In ultrasound imaging a motor needs to be attached to move the probe across the area of interest along the z-axis and acquire a stack of images (Fig. 3). The consecutive 2D slices can be compiled into a 3D view, displayed in various ways, and analyzed to obtain accurate volume information of the kidneys and also cysts or tumor masses within.

3.2 Visualization of Blood Flow When ultrasound waves are reflected by moving blood a frequency shift (Doppler shift) of the return ultrasound signal is induced that allows to display and measure blood flow. Most ultrasound systems are equipped with several imaging modes that use this Doppler effect to display blood flow in 2D images and allow for measuring blood flow velocities.

> Color Doppler mode, overlaid on top of a B-Mode image for anatomical reference, displays blood flow in the area of interest in gradients of red and blue (Fig. 4b). The color gradients indicate both the direction and mean velocity of the Doppler shift: blood

Fig. 3 Schematic workflow for 3D US image acquisition and processing

flowing toward the transducer is represented by a gradient between red (lowest velocity) and white (highest velocity) and blood flowing away from the transducer is represented by a gradient between blue (lowest velocity) and white (highest velocity). BART is an acronym commonly used to remember the colors associated with the direction of flow: Blue Away, Red Toward. Color Doppler mode is a very useful tool to identify the major and cortical vessels in the kidney and accurately guide the positioning for flow measurements, which are performed in pulsed-wave (PW) Doppler mode. To obtain a measurement, a Doppler sample volume (point where the blood flow should be measured) is placed within the vessel and the system generates a PW Doppler spectrum that displays time on the x-axis and flow velocity on the y-axis (Fig. 4d). Once PW Doppler spectrums of the renal artery and vein have been recorded, velocity measurements such as peak systolic velocity (PSV), end-diastolic velocity (EDV), and the velocity time integral (VTI) can be added. Based on these measurements, relevant parameters will be calculated as an indicator for vascular function.

l The resistive index (RI) is an example of additional analysis that may be useful. A RI of 0 is calculated for continuous flow patterns and an RI of 1 indicates a strongly pulsating flow with no flow during diastole. In the clinic an RI of the renal

Fig. 4 Assessment of blood flow; (a) power Doppler signal in transverse view of the left kidney is overlaid with the B-mode image, (b) Blood flow in the interlobal vessels as visualized in Color Doppler Mode, in this image red and blue colors indicate arterial and venous flow, respectively; (c) three-dimensional rendering of Power Doppler signal of a healthy kidney, (d) pulse-waved Doppler profile of renal artery flow, peak systolic (PSV), and lowest diastolic velocity (LDV) measurements are shown, the sample positioning is indicated in the scout image above the flow profile

artery ¼ 0.6 is observed in healthy adults and a rise in RI is seen as an indicator for renal disease.

$$\text{RI} = \frac{\text{PSV} - \text{EDV}}{\text{PSV}}$$

l The pulsatility index (PI) is another parameter often calculated in clinical renal diagnosis.

$$\text{PI} = \frac{\text{PSV} - \text{EDV}}{\text{VTI mean velocity}}$$

Power Doppler mode is another tool used to display and analyze blood flow within an area of interest. It is similar to color Doppler mode in detecting the occurring Doppler shift, but instead of displaying blood flow direction, it color-codes flow from orange to yellow according to signal intensity and is most useful for smaller vessels having slower blood flow (Fig. 4a).

It is crucial for PW Doppler mode and all other Doppler imaging modes, that an appropriate angle is achieved between the transmitted ultrasound beam and the blood flowing through the vessel; the imaging plane should be adjusted such that the angle between these two is less than 60 to provide an accurate velocity measurement.

Color and power Doppler information can also be acquired in 3D, visualizing the blood vessel structure in the entire organ (Fig. 4c). The amount of Doppler signal in a 3D volume can then be calculated. With the right animal setup, imaging both kidneys at the same time is feasible.

Imaging While Doppler imaging can be used to visualize blood flow in the renal artery and vein and also in the interlobar and arcuate vessels, monitoring perfusion of the kidney at capillary level requires the use of ultrasound contrast agents. Ultrasound contrast agents are very small gas bubbles (microbubbles) about 2–3 μm in diameter and encapsulated by a polymer or lipid shell (Fig. 5a). Several contrast imaging modes can be used for the detection of these microbubbles in the area of interest. Linear contrast enhanced ultrasound, the first contrast mode used at frequencies above 15 MHz, primarily relies on subtraction of background information to display the increase in echo due to the wash-in of contrast agent. Newer generation systems allow to monitor perfusion in real-time based on nonlinear fundamental detection with amplitude modulation, resulting in a higher sensitivity [7].

While ultrasound imaging to monitor perfusion requires a contrast agent and therefore intravenous access, photoacoustic imaging allows for the detection of hemoglobin and its oxygen saturation status (Fig. 2). Based on the change of hemoglobin's optical properties dependent on oxygen bound to the heme group, one can assess tissue oxygenation at a level matching BOLD (blood oxygen level–dependent) MRI [8]. It should be noted that ultrasound contrast agent and photoacoustic imaging of hemoglobin are nontoxic and harmless to the animal.

3.4 Molecular Information Ultrasound and photoacoustic contrast agents can also be used for molecular imaging. Ultrasound contrast agents can be targeted to assess biomarker expression (Fig. 5c) but are exclusively tailored to vascular endothelium as the contrast agent does not leave the vasculature. Photoacoustic contrast agents, dyes or nanometersized particles, can be used for specifically targeting extravascular targets or for assessing glomeruli function and to study biodistribution [9].

### 3.3 Perfusion

Fig. 5 Perfusion assessment of both mouse kidneys; (a) schematic model of a microbubble used for US contrast imaging; (b) animal and transducer position, (c) B-mode image, left and right kidney are outlined; (d) schematic binding of targeted contrast agent to a biomarker; Non-linear contrast image of both kidneys before (e) and after contrast injection (f). After contrast injection; for analysis, regions of interest in cortex and medulla of left (blue) and right (red) kidney are selected (g); a graph displays the contrast signal in these regions of interest (h)

### 3.5 Interventional Procedures

In contrast to other imaging modalities, in ultrasound and photoacoustic imaging the animal is accessible throughout the imaging session, allowing for various interventional procedures to be performed during an imaging session (Fig. 6). Using a so-called in-plane needle visualization technique, it is possible to follow the movement of a needle inside the animal in B-mode and ensure accurate placement of the needle tip for injections or aspiration. Image guided injections allow to place cells to create orthotopic tumor models or apply drug treatments precisely without the need for more invasive surgery [10]. Switching to image guided injections is an important step in the refinement of animal procedures and helps to shorten recovery time and minimize discomfort and

Fig. 6 Image guided injection into the left kidney of a mouse; (a) animal and needle positioning, (b) penetration of the skin, (c) the needle tip (marked with an asterisk) is positioned in the medulla of the kidney, (d) after injection of US contrast agent and needle retraction, injection site is marked with an asterisk

> complications as the procedure can be performed with only light anesthesia and the procedure itself does not require postoperative analgesia. In addition the success of any given injection or aspiration can be verified in real time.

### 4 Protocols

4.1 Standard Renal Exam The protocol outlined below depicts a standard renal ultrasound exam that can be performed on various ultrasound systems. If available, a system dedicated for small animal imaging using highfrequency linear transducers is be preferred due its superior resolution and sensitivity. For rats a transducer with a 50 or 75 μm resolution is appropriate and for mice a transducer with 40 μm resolution should be used. Despite the similar principle of various ultrasound systems, the availability of imaging modes, settings and analysis tools may vary. Please refer to instructions provided by the manufacturer for system-specific information and tailored work protocols.

While some brief ultrasound scans (e.g., for linear measurements, screening for pathologies or anatomical alterations) can be performed on awake animals, more comprehensive renal exams including flow measurements or volumetric analysis normally require the use of anesthesia. A brief description of isoflurane anesthesia is included below, but researchers should always carefully follow local rules and regulations.

In order to ensure reproducibility and to streamline workflow in longitudinal studies, it is recommended to store settings for all imaging modes and applications, stick to the same anesthesia type and level, as well as monitor the animal's physiologic parameters.

located caudally of the stomach and in close proximity to spleen and the tail of the pancreas (Fig. 1a).

Doppler should be activated before starting the 3D acquisition and the size of the window should be enlarged to cover the region of interest, the use of respiration gating and a wall filter is recommended to avoid respiration artifacts.

13. Confirm all images required for analysis have been saved. Wipe off the ultrasound gel, turn off the isoflurane and release the animal from the mouse handling table. Monitor the animal until fully recovered.

4.2 Perfusion Studies Studying kidney perfusion in ultrasound imaging requires the intravenous injection of a contrast agent. Administration can be performed either as a continuous infusion using a pump, or as bolus injection, which can be delivered by hand but using a pump is recommended to ensure reproducibility. In addition it is possible to study reperfusion by applying a short ultrasound pulse to destroy the contrast agent within the field of view and record subsequent reperfusion. Here we would like to outline a basic perfusion study that can be included in an imaging session in combination with the standard exam. To allow for comparison, the imaging procedure should be standardized as much as possible. For a detailed review of contrast imaging in small animals and various parameters that should be considered, please refer to [11].

length to allow for sufficient acquisition time (e.g., 30 s), locate the prospective saving option on the system. If required, acquire a 3D data set before contrast injection and do not move the animal or transducer. Once the settings have been optimized it is advised to store the settings as presets.

In addition to 2D and 3D ultrasound imaging showing renal anatomy, blood flow, and perfusion, photoacoustic (PA) imaging can add additional information by visualizing endogenous as well as exogenous contrast agents.

4.3 Photoacoustic Exam to Monitor Oxygen Saturation

4.4 Image Guided Needle Injections Image guided interventions can be added to any imaging session as an additional step to locally administer a drug or implement cells. This procedure can be performed independently and is a fast and reliable procedure to replace a surgical intervention. Please refer to Subheading 4.1 for instructions on anesthesia, animal preparation, and imaging and follow these additional steps:

positioned correctly, the whole length and tip of the needle should be visible on the live image as a bright line. Once a good alignment has been achieved, movement of the needle and probe in the Y-axis should be avoided.

analysis software. Please refer to the user instructions to ensure all data required for analysis is saved. Modern preclinical ultrasound and photoacoustic equipment allow for extracting the image data as DICOM or raw formats, making the data accessible in nonproprietary image analysis programs.

coding of anatomical structures and further editing for educational purposes or publishing.

### Acknowledgments

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

systemic hypoperfusion evaluated by micro-US, color and PW-Doppler. Radiol Med 124 (5):323–330

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 8

### Hardware Considerations for Preclinical Magnetic Resonance of the Kidney

### Paula Ramos Delgado, Ekkehard Ku¨ stermann, Andre´ Ku¨ hne, Jason M. Millward, Thoralf Niendorf, Andreas Pohlmann, and Martin Meier

### Abstract

Magnetic resonance imaging (MRI) is a noninvasive imaging technology that offers unparalleled anatomical and functional detail, along with diagnostic sensitivity. MRI is suitable for longitudinal studies due to the lack of exposure to ionizing radiation. Before undertaking preclinical MRI investigations of the kidney, the appropriate MRI hardware should be carefully chosen to balance the competing demands of image quality, spatial resolution, and imaging speed, tailored to the specific scientific objectives of the investigation. Here we describe the equipment needed to perform renal MRI in rodents, with the aim to guide the appropriate hardware selection to meet the needs of renal MRI applications.

This publication 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 chapter on hardware considerations for renal MRI in small animals is complemented by two separate publications describing the experimental procedure and data analysis.

Key words Magnetic Resonance Imaging (MRI), Kidney, Mice, Rats, Hardware, RF coils, Renal MRI

### 1 Introduction

The detection, staging, and monitoring of kidney diseases (KD) using measurements of renal function and structure is critical in clinical nephrology. The use of animal models in kidney research promotes the understanding of disease and therapy mechanisms. While current imaging modalities such as computed tomography or ultrasound can provide some information on structural changes in the kidney, these approaches offer limited insights into functional changes.

Magnetic resonance imaging (MRI) is a noninvasive imaging technology that offers an unparalleled level of anatomical and functional detail, along with diagnostic sensitivity. MRI is suitable

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\_8, © The Author(s) 2021

for longitudinal studies due to the lack of exposure to ionizing radiation. Together with established imaging techniques, novel approaches such as dynamic contrast-enhanced MRI (DCE-MRI), arterial spin labeling (ASL), blood oxygen level dependent MRI (BOLD), and diffusion-weighted imaging (DWI) can provide a broad characterization of the kidney, including renal vascular perfusion, oxygenation, and glomerular filtration rate.

For performing dedicated MR measurements of the kidney, the appropriate MRI hardware should be carefully chosen to balance the competing demands of image quality, spatial resolution and imaging speed, tailored to the specific scientific objectives of the investigation. This chapter is intended to outline the key hardware components of an MRI scanner and to provide an overview of currently established technologies, with the goal to guide the appropriate hardware selection that meets the needs of preclinical MRI of the kidney.

This chapter on hardware considerations for renal MRI in small animals is part of the book Pohlmann A, Niendorf T (eds) (2020) Preclinical MRI of the Kidney–Methods and Protocols. Springer, New York.

### 2 Static Magnetic Field Strength (B0)

A superconducting magnet that generates a constant homogeneous magnetic field is the core hardware component of an MRI scanner. This strength of this static magnetic field (B0) is proportional to the net polarization produced within the sample. The detected MR signal on a receive RF coil depends on B0 2 , whereas the noise increases linearly with B0 [1]. B0 governs the signal-to-noise ratio (SNR), which is the currency for MR image quality. This provides strong motivation and momentum to move to increasingly higher magnetic field strengths for preclinical and clinical renal MRI applications [2–4]. The proportionality of SNR with B0 is linear in the quasi-static regime, as demonstrated theoretically [5, 6] and experimentally [7]. For higher magnetic field strengths, wave propagation and the increasing resonance frequencies must be considered [2]. Additionally, an experimental study on intrinsic SNR behaviour with increasing B0 field strengths (3.0, 7.0, and 9.4 T) in the human brain reported a superlinear increase in SNR (SNR / B1:65 0 ) [7]. This observation was also confirmed in a simulation study [8].

Relaxation time constants of biological tissues also change with the magnetic field strength, thus influencing the SNR as well as the contrast-to-noise ratio (CNR), which ultimately has an impact on the image quality. The longitudinal relaxation time (T1) increases nonlinearly with increasing B0 [1, 9–11]. The transversal relaxation (T2) and T2\* decrease at different slopes linearly [1, 9, 12– 14]. These dependencies on B0 are responsible for the reduced T1-weighted contrast and increased BOLD effect at higher magnetic field strengths (see Note 1). MRI methods based on T2\* weighting (e.g., susceptibility weighted imaging, BOLD) will benefit from the T2\* shortening, and thus increase in sensitivity at high magnetic field strengths, which underlies their broad application in renal MRI. An example of renal T2\*-mapping (R∗ 2 ¼ 1=T∗ 2 Þ is the study of oxygenation and haemodynamics in renal disorders (for a review see [15]). MR spectroscopy also benefits from the enhanced signal dispersion achieved at higher magnetic field strengths, which results in better resolved spectra, and permits improved accuracy of metabolite identification and quantification [1, 16].

However, MRI at high magnetic field strengths is susceptible to more severe B0 and B1 inhomogeneities, compared to lower field strengths. These inhomogeneities are primarily due to differences in the magnetic susceptibility at tissue interfaces. Susceptibility artifacts present themselves as T2\*-induced signal losses (or even signal voids) in areas with very high B0 gradients, low phaseencoding bandwidth-related image distortion and off-resonance effects caused by ΔB0-induced frequency dispersions. MRI of the kidney poses additional challenges due to the anatomical context. The kidneys are adjacent to bowels, in close proximity to skin/fat/ muscle boundaries (e.g., perirenal fat) or in areas containing cavities might be particularly prone to susceptibility artifacts.

Artifacts are visible as signal losses at interfaces between fat and tissue or air and tissue (susceptibility artifacts), image blurring, image distortion, and broadening of spectral lines in MRS, which interferes with the reliable differentiation of metabolites. B0 inhomogeneities can be mitigated using B0 shimming (see Subheading 3).

B1 inhomogeneities lead to variations in the flip angle driven by the RF pulse excitation. B1 inhomogeneities affect tissue contrast and SNR, appearing as signal intensity variations or losses across the field of view. Nonuniformities in B1 are caused by the RF coil (e.g., surface RF coil, see Subheading 4) or by the creation of standing waves and constructive/destructive interferences within the object under investigation, due to the reduction of the effective wavelength which at higher field strengths becomes shorter, and can approach the dimensions of the specific anatomical structures under study [17]. While the RF wavelength inside tissue at 3.0 T (~128 MHz) is about 30 cm, at 7.0 T (~300 MHz) it is reduced to about 13 cm [18] and to 6.5 cm at 14.0 T (~600 MHz). In preclinical renal MRI, this is not a major concern, as the size of the abdomen in small animals is typically much smaller than the RF wavelength.

RF power requirements for spin excitation increase quadratically with B0. This imposes major concerns, especially for human applications, since the RF power deposition can easily reach the regulatory limits specified by IEC guidelines for the specific absorption rate (SAR) [16]. This constraint can be adressed by modification and adaptation of MR pulse sequences, acquisition strategies, and RF coils [18, 19]. In the case of preclinical studies, these issues have received less attention since there are no enforced SAR guidelines for small animals. Nevertheless, the physiological effects of tissue heating due to heavy RF deposition in experimental animals could be a serious confounding factor in preclinical studies that should not be ignored.

2.1 SNR Comparison for Magnetic Field Strength The SNR gains resulting from increasing B0, field have been computationally estimated for brain MRI to range from B0 1.2 (close to the brain surface) up to B0 2.1 (at the center of the head) [8]. Pohmann et al. [7] investigated the field dependence of SNR for magnetic field strengths ranging between 3.0 T and 9.4 T using gradient echo techniques, and reported that SNR increases superlinearly (B0 1.65). Using this relationship, the calculated relative SNR (normalized for the typical case of a 3.0 T scanner) and resolution gains (in μm, in-plane resolution 1 mm) corresponding to all static magnetic field strengths from 1.5 T to 21.1 T are shown in Table 1.

> The increase in SNR provided by higher magnetic field strengths can be invested in enhanced spatial resolution or improvements in imaging speed. The triad of B0, spatial resolution, and imaging speed form the so-called SNR conundrum in which the competing constraints of SNR, CNR, spatial resolution, and imaging speed govern the image quality. For example, increasing the matrix size for a certain field of view will increase the spatial resolution—which will permit the visualization of more subtle anatomical structures in the kidney—but at a cost of reducing the SNR, or signal intensity per pixel/voxel. Although resolution in the frequency-encoding direction can be conveniently enhanced by increasing the matrix size without major scan time penalty, increasing spatial resolution along the phase-encoding direction substantially prolongs the scan time. Boosting the SNR is also possible by modifying the size of the field of view or the slice thickness: the larger these parameters are, the more signal is acquired. Another way to further increase SNR is signal averaging, which reduces the noise but comes at the cost longer scan times. Depending on the individual experimental context, longer scan times may be a reasonable price to pay (especially for ex vivo studies). For in vivo studies, animal welfare considerations call for the minimum scan time possible to achieve the scientific objective, and shorter scans promote higher throughput of individual animals to improve the experimental sample size.

> Additional strategies to gain SNR include the use of SNR-promoting rapid imaging techniques such as RARE imaging, and dedicated RF coil technology, including cryogenically cooled RF coils (see Subheading 4).

Table 1 Relative SNR and resolution gain with increasing static magnetic field strength

SNR values were computed as proportional to B1.65. The relative SNR gain was calculated relative to that of 3.0 T

### 2.2 Magnetic Field Strengths Used in Recently Published Studies

To provide guidance as to which B0 is necessary or sufficient for preclinical MRI of the kidney, a total of 121 publications published over the past 10 years were reviewed. Of these, 66 publications reported on renal MRI in rats, and 55 in mice. Figure 1 shows the number of publications found per magnetic field strength for renal MRI in rats and mice. In rats, a total of 25 publications (38%) were performed at 3.0 T, closely followed by 9.4 T with 20 papers (30%). In mice the preferred magnetic field strength was 7.0 T (25 publications, 46%). The magnetic field strengths of 9.4 T (8 papers, 15%), 3.0 T (7 papers, 13%), and 11.7 T (6 publications, 11%) were also used for MRI studies of mouse kidney.

Figure 2 illustrates the number of publications per magnetic field strength, listed by year for rats and mice. In rats, the number of studies performed at 3.0 and 9.4 T was quite stable for the last 7 years. Studies at 4.7 T (7 papers, 11%) were also widely reported, as shown in the trends displayed since 2009. The use of field strengths of 1.5 and 7.0 T (7 and 6 publications, respectively) varied during the period under investigation. In contrast, 7.0 T was established as the typical magnetic field strength for preclinical MRI of the kidney in mice during this period. In the last few years, 11.7 T and 16.4 T have also emerged as MR systems suitable for the study of the mouse kidney.

Fig. 1 Number of publications per magnetic field strength for renal MRI in rats and mice, within a total of 121 reports on preclinical renal MRI published in the past 10 years

### 3 Gradient Systems

To enable the spatial localisation of the MR signal, magnetic field gradients are used, which are provided by gradient coils. Gradient coils produce a linear variation of the magnetic field as a function of the spatial position (X-, Y-, and Z-direction) modifying the resonance frequency for spatial encoding of the MR signal.

Typical parameters governing the performance of the gradient system are the peak gradient strength, the rise time, and the slew rate. The peak gradient strength is the maximum achievable gradient strength (in mT/m). Typical peak gradient amplitudes in preclinical scanners are in the range of 200–1000 mT/m. Clinical scanners used for preclinical MRI of the kidney provide maximum gradient amplitudes ranging between 45 mT/m and 82 mT/m. The gradient rise time is the time needed to change the gradient field from zero to the peak amplitude (in ms). The gradient slew rate is the most important parameter, and refers to the speed at which a gradient can be switched. It is defined as the peak gradient strength divided by the gradient rise time (in T/m/s). This defines the maximum scanning speed, since it influences the minimum achievable repetition time (TR) and echo time (TE), as well as the echo spacing in fast imaging techniques such as RARE or EPI (see Note 2). Typical slew rates in preclinical scanners are in the range of 640–9000 T/m/s. Clinical scanners used for preclinical MRI of the kidney offer slew rates ranging between 150 T/m/ s and 200 T/m/s. Gradient coils can be driven in two modes: (a) constant rise time, in which the slew rate is adapted for each gradient amplitude and (b) maximum slew rate, where the rise is adjusted for each gradient amplitude.

The duty cycle defines the amount of time during which the gradient system can be run at maximum power amplitude (in % time). This parameter depends not only on the gradient system specifications mentioned above but also on the specific MR imaging technique used. Typical duty cycles in existing preclinical scanners can reach almost 100%. For this purpose, gradient coils are watercooled to ensure proper heat dissipation and temperature management.

The accuracy of the spatial representation depends on the quality of the linear magnetic field gradients and on the homogeneity of the static magnetic field (typically given in parts per million, ppm). Any deviation will appear as object distortion in the image, and thus a correction procedure called "shimming" is typically performed. During system installation, the manufacturer places small metal plates inside the scanner bore to improve the uniformity of the main magnetic field in a static and permanent manner—this is "passive" shimming. The distribution of these ferromagnetic pieces is adjusted over several iterations until the achieved magnetic field uniformity meets the vendor's specification. In addition to passive shimming, resistive shim coils are used for B0 homogenization. These adjustments are performed with a shim calibration tool on an empty magnet, and therefore the B0 homogeneity becomes impaired as soon as a sample or subject (human or animal) is placed inside the magnet bore.

An "active" shimming procedure can be conducted during measurement adjustments with every MR scan. This is crucial for techniques such as fat suppression or MR spectroscopy, since the introduction of any matter—an object or animal—inevitably introduces local magnetic field perturbations of the main static magnetic field, which then must be actively compensated for. Especially at the boundaries between tissue and air-filled areas, where magnetic field distortions are pronounced due to the different magnetic susceptibility which can cause severe signal losses and image distortions. Corrective magnetic fields produced by electrical RF coils, designated as shim gradients [20] which are superimposed upon the main magnetic field. Typically, the shim gradient system contains a set of first-order gradients which impose linear field changes in the X-, Y-, and Z-directions, and some second order gradients, which create spherical harmonics in the Z2 , ZX, ZY X2 -Y2 , and XY directions [21]. Some MR systems are also equipped with more sophisticated shim gradients which provide third and fourth order correction terms [22].

Depending on the specific pulse sequence and imaging technique used, the scientist has to adjust these magnetic field corrections to meet certain quality prequisites. Especially with T2\*-weighted sequences (e.g., gradient-echo or EPI sequences) the image quality improves with improved B0 homogeneity. MR spectroscopic methods are even more demanding [23]. Usually, state-of-the-art MR systems provide automated methods to perform the B0 shimming task, but the user still needs to check the performance and correct the shim settings manually if necessary.

For B0 shimming a global approach covering the entire FOV, or a localized shim covering a target area can be used. In general, the smaller the region of interest the better the achievable B0 homogeneity. In the case of renal MRI, the B0 homogeneity achieved after shimming across one kidney will be better than that for the whole abdomen, with air-filled lungs and gut. In the case of tissues that are moving due to respiration, cardiac activity or peristaltic motion, the magnetic field distortions change over time, which require different strategies. The time-averaged global shim would be the fastest and easiest solution, if this is sufficient. Alternatively, shimming—and also data acquisition—would be performed in synchrony with the breathing by using a trigger signal from a physiological monitor [24].

### 4 Radio Frequency (RF) Coils

Besides the static magnetic field and the spatial encoding with the gradient system, MRI requires transmission of RF pulses and reception of the MR signal from the sample or subject.

For signal excitation, the RF transmit resonator, or RF coil, sends RF pulses with a particular frequency, bandwidth, shape, duration, amplitude, and phase modulation to tilt the net magnetization vector oriented in parallel with the main magnetic field during equilibrium. B0 (typically assigned as the Z-direction), is tilted into the perpendicular transverse (X-Y) plane by a given flip angle α. After excitation, the magnetization precesses back to equilibrium and induces a signal in the RF receive coil (for a comprehensive review see [25]).

Depending on the transmission and reception capabilities, three general types of RF coils are commonly used: (1) transmitreceive (Tx/Rx) or transceiver RF coils, (2) transmit-only (Tx) RF coils, and (3) receive-only (Rx) RF coils. These generally have two types of geometry: (1) volume RF coils and (2) surface RF coils. Both RF coil types and geometries determine the amount of tissue or organ coverage and sensitivity to the MR signal.

Volume RF resonators are designed to produce a uniform RF field over a large field-of-view (FOV), and are typically used as transmit-receive RF coils or transmit RF coils (using a surface RF coil for reception). The most widely used type of volume RF coils in both human and preclinical MRI are birdcage resonators [26, 27] driven in quadrature mode. This means that two spatially orthogonal RF channels are used for RF transmission as well as signal reception. This RF coil configuration provides a 41.4% SNR gain over simple single channel or "linear" RF coils, due to the coherent signal increases which scale linearly with the number of channels, while the incoherent noise increases with the square root of the number of channels. For example, Chang et al. [28] used this RF coil setup in a renal hypoxia study in a mouse model after myocardial infarction of different sizes using BOLD MRI. Other types of volume resonators include saddle RF coils, Helmholtz RF coils, and the transverse electromagnetic (TEM) resonators, among others.

Smaller anatomical coverage and FOVs, but enhanced SNR, are provided by surface RF coils. Due to the principle of reciprocity, the closer the position of the RF coil is to the tissue of interest, the higher is the attainable SNR. However, the sensitivity drops with increasing distance to the RF coil surface, which thus results in a reduced FOV, and an inhomogeneous sensitivity profile [29]. Within a short distance of typically up to ~1 cm, a surface RF coil outperforms a volume RF coil in terms of sensitivity. Since the kidneys are located beneath the skin, surface RF coils positioned directly above are used for optimal signal reception in renal MRI. This is advantageous not only for X-nuclei MR imaging (see Subheading 5) but also for other applications with low SNR, such as perfusion or diffusion-weighted imaging.

Surface RF coils may be used as transmit/receive devices or as receive-only RF coils in conjunction with a volume RF coil used for homogeneous RF transmission. The first option is technically very simple since only one RF coil is necessary, which occupies less space. In the case of small animal imaging, it is most likely a simple singleturn loop with at least two tunable capacitances for tuning and matching. Images will show a strong spatially varying SNR decay, due to the addition of the inhomogeneous excitation and reception profiles [30]. This effect may be mitigated, for example, by using adiabatic excitation RF pulses [31].

The second option involves the use of two RF coils (Tx: volume RF coil, Rx: surface RF coil) which are tuned to the same resonant frequency. To avoid RF coupling between both coils, the surface RF coil must be detuned during transmission while the volume RF coil must be detuned during reception. Thus, both RF coils must include an option, using active and passive diodes, to allow for tuning and detuning during the MR experiment. This is the most common setup, since it improves the SNR in comparison with a Tx/Rx volume RF coil, while still maintaining a homogeneous B1. For example, Hu¨per et al. [32] used a four-element quadrature surface receive RF coil in combination with a transmit volume RF coil to investigate renal perfusion in acute and chronic renal allograft rejection in translational mouse models. The use of the surface RF coil improved the SNR from 15 (with a standard transceive volume RF coil) to 35.

Further SNR gains in small animal MRI are achieved with arrays of surface RF coils for signal reception [33], which are typically used together with a volume RF coil for signal excitation (see Notes 3 and 4). The use of RF coil arrays for brain imaging is especially widespread, and these are routinely offered by manufacturers. State-of-the-art small-animal MRI systems are usually equipped with 4–16 Rx channels, to which custom-tailored RF coils may be connected. Less popular are metamaterial-inspired RF coils, which have been developed for superficial tissue imaging in order to provide the best compromise of high SNR over a large FOV, while using only one receive channel [34].

Depending on the size of the sample and the RF coil, and their corresponding temperatures and resistances, their noise electronic contribution varies [35–37]. This factor is even more crucial in preclinical applications where the noise contribution of the RF coil, capacitors, sample, shield, receive electronics, and transmission lines are important [26]. This fact is exploited in small-animal imaging by the use of cryogenically cooled (cryo-cooled) RF coils. By lowering the temperature of the receiver chain, the noise contributions corresponding to temperatures and resistances can be reduced [38–40], thus achieving a significant increase in SNR, which has been reported in the literature (for a review see [15]). Unfortunately, only negligible effects are achieved in human MR [15, 41]. With increasing size of the sample, the relative contribution to the noise increases, and eventually becomes the dominating noise source—this is the default situation for human MR. Thus, only small animal MR applications will benefit from this cooling approach, since the noise contributions from the sample and RF coil are in the same order of magnitude. The SNR gain achieved with such cryo-cooled RF coils is higher at lower static magnetic fields [15, 26, 42]. A recognized limitation of cryo-cooled surface RF coils is that only transceive capabilities are supported, resulting in nonuniform excitation and reception profiles, which might hamper quantitative measurements.

Another option for the small-animal MR imaging of the kidney is the use of clinical scanners together with RF coils tailored for the human hand/wrist or knee. These are typically 8- to 16-element RF coil arrays which provide quasi-homogeneous excitation and reception profiles. Although these RF coils have a small size, they are not optimized for small-animal imaging and are not designed with matching/tuning capabilities. Moreover, RF coils tailored for rodent imaging are designed for a broad range of loadings (i.e., from almost empty to full loading, with a typical frequencydiameter product in the range 2–30 MHz · m [26]), in comparison to the clinical setups. Therefore, these clinical RF coils will achieve a lower efficiency than that of dedicated RF coils designed for preclinical MR systems. Alternatively, RF coils tailored for renal MRI of small rodents provided by RF coil manufactures can be connected with clinical MR scanners to enhance the fill factor and SNR.

In the following subsections, examples of possible RF coil setups are shown from different vendors. A list of the available MR system and RF coil vendors can be found in Table 2.

4.1 Commercially Available RF Coils Suitable for Kidney MRI in Rats

Due to the considerable anatomical homology with humans, rats have become one of the most commonly use rodents in animal research, particularly, in preclinical MRI.  Their genetic similarity to humans has made them extremely useful, allowing the development of mouse models of many human diseases. Mice have an average weight of 15–40 g, making them too small for standard human RF coils:

Clinical System

– There are no standard human RF coils that are small enough to be suitable for mice MRI. Alternatively, a Tx/Rx quadrature volume resonator (MRI.TOOLS GmbH, Berlin, Germany), supporting an ID ranging from 30 mm to 45 mm and magnetic field strengths of 3.0 and 7.0 T can be used for renal MRI with clinical MR scanners (Fig. 4A).

4.2 Commercially Available RF Coils Suitable for Kidney MRI in Mice

#### Table 2 List of vendors of MR equipment and RF coils

Preclinical System

Fig. 3 Example RF coils suitable for kidney MRI of the rat. (a) Siemens 8-channel wrist RF coil, (b) Siemens 15-channel knee RF coil, (c) Bruker quadrature rat body RF coil with inner diameter (ID) 72 mm, (d) MRI.TOOLS GmbH quadrature rat body RF coil with ID of 72 mm, (e) 4 - 1 element Bruker rat cardiac Rx surface array RF coil, (f) Bruker rat body Tx volume resonator with ID 72 mm.

### 5 Multinuclear Imaging

The most frequent MRI method in biomedical research is proton ( 1 H) imaging. Hydrogen is the most abundant nucleus in living organisms. It also is the nucleus which is the most sensitive to interactions with an external magnetic field, due to its high gyromagnetic ratio (γ ¼ 42.576 MHz/T).

Fig. 4 Example RF coils suitable for kidney MRI of the mouse. (a) MRI.TOOLS GmbH quadrature mouse body RF coil with an ID ¼ 42 mm, (b) Bruker quadrature mouse body RF coil with 35 mm ID, (c) Bruker 2 - 2 element mouse cardiac surface array RF coil, (d) Tx/Rx quadrature surface RF coil with cryo-cooling technology (CryoProbe)

In preclinical renal MRI, structural and functional information is typically investigated using in vivo 1 H MR imaging of the kidney. Techniques such as T2/T2\* mapping, arterial spin labeling (ASL), dynamic contrast-enhanced imaging (DCE-MRI) or diffusionweighted imaging (DWI) are performed in order to study the oxygenation, perfusion, clearance rate, and water and metabolite motion within the kidneys.

Physiometabolic probing is supported by imaging and spectroscopic applications of other nuclei such as 13C, 17O, 23Na, 35Cl, 39K, 31P, and 19F [43]. The detection of X-nuclei is challening because of the low intrinsic SNR due to the smaller γ, compared to 1 H, and by the lower abundance of these nuclei in biological tissues.

Increasing B0 is the most effective way to increase sensitivity the benefits of higher B0 are unique metabolic and functional information, as well as outstanding anatomical detail [44– 47]. X-nuclei MR also benefits from the increase in SNR achieved with cryo-cooled transceive surface RF coils, which are currently available for 13C, 31P, 19F, and 1 H [48, 49].

Changes in the biodistribution of endogenous metabolites that can be detected with 23Na- or 31P-MR have been demonstrated to be markers of renal disease [50–53]. For more details on preclinical 23Na-MRI protocols, see the chapters by Grist JT et al. "Sodium ( 23Na) MRI of the Kidney: Basic Concept" and "Sodium (23Na) MRI of the Kidney: Experimental Protocol." Similarly, renal perfusion, oxygenation, and inflammatory cell infiltration can be investigated using 19F-MRI [54–57]. See the chapters by Grist JT et al. "Sodium (23Na) MRI of the Kidney: Basic Concept" and "Sodium ( 23Na) MRI of the Kidney: Experimental Protocol" on how to perform 19F-MRI renal studies in rodents.

Hyperpolarization techniques have great potential for explorations into renal metabolic diseases (e.g., hyperpolarized 13C-MRI), where SNR is boosted up to a 20,000-fold compared to conventional 13C-MRI [58]. Experimental protocols to perform hyperpolarized 13C-MRI in preclinical studies are provided in the chapters by von Morze C et al. "Hyperpolarized Carbon (13C) MRI of the Kidneys: Basic Concept" and by Laustsen C et al. "Hyperpolarized Carbon (13C) MRI of the Kidney: Experimental Protocol."

### 6 Physiological Monitoring

A major difference between clinical and preclinical MRI is the need to anesthetize the animals prior to imaging. Since most anaesthetic drugs produce physiological changes (temperature decrease, depression of cardiorespiratory function, etc.), monitoring the anaesthetized animals is a very crucial aspect in preclinical MRI studies. Even in the most simple case, that is, the scanning of spontaneously breathing animals without any further intervention or treatment, respiration and temperature of the animal has to be monitored continuously during the scan in order to guarantee the appropriate anaesthetic status and to avoid hypothermia. In particular, hypothermia is a critical issues in rodent imaging, since mice have a low mass vs. a large surface which promotes heat loss [59, 60]. This problem is less severe in rats due to their larger weight, but it is not negligible [61]. Heating systems must be therefore used to maintain temperature homeostasis.

Physiological monitoring of the animal is even more essential in functional MR imaging, since changes may produce confounding effects on the outcome of the MR experiment [62]. Careful monitoring and recording of physiological data should be performed at all times during the experiment to ensure the validity of the results and their reproducibility. The frequency of physiological monitoring depends on the invasiveness of the experimental procedure as well as the time after induction of anaesthesia. As a rule of thumb, the more invasive the procedure and the longer the time under anaesthesia, the more frequent the physiological monitoring of the animal should be performed. Additionally, the correct positioning of the animals on the animal bed systems, along with careful intubation (if needed) are essential to ensure respiratory and circulatory comfort during imaging, and to promote recovery of the animal after the experiment.

Typical physiological parameters monitored during MR experiments include temperature, heart rate, and respiration rate. Many manufacturers already offer these three monitoring channels, and are ready to use these signals for the purpose of gating or triggering the MR imaging techniques. Additionally, information on blood pressure, blood oxygen saturation, transcutaneous pCO2, or pH can also be collected. In some cases, invasive withdrawal of blood samples is necessary—this is a very difficult task since the sampling has to be executed using long tubes with considerable dead volumes and very limited visual control.

Finally, all equipment must allow for remote physiological monitoring, and strictly comply with MR safety requirements. Several integrated animal handling solutions for physiological monitoring and anaesthesia are commercially available. Examples are given in Table 3.

### 7 Practical Points to Consider When Planning a Study

The aim of this chapter is to outline the main hardware components required for preclinical MRI of the kidney and to guide the appropriate hardware selection to meet the individual needs.

This section is intended to provide a summary of the prequisites needed and to help prepare for preclinical renal MR applications:

Preclinical?

ble, we recommend to use a dedicated ultrahigh field MR system tailored for small-animal imaging with a magnetic field strength of 7.0 or 9.4 T. The reason is that the SNR gain achieved with preclinical MR scanners compared to that at lower B0 is instrumental for achieving good spatial and temporal resolution, as well as for reducing the scan times needed to perform MR examinations.

7.1.2 Should I Consider Moving to a High Magnetic Field Strength? As discussed above (see Subheading 2), the use of a higher magnetic field strength results in an increase in SNR proportional to B0 (SNR / B1:65 0 ). This means that renal MR images will achieve greater SNR that can be traded off for to achieve higher spatial resolution or reduced scan time as needed (see Table 1). Also, higher B0 increases the BOLD effect. In general, all methods based on T2\*-weighting will benefit from an increase in the magnetic field strength. MR spectroscopy applications also profit from the increased B0. Therefore, we strongly recommend the use of a high B0 for these MR applications. However, not all facilities have such high field MR scanners readily available, in which case it might be advisable to perform the experiments using a lower magnetic field system, as opposed to moving animals, equipment, and personnel to another facility.