Layer-by-layer capsules for magnetic resonance imaging and drug delivery

Advanced Drug Delivery Reviews 63 (2011) 772–788

Contents lists available at ScienceDirect

Advanced Drug Delivery Reviews j o u r n a l h o m e p a g e : w w w. e l s ev i e r. c o m / l o c a t e / a d d r

Layer-by-layer capsules for magnetic resonance imaging and drug delivery☆ Hua Ai ⁎ National Engineering Research Center for Biomaterials, Sichuan University, Chengdu, 610064, China Department of Radiology, West China Hospital, Sichuan University, Chengdu, 610041, China

a r t i c l e

i n f o

Article history: Received 9 November 2010 Accepted 30 March 2011 Available online 27 April 2011 Keywords: Layer-by-layer Self-assembly Polyelectrolyte capsule Magnetic resonance imaging Paramagnetic Superparamagnetic Gadolinium Iron oxide Drug delivery Molecular imaging

a b s t r a c t Layer-by-layer (LbL) self-assembled polyelectrolyte capsules have demonstrated their unique advantages and capability in drug delivery applications. These ordered micro/nano-structures are also promising candidates as imaging contrast agents for diagnostic and theranostic applications. Magnetic resonance imaging (MRI), one of the most powerful clinical imaging modalities, is moving forward to the molecular imaging field and requires the availability of advanced imaging probes. In this review, we are focusing on the design of MRI visible LbL capsules, which incorporate either paramagnetic metal-ligand complexes or superparamagnetic iron oxide (SPIO) nanoparticles. The design criteria cover the topics of probe sensitivity, biosafety, longcirculation property, targeting ligand decoration, and drug loading strategies. Examples of MRI visible LbL capsules with paramagnetic or superparamagnetic moieties were given and discussed. This carrier platform can also be chosen for other imaging modalities. © 2011 Elsevier B.V. All rights reserved.

Contents 1. 2.

Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Design criteria of multifunctional imaging capsules: general considerations . . . . . . . . . 2.1. Brief introduction of LbL capsules as drug carriers . . . . . . . . . . . . . . . . . . 2.2. Design considerations of imaging enabled capsules . . . . . . . . . . . . . . . . . 3. LbL capsules for magnetic resonance imaging . . . . . . . . . . . . . . . . . . . . . . . 3.1. Magnetic resonance imaging (MRI) . . . . . . . . . . . . . . . . . . . . . . . . 3.2. MRI contrast agents . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2.1. Small molecule paramagnetic agent . . . . . . . . . . . . . . . . . . . . 3.2.2. Superparamagnetic iron oxide nanoparticle based agents . . . . . . . . . . 3.3. Design considerations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.3.1. Design considerations of paramagnetic based macromolecular structures . . 3.3.2. Design considerations of superparamagnetic based macromolecular structures 3.4. LbL capsules with paramagnetic imaging moieties . . . . . . . . . . . . . . . . . . 3.5. LbL capsules containing superparamagnetic nanoparticles . . . . . . . . . . . . . . 4. Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Acknowledgments . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abbreviations: BLI, bioluminescence imaging; CEST, chemical exchange saturation transfer; CH, chitosan; CT, computed tomography; daer, aerodynamic diameter; FLI, fluorescence imaging; HA, hyaluronic acid; LbL, layer-by-layer; MF, melamine formaldehyde; MR, magnetic resonance; MRI, magnetic resonance imaging; NSF, nephrogenic systemic fibrosis; q, hydration number; PAA, poly(acrylic acid); PAH, poly(allylamine) hydrochloride; PAS, poly(anetholesulfonic acid); PASP, poly-L-aspartic acid; PE, polyelectrolyte; PEI, polyethylenimine; PEG, poly(ethylene glycol); PET, positron emission tomography; PGA, poly-L-glutamic acid; PLL, poly-L-lysine; PSS, poly(styrenesulfonate); r1, longitudinal relaxivity; r2, transverse relaxivity; RES, reticuloendothelial system; SPECT, single photon emission computed tomography; SBM, Solomon–Bloembergen–Morgan; T1, longitudinal relaxation time; T2, transverse relaxation time; τR, rotational correlation time; τM, residence time of the coordinated water molecule. ☆ This review is part of the Advanced Drug Delivery Reviews theme issue on “Layer-by-Layer Self-Assembled Nanoshells for Drug Delivery”. ⁎ National Engineering Research Center for Biomaterials, Sichuan University, 29 Wangjiang Rd., Chengdu 610064, China. Tel./fax: + 86 28 85413991. E-mail address: [email protected]. 0169-409X/$ – see front matter © 2011 Elsevier B.V. All rights reserved. doi:10.1016/j.addr.2011.03.013

H. Ai / Advanced Drug Delivery Reviews 63 (2011) 772–788

1. Introduction Medical imaging has long served as an important tool for diagnosis and therapeutic efficacy monitoring. During the last ten years, with the advances in molecular and cell biology discoveries, along with the development of noninvasive and high-resolution in vivo imaging technologies, molecular imaging has emerged as a new tool for understanding pathological changes at cellular and molecular levels [1–6]. This capability is appreciated in cancer and cardiovascular disease diagnosis [7–11], drug discovery [12–14], tissue engineering [15], drug delivery [16–21] and theranostics [22,23]. Different imaging modalities were involved in this development, including magnetic resonance imaging (MRI), X-ray computed tomography (CT), positron emission tomography (PET), single photon emission computed tomography (SPECT), bioluminescence imaging (BLI), fluorescence imaging (FLI), and ultrasound imaging. Each imaging modality has its own advantages and disadvantages in terms of spatial resolution, target, and penetration depth (Table 1). For example, MRI is safe and has very high spatial resolution around 25–100 μm at different magnetic fields, however, the sensitivity is much behind nuclear imaging such as PET, SPECT, and optical imaging methods. The introduction of contrast media can improve image contrast and delineate small changes which may be difficult to be discovered under regular scans, and combining of two or three imaging modalities can compensate the disadvantages of each other and provide a lot of information that may not be collected through one modality [24–26]. Gadolinium (Gd) or manganese (Mn) based small paramagnetic molecules, and superparamagnetic iron oxide (SPIO) based nanoparticles, are two major categories of MRI contrast agents approved for clinical applications. These contrast agents have shown their great values in diagnosis of cardiovascular, cancer and many other diseases. Rational design of advanced and multifunctional imaging probes can prolong their blood half-life, increase sensitivity, and recognize targets at molecular level, leading to further enhancement of their diagnosis capability, and expanding application portfolios to drug delivery and other fields [21,24,27,28]. A wide range of nanosystems have been used for carrying Gd or SPIO based agents, including dendrimers [29–31], carbon nanotubes [32,33], fullerenes [34,35] and viral capsids [36–38], and well established drug carriers such as liposomes [39–41] and micelles [28,42–44]. The combination of therapeutic and diagnostic agents into one platform can be named as “theranostics” [22]. One of the important applications of such a system is to monitor the drug delivery efficiency and optimize the therapeutic regime accordingly. Polyelectrolyte capsules developed from layer-by-layer (LbL) selfassembly have shown great applications in storing, protection, release, and delivery of different functional agents [45–54]. The advantages of LbL capsules include ease of size control, wide selection of materials for assembly, fine design of self-assembly architecture, mild loading environment, and controlled permeability. For biomedical applications, they have been used for controlled encapsulation and release of small molecule drugs [55,56], enzymes [57–61], protein drugs [62], and DNA [63]. Besides, LbL capsules have been engineered

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for studying their behaviors when interacting with biological cells through either passive or active means [64,65]. Adding imaging components into capsules will not only expand this platform technology to diagnosis and other applications, but also improve their values in controlled release and drug delivery. In this review, we will discuss how to design and develop multifunctional LbL self-assembled capsules with imaging functionality, and their potential applications for diagnosis, drug delivery and theranostics. In particular, we will discuss the design rationale on how to incorporate MR imaging components (e.g., Gd or iron oxide based ones) into the LbL capsule system, how to improve their imaging sensitivity, and how to load sufficient amount of drugs at the same time. Some examples of MRI visible capsules will be given and discussed. Other imaging modalities such as optical imaging, ultrasound imaging or nuclear imaging are also important tools, but not the focus of this review. Besides the imaging issues, we will also cover two other important design criteria: targeting efficiency and long circulation capability. 2. Design criteria of multifunctional imaging capsules: general considerations 2.1. Brief introduction of LbL capsules as drug carriers Electrostatic LbL self-assembly, introduced in the early 1990s, now is a versatile technology used to design and develop biomaterials through the bottom-up strategy [47,48,66–68]. When used for drug delivery systems, LbL capsules are recognized as one of the nanotechnologies that advanced the field of drug delivery since its introduction in 1998 [51,52,69]. The conventional drug delivery platforms such as polymer micelles or liposomes were developed prior to most imaging probes. There are smart design and development strategies that have been successfully demonstrated in clinical medicine for either drug delivery (e.g., liposomal formulation Doxil® for delivery of anticancer drug doxorubicin) or medical imaging (e.g., lipid vesicle formulation called SonoVue® as ultrasound imaging contrast agent) applications. When designing LbL capsules for imaging and drug delivery devices, it would be helpful to learn from the characteristics of those well established delivery systems with approved in vivo performance. Polyelectrolyte capsules present unique advantages for delivery of peptide and protein based drugs because: 1) easy fabrication process; 2) no chlorinated organic solvent is needed; 3) precise control of permeability through membrane thickness and shell wall pore size; 4) broad selection of shell composition materials; 5) capsules can be switched between “open” and “closed” states for controlled release. There are two commonly used methods for encapsulation of therapeutic agents inside polyelectrolyte capsules: 1) direct coating of oppositely charged polyelectrolytes onto drug micro/nanoparticles; and 2) loading of drug molecules into preformed “hollow” polyelectrolyte capsules. The second strategy is more popular because not all drugs can form into nanoparticle suspensions, and capsule size and monodispersity can be easily controlled by choosing appropriate

Table 1 Comparison of commonly used biomedical imaging modalities [6,13,16,22]. Modality

Spatial resolution

Depth

Time

Quantitative

Target

Imaging probes

MRI

25–100 μm

No limit

Min–hr

Yes

Anatomical, physiological, molecular

PET SPECT CT Ultrasound Optical imaging

1–2 mm 0.5–1 mm 50–200 mm 50–500 mm 1–3 mm

No limit No limit No limit mm–cm mm

Min Min Sec–min Sec–min Sec–min

Yes Yes Yes Yes No

Physiological, molecular Physiological, molecular Anatomical, physiological Anatomical, physiological Physiological, molecular

Paramagnetic: Gd, Mn chelates Superparamagnetic: Fe3O4, gamma-Fe2O3 18 F, 11C, 15O 99m Tc, 111In chelates Iodine SF6, C3F8 microbubbles Bioluminescence: Luciferin Fluorescence: a) dyes: FITC, Alexa, etc. b) quantum dots: CdSe, CdSe/ZnS, etc.

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templates. However, the first strategy ensures a much higher loading capacity than the second one. The capsule assembly procedure is simple and straightforward as illustrated in Fig. 1. First, one needs to choose a template with desired composition and size. A wide range of sacrificial templates including weakly cross-linked melamine formaldehyde (MF) particles [51,70], organic [71] and inorganic crystals [72], metal nanoparticles [73], organic/inorganic hybrid particles [74,75] and biological templates [76,77] are available. In this scheme, hybrid micro/nanoparticles composed of polymers and inorganic component are used as templates for subsequent polyelectrolyte coating. The first step is co-incubation of templates with excessive amount of polyelectrolytes, usually the oppositely charged polymer against the template. However, choosing the same charged polymer has proven to be effective in some cases, such as using PSS as the first layer to coat MnCO3 and MnCO3/polysaccharide microtemplates in order to induce higher negative charges on particle surface, and this phenomenon is probably mainly attributed to the surface-active properties of PSS [75,78]. Then a washing step is required to remove free polymers before coating the second layer. This can be easily done through appropriate centrifugation or filtration. By repeating the coating and washing steps, multilayers of polyelectrolytes can be coated on micro/ nanotemplates with precise control of thickness and molecular structures. Besides the electrostatic interactions, one may consider to use alternate adsorption methods to combine multilayers of polymers together and then form capsules. It has been suggested that the short-range hydrophobic forces should be considered as a driving force of LbL adsorption of dyes, proteins, polymers, and nanoparticles [79]. Specific interactions between biomolecules are also important in LbL self-assembly, such as alternate assembly of Con A and glycogen (branched polyglucose) as well as biotinylated polylysine and streptavidin [80,81]. Another well investigated assembly method is based on hydrogen bonding [82–85]. Beyond the above mentioned conventional methods, unconventional methods such as inclusion complexes, noncovalent modification, coordination polyelectrolyte, and block polymer micelles can also be used to build LbL capsules [86]. Based on this microencapsulation and further removal of templates, a unique micro/nano carrier system “hollow capsules” can be

Fig. 1. Schematic illustration of LbL self-assembly of polyelectrolyte capsules with imaging and other functionalities: I) LbL self-assembly of polyelectrolytes onto hybrid templates; II) template decomposition; III) purification of polymer matrix containing polyelectrolyte capsules; IV) loading of therapeutic agents into the capsules; and V) addition of imaging, targeting, and protection moieties into capsule systems.

developed (Fig. 1). After multilayered polyelectrolyte capsules formed on templates, the core can be removed by exposing to acidic or basic solutions (step III in Fig. 1). The capsules are monodisperse in size and the capsule wall serves as a permeable barrier. The interior environment of a polyelectrolyte shell is aqueous, similar to that of liposome and polymer vesicle. However, the loading efficiency in preformed hollow capsules is generally not high enough for practical use, since the mechanism for this encapsulation is a simple diffusion process mainly driven by concentration gradient. Different strategies have been applied to increase the loading efficiency. It has been found that drugs can be effectively loaded into MF residue containing capsules. The negative charge of polyanion/MF oligomer complex inside the capsule can preferentially induce the deposition of oppositely charged water-soluble substances into the capsules and result in relatively high loading efficiency [60,62]. This is the phenomenon of “spontaneous deposition” driven by electrostatic interaction. Similarly, charged polyelectrolyte complexes were created in capsules by adsorption of polyelectrolytes into the porous CaCO3 microparticles to improve the loading performance of the resulted capsules [74,87]. In another example, hybrid microparticles composed of MnCO3 and ionic polysaccharide were used for preparation of matrix capsules with a much better loading capacity [75]. Polyelectrolyte capsules are appropriate carriers for biomacromolecules because organic solvent is not involved in the loading process, and resulted in less damages to their bioactivities. Multifunctional LbL capsules can be developed through the incorporation of targeting ligands, protection polymers, and imaging moieties (step V in Fig. 1). How these components are added into LbL capsules depends on the designed structures, and more details will be discussed later in the design criteria and examples. 2.2. Design considerations of imaging enabled capsules Addition of imaging moieties (active components of contrast agents) into capsule systems enables them to be used either as diagnostic probes or imaging visible drug delivery carriers. There are some general rules we need to consider for design of high performance multifunctional nanostructures. For diagnostic and drug delivery applications, the efficacy of systems needs to be tested in vivo eventually. So their in vivo biocompatibility and stability are two important questions we need to address. The biosafety of a multifunctional system mainly depends on the imaging component and capsule materials. Capsules can be formed based on pairs of oppositely charged polyelectrolytes, which could be selected either from synthetic or natural biocompatible and biodegradable polymers. A wide selection of polypeptides (e.g., polyL-lysine, poly-L-glutamic acid, and poly-L-aspartic acid) and polysaccharides (e.g., hyaluronic acid, heparin, dextran, chondroitin, and chitosan) are readily available for self-assembly of capsules [53]. The composition and the structure of imaging components decide their in vivo safety. Gadolinium is considered as toxic, once released from the ligand and it can potentially cause Nephrogenic Systemic Fibrosis (NSF) [88,89]. The small size of nanoparticles such as iron oxide nanocrystals (4–16 nm of the inorganic core) needs to be carefully evaluated as the kinetics of nanomaterials is one of the key issues in safety evaluation [90,91]. Long circulating nanocarriers can have enough time windows to accumulate at the target site instead of being quickly cleared by the reticuloendothelial system (RES) after administration, generating higher imaging contrast between a target site and background tissues, ensuring higher accuracy for diagnosis and drug delivery. The size and surface properties of LbL capsules are key parameters that influence the circulation time. For most current nanoparticle systems (e.g., liposome, micelle), their size is preferably below 200 nm, because larger particles have difficulties to pass through the splenic filtration and can be cleared in spleen [92]. However, for human lung drug

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delivery, carriers with aerodynamic diameter lower than 3 μm tend to be exhaled, while larger particles with daer = ~ 3 μm can get the maximal deposition of monodisperse aerosol particles in the alveolar region without obvious clearance by alveolar macrophages [93]. LbL capsules usually have diameters of a few microns because of the wide availability of micron templates (e.g., MF, CaCO3, MnCO3), so choosing smaller templates is a key to develop nano-sized capsules for longer circulation in the blood stream. Gold and silica nanoparticles are most commonly used nanotemplates because of their uniform size distribution [94–96], but dissolution conditions (e.g., hydrofluoric acid is used to dissolute SiO2 templates) should be mild and not damaging the integrity of biodegradable capsules. Poly(ethylene glycol) (PEG), a non-charged hydrophilic polymer, is well accepted in pharmaceutical industry for drug formulation [21,97]. When incorporated onto liposome surface or used as micelle corona, PEGylated nanocarriers have shown prolonged blood circulation time with improved pharmacokinetics [98]. Polyelectrolyte capsules usually have charged surfaces, which can easily adsorb plasma proteins and then form aggregates and cleared by the RES system. PEGylation on LbL capsule surface serves as a simple and efficient method to make drug carriers resistant to protein adsorption [64], reducing unspecific cellular uptake [99] and potentially leading to prolonged blood circulation time. To make capsules visible under MRI scans, imaging components can be either conjugated on shell surface, in the shell wall or inside capsules. The exact design depends on which imaging probe to use, what are the specific applications, and the feasibility of labeling procedure. For drug delivery, the imaging moiety can be conjugated to a drug, or into the capsule system, or both of them. Conjugation of imaging component to polyelectrolytes before capsule formation would be one option to consider. For one example, Gd-DTPA was first conjugated to PEI, and then self-assembled with PSS to form MRI visible capsules onto silica nanoparticles [100]. If the imaging moieties such as Gd-ligand complexes are incorporated through passive adsorption instead of chemical conjugation, it's important to make sure most of them stay in capsule system instead of leaking out during the multiple steps of purification. For either diagnosis or therapeutic application, the accuracy of delivery of nanocarriers to the site of interest determines the diagnosis and therapeutic efficacy. A wide selection of targeting moieties including monoclonal antibodies, peptides, and aptamers can be considered. Incorporation of targeting moieties on capsule surface can be either through chemical conjugation or passive adsorption. For antibodies and peptide ligands, their spatial configuration on capsule surface is important to decide the recognition efficiency of the target. Chemical conjugation is preferred over passive adsorption because the latter strategy has limited control on ligand or antibody's configuration when attached to capsule surface, leaving less chance for optimal binding to targets [101,102]. However, LbL self-assembly is still a simple method to use without complicated bioconjugation procedures, and electrostatic adsorption of monoclonal antibody or cell penetrating peptide on LbL polyelectrolyte coated templates has shown either better targeting efficiency or increased cell uptake rate [65,103]. Besides labeling methods, it also requires careful arrangement of targeting moieties (ligand or antibody) and imaging component on the capsule surface. They can be separately added on capsules with an optimal ratio to ensure good targeting and enough imaging sensitivity. If imaging moieties and targeting ligands are conjugated before LbL assembly, attention should be paid because labeling a peptide with an imaging moiety may potentially interfere with its binding region. To retain the targeting activity of peptides or antibodies, spacer can be incorporated between the active site of the targeting and the imaging moiety, or the imaging moiety should be conjugated to a peptide sequence position that is not actively involved in receptor recognition [104].

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3. LbL capsules for magnetic resonance imaging 3.1. Magnetic resonance imaging (MRI) MRI is a powerful non-invasive imaging tool which has shown great value in early cancer diagnosis, brain functional imaging, vasculature imaging, implant monitoring, and drug discovery. The advantages of MRI include, safe imaging without using high-energy radiation, deep tissue penetration, and high spatial resolution (down to 25–100 μm at higher magnetic fields) for soft tissues [13,16,22]. Not limited to providing anatomic information of tissues, MRI can bring us important functional information under one scan as well [105]. With the help of advanced imaging probes, it is able to image pathological changes at cellular and molecular levels [9–11]. Besides, application of MRI for drug discovery has proven to be effective, especially when used for characterization of disease models and therapeutic efficacy [13,14]. When used for drug delivery monitoring, MRI has demonstrated that it can track the drug carriers effectively both in vitro and in vivo [28,106]. The fundamental principle underlying MRI is that unpaired nuclear spins (such as hydrogen atoms in water and organic compounds) align themselves when placed into a magnetic field [107]. During clinical imaging, what is primarily observed are hydrogen atoms from water that are present in tissue at ~90 M [108]. To enhance the differentiation between tissues, it is important to adjust the longitudinal (1/T1) and transverse (1/T2) relaxation rates of water at different compartments, thus leading to better image contrast to discover minor changes. The availability of unlimited pulse sequences guarantees the broad selection of scanning parameters during MRI scan, giving us a lot of freedom to choose the optimal image contrast based on different tissue composition, location and surrounding tissues. The overriding challenge with MRI for molecular imaging is its relatively low sensitivity comparing to PET and optical imaging systems [2,108,109]. The introduction of MRI contrast agents, especially customized probes, is essential to discover pathological changes at the molecular level such as apoptosis [110,111], gene delivery [112], and molecular interactions [113]. The performance of a contrast agent is measured as its relaxivity, which is the change in proton relaxation rate normalized to the concentration of the agent in 1 mM [114,115]. 3.2. MRI contrast agents Two kinds of MRI contrast agents are available for both basic and clinical applications: they are paramagnetic small molecule agents and superparamagnetic iron oxide nanoparticles. They are able to shorten either longitudinal or transverse proton relaxation time and thus provide higher image contrast. The summarization of clinically approved MRI contrast agents with detailed information can be found in a few excellent reviews [114,116,117]. 3.2.1. Small molecule paramagnetic agent Among the different paramagnetic metal ions (Mn2+, Mn3+, Fe3+, and Gd3+) used in contrast agents, Gd3+ is the most frequently used one, because of its seven unpaired electrons and long electronic relaxation time caused by the symmetric S-state of Gd(III) [118,119]. Besides Gd(III), Manganese(II) is the only other ion approved as paramagnetic agent for clinical imaging named as Mn-dipyridoxyldiphosphate [Mn(DPDP)]4− (Telsascan®, GE) for liver imaging. Mn(II) has five unpaired electrons and fast electronic relaxation rate. Gadolinium (Gd) or Mn(II) based MRI contrast agent, is of paramagnetic property, and can enhance the MRI signal intensity by shortening the longitudinal relaxation time (T1) of the protons from surrounding water molecules. The relationship between the observed solvent relaxation rate, 1/(Ti)obs, the intrinsic diamagnetic solvent

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relaxation rate, 1/(Ti)0, and the relaxivity of a paramagnetic agent, (ri)agent is shown in Eq. (1)[114,120]: 1 1 = + ðri Þagent ⋅ Cagent ðTi Þobs ðTi Þ0

i = 1; 2 :

ð1Þ

It is noteworthy to point out that paramagnetic agents can shorten both T1 and T2 relaxation rates, but more prominent for the T1 effect. For most commercial Gd based agents, the average T1 relaxivity (r1) is around 4 Gd mM− 1 s− 1. However, because of the toxicity of free Gd ions, they have to be formed into stable complexes with chelators and excreted intact after intravenous administration. There are a few Gdand Mn-based agents approved for clinical applications with three major categories of chelating ligands (Chart 1): 1) DTPA (diethylenetriaminepentaacetic acid) for Magnevist®, and derivatives for Omniscan®, OptiMark®, MultiHance®, Primovist®, and Vasovist®; and 2) DOTA (1,4,7,10-tetraazacyclododecane-1,4,7,10-tetraacetic acid) for Dotarem®, and derivatives for ProHance®, and Gadovist®; 3) DPDP (dipyridoxyldiphosphate) for Teslascan®. Complexes of GdDOTA are relatively stable in vivo because the ligands can protect decomplexation of Gd(III) and effectively prevent transmetallation by Zn(II) or Ca(II) [121,122]. The recommended clinical dosage is usually between 15.7 and 47.1 mg Gd/kg B.W., resulting in injection of more than 1 g of Gd into a 70-kg person. Extreme caution should be applied for patients with low kidney perfusion rates, because inefficient discharge can lead to drug accumulation inside the body, causing NSF [88,89]. Another unique type of paramagnetic contrast agent is called chemical exchange saturation transfer (CEST) agents, and it can selectively saturate a proton pool in slow-to-intermediate exchange with the bulk water [27,123,124]. The saturated magnetization of the irradiated spins is then transferred to the water resonance, thus decreasing its signal intensity [27]. The CEST probes are not the focus of this paper and there a few excellent reviews available for reference [125–127]. 3.2.2. Superparamagnetic iron oxide nanoparticle based agents Different from small molecule paramagnetic agents, superparamagnetic iron oxide (SPIO) nanoparticles themselves are the assembly of thousands of Fe(II), Fe(III), and O atoms into an organized nanostructure. SPIO nanoparticles can shorten transverse relaxation time (T2) and bring negative contrast, resulting in hypointense images. One prominent characteristic of SPIO nanoparticle based contrast agents is that they have high T2 relaxivity and can go beyond 900 Fe mM− 1 s− 1[128], which is essential in imaging of small pathological changes and tracking drug delivery vehicles in vivo. For ultrasmall sized SPIO nanoparticles, they can be used as T1 agents, for example, Clariscan (NC100150 injection) has a low r2/r1 ratio of 1.83 and effective T1 shortening capability [129,130]. SPIO nanoparticles have crystalline structures with the general formula of Fe3+2O3M2+O, wherein M2+ is a divalent metal ion such as iron, manganese, nickel, cobalt, zinc, or magnesium. Particle size usually ranges from 4 to 16 nm in diameter for inorganic core. Among a few choices of M2+ such as Fe2+, Zn2+, Mn2+, Co2+, and Ni2+,

MnFe2O4 has the highest magnetic susceptibility, with approximate magnetic spins of 5 μB. In comparison, the magnetic susceptibility gradually decreased as M2+ changed from Fe2+ to Co2+ to Ni2+: magnetic spin magnitude changed from approximately 4 μB to 3 μB to 2 μB[131]. Mn-doped MnFe2O4 showed the strongest MR contrast effect with the highest T2 relaxivity value, comparing to other metaldoped nanoparticles with the same size. The MnFe2O4 nanoparticles have demonstrated their applications in tumor targeting [131] and liver imaging [43]. Other magnetic nanoparticles such as FePt alloy nanoparticles, FePt/Fe3O4 core/shell nanoparticles, and Au–Fe3O4 dumbbell nanoparticles are also potential candidates as MRI probes [132]. Current clinically approved SPIO based MRI contrast agents are prepared through aqueous precipitation of ferric and ferrous ions in basic environment in the presence of dextran or carboxydextran. Comparing to Gd-ligand small molecule paramagnetic agents, SPIO agents are much safer because the human body has a big iron pool with the average amount of 35 to 45 mg iron per kilogram of body weight and usually reaches to 3 to 5 g per person [133,134]. The recommended dosages of Feridex and Resovist are 0.56 and 0.45– 0.9 mg iron/kg B.W. respectively, which is a much lower dose comparing to Gd agents (15.7 mg/kg B.W.) and only contribute to a small portion of body iron pool. The iron from SPIO degradation can enter the body iron pool and participate physiological iron homeostasis through macrophage uptake and subsequent degradation [135]. A biocompatible coating (e.g., dextran or carboxydextran) on particle surface is important for stabilization as colloidal suspension. There are a few SPIO agents for clinical applications: Feridex IV® for liver and spleen imaging, Resovist® for liver imaging, and GastroMARK® for bowel imaging. 3.3. Design considerations 3.3.1. Design considerations of paramagnetic based macromolecular structures Gd(III) and Mn(II) based paramagnetic agents have been successfully used in clinical applications with millions of cases. A standard clinical injection dose of 0.1 mmol/kg corresponds to an average extracellular tissue concentration of ∼0.5 mM and, given a typical relaxivity of 4 mM− 1 s− 1 for a clinical Gd3+-based agent, this corresponds to an increase in water proton relaxation rate of 2 s− 1, a value that is easily detected at clinical imaging fields [125]. Their capability to shorten the T1 time and increase T2 rate is recognized at tissue levels, but when moving to cellular and molecular level imaging, the sensitivity still needs to be further improved. There are two major reasons behind this, one is the low concentration receptors on cell membrane for the probe to recognize, and the other is the fast clearance of probes in vivo because of the small molecular weight. The T1 relaxivity (r1) reflects how the probe can shorten the T1 relaxation time of the surrounding water. Based on the Solomon–Bloembergen–Morgan (SBM) theory [136], the inner-sphere relaxation is related to a few key parameters, which play important effects on the T1 relaxivity of a probe, and they are 1) q: the number of labile water molecules coordinated to the metal, also

Chart 1. Clinically approved T1 MRI contrast agents based on the ligands of DTPA, DOTA, and DPDP.

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named as “hydration number”; 2) τM: residence time of the coordinated water molecule, briefly as “residency time”; 3) τR: the tumbling motion of the complex, also called rotational correlation time; and 4) TiE: the relaxation characteristics of the unpaired electrons of the metal ion. The q value has a significant impact on the inner-sphere relaxivity and all clinically approved contrast agents having a value of one (q = 1) [117]. From the following Eq. (2), we can see that the inner sphere T1 relaxivity (rIS 1 ) can be improved with a higher number of q value: IS

r1 =

q = ½H2 O : ðT1m + τM Þ

ð2Þ

Here, T1m is the T1 of the hydrogen in the inner-sphere and [H2O] is the water concentration in mM [108]. It is desired that an optimized ligand is able to have q = 2, that is two water molecules coordinated to the paramagnetic metal such as Gd or Mn, requiring design of new ligands with less steric constraints. Another important parameter is τM, and in general, it should be reasonably short in the range of 10– 30 ns, to give an overall increased r1 as shown in Eq. (2)[108]. Careful consideration should be given to the design of T1 agent between stability and relaxivity. The requirements for in vivo stability suggest that kinetic stability, also known as kinetic inertness, of the complexes is much more important than their thermodynamic stability [117]. Increasing the hydration number can increase the inner sphere relaxivity to some degree, but it is often accompanied by a decrease in thermodynamic stability and/or kinetic inertness of Gdligand associated with toxicity issues [114]. A lot of work focused on improvement of T1 relaxivities is related to the increase of τR, which will be used to slow the tumbling motion of the metal–water proton vector of the system. Early attempts include using conjugation of Gd-ligand to polymers such as poly-Llysine [137,138] and polyethylene glycol [139], or proteins such as BSA [140,141]. Later strategies usually involve nanocarriers such as dendrimers [29,100], micelles [44], liposomes [40,41], fullerenes [34,35], viral capsids [36–38], carbon nanotubes [33,142], and proteins [143], resulting in greatly improved T1 relaxivity as high as 170 Gd mM− 1 s− 1[142]. The rotational correlation time is closely related to the size and rigidity of the overall system [144]. Usually, larger supramolecular aggregates with higher rigidity can slow the tumbling motion efficiently, leading to prolonged τR. Additionally, the multimeric properties of Gd-ligand conjugated on the same sized molecule also influence the T1 relaxivity [145]. Based on the experience with the above systems, there are a few suggestions that have been proposed for design of better imaging moiety carriers: 1) placement of Gd or Mn ions at the barycenter of the macromolecule or nanosystem, so the tumbling of such a carrier can transfer into rotation of metal–water proton vector [27,117]; 2) spherical carriers are preferred because they may ensure an isotropic and slow tumbling motion [27,117]; 3) the linking spacer between metal-ligand and the system should have a certain degree of rigidity, because a flexible linkage usually lets the metal–water proton vector rotate independently and faster than the whole complex [117,146]. There are different ways of incorporation of Gd-ligands into the capsule system: 1) Gd-ligands can be loaded into the inner compartment of the capsule. For liposomes, it has been found that the relaxivity of the entrapped paramagnetic agents appears to be lowered because of the limited exchange of bulk water with the contrast agents [147,148]. For LbL capsules, the capsule wall usually has good permeability for small molecules, so the exchange between bulk water and contrast agents should not be excessively affected. Another issue is how to anchor Gd-ligand in the center of a capsule instead of allowing them move freely inside capsules. 2) The second strategy is to incorporate Gd-ligands into the capsule membrane. And this can be done by conjugation of contrast agents to one kind of

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polyelectrolyte, and then LbL assemble it with an oppositely charged polymer. As shown in Fig. 2, the LbL capsules are decorated with Gdligands at the peripheral part, and the rotational correlation time of the metal–ligand complex (τR1) may be different from that of the capsule (τR2) and the linking spacer plays an important role. It is preferred to use linking spacers that have good rigidity to bridge the ligand and the polymer, thus reducing the difference between τR1 and τR2. During “hollow” capsule fabrication, the removal of template core usually requires using acidic solution at pH 1 or similar values to ensure the complete dissolution of core materials such as MF, CaCO3, MnCO3, etc. Because the stability of Gd-ligand is reduced at lower pH values, for example, for Gd(III)-DTPA complexes, the formation of kinetically unstable protonated [Gd(Hdtpa)(H2O)]-species can be found when pH drops to ~4.5 [117]. So it would be better to use mild acidic conditions to remove core materials without damaging the metal-ligand complex during the core dissolution process. Another option is to complex the metal ion with ligand after core removal, but this should be carried out with care to avoid unnecessary binding of metal ion with charged polyelectrolytes. The relationship between τR, τM, and T1 relaxivity has been simulated and discussed for better understanding of the design of MRI probes at different magnetic fields [108,117,149]. Increasing τR is not always the best strategy and it usually works in the magnetic field with the range of 0.24–1.65 T as suggested by Hermann et al. from their simulation studies [117]. As illustrated in Fig. 3 [117], in the frequency range of 10–70 MHz, higher τR is preferred for stronger T1 relaxivity with τM of 10 ns. However, for higher magnetic field of 2.35–9.4 T (frequencies 100–400 MHz), τR should have an intermediate value such as 1 ns and τM has a shorter one. Clinical MRI scanners are usually operating at 1.5 or 3 T fields, and 3 T is preferred in many applications for the excellent spatial resolution of soft tissues. Based on the simulations, one should design the MRI probe with reasonable τR value, which can be tuned by the size, shape, and multimeric properties of nanocarriers. 3.3.2. Design considerations of superparamagnetic based macromolecular structures Different from paramagnetic agents, which are mostly dependent on the inner-sphere contribution and requires the close contact between metal center and water molecules, superparamagnetic agents do not necessarily need the direct contact between metal ion

Fig. 2. Schematic illustration of the design of LbL capsules equipped with Gd-ligand components.

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Fig. 3. Simulations of T1 relaxivity as a function of proton Larmor frequency (1H NMRD profile) with 298τM = 10 ns (black) and 500 ns (red) and 298τR = 100 ps, 1 ns and 10 ns. T = 37 °C, 298τV = 40 ps, Δ2 = 1019 s− 2, RGdH = 3.1 Å. The gray area shows the range of magnetic fields (0.5–3 T) in clinical practice. Reproduced with permission from the Royal Society of Chemistry [117] Copyright 2008.

and water molecules, and the outer-sphere relaxation is the dominant factor. The design considerations for SPIO agents are at the macroscopic level comparing to the small molecule paramagnetic ones. For well controlled nanostructures based on self-assembly, we need to choose SPIO nanoparticles with required size, crystal structure and good magnetic properties. There are a number of methods available to synthesize SPIO nanoparticles [150], including 1) coprecipitation in aqueous phase; 2) high temperature decomposition of metal complexes in organic phase; 3) sol-gel reactions; 4) hydrothermal method; and 5) flow injection syntheses. The coprecipitation method of Ferric (Fe3+) and Ferrous (Fe2+) ions in aqueous environment at basic pH values is probably the simplest way to obtain SPIO nanoparticles. But particles obtained from coprecipitation method usually have wide size distribution and poor control of particle crystal structure. In comparison, thermal decomposition of iron precursors has shown excellent capability in precise control of particle size and crystallinity [132]. This method can be used to synthesize SPIO nanoparticles with one nanometer precision, that is, one can get batches of 4, 5, 6 and up to 20 nm with stepwise one nanometer difference and good monodispersity [132,151]. These nanocrystals are soluble in organic solvents and it is necessary to have them well dispersed in water for any biological application. Besides, there are two important prerequisites that need to be satisfied: first, the water soluble particles need to preserve their original magnetic properties after phase transfer; second, the coating materials need to be biocompatible for in vivo applications. Different materials have been chosen for nanocrystal phase transfer including small molecule ligands [152,153] and PEG-phosphine oxide polymer [154]. The amphiphilic coating on SPIO nanoparticle is usually the hydrophilic corona of PEG, which makes it almost impossible for further LbL assembly with other polyelectrolytes. Using a charged hydrophilic coating is necessary for pairing with another oppositely charged material to form LbL capsules. When designing SPIO based MRI probes, there are a few important aspects that need to be considered: 1) T2 relaxivity, this is directly related to the probe sensitivity; 2) colloidal stability; 3) biosafety, the formulation should be carefully planned because most applications require the direct intravenous injection of contrast agents; and 4) pharmacokinetics. There are a few factors that govern the T2 relaxivity of SPIO agents, including particle composition, particle size and self-assembly structure. On the composition side, Fe3O4 based nanoparticles are generally preferred because of their excellent biosafety, high magnetization and strong T2 relaxivity. In general, larger SPIO nanoparticles have higher T2 relaxivity, but particles with inorganic core diameter beyond 20 nm are usually no longer superparamagnetic

at room temperature [132]. One of the interesting discoveries is that when multiple SPIO nanoparticles are gathered together and formed small clusters, their T2 relaxivity is greatly improved over single SPIO nanoparticles, leading to a much better signal contrast enhancement [28,42,43,155–157]. The exact reason behind this phenomenon is not clear, and similar discovery has been reported by Taktak et al. for the SPIO clusters formed by mixing SPIO-biotin and avidin [158]. They found that T2 relaxation rate is proportional to the cluster size, larger ones led to higher T2 relaxivity. Polymer micelle encapsulated SPIO clusters have been demonstrated as sensitive probes for tumor imaging in vitro and in vivo[28,106], in vitro cartilage cell labeling, and in vivo stem cell monitoring [155]. This clustering concept can be applied for designing of the SPIO containing LbL capsules, and we will have detailed discussion in Section 3.5. SPIO nanoparticles and related self-assembly structures are nanocolloids, which require good dispersity in water before any biological applications. As mentioned above, larger SPIO self-assembled structures (e.g., clusters) have higher T2 relaxivity, however, if the overall size is too large they can precipitate out from solution. So a balance between the probe sensitivity and colloidal stability requires rational design. The optimal size of overall nanostructures is dependent on specific applications, for example, in vivo injection of targeting nanoprobes should have diameters smaller than 200 nm [92], preferably less than 50 nm in some examples [159,160]. When SPIO nanoparticles (ρFe3O4 = 5.2 g/cm3) are incorporated into LbL capsules, the density of the whole capsule is increased, and how to control the stability is important for either imaging or imaging/drug delivery applications. For in vivo pharmacokinetics design, how the particles interact with blood and other tissues needs comprehensive study and indepth understanding. As a general consideration for design, the coating materials on SPIO nanoparticles have to be biocompatible and biodegradable, and they won't generate immune response or cause nonspecific adsorption of serum proteins when circulating in the blood vessel because unwanted aggregates of these particles will not only damage their imaging contrast performance, but also potentially lead to toxicity in vivo. There are excellent reviews with comprehensive discussions regarding nanotoxicology and will serve as insightful references for design of nanocarriers [161–164]. The overall in vivo performance of SPIO agents requires careful considerations because there are many in vivo situations one cannot predict or simulate outside of the human body. Under regular methodology, probes are usually tested first with in vitro setups, but it is highly possible that a sensitive probe that works well for in vitro experiments may not have good performance in vivo. SPIO agents can be easily cleared by RES system if their size is too big or their surface coating is inappropriate, leading to poor imaging contrast effects regardless of their original high T2 relaxivity. Besides particle size, another important parameter that controls the agent's in vivo performance is the particle surface coating material. For the same sized nanoparticles, the coating materials, instead of the nanoparticle core, play a much more important role when interacting with cells and will have influence on the in vivo pharmacokinetics [19,150,165]. Self-assembly of nanoparticles and polymers is similar to using “bricks” and “mortar” to build designed structures with fine control [166,167]. The polyelectrolyte layer between nanoparticles is also called “electrostatic glue”, which means it serves as a bridging interlayer to hold nanoparticles into a controlled structure [168]. To obtain LbL polyelectrolyte capsules carrying SPIO agents, a simple method is to self-assemble charged SPIO nanoparticles and polyelectrolytes as part of capsule membranes (Fig. 4). Capsules with the wall structures of (nanoparticle/polyelectrolyte)n has been reported before for different systems including magnetic nanoparticles [169,170] and quantum dots [171]. Comparing to paramagnetic ion ligand complexes, SPIO nanoparticles do not require close contact with water in the capsule membrane. In the absence of external

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decides how SPIO nanoparticle should be incorporated into such a system. 3.4. LbL capsules with paramagnetic imaging moieties

Fig. 4. Schematic illustration of LbL capsules functionalized with SPIO nanoparticles.

magnetic field, the magnetic domains of SPIO nanoparticles are free to rotate from thermal motion, and are randomly oriented with no net magnetic field [116]. The external magnetic field supplies energy to overcome the anisotropy energy barrier and causes the magnetic domains to reorient as the direction of external field B0[114,116]. The SPIO based agents are also called susceptibility agents because they have much higher magnetic moment and magnetic susceptibility than the body tissues [172]. They can create local magnetic field inhomogenieties (Fig. 4) and cause dephasing of water molecules that are diffusing through [173]. Capsule size, the way and the amount of SPIO nanoparticles incorporated into the LbL capsule system may all play important roles on T2 relaxivity. In another strategy, one may load SPIO nanoparticles inside capsules instead of in the capsule membrane [174]. The specific application of multifunctional capsules

There is a wide pool of commercially available polyelectrolytes to choose for fabrication of LbL capsules. However, one has to prepare Gd-ligand modified polymers before self-assembling of capsules with MRI visibility. To make customized polymers with paramagnetic components, there are Gd-ligands with different reactive groups available for chemical conjugation. Taking the example DOTA ligand, there are some commonly used forms with NHS, azide, malemide, NH2 and COOH groups at the periphery for further reactions (Chart 2). These ligands can be conjugated to most active groups (e.g., NH2, alkyne, COOH, etc.) of a polyelectrolyte and used for subsequent LbL self-assembly. Examples of Gd-ligand conjugated polymers are: 1) non-charged synthetic polymers: polyethylene glycol (PEG) [139]; 2) proteins: albumin [140,141]; 3) polypeptides: poly-L-lysine [137,138], poly-Lglutamic acid [190,191]; 4) polysaccharides: chitosan [192], dextran [193,194], hyaluronic acid [195]; 5) poly(Gd-ligand)n [196]; 6) branched polymers: polyethylenimine [100]; and 7) copolymers: PEG-PLL [44,197], PEG-PASP [198], PAA-PMA [199]. Some of them can be used for LbL self-assembly of polymer capsules because they or their salt formulations present charges on the molecular chains (Chart 3). PLL was an early example of polymer used for conjugation of Gd-ligand and has shown improved T1 relaxivity and longer circulation time [137,138]. Polyelectrolytes can also carry drugs through conjugation method. Paclitaxel has been conjugated to PGA [200] and hyaluronic acid [201], and PGA-paclitaxel conjugates have entered the clinical trials for cancer therapy. It is feasible to design polymers carrying both drug and imaging agents, and such a polymer can be used to build multifunctional LbL capsules. Gd-ligand conjugated copolymers are useful for surface modification of LbL capsules, because it will not only introduce the paramagnetic components into the system but also provide a protection layer (e.g., PEG) for in vivo applications.

Chart 2. The list of commonly used DOTA ligand with chemical groups for conjugation: Carboxyl [175–177]; NHS [178–181]; Maleimide [182]; Amine [183]; Azide [184]; Nitrobenzyl [183]; Aminobenzyl [185–187]; and Isothiocyanatobenzyl [185,187–189].

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Chart 3. List of polymers conjugated with Gd-ligand: (a) polysaccharides; (b) polypeptides; (c) block copolymers; (d) branched polymer and poly(Gd-ligand)n.

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LbL self-assembly of capsules containing paramagnetic components have been reported in a few studies [100,196]. During the LbL self-assembly, quartz crystal microbalance (QCM) can be used as a useful tool to estimate the LbL buildup process. In one study with GdDTPA-PEI and PAS pairs, the thickness of a Gd-DTPA-PEI/PAS bilayer is about 5.2 nm, slightly lower than a PEI/PAS bilayer (6.9 nm). The increase of film thickness of (PEI/PAS) multilayers is mainly due to higher amount of PAS adsorption (average Δf = 36.2 Hz) on every PEI layer. It is possible that Gd-DTPA conjugation on primary amines of PEI provides fewer binding sites for polyanions, resulting in a reduced PAS adsorption [100]. In another study of chitosan (CH) and Gd-DTPA modified hyaluronic acid (HA), similar film growth was observed for the first five layers between Gd-DTPA-HA/CH and HA/CH pairs, but slight increment in film thickness of Gd-DTPA-HA/CH pairs with further assembly [195]. It has been found that the relaxivity values for LbL self-assembled capsule containing poly(Gd-DOTA)n on silica nanoparticles are about 19.0 ± 1.7 mM− 1 s− 1 on a per Gd basis. This is greatly improved for Gd-DOTA (4.2 Gd mM− 1 s− 1) and Gd-DOTA oligomers (6.9 Gd mM− 1 s− 1) [118,196]. Similarly, LbL capsules composed of Gd-DTPA-PEI and PAS also have enhanced T1 relaxivity reaching to 15.1 mM− 1 s− 1, comparing to 4.8 mM− 1 s− 1 for Gd-DTPA [100]. The T1-weighted image brought by this formulation has much stronger contrast than Gd-DTPA when against water samples at a clinical field of 1.5 T (Fig. 5). The relaxivity of a nanosystem can also be expressed on a per particle basis, NP mM− 1s− 1. It was noted that the per particle r1 values increase from 1.94 × 105 to 5.34 × 105 mM− 1 s− 1 with the coating shifted from one to eight Gd-ligand layers on silica particles [196]. This indicates T1 relaxivity is greatly improved for individual nanoparticles when compared to single metal complexes. Considering the potential cytoxicity of Gd ion, the relaxivity on a per Gd basis is more meaningful when the probe is designed for in vivo applications. The detailed mechanism of increased T1 relaxivity of LbL capsule is unknown, but very likely related to the prolonged rotational correlation time. Because Gd-DTPA is part of the LbL capsule system, the rotational correlation time of the capsule should have some effects on the tumbling motion of the metal–water proton vector. As discussed before and shown in Fig. 3, at lower magnetic field (0.24– 1.65 T), prolonged τR is effective to increase T1 relaxivity, but it may decrease the relaxivity at higher fields (2.35–9.4 T). For the study carried out on Gd-DTPA-PEI covered SiO2 nanoparticles, the relaxivity was measured in clinical field of 1.5 T, and belongs to the lower field range. NMRD profile can be insightful to understand the performance of a macromolecular MRI probe, because their T1 relaxivity behaviors can be compared to the small molecule agents at different magnetic fields. As discussed before, it is a general strategy to prolong the rotational correlation time by covalent or noncovalent binding of the metal complex to macro-/supramolecular structures, but the actual τR of the paramagnetic agent can be different from the carrier because of the flexibility of the linking spacer [27]. Covalent binding is usually preferred than passive adsorption because it provides a more stable bond. Surprisingly, the electrostatic adsorption of multiple Poly

Fig. 5. T1-weighted MRI images (1.5 T, spin-echo sequence: TE=5.3 ms, TE=50 ms) of water, Gd-DTPA (0.15 Gd mM) and Gd-SiO2 (Gd-DTPA-PEI/polyanion covered SiO2, 0.15 Gd mM). Adapted with permission from American Scientific Publishers [100] Copyright 2010.

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(Gd-DOTA)n polymer layers has higher T1 relaxivity on a per Gd basis than covalently attached multilayers of Gd chelates on silica nanoparticles [202]. It was suggested that the highly disordered and hydrophilic nature of Gd-DOTA oligomers and PSS on nanoparticles allows ready accessibility of water molecules to the Gd centers for efficient water proton relaxation [196]. For further increase of T1 relaxivity of LbL capsules, one may consider to fix the metal–water proton vector at the molecular barycenter, because the tumbling of such a carrier can transfer into the vector and have a longer τR [27,117]. Another interesting phenomenon is that with increased number of Gd-polymer coating, the relaxivity cannot be further increased on a per Gd basis [100,196]. For example, the T1 relaxivity is 20.0, 20.5, 16.4, and 18.2 Gd mM− 1 s− 1 for nanoparticles covered with one, two, three, and five layers of Gd-DOTA oligomer coating [196]. And for the Gd-DTPA-PEI containing capsules, one, two or three layers of GdDTPA-PEI polymer coatings on silica nanoparticles generate very similar imaging contrast against water (data not shown). Because the LbL capsule has good water permeability, so the contact between water molecules and Gd centers should not be restricted, and one possible reason may be that the rotational correlation time of Gdligands within those systems is not further increased with higher number of coating layers. Interestingly, when a piece of medical-grade polyurethane (PU) sample was coated with multiple layers of GdDTPA-hyaluronic acid/chitosan, higher number (24) of coating layers displayed better signal contrast comparing to lower ones (12 or 6) [195]. It was not described that if the T1 relaxivity was increased for the coating of 24 layers, the accumulation of higher amount of Gd in the coating film definitely contributed to the better image contrast. It is of great interest that one can control the carrier relaxivity by varying the number of assembled Gd-containing polymer layers. Such a design requires considering how Gd-ligands be conjugated to polymers, which kind of linking spacer should be used, and if crosslinking between adjacent layers would be necessary, etc. 3.5. LbL capsules containing superparamagnetic nanoparticles Magnetic nanoparticle containing LbL capsules can be considered as functional nanodevices, mainly composed of SPIO nanoparticles and polyelectrolytes, aiming at diagnosis, drug delivery or other biomedical applications. The prerequisites for LbL capsules with well controlled magnetic properties are that the magnetic nanoparticles have uniform size distribution, good crystallinity, good monodispersity in water, and charged surface. As demonstrated in many examples, nanoparticles synthesized through the high temperature decomposition of iron precursors can generate high quality iron oxide nanoparticles [132,203–207]. Different charged molecules including N-alkyl-PEI [155,208], (3-carboxypropyl)trimethylammonium chloride [153], 2-carboxyethyl phosphonate [153], and tetramethylammonium 11-aminoundecanoate [203] were used to transfer them into aqueous phase with either positive or negative charges. We will focus on N-alkyl-PEI covered SPIO nanoparticles in the following discussion because PEI is one of the most popular polyelectrolytes used in the LbL self-assembly process. Amphiphilic N-alkyl-PEI2k or N-alkyl-PEI25k was introduced for stabilizing SPIO nanoparticles in water and the composites have shown good stability and magnetic properties [155,208]. One of the most important biomedical applications for PEI is gene transfection, and PEI25k is well known for both of its high transfection efficiency and high toxicity [209]. In comparison, lower molecular weight PEI has much better biocompatibility, but suffered from low efficiency to form stable complexes with negatively charged DNA [209] or magnetic nanoparticles [210]. From our recent studies, we found that alkylated amphiphilic PEI2k was able to encapsulate one or multiple SPIO nanoparticles and form stable composites in aqueous environment [43], and they can be chosen as appropriate

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Fig. 6. TEM images of SPIO nanoparticles before (left) and after (right) transferring into aqueous phase with the help of N-alkyl-PEI2k. (Scale bar = 100 nm).

candidates for further LbL self-assembly of more complicated nanostructures. N-alkyl-PEI2k/SPIO nanocomposites have demonstrated the following properties: 1) highly positively charges with zeta-potential more than 40 mV. The highly charged surface keeps them well dispersed even after a 2 hour centrifugation at a speed of 50,000 g (Fig. 6). This strong positive charge can be advantageous for self-assembly with negatively charged polyions to build LbL particle/polymer structures; 2) good stability against pH (from 2 to 13) or ionic strength shifts. Addition of NaCl salt to such solution with a final concentration of 1 M didn't precipitate out any nanoparticles. The less salt sensitivity of N-alkylPEI2k/SPIO nanocomposites may be related to reversible protonation and chain flexibility of PEI [211]; and 3) good biocompatibility for cell labeling. When low molecular weight N-alkyl-PEI2k was used to stabilize SPIO nanoparticles and subsequent cell labeling, no obvious cytotoxicity was observed [155,212], and labeled stem cells are unaffected in their viability, proliferation, or differentiation capacity [155]. SPIO nanoparticles can be self-assembled onto any micro/nanotemplate pairing with polyelectrolytes, and a precursor film of a few polyelectrolyte bilayers would make the process easier. The anchoring density or the inter-particle distance of SPIO nanoparticles per template can be controlled by varying the coating conditions such as pH, ionic strength, and the presence of free polyelectrolyte. From our preliminary observation, higher salt concentration can lead to more SPIO nanoparticle adsorption onto SiO2 templates (Fig. 7a), and the presence of free N-alkyl-PEI2k resulting in much lower anchoring density (Fig. 7b). Silica nanotemplates covered with SPIO/polymer

capsule led to increased T2 relaxivity than single SPIO nanoparticles, SiO2 nanotemplates covered with higher SPIO nanoparticle density (Fig. 7a) displayed a 70% increase in T2 relaxivity compared with the lower density ones (Fig. 7b) and about 2.5 times higher than single N-alkyl-PEI2k/SPIO nanocomposites. The mechanism behind this increased T2 relaxivity is still under investigation. Similar T2 enhancement was also observed for other forms of small aggregates, for example, SPIO nanoparticle clustering with the help of amphiphilic or neutral-PEG polymers [28,42,43,155,157], or clusters formed by SPIO-biotin and avidin [158]. The common feature of these aggregates is that multiple SPIO nanoparticles are physically held together in a small spatial compartment. The interaction between SPIO nanoparticles is either hydrophobic, electrostatic, or biospecific interactions, which can hold nanoparticles and carrier materials together. The tumbling motion of SPIO nanoparticles is associated with the carrier, and thus has higher influence on the outer-sphere relaxation. The advantages of higher relaxivity can be expressed in two ways: 1) at the same iron concentration, the SPIO containing capsules have better sensitivity; 2) or use much smaller dosage, achieving similar image contrast as single SPIO nanoparticles, resulting in better biocompatibility. The templates of SPIO containing capsules can be removed and shells can then be used as drug carriers with MRI visibility. Either matrix-containing or matrix free capsules were developed with SPIO nanoparticle in capsule membranes. And for longer circulation time in vivo, PEG-polyelectrolyte can be added as the outermost layer for protection. In one example, SPIO nanoparticles were first coated with PEI to generate the positive charge, followed by PEG-block-poly

Fig. 7. N-alkyl-PEI2k/SPIO nanocomposites adsorbed on polyelectrolyte covered SiO2 nanotemplates with a) higher and b) lower anchoring density. (Scale bar = 100 nm).

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(glutamic acid) adsorption with a final slightly negatively charged surface of −3 mV [213]. This nanocomposite has demonstrated good in vivo liver contrast imaging effects as compared to a commercial product Resovist®. The loading efficiency in preformed hollow capsules can be improved by adding charged matrix inside capsules. Recently, hybrid microparticles composed of MnCO3 and ionic polysaccharide including sodium hyaluronate, sodium alginate or dextran sulfate sodium with narrow size distribution were developed as templates (Fig. 8a) [75]. Confocal image of capsules in Fig. 8b indicates the existence of HA matrix in capsules after removal of MnCO3. The polysaccharide matrix containing capsules has much higher loading efficiency up to four times as the matrix-free capsules [75]. Similarly, polycation/ inorganic particle templates can be fabricated, and PEI matrix containing capsules were obtained after dissolving of inorganic part. SPIO nanoparticles can be incorporated into the capsule membrane (Fig. 9a) and used as a T2 based MRI probe. We have embedded SPIO containing capsules in collagen gels and injected them subcutaneously in mice. It has been shown that LbL SPIO capsules with a concentration of 8.8 × 104 capsule/mm3 can generate strong image contrast against the same number of SPIO-free capsules, with a signal intensity ratio between them around 5.0 (Fig. 9b). The applications of SPIO capsules can be categorized as 1) dual functional carriers having both imaging and drug delivery functions, the imaging function is mainly used to track capsules in vivo; 2) molecular imaging for diagnosis: for example, nanosized SPIO capsules with targeting ligands can be injected in vivo and preferentially accumulate at specific sites for discovering small

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pathological changes; 3) biosensor for measuring trace amount of small molecules or biomacromolecules; 4) cell labeling for tracking their behaviors in vivo: the SPIO capsules can be used to label stem cells, T cells, or dendritic cells and re-inject them back to the body. The non-invasive and dynamic imaging tracking has been demonstrated as a very informative method to evaluate the efficacy of cell therapies [214]. Taking the example of cell labeling, SPIO capsules were chosen to label HEK293 cells in vitro, and then collected and mixed with collagen gel and placed underneath of mouse skin. Comparing to the same amount of cells without MRI probe labeling, the labeled cells (1 × 104 cells/mm3) display hypointense signals and have strong contrast against cells without labeling (Fig. 9c). This cell density is about ten times lower than most tissues, so it may have great potentials for tracking a small number of cells noninvasively in vivo. Non-invasive imaging tracking certainly has advantages over histological studies, which require sacrificing a lot of animals at different time points and encounter individual difference. One general disadvantage of cell MRI labeling is that the imaging contrast can be diluted with every cycle of cell division, and after a few cycles of division, the remained MRI probe in cells may not be strong enough to generate detectable signals. One solution is to use relatively large magnetic particles (0.76–1.63 μm) for cell labeling, because even after many cell divisions, daughter cells still carry individual particles, that is, only one particle inside a cell and cell divisions cannot dilute the label [215]. And these single particles can generate strong signals to be detected at a resolution of 50 μm under MRI scans [215]. Besides the dilution factor and particle size, attentions should also be paid to the

Fig. 8. Confocal laser scanning microscopy (CLSM) images of a) hybrid microparticles (composed of MnCO3 and 5-([4,6-dichlorotriazin-2-yl]amino)-fluorescein hydrochloride (DTAF) labeled HA) and b) HA matrix containing microcapsules; c) fluorescence intensity profile of a polyelectrolyte capsule; d) SEM image of HA containing capsules. Adapted with permission from Elsevier [75] Copyright 2011.

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iron metabolism after the iron released from particle degradation. Iron is a functional component of oxygen-carrying globin proteins, cytochromes, and enzymes that transfer electrons, and disrupting of iron homeostasis could lead to the development of various disorders [133,217–220]. The excess accumulation of intracellular iron can oxidize and damage the protein and nucleic acid components of cells. Under conditions where intracellular iron is high, iron-binding proteins bind to iron responsive elements and regulate mRNA stability and translational activity of TfR-1 and ferritin, respectively [221,222]. RT-PCR results have shown that N-alkyl-PEI2k/SPIO composite labeled MSCs resulted in a significant decrease of TfR-1 mRNA expression than unlabeled ones as early as one day after labeling [155], possibly due to a feedback of the intracellular iron increase [223,224]. An increasing trend in ferritin gene transcripts was observed with a significant increase in ferritin mRNA levels from labeled MSCs after one and three days [155,223,224]. It is well known that ferritin synthesis can be induced with an increase in cytoplasmic iron content, resulting in sequestration of extra iron into the ferritin molecules [225]. Up-regulated ferritin expression at protein level was also detected for labeled chondrocyte cells [212]. Down-regulation of transferrin receptor and up-regulation of ferritin in labeled cells may protect cells from possible cytotoxicity led by higher intracellular iron concentrations. Other iron cellular import and export regulating proteins such as divalent metal transporter 1 (DMT-1), ferroportin, natural resistance associated macrophage protein 1 (Nramp1) should be investigated in further studies. Iron metabolism of SPIO probes should be taken into consideration as part of the probe design. 4. Conclusions LbL capsules with MR imaging capability have been covered in this review from the point of design criteria to examples. In general, the LbL self-assembly technique is a versatile tool and gives one freedom to develop capsules with a wide range of materials including polymers, lipids, and nanoparticles. Important physical properties such as capsule size and capsule wall molecular structure can be well designed and precisely tuned. Either paramagnetic-ligand molecules or superparamagnetic nanoparticles have been incorporated into the capsule system and resulted in increased T1 or T2 relaxivities. Further utilization of these sensitive probes for in vivo diagnosis, drug delivery or theranostics requires rational design and comprehensive studies.

Fig. 9. a) TEM image of SPIO nanoparticle decorated LbL capsules; b) LbL capsules were mixed with collagen and implanted subcutaneously of a mouse with a capsule concentration of 8.8 × 104/mm3 for 1) without (dashed blue line circled area) and 2) with SPIO nanoparticles in capsule wall (dashed yellow line circled area) under T2-weighted scan; c) HEK293 cells were mixed with collagen and implanted subcutaneously of a mouse with a cell concentration of 1 × 104/mm3 for 3) without (dashed blue line circled area) and 4) with SPIO capsule labeling (dashed yellow line circled area) under T2-weighted scan.

fact that we are imaging the labeled cells instead of the probes themselves [216]. Because imaging probes can be detached from cell surface if they are not internalized during labeling or leaked out after cell dies, leading to uptake by neighbored cells or left in extracellular matrix. This will make the judgment of cell locations as a difficult job and possibly lead to wrong conclusions. Comprehensive understanding of the clearance pathways and timelines of probes beforehand will be helpful to design labeling strategies with higher tracking accuracy [216]. One of the most important biosafety issues related to SPIO capsules is how the iron oxide nanoparticles will interfere with physiological

1) Nanotoxicology: the comprehensive biosafety studies of LbL capsules are the foundations for their biomedical imaging applications. Most LbL capsules are currently studied at in vitro environment, which cannot satisfy the preclinical requirements. Further investigation of their in vivo pharmacokinetics, pharmacodynamics, and metabolism in animal models is helpful for better design of highly sensitive imaging probes. 2) Multimodality imaging probes: the LbL capsule platform can be easily adapted to other imaging modalities besides MRI. Incorporation of both MRI and optical/nuclear imaging moieties enables one to obtain anatomical, functional and molecular information with high sensitivity. It is important for one to design capsules that each probe's imaging properties are not compromised or interfered with each other, and besides special care should be given to optical probes because self-assembled fluorescence moieties such as incorporation of quantum dots into aggregated structures may potentially lower the quantum yield. PET/MRI dual modality capsules can be used to discover small pathological changes in deep tissues with good sensitivity and precise locations. 3) Smart diagnostic probes: to discover small changes at cellular and molecular levels require sensitive probes with targeting function and signal amplification capability. LbL capsules responsive to specific pathological environment such as pH difference, higher macrophage accumulation at vascular plaques, or high matrix

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metalloproteinase expression will trigger signal intensity change and lead to better contrast. 4) Multifunctional nanocarriers: LbL capsules are promising carriers for controlled drug release. Combining imaging, targeting and drug delivery functions into one carrier will expand its functionality and provide in-depth understanding of the delivery mechanism, leading to a better design of drug carrier systems. Special attention should be paid to drug/imaging probe interactions, controlled drug delivery and release strategies, and accurate imaging of drug and drug carriers. Acknowledgments The work was supported by Program for New Century Excellent Talents in University (NCET-06-0781), Doctoral Fund of Ministry of Education of China (20090181110068), Distinguished Young Scholars Project of Sichuan Province (06ZQ026-007), National Natural Science Foundation of China (30570514, 20974065, 50603015 and 50830107) and Sichuan University 985 Platform of Biomedical Engineering (Ministry of Education). References [1] C.M. Tempany, B.J. McNeil, Advances in biomedical imaging, JAMA 285 (2001) 562–567. [2] R. Weissleder, U. Mahmood, Molecular imaging, Radiology 219 (2001) 316–333. [3] H.R. Herschman, Molecular imaging: looking at problems, seeing solutions, Science 302 (2003) 605–608. [4] F.A. Jaffer, R. Weissleder, Molecular imaging in the clinical arena, JAMA 293 (2005) 855–862. [5] J.M. Hoffman, S.S. Gambhir, Molecular imaging: the vision and opportunity for radiology in the future, Radiology 244 (2007) 39–47. [6] R. Weissleder, M.J. Pittet, Imaging in the era of molecular oncology, Nature 452 (2008) 580–589. [7] M.G. Harisinghani, J. Barentsz, P.F. Hahn, W.M. Deserno, S. Tabatabaei, C.H. van de Kaa, J. de la Rosette, R. Weissleder, Noninvasive detection of clinically occult lymph-node metastases in prostate cancer, N. Engl. J. Med. 348 (2003) 2491–2499. [8] S.S. Gambhir, Molecular imaging of cancer with positron emission tomography, Nat. Rev. Cancer 2 (2002) 683–693. [9] S.Y. Shaw, Molecular imaging in cardiovascular disease: targets and opportunities, Nat. Rev. Cardiol. 6 (2009) 569–579. [10] J.C. Wu, F.M. Bengel, S.S. Gambhir, Cardiovascular molecular imaging, Radiology 244 (2007) 337–355. [11] R. Weissleder, Molecular imaging in cancer, Science 312 (2006) 1168–1171. [12] P. Lang, K. Yeow, A. Nichols, A. Scheer, Cellular imaging in drug discovery, Nat. Rev. Drug Discov. 5 (2006) 343–356. [13] M. Rudin, R. Weissleder, Molecular imaging in drug discovery and development, Nat. Rev. Drug Discov. 2 (2003) 123–131. [14] J.K. Willmann, N. van Bruggen, L.M. Dinkelborg, S.S. Gambhir, Molecular imaging in drug development, Nat. Rev. Drug Discov. 7 (2008) 591–607. [15] A. Watrin-Pinzano, J.P. Ruaud, Y. Cheli, P. Gonord, L. Grossin, I. BettembourgBrault, P. Gillet, E. Payan, G. Guillot, P. Netter, D. Loeuille, Evaluation of cartilage repair tissue after biomaterial implantation in rat patella by using T2 mapping, MAGMA 17 (2004) 219–228. [16] D. Kozlowska, P. Foran, P. MacMahon, M.J. Shelly, S. Eustace, R. O'Kennedy, Molecular and magnetic resonance imaging: the value of immunoliposomes, Adv. Drug Deliv. Rev. 61 (2009) 1402–1411. [17] J.R. McCarthy, R. Weissleder, Multifunctional magnetic nanoparticles for targeted imaging and therapy, Adv. Drug Deliv. Rev. 60 (2008) 1241–1251. [18] A.M. Smith, H. Duan, A.M. Mohs, S. Nie, Bioconjugated quantum dots for in vivo molecular and cellular imaging, Adv. Drug Deliv. Rev. 60 (2008) 1226–1240. [19] C. Sun, J.S. Lee, M. Zhang, Magnetic nanoparticles in MR imaging and drug delivery, Adv. Drug Deliv. Rev. 60 (2008) 1252–1265. [20] V.P. Torchilin, PEG-based micelles as carriers of contrast agents for different imaging modalities, Adv. Drug Deliv. Rev. 54 (2002) 235–252. [21] V.P. Torchilin, Multifunctional nanocarriers, Adv. Drug Deliv. Rev. 58 (2006) 1532–1555. [22] S.M. Janib, A.S. Moses, J.A. Mackay, Imaging and drug delivery using theranostic nanoparticles. Adv. Drug Deliv. Rev. 62 (2010) 1052–1063. [23] V. Ozdemir, B. Williams-Jones, S.J. Glatt, M.T. Tsuang, J.B. Lohr, C. Reist, Shifting emphasis from pharmacogenomics to theragnostics, Nat. Biotechnol. 24 (2006) 942–946. [24] W.J. Mulder, G.J. Strijkers, G.A. van Tilborg, D.P. Cormode, Z.A. Fayad, K. Nicolay, Nanoparticulate assemblies of amphiphiles and diagnostically active materials for multimodality imaging, Acc. Chem. Res. 42 (2009) 904–914. [25] A. Fatemi-Ardekani, N. Samavati, J. Tang, M.V. Kamath, Advances in multimodality imaging through a hybrid PET/MRI system, Crit. Rev. Biomed. Eng. 37 (2009) 495–515.

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