Keywords

1 Introduction

Magnetic resonance imaging (MRI), computed tomography (CT), ultrasound (US) and certain specialised optical imaging (OI) techniques can produce valuable medical images on the basis of endogenous contrast mechanisms. The nature and scope of such intrinsic contrast-generating properties is largely determined by the physical principles of the imaging technique and the way in which the imaging signal is generated and detected. However, many clinical imaging protocols involving MRI, CT, US and OI make use of contrast agents to enhance the specificity and/or the sensitivity of disease detection. Contrast agents are essential for nuclear imaging with positron emission tomography (PET) and single-photon emission computed tomography (SPECT), since these techniques rely exclusively on the measurement of the spatial distribution of gamma (γ) rays generated by radioactive decay of injected tracer compounds. Without these tracers, no (PET or SPECT) signal is generated, which also explains why nuclear imaging necessitates the use of an additional imaging device to provide anatomical and, therefore, x-ray attenuation information. In most cases, CT is used for that purpose. Most optical imaging in biomedical research also requires the injection of exogenous contrast material to generate usable signal.

The molecular properties of contrast materials are dictated by two main factors, i.e. the imaging method for which they are going to be used and the biomedical application, which they should address. Each imaging method has its own characteristics in terms of how signal is generated and detected, and this has major consequences for the optimal design of the contrast material. On the other hand, the molecular properties of the contrast agent, including its size, shape, chemical composition and charge, have a major influence on their in vivo pharmacokinetic and pharmacodynamic properties. For these reasons, the optimisation of contrast material design should involve an integrated approach in which all above factors are considered in a balanced manner, so as to arrive at a design that is most suitable for a (range of) specific application(s). Design criteria must include the use of materials that are safe and affordable if the agent is ultimately to be used in humans.

This chapter will provide an overview of the prime design criteria for diagnostic contrast materials. In view of the overwhelming breadth of contrast-enhanced (bio)medical imaging, we have decided to focus the overview on MRI contrast agents and, more specifically, on a particular class of MRI contrast agents, i.e. dendrimers. Dendrimers have been extensively studied, and they represent a unique class of materials in that they can be synthetised in a very broad range of sizes while keeping the chemical composition essentially the same. By tuning the size and other molecular characteristics, the dendritic agents can either be designed to act as low-molecular-weight contrast materials or behave like high-molecular-weight nanoparticles. The specific focus on dendrimers aims to provide a basic understanding of the relation between molecular properties and in vivo behaviour, which offers important handles for design optimisation. Many findings with dendrimers have broad implications and can be generalised to other contrast agents, not just for MRI but also for the other imaging methods. When appropriate, we will make a side step and describe hallmarks of contrast materials for other imaging modalities as well. James and Gambhir (2012) have provided an excellent in-depth overview of recent advances in the design of contrast materials and the strengths and limitations of the different biomedical imaging modalities. For more specific reviews of the state of the art in the design of contrast materials for biomedical imaging, see Geraldes and Laurent (2009), Mulder et al. (2013) and Nicolay et al. (2013) for MRI; Li et al. (2014) for CT; Licha and Resch-Genger (2011) and Khemthongcharoen et al. (2014) for OI; Huang et al. (2015) for PET; Gnanasegaran and Ballinger (2014) for SPECT; and Kiessling et al. (2014), Wen et al. (2014) and van Rooij et al. (2015) for US.

2 Desired In Vivo Behaviour of Diagnostic Contrast Materials

Ideally, shortly after injection, diagnostic imaging probes would very specifically distribute to the target of the imaging study, e.g. sites of disease activity, and otherwise be excreted reaching low levels in the background. This would create the best opportunities for sensitive as well as specific detection of the location and severity of the disease process. Following the imaging study, the probe should preferably be eliminated rapidly from the entire organism including the target. Clearance ideally occurs via the kidneys into the urine as this represents the fastest elimination pathway. It goes without saying that the contrast material should be biocompatible, should be stable for the entire duration of the imaging study and should not cause any adverse effects.

In reality, the above ideal situation does not hold. In fact, the optimal design of contrast materials for diagnostic biomedical imaging depends on many parameters, in a very complex fashion, and usually represents a compromise between mutually conflicting factors. Unfortunately, most of these factors can only be dealt with in an empirical (partly trial and error, and therefore time consuming) fashion, as there is currently a lack of quantifiable criteria that can rationally and time efficiently steer the process of contrast agent development. This paragraph is intended to provide some feeling for the parameters that play a dominant role in determining the efficacy of a contrast agent for a certain diagnostic task. For further reference, the reader is referred to specialised literature. As examples, Geraldes and Laurent provide a comprehensive classification of MRI contrast agents (Geraldes and Laurent 2009), while Chen and Chen (2010), Choi and Frangioni (2010), Rosenblum et al. (2010), Kobayashi et al. (2011) and Wu et al. (2015) present an in-depth description of the design criteria for guiding the development of probes for (molecular) imaging. James and Gambhir (2012) also portray the dominant probe design considerations.

First and foremost, the probe should be capable of generating sufficiently strong signal (or signal change) on the diagnostic images, as acquired with the modality of choice. This criterion is typically expressed as the signal-to-noise ratio (SNR) and/or contrast-to-noise ratio (CNR), depending on the imaging modality, which is achieved upon probe injection. The concentrations of the probe, which are needed to produce a sufficiently high SNR or CNR, differ strongly among the different imaging modalities. This is caused by the very different intrinsic sensitivities of each modality. PET imaging can detect picomolar concentrations of an agent, while MRI usually needs on the order of micromolar concentrations. PET is therefore considered a tracer technique, while MRI definitely is not. The multiple orders of magnitude difference in intrinsic sensitivity among modalities implies that biocompatibility and stability of contrast material is much less of an issue in the case of PET than with MRI since the doses of the agent will be dramatically lower. However, PET is associated with exposure to ionising radiation, in particular when combined with CT, as is common practice since PET lacks anatomical information. In contrast, MRI does not involve the use of ionising radiation. For these reasons, the optimal choice of imaging probe and imaging technology is typically a balance between several pros and cons. Obviously, not only the above factors but also the information sought after by the imaging study plays a key role in guiding the choice of probe and imaging modality.

The above issue of detecting sensitivity is also strongly influenced by the question whether sufficient quantities of contrast material can reach the intended target site. As contrast agents are most often injected intravenously, they will first be distributed in the vascular space. What ideally happens next is largely dependent on the goal of the imaging study. Most imaging-based tumour diagnoses require that the contrast agent leaks from the vascular compartment into the extravascular, intratumoural space. This implies that the probe should (i) be sufficiently small to cross the tumour vascular endothelium, (ii) circulate in the blood sufficiently long to allow suitably high intratumoural levels to be reached for generating a significant imaging signal and (iii) persist in the tumour compartment sufficiently long while being essentially completely cleared from the circulation, in order to achieve an as high as possible tumour-to-background ratio on the images. These factors are governed by the biodistribution of the imaging agent over time and are often collectively considered under the absorption, distribution, metabolism and excretion (ADME) concept, which is a leading theme in pharmacology. ADME describes the disposition of a foreign compound within a living organism, and, as such, knowledge of the ADME properties of a contrast material may provide strong leads for optimising its molecular properties for the desired application. As will be shown further on, ADME information typically aids to define rules of thumb for guiding contrast agent design rather than providing quantitative handles for predictable, fully rational design optimisation.

Most of the insights in in vivo contrast agent properties have been collected by extensive studies in small animals, particularly mice. However, contrast agent behaviour is different in humans. Distribution volumes are obviously much bigger in humans, while the rates of physiological processes are usually much slower. Nevertheless, the extensive small animal studies have yielded a wealth of information on many aspects of contrast agent behaviour and as such offer an essential starting point for the translation of novel contrast agent concepts from the preclinical to the clinical setting. The most relevant factors governing diagnostic probe design will be highlighted next.

3 Basic Design of Dendrimers for Diagnostic Imaging and Therapy

As stated above, the focus of this chapter is on dendrimer-based contrast materials. Dendrimers have been very extensively studied, and this has led to a reasonably coherent set of design criteria. This paragraph describes the molecular architecture of dendrimers and features general aspects of their usefulness in diagnostic imaging, with an emphasis on MRI, as well as therapeutic applications. More specific qualities are presented in upcoming sections. For papers on the basics of dendrimer design and biomedical uses, see Bosman et al. (1999), Boas and Heegaard (2004), Barrett et al. (2009), Bumb et al. (2010), Menjoge et al. (2010), Mintzer and Grinstaff (2011) and Longmire et al. (2011, 2014).

Dendrimers are a class of macromolecules that are characterised by branching polymerisation, which results in a three-dimensional structure (Fig. 11.1). The term “dendrimer” is inspired by the Greek words “dendron”, which means “tree”, and “meros” meaning “part” and refers to the basal structure of dendrimers: a central core from which numerous polymerised branches extend out to the molecule’s periphery. The structure of dendrimers consists of three main elements: the core, interior and shell. The core has major effects on the three-dimensional shape of the dendrimer, i.e. spherical, ellipsoidal or cylindrical. The architecture and composition of the interior part affects the host–guest properties. The surface can be further polymerised or modified with an enormous diversity of functional peripheral groups. The stepwise addition of polymeric elements from the core outwards creates monodisperse molecules that differ only in size with each successive layer, which are referred to as a “generation” (i.e. G1, G2 and G3) (Fig. 11.1). This remarkable flexibility in the molecular structure, composition and size of dendrimers explains the enormous interest they have generated for use as carriers for imaging and therapeutic agents. A very important attribute of dendrimers is their multivalency, with the number of surface end groups exponentially growing with each generation. Dendrimers for imaging are most often prepared by chemical conjugation of imaging labels to the end groups. Paramagnetic dendrimers, suited for 1H-MRI, are synthesised by coupling DOTA or DTPA moieties to the end groups, followed by incorporation of Gd3+ ions (Fig. 11.1). Chelation of the paramagnetic Gd ion via high-stability complex formation with, for example, DOTA or DTPA is needed, since free Gd3+ ions are highly toxic (Aime et al. 1998; Sherry et al. 2009). The application of Gd-labelled dendrimers as MRI contrast agents is primarily based on shortening of the T1 relaxation time, one of the prime MRI contrast parameters. The presence of the paramagnetic agent can either be detected with the use of T1-weighted MRI (causing a higher signal intensity as compared to the pre-contrast scan) or more quantitatively assessed with T1 parameter mapping. In the latter case, multiple MR images are acquired with different degrees of T1 weighting, followed by pixel-wise fitting of the data to estimate the T1 relaxation time. The degree of T1 shortening can be used to estimate the local contrast agent concentration. This conversion of T1 change to contrast agent concentration requires knowledge of the molar relaxivity, r1, in units of mM−1s−1, which can be obtained by measuring the T1s of a contrast agent dilution series in aqueous buffer (Strijkers et al. 2007). Gd-DTPA, Gd-DOTA and related low-molecular-weight Gd chelates have very similar r1s in the order of 1–4 mM−1s−1, depending on the magnetic field strength (for further details, see Sect. 11.4) (Nicolay et al. 2013; Aime et al. 1998). Dendrimers tend to have higher relaxivities than simple Gd chelates.

Fig. 11.1
figure 1

Architecture of dendrimers. (Top) Schematic structures of different generation dendrimers, starting with the basic core element (= G0, here with 3 functional groups) and branching further out towards the generation 5 material (= G5). For each generation, two new branches are added to each existing arm doubling the number of functional end groups. (Bottom, left) Build-up in three dimensions of a dendrimer, starting from the inner core and increasingly extending the dimensions and the number of surface functionalities with every next generation. (Bottom, right) Schematic example of a paramagnetic Gd-DTPA- or Gd-DO3A-conjugated G3 PPI dendrimer for contrast-enhanced MRI. The Gd-containing units are conjugated to the distal ends of the dendrimer. A variety of conjugation methods and a range of conjugated moieties (e.g. PEGs) can be chosen. The structures shown are schematic example pictures only. For details on employed dendritic molecular structures, see the specific literature references

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The multivalent nature of dendrimers (Fig. 11.1) is also widely exploited to synthesise dendrimers with multiple appending ligands (for achieving high-affinity binding to specific molecular targets). The options are essentially endless. Furthermore, dendrimers are suitable as drug carrier devices, as a wide range of therapeutic agents, including DNA and RNA, can be incorporated in the molecule’s interior (Barrett et al. 2009; Baker 2000; Kannan et al. 2014). Several dendrimer formulations of therapeutic agents are in clinical trials (Kannan et al. 2014). A docetaxel–dendrimer conjugate is in a phase 1 clinical trial for advanced or metastatic cancer. The dendrimer’s molecular architecture also readily allows for combinations of imaging and therapy (Kannan et al. 2014) in theragnostic (or theranostic) applications, in which diagnostic imaging converges with therapeutics (Loudos et al. 2011). The focus in the rest of the chapter will be on the use of dendrimers as contrast agents for diagnostic imaging.

4 The Effects of Contrast Agent Size

The ability to control the molecular structure of dendrimers generates quite unique flexibility to tune their size while minimally affecting the overall chemical composition. This aspect continues to be one of the main assets of the use of dendrimers for diagnostic imaging, as contrast agent size is a major factor controlling its behaviour in vivo and hence its usefulness in diagnostic imaging. Below we will highlight some of the main manifestations of dendrimer size effects in diagnostic MRI. There are many eloquent reviews on this topic (Menjoge et al. 2010; Longmire et al. 2008, 2011, 2014; Kobayashi et al. 2010a).

A very illustrative example of the consequences of the variation of the size of paramagnetic Gd-DTPA-labelled dendrimers for contrast-enhanced MRI is shown in Fig. 11.2. Langereis et al. (2006) report the preparation of G0, G1, G3 and G5 poly(propylene imine) (PPI) dendrimers and their use for in vivo whole-body contrast-enhanced MRI in mice. The dendrimer contrast materials, which had molecular weights ranging from 0.7 to 51.2 kDa for the G0 and G5 dendrimers, respectively, were compared to Gd-DTPA (0.6 kDa) as the non-dendrimer control. The number of Gd ions per molecule ranged from 1 for Gd-DTPA and the G0 dendrimer to 4, 16 and 64 for the G1, G3 and G5 dendrimers, respectively. First, the materials were characterised by ionic relaxivity measurements at a field strength of 1.5 tesla (Table 11.1). Measurements were done in mouse plasma to mimic the conditions in the blood compartment. Interestingly, on a per Gd ion basis, the ionic r1 relaxivities increased strongly from circa 4 mM−1s−1 for Gd-DTPA to about 19 mM−1s−1 for the G5 material. The ionic r2 relaxivities, which describe the contrast agent-induced shortening of the T2 relaxation time (like T1 a fundamentally important MRI contrast parameter) (Strijkers et al. 2007), were shown to increase in similar proportions with increasing molecular weight. The r2 increased, however, even more significantly, reaching values of about 25 mM−1s−1 for the G5 agent, where this value is on a per Gd ion basis. These pronounced increases in r1 and r2 with dendrimer generation can be explained on the basis of their higher molecular weights. The larger molecular size causes the correlation times characterising the rotational mobility of the molecules to increase in proportion. This enhances the efficacy of the T1-shortening magnetic interactions, which reaches an optimum when the time scale of the dendrimers’ molecular tumbling is similar to the resonance frequency of the MRI scanner, i.e. around the Larmor frequency. It turns out that in many macromolecular systems, including higher-generation dendrimers, the molecular mobility quite closely matches the Larmor frequency corresponding to a magnetic field strength around 1 tesla. This explains why graphs depicting the ionic relaxivity r1 as a function of magnetic field display a characteristic (macromolecular) peak around this field strength (Nicolay et al. 2013; Aime et al. 1998; Strijkers et al. 2007). Longmire et al. (2014) have reported r1 and r2 relaxivity profiles of G6 PAMAM dendrimers as a function of magnetic field strength. The authors indeed observed the characteristic peak in the r1 profile around 1 tesla, followed by a drastic reduction at higher magnetic fields. The maximal r1 values at the optimal field strength as reported by Longmire et al. (2014) are similar to those found by Langereis and coworkers (Langereis et al. 2006). A very important study by Ogawa et al. (2010) reported that the architecture of paramagnetic Gd-loaded dendrimers has a major effect for their T1-shortening efficacy. Similarly sized lysine-grafted dendrimers, in which the Gd chelates are partly buried in the dendrimer interior, had lower r1 values than those observed in traditional poly(amidoamine) (PAMAM) dendrimers. The Gd chelates are localised on the dendrimer surface in the latter case, enabling the unhindered interaction between bulk water protons and the paramagnetic centres (Ogawa et al. 2010) that is a prerequisite for maximising the ionic r1 (Strijkers et al. 2009). Such differences in molecular arrangement of the Gd chelates in the lysine-grafted versus more regular PAMAM dendrimers also explain the different responses of the ionic r1 relaxivities to an increase in temperature. At higher temperatures, lysine-grafted paramagnetic dendrimers have higher r1 values, which are attributed to the enhanced accessibility of water protons to the internal Gd chelates. Temperature elevations rather cause a drop in r1 for PAMAM dendrimers with similar size due to a decrease in the rotational correlation time (Ogawa et al. 2010).

Fig. 11.2
figure 2

Whole-body contrast-enhanced MRI of mice at 1.5 tesla. Maximum intensity projections of digitally subtracted contrast-enhanced whole-body MR images show good visualisation of the cardiovascular system and the renal parenchyma directly after paramagnetic agent injection (8 s). In time, the lower-generation PPI dendrimer appears to accumulate in the kidney, whereas the higher generations do so in the liver and spleen. Eight minutes after administration of MRI contrast agent, the higher-generation PPI dendrimers accumulate in the renal collecting system of the kidney and to some extent in the bladder (Modified with permission from Langereis et al. 2006)

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Table 11.1 Longitudinal (r1) and transverse (r2) relaxivities of Gd-DTPA, the reference Gd-DTPA complex (G0) and different generations of Gd-DTPA-terminated PPI dendrimers (G1, G3 and G5) in mouse plasma at 1.5 T and 20 °C
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As indicated above, the macromolecular advantage of higher relaxivity r1 is lessened at high field strengths, which are most often used for small animal MRI (i.e. at magnetic fields of 7 tesla and beyond); the difference between paramagnetic Gd-labelled low- and high-molecular-weight agents is smaller. The above pattern of the MR relaxation dispersion profile, as the r1 field dependence phenomenon is also called, explains the recent interest in the use of relatively low-field MRI scanners for small animal contrast-enhanced MRI.

Figure 11.2 demonstrates that the dendrimer-induced signal enhancement on T1-weighted MRI scans of mice differs strongly among different PPI dendrimer generations (Langereis et al. 2006). Low-molecular-weight compounds, such as the G1 material, initially enhance the blood pool followed by a rapid signal increase in the kidneys and sometime later the bladder. This is due to the fact that this agent is cleared via the renal clearance pathway leading to elimination from the organism via the urine. The higher-generation dendrimers have very different distribution profiles, in which the liver (as well as the spleen) shows strongly increased signal, while the kidney and bladder are enhanced considerably less and with a significant time delay as compared to the G1 agent. Similar pronounced effects of dendrimer size on dynamic whole-body MRI scans in small animals have been frequently reported (Longmire et al. 2011, 2014; Kobayashi et al. 2003; Kobayashi and Brechbiel 2005; de Lussanet et al. 2005).

Much of the above results on the size dependence of dendrimer behaviour can be understood in terms of the size dependence of blood clearance via the kidneys. Choi and Frangioni (2010; Choi et al. 2007) have done a series of seminal studies with the use of quantum dots (QDs), brightly fluorescent nanomaterials that can be prepared in a wide range of sizes. They noted that QDs with hydrodynamic diameters < 5.5 nm were rapidly and efficiently excreted from the body via the urine. Larger-sized particles persist for much longer and are mainly eliminated by the reticuloendothelial system (RES) predominantly involving the liver and spleen, as is known to generally hold for nanomaterials that escape renal clearance (Choi and Frangioni 2010). It goes without saying that clinical translation of RES eliminated agents can only occur in case that their constituents are completely nontoxic and are eliminated from the body in a reasonably short period of time.

In analogy to the QD studies, several groups have extensively evaluated the biodistribution of dendrimers by quantitatively measuring dendrimer content in various tissues and organs at several time points after intravenous injection. Thus, Gillies et al. (2005) have measured the biodistribution of dendrimers, with molecular weights ranging from circa 20,000 to 160,000, in mice. The authors used poly(ethylene oxide) (PEO)-coated 125I-radiolabelled dendrimers and reported that these had circulation half-lives between 8 and 50 h for the lowest- and highest-molecular-weight materials, respectively. However, clearance from other organs, such as the liver, can take substantially longer. The next section will provide more details on the consequences of specific changes in the surface properties of dendrimers on their biodistribution behaviour.

Maeda and collaborators (2010, 2013) have made vital contributions to the understanding of the effects of molecular size as well as composition on the in vivo biodistribution and in particular the tumour homing of macromolecular agents for imaging and therapy. Tumour penetration of different materials is to a large extent governed by what Maeda called the enhanced permeability and retention (EPR) effect. The EPR phenomenon describes the extravasation of materials from the blood compartment into the extravascular, interstitial space of the tumour and the prolonged retention of macromolecular materials in the tumour interstitial space as compared to the situation in normal tissues. The EPR phenomenon, which among others is caused by the incomplete endothelial lining of tumour blood vessels, is widely exploited for tumour diagnosis as well as for the delivery of therapeutic agents.

The pronounced difference in MRI enhancement pattern among different generation dendrimers, as depicted in Fig. 11.2, is widely exploited for the characterisation of diseased tissues and in particular for tumour diagnostics. De Lussanet et al. used differently sized PPI dendrimers for tumour characterisation (de Lussanet et al. 2005) and employed tracer kinetic modelling of dynamic MRI enhancement data. Low-molecular-weight dendrimers were found to primarily report on microcirculatory blood flow characteristics, whereas the use of high-molecular-weight agents provided handles for quantifying vascular permeability and vascular volume.

Recently, Jacobs et al. (2015) reported a conceptually novel design of dendrimer-enhanced MRI scanning for improved tumour characterisation. Instead of assessing MRI signal dynamics upon injection of different dendrimer generations in different animals, Jacobs and collaborators performed sequential injections of three differently sized contrast agents and monitored MRI responses to each of these within single animals. Subsequently, the MRI data were subjected to tracer kinetic modelling by jointly analysing the signal dynamics for all three consecutive injections in one go. All three dynamic response curves were sampled over a time span of less than 45 min. Therefore, it could be assumed that tumour physiology, including tumour microcapillary perfusion, was identical throughout the duration of the session. This allowed for significant reductions in the uncertainty of modelling estimates of vascular permeability and blood volume fraction as compared to the traditional approach. Figure 11.3 shows typical examples of dynamic contrast-enhanced MRI (DCE-MRI) curves from a single pixel in the MR images upon injection of Gd-containing G5 and G2 dendrimers as well as Gd-DOTA, the control low-molecular-weight clinical Gd chelate. For details, see the legend in Fig. 11.3. The above multi-agent approach can potentially be used for the detailed evaluation of anti-vascular tumour therapies. For these studies of Jacobs et al., PPI-based G2 and G5 materials were prepared with hexa-ethylene glycol-spaced Gd-DO3A end groups, as opposed to using materials with the Gd chelate directly attached to the dendrimer’s exterior groups.

Fig. 11.3
figure 3

Dynamic contrast-enhanced MRI of mouse tumours with the use of differently sized Gd-based contrast agents. (a) Schematic representation of the effect of contrast agent size on its leakage from the vascular compartment into the extravascular, extracellular space. In this study, G5 dendrimers, G2 dendrimers and the low-molecular-weight clinically approved gadoterate were compared in terms of contrast enhancement patterns. (b) Dendrimer schematic illustrating the conjugation of multiple Gd-DO3A moieties. (c) Typical changes in longitudinal relaxation rate, R1 (i.e. 1/T1), over time following the sequential injection of the three contrast materials. Data points were taken from a single pixel in the tumour periphery, as depicted on the T2-weighted image (see inset). Note that the different materials, as expected, cause distinctly different enhancement profiles. The G5 dendrimer is only very slowly cleared from the circulation, explaining the absence of significant reduction in R1 change following the abrupt increase upon injection. (d) Extraction fractions for the different materials as deduced from mathematical modelling of the dynamic MRI data. Three consecutive measurements (n = 5) were done at 2-day intervals. The data show the reproducibility of the measurements and modelling findings for each of the materials by itself and that the extraction fractions are significantly lower for each of the dendrimers compared to gadoterate. Extraction fractions tended to be lower for the G5 than the G2 dendrimer; however, this did not reach significance. Statistical tests: ### P<0.001 between measurement day 2 and measurement day 3. ** P<0.01, *** P<0.001 between the different contrast agents. For further details, see text. MRI measurements were made at a field strength of 7 tesla (Reproduced with permission from Jacobs et al. 2015)

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One of the merging clinical applications of dendrimer-enhanced MRI is in lymphatic imaging, which was also shown to benefit from a judicious choice of dendrimer size. Despite the vital role of the lymphatic system in health and disease, there is a lack of effective tools for high-spatial resolution lymphatic imaging. An example of the utility of paramagnetic dendrimers for this sort of application can be found in Longmire et al. (2014). In this case, the contrast material was injected interstitially, and serial MRI scans were made over time to monitor the agent’s accumulation in the lymphatics. Kaminskas and Porter (2011) have presented evidence that lymphatic imaging is in general improved with dendrimer generation and that the dendrimer’s molecular composition strongly affects the rate of drainage and retention profiles. Opina et al. (2014) have employed a Gd-loaded G5 PAMAM dendrimer preparation and have shown that this material is very effective for use in lymphatic imaging. However, the authors noted that the material is only very slowly cleared from the body and that this may obstruct its eventual clinical translation for lymphatic visualisation.

5 The Effects of Surface Characteristics

The effects of modifications of the surface properties of dendrimers on their in vivo characteristics and utility as contrast agents have been extensively studied. The consequences of surface modifications can to a large extent be ascribed to alterations in the interactions between dendrimers and blood constituents. It is important to stress that many of the findings reported below do not exclusively apply to dendrimers but are equally valid for a wide range of other nanoparticles, including liposomes, micelles and microemulsions that are also extensively studied for their diagnostic and therapeutic usefulness (Mulder et al. 2006, 2009; Hak et al. 2009; Langereis et al. 2013, 2014).

The best-studied surface modification of dendrimers is PEGylation, i.e. the covalent chemical coupling of poly(ethylene glycol) (PEG) chains to the distal ends of dendrimer side groups. PEG modification could enhance the utility of dendrimers for two reasons. First, it is known to reduce dendrimer toxicity, in particular dampening the haemolytic properties of certain dendrimer materials, particularly those that have primary amine end groups (Malik et al. 2000; Tack et al. 2006; Gajbhiye et al. 2007). Secondly, it causes the dendrimers to escape from RES recognition resulting in reduced plasma clearance and prolonged circulation half-life (Kobayashi et al. 2001a, b). Elimination of nanoparticles like dendrimers from the circulation by the RES is initiated by serum opsonins that attach to the surface of the agents. This is followed by phagocytosis by macrophages and sequestration of the foreign materials in the liver and spleen. PEGylation in a way camouflages the nanoparticles, thereby preventing adhesion of opsonins so that the nanomaterials remain in circulation and temporarily evade the RES. This is usually referred to as the stealth effect of PEGylation. PEG chains achieve the reduced opsonisation by a combination of steric and hydration effects. The lengths of PEG chains can essentially be varied at will, and the effects that such variations have on dendrimer properties have been widely studied (Kojima et al. 2011).

Figure 11.4a serves to illustrate the prominent effects of PEG conjugation on the blood clearance kinetics of a G4 PAMAM dendrimer material (Kobayashi et al. 2001b). Coupling of one or two 20,000 Da PEG chains causes a strong prolongation of the blood residence times. Likewise, the PEGylation has a remarkable effect on the biodistribution of the material over the main organs (Fig. 11.4b). Whole-body T1-weighted MRI scans of mice injected with non-PEGylated or two-chain PEGylated dendrimers (Fig. 11.4c, d, respectively) are in line with the biodistribution data: PEGylation causes the blood signal to be brighter for a longer time while showing a similar enhancement of the kidneys and a slightly increased disposition in other organs such as the liver and spleen.

Fig. 11.4
figure 4

The effect of PEGylation on dendrimer biodistribution properties. (a) Blood clearance of 153Gd-labelled G4 PAMAM dendrimers, either conjugated with no, one or two 20 kDa PEG chains, in normal nude mice. The data are expressed as the mean percentages of the injected dose per gram of blood ± SD (n = 4 per time point). PEGylation leads to prolonged blood circulation due to the fact that the material becomes more hydrophilic. (b) The distribution of the same three materials was determined at 15 min after injection. The data are expressed as the mean percentages of the injected dose per gram of material ± SD (n = 5 per time point). The dendrimer with two PEG chains remained in the blood significantly longer and accumulated significantly less in the liver and kidney than the other two materials. Overall, the positive effects of PEG conjugation on the dendrimer contrast materials were found to have prolonged retention in the circulation, increased excretion and decreased organ accumulation. (c, d) Typical examples of whole-body 3D MRI of mice injected with 0.033 mmol Gd/kg of dendrimer without (c) and dendrimer with two PEG chains (d). The images, which are negative displays with higher intensity darker than lower intensity, were obtained at 2 min (left) and 10 min (right) post-injection and are shown as maximum intensity projections. Darker blood and brighter kidneys were noted on both early and delayed images with the PEGylated versus the non-PEGylated material (Reproduced with permission from Kobayashi et al. 2001b)

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Kojima et al. (2011) have systematically investigated the effects of PEGylation of dendrimers on their MRI relaxivity properties. Generation 4 and 5 PAMAM dendrimers bearing Gd chelates were conjugated with two different PEG chains (i.e. 2 or 5 kDa). Conjugation of the 2 kDa PEG chains barely affected the r1 relaxivity, while 5 kDa PEG chains caused a reduction in r1 (and also r2). The highest r1 was measured for shorter PEG, higher-generation conjugates. These effects can most likely be explained by a combination of two counteracting factors. PEGylation may interfere with the access of water protons to the paramagnetic Gd3+ ions, reducing relaxivity. For the larger dendrimers, this is partly compensated for by the reduced molecular mobility that leads to higher r1. As indicated earlier, the positive effect on r1 of reduced mobility holds at clinical field strength up to circa 1.5 tesla and gradually vanishes towards the higher field strengths typical for most small animal MRI instruments.

It is important to note that not only the surface properties of dendrimers dictate their in vivo fate in terms of clearance pathways and biodistribution and thereby their utility for contrast-enhanced MRI. The chemical nature of the dendritic core also has major consequences. As an example, Kobayashi et al. (2001a) showed that G4 PPI dendrimers rapidly accumulate in the liver, whereas G4 PAMAM dendrimers act as vascular contrast material since they have much longer blood retention. It should be noted that the G4 PAMAM material is higher in molecular weight (and thus in size) than the G4 PPI material. The distal end groups of the two classes of dendritic materials were conjugated with chelating moieties to allow complexation of Gd3+ ions for MRI detection. The authors argued that the very different behaviour of the different dendrimer contrast materials is due to the overall more hydrophobic nature of the liver-seeking dendrimers.

It is worth noting that PEGylation of dendrimers causes increased solubilisation of a broad range of therapeutic agents in their core and also the PEG layers themselves afford drug-loading capacity (Gajbhiye et al. 2007). PEGylated dendrimers form stable complexes with plasmid DNA and achieve improved gene transfection compared to non-PEGylated dendrimers.

Zhang et al. (2014) have extensively studied the effects of conjugation of various functional groups to the surface of G5 PAMAM dendrimers, including the attachment of acetyl and succinamic acid groups, on the encapsulation and release of the anticancer drug doxorubicin. The nature of the functional group was shown to strongly affect these properties and thus represents an important handle for optimising the properties of dendrimers as drug delivery system.

6 The Effects of Shape and Flexibility

The shape of molecules in general is an important determinant of their functionality per se as well as their in vivo behaviour, and this similarly applies to dendrimers (Longmire et al. 2011). Indeed conjugation of two long, linear PEG chains to G4 PAMAM dendrimers led to the formation of an asymmetric elongated nanoparticle that underwent relatively rapid renal clearance. By contrast, when the same generation dendrimer was fully decorated with short PEG tails, no renal clearance was observed despite the similar physical size. Large-scale systematic studies into the shape phenomenon unfortunately are lacking as yet.

Gillies et al. (2005) have provided evidence that dendrimers with higher internal flexibility are excreted faster into the urine, as compared to materials with decreased flexibility and otherwise comparable properties. The authors prepared different generation PEO-conjugated dendrimers with the number of PEO arms ranging from two to eight. Comparison of the renal clearances for the four-arm versus eight-arm polymers indicated that the more branched polymers were excreted more slowly into the urine, a result attributed to their decreased flexibility.

As noted in Sect. 11.4, the translational and rotational mobility of a dendrimer-based MRI contrast agent as a whole affects its relaxivity properties, especially at lower magnetic field strength. However, also the local internal dynamics of a contrast agent can have profound effects on its relaxivities r1 and r2. One manifestation of this phenomenon is the aforementioned influence of temperature changes on r1. This partly has to do with the temperature dependence of water proton accessibility to the paramagnetic centres and partly with temperature-induced changes in local segmental rotational correlation times. Systematic studies that are backed up by measurements of the internal molecular dynamics of dendrimer contrast agents by independent techniques are currently scarce. Consequently, in-depth exploitation of these phenomena, let alone using them for directed molecular design, is lacking as well.

7 The Effects of Charge

The charge of dendrimers and of nanoparticles in general can significantly alter their in vivo characteristics. In the setting of drug delivery, for example, cationic nanoparticles typically deliver their therapeutic cargo more efficiently. This is partly related to the improved ability of cationic drug delivery systems for cellular internalisation and subcellular routing as compared to anionic and neutral delivery systems (e.g. see Asati et al. (2010)).

However, cationic dendrimers are known to be more cytotoxic (among others causing haemolysis and changes in red blood cell morphology) than anionic species, which generally are well tolerated in particular when equipped with PEG or other polymeric moieties (Malik et al. 2000). These effects are also dependent on dendrimer generation. For these reasons, anionic dendrimers are usually preferred for diagnostic imaging applications.

Jaszberenyi et al. (2007) have studied a number of different PAMAM and hyper-branched dendrimers that had either been conjugated with negatively charged Gd-DOTA or neutral Gd-DO3A chelates for MRI detection. Surprisingly, at field strengths below 1.5 tesla, the Gd-DOTA-conjugated material had twice as high r1 as the Gd-DO3A-conjugated counterpart. The authors attributed this remarkable difference to differences in the internal dynamics of the two materials. The Gd-DOTA agent has a higher rigidity, possibly due to electrostatic repulsive forces between the negatively charged end groups, which also increase the effective hydrodynamic size of the macromolecule. Interestingly, these authors found that the attachment of PEG chains has little effect on the r1.

The above observations demonstrate that subtle alterations in the molecular composition and architecture may have major consequences for the efficacy of dendrimers as MRI contrast agents, which usually are hard to predict. It should be noted, however, that some of these relaxivity modulations might be pronounced at relatively low magnetic fields but without practical consequences at the higher fields that are typically used in preclinical MRI.

8 Passive Versus Active Targeting

An impressive body of research is devoted to improving the specificity of diagnostic imaging read-outs by using ligand-conjugated contrast materials. In the setting of cancer diagnostics, the idea is to equip the contrast agent with a ligand (e.g. a peptide, synthetic ligand or monoclonal antibody) that allows enhanced sequestration of the contrast agent at those sites where the target of the ligand is abundantly present. In the case of cancer imaging, the target could, for example, be a cell surface receptor (either overexpressed on tumour cells or on endothelial cells lining tumour blood vessels) and an extracellular matrix component (such as fibrin), or one could report on the presence of a disease-specific enzyme (such as a member of the matrix metalloproteinase isoenzyme family). Obviously, ligand attachment can be expected to alter the shape, the size and the global and local motional dynamics of contrast materials.

The added value of ligand conjugation of dendrimers for diagnostic imaging has mainly been explored with the use of receptor-targeted compounds. Examples include the HER2, transferrin and folate receptors (for review, see Agarwal et al. (2008)). In most cases, these materials were also equipped with fluorescent moieties for optical detection. Increasingly, however, the utility of target-specific MRI with paramagnetic ligand-conjugated dendrimers is being explored (for overview, see Kobayashi et al. 2011; Godin et al. 2011; Cheng et al. 2011). Also combinations of dendrimers with superparamagnetic iron oxide nanoparticles for targeted tumour MRI are studied (e.g. see Yang et al. 2015). For the sake of clarity, the present description will be restricted to the use of targeted paramagnetic Gd-based systems.

Han et al. (2011) have conjugated a seven-amino acid peptide ligand for recognition of the human transferrin receptor to PEGylated G5 PAMAM dendrimers that were also equipped with Gd-DTPA for in vivo tumour imaging with MRI. The peptide had been identified with the use of the phage display technique. In a very complete study, the authors extensively characterised the dendrimer contrast material in vitro and used it for in vivo MRI detection of tumours growing subcutaneously in mice. Both targeted and nontargeted versions of the paramagnetic G5 dendrimers had relatively high r1 relaxivities at 4.7 tesla (i.e. 10.7 mM−1s−1 versus 4.8 mM−1s−1 for the single Gd-DTPA molecule), the field strength at which the authors did all MRI scanning in their study. This r1 relaxivity per Gd3+ ion corresponded to a value of 984.4 mM−1s−1 for the entire G5 dendrimer molecule as a whole. The paramagnetic dendrimers showed negligible haemolytic toxicity up to high concentrations. Tumour cells readily internalised the transferrin-targeted dendrimers, as deduced from fluorescence microscopy analysis of isolated cells. Very little uptake occurred for the nontargeted material. Whole-animal fluorescence imaging studies in tumour-bearing mice showed that both targeted and nontargeted dendrimers mainly distributed towards the liver, kidney and spleen, in agreement with biodistribution studies alluded to above. Tumour enhancement was most prominent for the targeted material. Interestingly, the level of tumour homing was very low in the case of C6 glioma growing in the mouse brain, both for targeted and nontargeted agents. The promising results of the qualitative fluorescence studies led the authors to embark on in vivo MRI studies of both tumour models and both types of agents. Examples of MRI scans of subcutaneous xenografts in the mouse hindleg are depicted in Fig. 11.5. The authors employed T1-weighted MRI for detecting the presence of paramagnetic material. The latter leads to a local T1 shortening, which causes a signal intensity increase on the MR images. This so-called positive contrast change was strongest with the targeted dendrimer and amounted to 187 % compared to pre-contrast signal values at 24 h after contrast injection (Fig. 11.5b). Maximal signal changes for Gd-DTPA and nontargeted dendrimers (Fig. 11.5a, b) were much smaller and occurred earlier (i.e. 121 and 130 % at 30 min after injection) and totally disappeared within hours. The entire time course of MRI signal changes with all three agents is depicted in Fig. 11.5b. Similar to the above fluorescence findings, negligible MRI signal changes on T1-weighted scans were noted in case of the C6 glioma with any of the dendrimer types (Han et al. 2011). The reason for the low level of contrast enhancement in case of the brain tumour is unclear but is possibly related to the presence of the blood–brain barrier (BBB) limiting extravasation of the nanoparticulate dendrimers. Brain tumour detection with contrast-enhanced imaging may thus require more elaborate materials that are capable of crossing the (intact) blood–brain barrier.

Fig. 11.5
figure 5

Tumour-associated transferrin-specific MRI with ligand-conjugated paramagnetic dendrimers. MR was done on mice with subcutaneous Bel-7402 xenografts. (a) T1-weighted MR images of nude mice were measured at various time points after the intravenous injection of Gd-DTPA (top row), Gd-DTPA-PAMAM-PEG (second row; representing nontargeted contrast material) and Gd-DTPA-PAMAM-PEG-T7 (bottom row; this is the transferrin-targeted dendrimer preparation). Scans were made using a 4.7 T small animal MR instrument. Images were acquired pre-injection, 5 and 30 min, as well as 1, 2 and 24 h post-injection. The white arrow indicates the location of the xenograft. (b) Quantitative analysis of MR images. The average MR signal intensity was measured for each tumour, and the relative signal intensity was then calculated as the ratio of the signal intensity at the different time points in the post-contrast images compared to the pre-contrast image, indicated in percentage. Statistical tests: *P < 0.05, **P < 0.01 and ***P < 0.001. The red asterisks indicate Gd-DTPA-PAMAM-PEG-T7 versus the other groups. Black asterisks represent Gd-DTPA versus Gd-DTPA-PAMAM-PEG (Reproduced with permission from Han et al. 2011)

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Chen et al. (2012) have conjugated cyclic RGD tripeptide and Gd-DTPA to generation 3 dendrimers for DCE-MRI of tumour angiogenesis and for monitoring the early responses to anti-angiogenic therapy with bevacizumab. The DCE-MRI data measured at 30 min post-contrast injection provided evidence for specific targeting and allowed detection of the response to therapy in terms of reduced MRI contrast enhancement.

Target-specific MRI is very challenging because of the low intrinsic sensitivity of the technique and the fact that many interesting tissue biomarkers occur in relatively low levels. Detection of molecular disease markers with MRI will therefore never become routine, especially when it comes to detecting sparse markers. Nuclear imaging techniques have a better chance of clinical translation in this setting, since these only need tracer levels of probes. As an example, Sato et al. (2001) have used dendrimers that were loaded with anti-sense oligo-DNA and that were labelled with 111In as a nuclear imaging tracer, for tumour imaging with scintigraphy and for monitoring of gene delivery. Ligand-conjugated dendrimers are also being explored for enhancing the specificity of the delivery of therapeutic cargo to disease sites, in particular tumours (Kesharwani and Iyer 2015; Prabhu et al. 2015).

9 Biocompatibility and Safety

The in vitro and in vivo toxicity of dendrimers has been extensively studied (for comprehensive reviews, see Cheng et al. (2011); Jain et al. (2010)). The main issue here is the presence of many primary amine end groups in the most widely employed PAMAM and PPI type of dendrimer materials (see Fig. 11.1), leading to a high density of cationic surface charges for these molecules. This design feature constrains direct in vivo use, since it causes these dendrimers to be toxic. This toxicity is primarily attributed to interactions between the positive charges at the dendrimer’s exterior and negatively charged biological membranes in vivo (Jain et al. 2010). Such interactions may lead to membrane disruption, via pore formation, and explains the toxic effects that dendrimers exert on cells in the blood as well as the general cytotoxicity. In order to eliminate or at least minimise this toxicity, two main strategies are being employed. The first approach is to synthesise dendrimers that are fully biocompatible, consisting of a biodegradable core and with branching units that consist of molecular units that are also biocompatible (e.g. intermediates of metabolic pathways). The second, most widely used approach is to mask the positive charges on the dendrimers’ exterior by surface engineering, e.g. by amidation of the amine end groups. A prime example of this masking approach is PEGylation (see Sect. 11.5), which aids to conceal the positive charges on cationic dendrimers, thereby reducing their interactions with negatively charged membranes. This surface engineering much improves the biological inertness and thereby the biocompatibility of the material.

Biodegradability of contrast agents is yet a completely different issue, as it would necessitate the use of materials that can be taken apart by enzymatic reactions or under the influence of harsh conditions (such as low pH) in some tissue compartments. An appreciable level of biodegradability seems advantageous at first sight, as it may help to more rapidly eliminate materials that might otherwise reside in the organism for a prolonged period of time. However, materials that are highly biodegradable may fall apart in many small pieces, each of which could have very different biodistribution, as well as safety and biocompatibility profiles. For these reasons, it is attractive that the elimination of foreign materials from the body, including contrast materials for diagnostic imaging, occurs relatively rapid and in intact form. This notion explains that relatively fast renal clearance, leaving little time for material processing, is the preferred elimination route. It is evident that this criterion may not always be fulfilled as it sets very restrictive boundary conditions, in terms of size, surface architecture and other molecular characteristics of dendrimer-based contrast materials that may be at variance with their intended use.

10 Probes for Multimodality Imaging

Routinely, dendrimer contrast agents for MRI are labelled with Gd chelates, as detailed above. However, a very interesting alternative is to equip the dendritic structures with fluorine atoms instead, to allow for 19F-MRI (Yu 2013). 19F is the second most sensitive nonradioactive nucleus for MRI, the 1H nucleus being first. In addition, because there is no endogenous source of signal, 19F-MRI can be used as a hotspot imaging modality (Chen et al. 2010), which greatly simplifies image interpretation. Higher-generation dendrimers can readily be equipped with many chemically equivalent 19F-atoms to enhance the detection sensitivity, which continues to be a major challenge for MRI. One of the prime virtues of 19F-MRI is that it directly detects the fluorinated contrast materials. In case of using Gd-based contrast agents for 1H-MRI, the contrast materials are indirectly detected via their effect on the magnetic properties of nearby water protons. This may have major consequences for image interpretation and complicate the quantitative analysis of contrast changes in 1H-MRI images in terms of contrast agent concentrations, as we have previously shown upon cell internalisation of Gd-containing paramagnetic liposomes (Kok et al. 2009). 19F-MRI does not suffer from this effect, again because of its direct probe detection properties (Kok et al. 2011).

Criscione et al. have explored the utility of fluorinated PAMAM dendrimers, which exhibit spontaneous self-assembly into nanoscopic and microscopic particulates (Criscione et al. 2009). The nanostructures disassemble at low pH, creating options for controlled release of encapsulated agents. Huang et al. have similarly prepared fluorinated dendrimers for use with 19F-MRI, except that they co-conjugated Gd-DOTA chelates (Huang et al. 2010). This approach will improve the time efficiency of 19F-MRI, as images can be obtained at shorter repetition times due to the intramolecular T1-shortening effect of Gd on nearby 19F-nuclei. Jiang et al. (2009) have proposed a dendrimer-like low-molecular-weight fluorinated contrast agent for 19F-MRI, which is also rapidly excreted. The molecule produces a relatively strong signal from 27 magnetically equivalent fluorine atoms and was shown to allow for whole-body 19F-MRI of time-dependent changes in its biodistribution. Finally, G1, G3 and G5 dendrimers with hydrophobic CF3-labels buried at the interface between the PPI dendritic core and the PEG corona have been introduced recently (de Kort et al. 2014). These materials are water soluble (and non-aggregating) such that they could be employed for probing hydrogels with the use of NMR diffusometry. The suitability of these novel nanoprobe systems for 19F-MRI remains to be established.

Many different dendrimer-based probes for imaging with non-MRI methods have been prepared over the years. Kobayashi and coworkers have explored the utility of radioactive labelling of dendrimers for nuclear imaging (Kobayashi et al. 1999, 2000). The main virtue of nuclear imaging techniques is that they afford very high detection sensitivities, enabling the use of tracer quantities of contrast materials. PAMAM dendrimers incorporated with nuclear labels, such as 111In, allowed for quantitative whole-body scintigraphy studies of their biodistribution, including their tumour-homing capabilities (Kobayashi et al. 1999, 2000). Liu and others (2013) have employed dendrimer-stabilised gold nanoparticles for targeted tumour imaging with CT imaging. Tumour homing was enhanced by conjugation of ligands for binding to the folate receptor, which is overexpressed on a variety of tumour types. In line with findings for other nanomaterials, the liver and spleen took up the majority of the gold nanoparticles. Importantly, the delivery of folate receptor-targeted CT contrast material at the tumour site was two times higher than that of nontargeted nanoparticles. Equipping dendrimers with labels for detection by optical techniques provides many opportunities for in vivo and ex vivo microscopy analysis of their biodistribution and tumour delivery properties in diagnostic and/or therapeutic studies. As an example, Kobayashi et al. (2007) have conjugated G6 PAMAM dendrimers with five different optical labels and used these sophisticated materials for multicolour lymph node imaging in tumour-bearing mice. Specific detection of individual dendrimers is possible by spectrally selective excitation of the different optical probes (see also Kaminskas and Porter 2011).

Multimodality imaging is an approach whereby two or more imaging modalities are used in a complimentary fashion to compensate for the limitations of each imaging method while exploiting their individual strengths. A well-established example of multimodal imaging of this sort is the combination of PET and CT imaging, where PET delivers contrast-enhanced functional and molecular imaging data and CT provides the anatomical information that is essential for the interpretation of the PET scans. A potentially valuable approach towards contrast-enhanced diagnostic imaging is enabled by the use of contrast materials that can be detected by two or more different imaging modalities (James and Gambhir 2012). Dendrimers are ideally suited for this type of multimodality imaging, since the large number of end groups at the periphery of higher-generation dendrimers allows for controlled sequential conjugation of several different imaging labels. As an example, G6 dendrimer for dual-modality MRI–fluorescence imaging was conjugated with Gd-DTPA and the Cy5.5 fluorophore (Kobayashi et al. 2011; Talanov et al. 2006) and was shown to afford effective sentinel lymph node imaging in mice (Fig. 11.6). The most relevant attribute of such dual-labelled dendrimers for MRI and optical detection is that they are ideally suited for cross modality validation. In this case, MRI is employed for in vivo analysis of the spatial distribution of contrast enhancement patterns (e.g. in three dimensions in a tumour mass). The spatial resolution of MRI, however, is too low to precisely pinpoint the location of the contrast material at the (sub)cellular level. Following sacrifice and tumour slicing, fluorescence microscopy detection of the optical label in the dual-labelled dendrimer can be employed to determine its spatial distribution at the (sub)cellular level as well as its co-localisation with vital cellular markers (e.g. an endothelial cell surface marker in case of vascular targeted contrast material).

Fig. 11.6
figure 6

Dual-modality MRI and near-infrared (NIR) optical sentinel lymph node imaging of breast cancer in a mouse. As indicated schematically, G6 PAMAM dendrimer-based dual-labelled contrast material of approximately 10 nm in diameter, containing on average 172 Gd3+ ions and 2 Cy5.5 fluorophores, was employed. Both MRI and NIR optical images define neck (yellow arrow) and axillary lymph nodes (red arrow) as sentinel lymph node in this mouse model of breast cancer (Reproduced with permission from Kobayashi et al. 2011)

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In recent years the added value of dual-modality MRI–CT imaging for tumour diagnostics has been actively explored with the use of dendrimer-assisted formation of nanoparticles for dual-mode MRI–CT scanning. The MRI–CT combination is of interest since it joins the high soft tissue contrast and flexible image orientation abilities of MRI with the high spatial and density resolution and relatively low cost of CT. Shi and coworkers have demonstrated that this dendrimer-enhanced MRI–CT approach provides effective tools for tumour detection in animal models (Wang et al. 2012; Chen et al. 2013, 2015; Li et al. 2013, 2014; Wen et al. 2013; Cai et al. 2015).

A particularly powerful multimodal combination is afforded by hybrid PET–MRI (Wehrl et al. 2015). This combination differs fundamentally from the ones mentioned above, in the sense that in this case image acquisition of the two modalities is done at the same time. To that end, the two scan devices are fully integrated, making use of MRI-compatible PET detectors. Hybrid PET–MRI instruments are available both for small animal and human scanning. It should be noted, however, that there are no pressing reasons why PET and MRI scans should always be done simultaneously to nevertheless benefit from their combination. Back-to-back MRI and PET scans, followed by multimodal image registration, can likewise be of great value. The combination of PET with MRI is almost ideally synergistic. PET has a very high detection sensitivity and specificity, and new PET tracers can relatively easily be translated to the clinic, since one needs only tracer amounts of material. MRI does not make use of ionising radiation (while CT that is now mostly used in combination with PET does) and produces images with a remarkable range of different soft tissue contrast options. With respect to contrast agents for hybrid PET–MRI, two options come to mind. As a first step, one could inject a cocktail of a PET and an MRI contrast agent. Single agents for hybrid PET–MRI scanning are presently mainly based on (super)paramagnetic nanoparticles for MRI detection that are radiolabelled with 64Cu or 18F for PET detection (de Rosales 2014). When using these dual-modality contrast agents, one could imagine that PET provides whole-body analysis of contrast agent distribution and local concentration measurements, guiding the MRI scanning towards locations at which a high-resolution local analysis is carried out (exploiting all capabilities of MRI, including contrast agent detection). The added value of hybrid PET–MRI per se is being actively explored and will be defined in the years to come. The development of hybrid PET–MRI contrast materials is also still in its infancy. To the best of our knowledge, hybrid PET–MRI with dendrimer contrast materials has not been reported until now. In view of the orders of magnitude differences in sensitivity between PET and MRI, combined with the present limited medical need for assessing diagnostic agents with multiple imaging methods in vivo, the added value of hybrid PET–MRI probes is questionable. The same applies to the CT–MRI combination. Hybrid MRI–optical imaging agents have proven utility for in vivo–ex vivo validation purposes, as explained above.

11 Smart Imaging Probes

An interesting domain of contrast agent development concerns the design of so-called smart imaging probes. Although loosely defined, the annotation smart refers to the fact that these agents are designed to report on changes or spatial differences in physiological environment. Examples include the use of contrast materials that respond differently depending on the local pH; see, for example, the PPI dendrimer-based CEST agents that are able to probe the pH of the solution (Pikkemaat et al. 2007). A special class of smart agents is represented by activatable probes that allow one to control the magnitude of the imaging signal and minimise the background signal in relation to the presence of the biological activator that represents a special class of smart agents. Here, the activator can, for example, be a specific enzyme, in the presence of which imaging signal is generated and in the absence of which no signal occurs. This approach leads to higher target-to-background ratios than would be measured in case of conventional imaging agents that are “always on”. Smart agents are widely explored for optical detection and are particularly based on fluorescence quenching–dequenching schemes (Kobayashi et al. 2010b). To the best of our knowledge, no smart probes that are based on dendrimers have been described as yet. The flexible nature of dendrimer design, which offers many options for anchoring a broad range of compounds to the surface, provides many opportunities for exploring the added value of dendrimer-based responsive materials for diagnostic imaging. In view of the low intrinsic sensitivity of MRI and the principles of the detection of MRI contrast agents, smart probes for use with MRI are not expected to come to routine imaging application.

Conclusions

The design of powerful contrast agents is of vital importance for enhancing the specificity and sensitivity of diagnostic medical imaging. In general, the optimal molecular structure of a contrast agent will depend on a multitude of factors, including the imaging modality and application(s) for which it is intended, its biocompatibility and stability and safety profile. In recent years, a multitude of novel contrast agents has been designed and experimentally tested for all major imaging modalities. Many of these have shown to be useful for in vivo imaging of small animals in the context of biomedical research into novel diagnostic procedures, disease mechanisms and therapy development. The translation of novel contrast materials from the preclinical to the clinical setting obviously is a complex process. In recent years, very few contrast agents have made it from the preclinical to the clinical setting. Exploratory human studies are relatively easy to conduct for nuclear imaging agents, as these require relatively small amounts of contrast materials. Therefore, safety issues are less pressing with PET than with MRI, for example, as the lower sensitivity of the latter technique requires the use of substantial contrast agent quantities. It is improbable that a single imaging test will answer all clinical questions related to a certain diagnostic problem. Therefore, it may be expected that multimodality imaging agents are going to be widely explored in the near future and have a significant impact on the future of personalised medicine (Iagaru and Gambhir 2013). These developments will continue to require in-depth translational studies of contrast materials in a multidisciplinary setting, joining expertise ranging from chemistry to medical technology assessment.