Bioinspired Coordination Micelles Integrating High Stability, Triggered

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Bioinspired Coordination Micelles Integrating High Stability, Triggered Cargo Release, and Magnetic Resonance Imaging Keting Xin, Man Li, Di Lu, Xuan Meng, Jun Deng, Deling Kong, Dan Ding, Zheng Wang, and Yanjun Zhao ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.6b09425 • Publication Date (Web): 13 Dec 2016 Downloaded from http://pubs.acs.org on December 17, 2016

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ACS Applied Materials & Interfaces

Bioinspired Coordination Micelles Integrating High Stability, Triggered Cargo Release, and Magnetic Resonance Imaging Keting Xin, 1,† Man Li, 1,† Di Lu, 1 Xuan Meng, 1 Jun Deng, 1 Deling Kong,2,3 Dan Ding,2,3 Zheng Wang, 1 Yanjun Zhao*,1 1

School of Pharmaceutical Science & Technology, Tianjin Key Laboratory for Modern Drug Delivery & High Efficiency, and Collaborative Innovation Center of Chemical Science and Engineering (Tianjin), Tianjin University, Tianjin 300072, China

2

Key Laboratory of Bioactive Materials, Ministry of Education, College of Life Science, and Collaborative Innovation Center of Chemical Science and Engineering (Tianjin), Nankai University, Tianjin 300071, China 3

State Key Laboratory of Medicinal Chemical Biology, Nankai University, Tianjin 300071, China

ABSTRACT: Catechol-Fe3+ coordinated micelles show the potential for achieving ondemand drug delivery and magnetic resonance imaging (MRI) in a single nanoplatform. Herein, we developed bioinspired coordination-crosslinked amphiphilic polymeric micelles loaded with a model anti-cancer agent, doxorubicin (Dox). The nanoscale

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micelles could tolerate substantial dilution to a condition below the critical micelle concentration (9.4 ± 0.3 µg/mL) without sacrificing the nanocarrier integrity due to the catechol-Fe3+ coordinated core crosslinking. Under acidic conditions (pH 5.0), the release rate of Dox was significantly faster compared to that at pH 7.4 as a consequence of coordination collapse and particle de-crosslinking. The cell viability study in 4T1 cells showed no toxicity regarding placebo crosslinked micelles. The micelles with improved stability showed a dramatically increased Dox accumulation in tumors and hence the enhanced suppression of tumor growth in a 4T1 tumor-bearing mouse model. The presence of Fe3+ endowed the micelles T1-weighted MRI capability both in vitro and in vivo without the incorporation of traditional toxic paramagnetic contrast agents. The current work presented a simple “three birds with one stone” approach to engineer the robust theranostic nanomedicine platform.

KEYWORDS: drug delivery, micelle, stability, stimuli-responsive, magnetic resonance imaging. 1. INTRODUCTION: The rapid development of nanoscience and nanotechnology has revolutionized the field of drug delivery.1,2 Multiple physiological and transport barriers have to be overcome to realize efficient drug delivery and satisfactory therapeutic outcomes. These include nonspecific biodistribution, degradation within biological microenvironment, elimination by mononuclear phagocyte system, cellular internalization, endosomal escape, and nuclear transport. Nanomaterials or nanocarriers show unique physical, chemical, optical, magnetic, and electronic properties that could be facilely tuned by varying their size,


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shape, composition, and architecture.3 Hence, the above issues could be addressed by nanotechnology to achieve drug solubilization, controlled or sustained release, tight junction circumvention, cell- or tissue-specific targeting, intracellular delivery of biopharmaceuticals, co-delivery of multiple drugs, and combination of therapeutic agents with imaging modalities. There are currently over 200 nanomaterial-based medical products that have been approved by the Food and Drug Administration (FDA) or in different stages of clinical study.4 Precise particle engineering enables stimuli-responsive nanocarriers that make ondemand drug delivery possible in a spatial-, temporal- and dosage-controlled manner.5,6 Nanoscale triggered delivery systems are feasible at both tissue and cellular level via taking advantage of the specific microenvironmental changes associated with either physiological site/location or pathological situation. There are various endogenous stimuli available such as pH, redox potential, and enzymes; likewise, external physical stimuli can also be utilized including temperature, light, ultrasound, electric pulse, and magnetic field.7 Among these, pH-responsive nanocarrier is one of the systems that attracted most attention due to the abundance of pH-labile materials and the unique acidic microenvironment of endosomes/lysosomes.8 Integrating drug delivery and in situ imaging in one single stimuli-responsive nanocarrier is appealing due to its potential of providing real-time feedback on the pharmacokinetics, the target site localization and the off-target organ accumulation of nanomeidinces.9 Among these theranostic nanosystems, magnetic resonance imaging (MRI) is unique due to its non-invasive nature. The commonly used paramagnetic Gd3+based T1 contrast agents are known for their difficulty in metabolism and severe adverse



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effects.10 The superparamagnetic Fe3O4 nanoparticle-based T2 contrast agents have the problem of poor drug loading.11 Therefore, exploring alternative safe contrast agents is demanded and the paramagnetic Fe3+ is a good choice since it has five unpaired electrons as T1-based MRI agent and naturally exists in the living systems with good biocompatibility and known metabolism.12 The multifunctional theranostic nanocarriers are often characteristic of complicated particle architecture and tedious fabrication processes, which hinders their scaling-up manufacture, quality control, and hence clinical translation.13-15 Therefore it is imperative to develop theranostic nanosystems via a simple and robust means. In 2007, Lee et al. first reported that polydopamine could be used as a versatile coating material for various surfaces including metal surfaces.16 Inspired by this discovery, Holten-Andersen et al. found that the stable catechol-Fe3+ coordination is pH-dependent, i.e. mono- (pH < 5.6), bis- (5.6 < pH < 9.1), and tris- (pH > 9.1), and employed this phenomenon to construct non-covalent self-healing polymer network.17 Afterwards, Hwang et al. utilized this unique and strong cathecol-Fe3+ crosslinking in mussel to produce coordination nanostructure and achieved pH-dependent on-demand drug release.18 This study is quite interesting as it provides a method for fabricating reversibly ironically crosslinked nanocarriers for intracellular pH-triggered drug delivery (e.g. in endosomes/lysosomes) without premature dose leakage in blood circulation. Meanwhile, there have been growing interests in dopamine-based material and structure design in biomedical, energy, and environmental fields, which has been reviewed by Liu et al.19 In addition, Fe3+ can be easily incorporated within the nanocarrier with the aid of catechol for MRI purpose.20


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It was postulated that the distinctive cathecol-Fe3+ interaction could be utilized to construct robust nanocarriers with high stability, pH-triggered drug release, and MRI capacity. Amphiphilic polymeric micelles were selected as the model theranostic nanocarriers owing to their exceptional tunable physiochemical properties with ease.21 Hence the aim of the current work was to utilize a simple bioinspired approach to generate a robust “three birds with one stone” theranostic nanoplatform with high stability, pH-triggered drug delivery and in situ MRI ability. A self-assembling polymer conjugate was designed containing three domains, i.e. hydrophilic methoxyl poly(ethylene glycol) (mPEG), hydrophobic Trityl group, and catechol-presenting dopamine (Dopa) (Scheme 1). Doxorubicin (Dox) was chosen as the model drug and the in vivo performance of Dox-loaded micellar nanocarrier was assessed in a tumor-bearing mouse model.

Scheme 1. Illustration of the bioinspired multifunctional theranostic micelle. The bivalent catechol-Fe3+ coordination at neutral condition could produce core-crosslinked micelles with high systemic stability. Upon pH reduction, the core de-crosslinking process would



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enable pH-triggered drug release. Such design could also make possible T1-weighted magnetic resonance imaging. Dox represents the model drug doxorubicin.

2. RESULTS AND DISCUSSION 2.1 Synthesis and Characterization of Amphiphilic Polymer Conjugate. The synthesis of catechol-containing amphiphilic polymer conjugate primarily involves four major steps (Scheme 2). The conjugate was designed to incorporate three moieties; besides catechol, it is the mPEG and Trityl moieties that enable the polymer amphiphilic and thus endow it self-assembly property. In order to combine these three moieties together, the trivalent cysteine was selected as the linker molecule. Commercial available S-Trityl-Cys was picked up as the starting material; the amine group was first protected to generate the intermediate product, N-Boc-S-Trityl-Cys (Figure S1-S3) that was further conjugated with Dopa via amidation to produce the second intermediate, N-Boc-S-TritylCys-Dopa (Figure S4-S6). After deprotection of amine, the obtained S-Trityl-Cys-Dopa (Figure S7-S9) was linked to carboxyl-terminated methoxy PEG (mPEG-COOH, Figure S10-S12) via the amide bond to make the target product, mPEG-S-Trityl-Cys-Dopa (Figure S13-S15). All the intermediate and target products were purified via an appropriate method including precipitation, dialysis and flash chromatography. Successful synthesis was verified by nuclear magnetic resonance (1H NMR and NMR), and mass spectrometry (MS).


13

C

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Scheme 2. Synthetic route the amphiphilic polymer conjugate, mPEG-S-Trityl-Cys-Dopa. The use of PEG as the hydrophilic segment has been the golden standard for engineering amphiphilic polymers due to its biocompatibility, tunable molecular size, and the diversity of terminal groups (e.g. –OH, -COOH, and –NH2) for functionalization.22 Although a relatively low molecular weight (MW) mPEG (550 Da) was used in the current work, it still exhibited the nature of polydispersity regarding MW distribution and the polymerization degree ranged from 8 to 13 based on MS analysis (Figure S10). Hence, the target catechol-bearing polymer conjugate also inherited the polydisperse feature despite its comparatively low MW of less than 1,500 Da (Figure S10). As the biomedical application of high MW non-biodegradable polymers has been often questioned about their in vivo elimination and potential side-effects, such smaller conjugate would not present such hazard since its size is much below the threshold of kidney filtration for many typical polymers (at the magnitude of 10 k Da).23,24 It should be noted that the employment of trityl group was to simplify the synthetic and purification process as a “proof-of-concept” study. The hydrophobic segment could be easily replaced with other moieties via selecting appropriate trivalent linker and utilizing



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other biocompatible hydrophobic molecules such as tocopherol, cholesterol, and fatty acids. 2.2 Generation and Characterization of Polymeric Micelles. The preparation of polymer conjugate micelles employed the conventional dialysis method.25 The supplementation of Fe3+ in the dialysis medium followed by consequent purification process for removal of excess free Fe3+ generated catechol-Fe3+ coordination crosslinked placebo

micelles.

The

presence

of

nitrilotriacetic

acid

in

the

3-

morpholinopropanesulfonic acid (MOPS) buffer aimed to avoid Fe3+ precipitation.26 Drug-loaded micelles were produced in a similar way with sufficient Dox feeding in the dialysis tube. At neutral condition (pH 7.4), single Fe3+ could theoretically complex with two catechol group.17 The Dopa/Fe3+ molar ratio in both crosslinked micelles was determined at 2.1 ± 0.1 (placebo) and 1.9 ± 0.2 (drug-loaded), respectively. The iron content in the placebo micelle (Fe) and Dox-loaded micelle (Fe) was 18.7 ± 0.9 (mg/g) and 21.0 ± 2.6 (mg/g), correspondingly; there are no significant difference between iron mass in both samples (p > 0.05). The transmission electron microscopy (TEM) demonstrated that all four types of micelles exhibited sphere morphology with a core size ca. 70-100 nm (Figure 1). The hydrodynamic size of each sample was bigger than the corresponding TEM diameter (Table 1). The crosslinking generated smaller compact micelles compared to the ironfree samples. This might be a result of closer inter-polymer alignment induced by catechol-Fe3+ coordination. Irrespective of crosslinking, the presence of Dox in the micelles slightly increased the particle size due to the expansion of the core, which was a routine phenomenon for cargo-loaded micelles.27 Upon Dox encapsulation, the zeta


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potential of micelles increased, suggesting a less negatively charged surface, which could be attributed to the presence of cationic amine group in Dox. Such result occurred well with previous independent investigation.28

Figure 1. Transmission electron microscopy images (left) of un-crosslinked (a) placebo and (c) doxorubicin-loaded micelles, catechol-Fe3+ coordination crosslinked (b) placebo and (d) doxorubicin-loaded micelles, scale bar: 200 nm. The critical micelle concentration (CMC) plot of un-crosslinked placebo micelles was shown on the right. Table 1. The particle size and surface charge of four types of polymer conjugate micelles (n = 3). The hydrodynamic diameter and TEM diameter was obtained by dynamic lighter scattering and transmission electron microscopy analysis, respectively. Dox and (Fe) represent doxorubicin, and catechol-Fe3+ coordinated core crosslinking of micelles.

Hydrodynamic

TEM diameter

Zeta potential

diameter (nm)

(nm)

(mV)

Placebo micelle

111.3 ± 1.6

89.2 ± 9.3

-11.5 ± 0.7

Placebo micelle (Fe)

90.0 ± 0.7

69.2 ± 10.3

-18.1 ± 2.0

Dox micelle

116.7 ± 1.5

94.4 ± 7.9

-2.3 ± 0.5

Dox micelle (Fe)

95.5 ± 1.6

73.8 ± 9.0

-3.6 ± 1.4

Sample



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2.3 In Vitro Stability Assessment. The critical micelle concentration (CMC) is a key indicator of micellar nanocarrier’s thermodynamic stability. The CMC of un-crosslinked placebo micelle was 9.4 ± 0.3 µg/mL in accordance with a fluorescence probe-based method (n = 3) (Figure 1).29 The driving force of polymer self-assembly and micelle formation is the reduction of system free energy that is associated with the rearrangement of hydrophobic fragments within the inner core and the hydrophilic moiety being exposed to the external aqueous medium.30 The self-assembling performance of polymers lies in their amphiphilicity that are primarily determined by the polymer composition together with cargo encapsulation and environmental conditions. The reason to choose low MW PEG is to balance the conjugate’s hydrophilicity and hydrophobicity since the molecular mass of hydrophobic Trityl moiety is low. Longer PEG chain could render the conjugate too hydrophilic to lose the assembling capacity. As the micellar formulations generally experience a large ratio of dilution in the systemic circulation upon in vivo dosing, their physical stability is critical for drug delivery efficiency and hence therapeutic efficacy. Because the micelles’ in vivo stability could be reasonably predicted by their in vitro performance,31 we carried out a simple dilution experiment to assess the stability iron-crosslinked micelles (Figure 2). Either aqueous PBS (pH 7.4) or organic dimethyl formamide (DMF) was employed as the dilution medium. When being treated with 1,000 times dilution by PBS, the placebo micelles experienced considerable diameter change evidenced by the formation of large aggregates at micrometer scale. In contrast, the hydrodynamic size shift of ironcrosslinked placebo micelle (Fe) was less noteworthy. Since the obtained conjugate was


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fully soluble in DMF, the employment of this organic solvent as dilution vehicle could generate more distinct contrast between iron-crosslinked and iron-free samples.32 In fact, the DMF dilution (5 times) completely destroyed un-crosslinked placebo micelle, but the presence of catechol-Fe3+ coordination diminished such damaging effect on micelle’s integrity. All dilution curves clearly demonstrated that simple iron-mediated coordination could substantially enhance micelle stability without chemical crosslinking that was not only synthetically tedious and un-scalable, but also challenging in terms of product purification.33 In fact, the crosslinked micelles also exhibited pH-dependent particle size due to the variation of crosslinking degree (Figure S16). This was a consequence of the reduction of catechol/Fe3+ complexation ratio from 2 (at pH 7.4) to 1 (at pH 5.0),17 which induced de-crosslinking and particle swelling or disassembly.

Figure 2. Stability assessment of micelles via dilution with phosphate buffered saline (PBS) or dimethyl formamide (DMF). The hydrodynamic size of micelles was used as the indicator. The micelle concentration was set at 1 mg/mL prior to dilution for all



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samples; the ratio of dilution in PBS (a, b) and DMF (c, d) was 1000 and 5 times, respectively. The symbol (Fe) represents catechol-Fe3+ coordinated core crosslinking. 2.4 In Vitro Drug Release. The high performance liquid chromatography (HPLC) assay revealed a drug loading of 1.8 ± 0.1% (w/w) and 1.7 ± 0.2% (w/w) for un-crosslinked and crosslinked micelles (n = 3), respectively (p > 0.05). The corresponding mean encapsulation efficiency was calculated as 36% (un-crosslinked) and 34% (crosslinked). Typically the loading of hydrophobic active agents in nanocarriers was less than 5%. It has been reported that the polymer-drug miscibility is vital in determining the drug loading capacity in polymer micelles and a longer hydrophobic block is beneficial to enhance loading.34 In the current work, the Dox was converted to its hydrophobic version to facilitate its micellar encapsulation via hydrophobic interaction with the carrier. However, the small Trityl moiety could not contribute to a very high loading. Since the released compound was hydrophobic Dox, the in vitro drug release experiment employed sodium dodecyl sulfonate (SDS)-containing buffer as the release medium for solubility enhancement.35 The solubility of Dox in the above medium at ambient temperature was 402 ± 3 (µg/mL). The dosing in the donor compartment was controlled to aid the trouble-free drug quantification and allow the equilibrium Dox concentration in the release medium is fairly below 10% of its saturated solubility to maintain a sink condition. Both un-crosslinked and crosslinked micelles showed an acidity-dependent release profile; Dox release was favored at pH 5.0 other than at pH 7.4 (Figure 3). For the crosslinked micelles, the theoretical average coordination ratio between catechol and iron reduced from 2 (at pH 7.4) to 1 (at pH 5.0).19 Therefore, the extent of crosslinking dramatically weakened, leading to pH-triggered rapid Dox release


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as a consequence of steady particle de-assembly. The trend concurred well with previous work that utilized different catechol polymer to realize on-demand cargo release,18,36 implying that catechol-Fe3+ interaction could be a versatile mechanism for pH-responsive drug release. At the absence of iron, there was more cumulative drug at pH 5.0 in comparison to that at pH 7.4, which was largely a result of the enhanced drug solubility at pH 5.0 due to the ionization of amine group in Dox.

Figure 3. In vitro release profiles of doxorubicin (Dox) from two types polymer micelles at pH 7.4 or pH 5.0. The symbol (Fe) represents catechol-Fe3+ coordinated core crosslinking. The release medium contained 5% sodium dodecyl sulfonate to keep sink condition. Data were presented as the percentage of cumulative released drug against time (mean ± standard deviation/SD, n = 3). 2.5 Cytotoxicity and Cellular Uptake. The cell viability study revealed that the drugfree micellar nanocarriers showed no noticeable cytotoxicity in 4T1 cells irrespective of the presence of iron (Figure 4). The IC50 of both types of Dox-loaded micelles were similar at 1.8 ± 0.1 µg/mL (un-crosslinked) and 1.6 ± 0.2 µg/mL (crosslinked), respectively (p > 0.05). It seemed that the different in vitro drug release behavior of both



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micelles did not dominate their cellular toxicity. It was assumed that the majority of Dox was released from both nanocarriers at the time point of cytotoxicity assessment (at 24 h post dosing) due to the perfect intracellular sink condition.

Figure 4. The dose-dependent cell viability of 4T1 cells in response to four types of micelles (n = 6). The assessment was carried out using standard MTT assay. The symbol (Fe) represents catechol-Fe3+ coordinated core crosslinking. Dox indicates model drug doxorubicin. Taking advantage of the inherent red fluorescence of Dox, no additional drug labelling procedure is required for cellular uptake study. The confocal images proved that both drug-loaded micelles could be successfully internalized by 4T1 cells and the nuclei accumulation of Dox increased gradually from 2 h to 4 h (Figure 5). Although the detailed mechanism of cellular uptake was not investigated, both samples should undertake similar endocytosis approach since their size and surface charge were analogous. At 2 h, the Dox mainly located in the cytoplasm for crosslinked micelles, but a considerable amount of Dox transported into the nuclei from the un-crosslinked counterpart. This could be explained by the relatively poor drug release from the


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catechol-Fe3+ coordinated nanocarrier at the beginning, which was consistent with the in vitro release profile (Figure 3). As the nuclear entry is a multiple-step process involving cellular uptake, drug release, endosomal escape, and transport into nuclei,37 the crosslinked micelles slightly delayed the passage of Dox to nuclei.

Figure 5. Intracellular uptake of two types of nanocarriers: Dox micelle, and Dox micelle (Fe). The free Dox·HCl (doxorubicin hydrochloride) was used as the control. The symbol (Fe) represents catechol-Fe3+ coordinated core crosslinking. Images were taken by the confocal laser scanning microscope at 2 h and 4 h. The blue and red colors signify the location of nuclei and model drug (Dox). The color overlap indicates the nuclear entry of Dox. Scale bar: 20 µm. 2.6 Ex Vivo Biodistribution. To get a brief semi-quantitative information of drug location upon animal dosing, we performed an ex vivo biodistribution study (Figure 6). Due to the incapability of Dox in the near infrared (NIR) region, the whole animal imaging was not feasible without additional NIR probe labelling. Compared to the control (free Dox·HCl), both Dox-loaded micelles with or without crosslinking significantly delivered more drug



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to the tumor site (p < 0.05). Such result was presumed due to the enhanced permeability and retention (EPR) effect that was characteristic of enhanced permeability of tumor vasculature and impaired lymphatic drainage.38 The appropriate particle size and negative surface charge of both Dox-containing micelles contribute to prolonged systemic circulation, avoidance of non-specific binding, and hence improved EPR effect. Interestingly, at all three time points (2 h, 6 h, and 24 h), the Dox fluorescence in the tumor (as an indicator of Dox content) was significantly higher for the crosslinked micelle compared to the un-crosslinked control (p < 0.05). Such result was believed as a consequence of coordination-mediated micellar stabilization. First of all, the enhanced stability would minimize the premature drug release in the blood circulation, and thus there were more drugs available when the micelle reached the tumor. In addition, the high stability could aid the retention of both nanocarrier and Dox in the tumor. The elevated drug content in the disease site would apparently enhance the antitumor effect.

Figure 6. Semi-quantitative ex vivo fluorescence imaging of major healthy tissues and tumor collected from 4T1 tumor-bearing mice at 2 h, 6 h, and 24 h after intravenous


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injection of (I) Dox micelle (Fe), (II) Dox micelle, and (III) Dox·HCl (top). Quantitative analysis of fluorescence signals from the tumors and normal tissues was shown on the bottom (n = 3). The drug dose was controlled as the same for all samples. Dox and (Fe) represent doxorubicin, and catechol-Fe3+ coordinated micelle core crosslinking. The star represents a significant difference (p < 0.05) Apart from the tumor site, Dox also distributed in heart, liver, spleen, lung and kidney, particularly in the highly perfused organs. Over the time course of 24 h, the amount of Dox varied in different organs from crosslinked and un-crosslinked micelles (Figure 6). Quantitative analysis of Dox content in tissues by HPLC-MS agreed well with the fluorescence analysis (Figure S17). However, at all three time points, both micelles delivered more Dox to the tumor compared to the free drug (p < 0.05) and crosslinked micelles were significantly superior to the un-crosslinked control regarding drug content in tumor (p < 0.05). In addition, the uptake by mononuclear phagocyte system in these organs was a factor of this phenomenon. The drug accumulation in all these healthy organs is inevitable for antitumor nanomedicine as an outcome of biodistribution.39 Usually the majority of therapeutic agents located in these non-target organs via blood circulation and extravasation after intravenous administration, and only less 5% of total formulation (in most cases) deposited in the tumor.40 At 2 h, the fluorescence signal of both micelles in liver and kidney was high since both organs were responsible for nanoparticle elimination.41 The extensive lung accumulation of un-crosslinked micelle was thought as a consequence of its poor stability. The in vitro PBS dilution experiment showed that the size of iron-free nanocarrier could dramatically increase up to micrometer magnitude, which inferred that its in vivo aggregation upon blood dilution



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was highly possible. Since large particles could cause microemboli when stuck in capillaries and preferentially stay in the lung,42 it is sensible to see the high accumulation of un-crosslinked micelles in lung. Interestingly, both types of micelles exhibited a considerable drug accumulation in kidney at 2 h, which might be a consequence of premature drug release. The pharmacokinetics data showed that both micelles displayed extended blood circulation compared to free Dox and the crosslinked micelles were more stable than the iron-free micelles (Figure S18). Usually ultra-small particles (< 5 nm) are subject to quick kidney filtration. However, this was not the case for both micelles in the current study. Hence, the relatively high kidney deposition of Dox from both micelles was assumed as a consequence of premature drug release. The blood circulation is a perfect sink condition and the drug release during such period is inevitable. The short hydrophobic moiety of micelles might make a contribution to such effect since the carrier-drug affinity was compromised. 2.7 In Vivo Efficacy Study. To evaluate the therapeutic efficacy of catechol-Fe3+ coordination crosslinked micelles, the in vivo efficacy study was carried out using 4T1 tumor-bearing mice model with twice intravenous (i.v.) dosing. The tumor size and body weight of mice were monitored continuously. The crosslinked Dox micelle demonstrated the best in vivo antitumor efficacy among all six tested samples (Figure 7). The negative controls (including saline, placebo micelles with and without crosslinking) cannot effectively inhibit tumor growth since there was no active agent in these formulations. Another control (free Dox) showed certain degree of tumor growth suppression, but its inhibitory performance was much worse compared to both Dox-loaded micelles, particularly at the later stage of experiment. This was mainly because the low


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accumulation of small molecule (Dox) in the tumor, as evidenced by the ex vivo biodistribution result (Figure 6). Although the in vitro cytotoxicity of both drug-loaded micelles was similar regardless of the presence of iron, the un-crosslinked Dox micelles indeed displayed significantly lower therapeutic efficacy compared to the coordinated ones. Such discrepancy was presumed to correlate to the enhanced drug exposure to the tumor by the crosslinked micelles. The nanomedicine is subjected to a multiple-step process before a pharmacological action can be elicited, involving blood circulation, tumor extravasation, tumor penetration, cellular internalization, endosomal escape, drug release and nuclear entry.43 The introduction of catechol-Fe3+ coordination enabled more stable micelles, which resulted in efficient Dox delivery to the tumor via minimizing drug loss during blood circulation, reducing non-specific biodistribution, and enhancing tumor accumulation, which was confirmed by the Dox fluorescence (or content) in tumor at 2 h, 6 h, and 24 h (Figure 6). Moreover, the crosslinking is reversible; within the acidic endosomes/lysosomes, the de-crosslinking could speed up drug release with no delay of therapeutic action. The micelle stability-regulated antitumor efficacy enhancement has been also observed previously.44 At the end of experiment, the volume of excised tumor from the mouse treated by Dox-micelles (Fe) ranked at the bottom. As an indicator of systemic toxicity, mice body weight was measured simultaneously. All samples except free Dox exhibited almost constant body weight over time. However, the mice treated with free drug exhibited severe weight loss post the second dose (i.e. from the 4th day), which denoted severe systemic toxicity. Such effect was also seen in other independent investigations.45,46



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Figure 7. In vivo antitumor efficacy assessment. (a) The 4T1 tumor growth curves after intravenous administration of 6 types of formulations; (b) The body weight variation of 4T1 tumor-bearing mice during the course of treatment; (c) Representative images of 4T1 tumors collected from the mice at the end of experiment, and the quantitative summary of tumor weights after treatment with 6 different formulations at day 16. All data were presented as the mean ± standard deviation (SD) (n = 6).* p < 0.05 (ANOVA test). Dox and (Fe) represent doxorubicin and iron-mediated micelle core crosslinking. 2.8 Histological and Apoptosis Analysis. To test the toxic effects of micellar formulations on normal organs, hematoxylin and eosin (H&E) staining was performed at the end of in vivo efficacy study. The histological analysis showed that both Dox-loaded micelles produced no evident organ damage or inflammatory lesion in all five types of tested organs (Figure 8). It is well known that Dox can induce cardiotoxicity as its major side effect,47 however, heart tissue treated by both Dox micelles showed that the compact cardiomyocytes arranged in a well-defined structure, indicating the absence of cardiotoxicity. Despite the high accumulation of Dox in lung via the micellar vehicle, no obvious lung injury was noticed. All these demonstrated that the Dox micellar formulations displayed no apparent toxicity to the normal tissues at the treatment dose. To further detect cell apoptosis, the terminal deoxynucleotidyl transferase dUTP nick end


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labeling (TUNEL) assay was employed for tumor tissue analysis. Irrespective of crosslinking, Dox micelles showed clear cell apoptosis as verified by the strong green fluorescence (Figure 9). The catechol-Fe3+ coordination enhanced the extent of apoptosis because of the relatively high drug content in tumors treated by crosslinked micelles, which agreed well with the ex vivo biodistribution and in vivo efficacy results (Figure 67). The free Dox control only induced minor DNA fragmentation due to the poor drug accumulation in tumors. All three drug-free samples did not induce any significant apoptotic event, which signified that Dox was the major contributor to induce cell death. It is interesting that the presence of Fe3+ in the crosslinked micelles did not bring additional toxicity in spite of its strong oxidation capability. First, the delivered iron dose might be not sufficient enough to induce toxicity. Second, within the acidic sub-cellular organelles (e.g. endosomes/lysosomes), the pH-dependent catechol-Fe3+ coordination could still retain some iron via 1:1 ratio for iron trapping.17 Thus, the released free Fe3+ would be reduced. Third, the iron metabolism process could handle the excess Fe3+ to maintain homeostasis of iron at both systemic and cellular level.48 Upon the cellular entry of Fe3+, the intracellular iron reductases could convert Fe3+ to Fe2+ that would be either expelled out of the cell via transporters (e.g. ferroportin),49 or circulated into the labile iron pool.50 In addition, the Fe3+ could also be stored in ferritin and take part in a series of cellular process such as protein biosynthesis.51



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Figure 8. Histological observation of the healthy tissues after Dox-loaded micellar nanocarrier treatment with saline as the control formulation. All tissues were stained with haematoxylin and eosin (H&E). Dox and (Fe) represent doxorubicin and iron-mediated micelle core crosslinking. Scale bar: 100 µm.

Figure 9. Apoptosis analysis of the tumor tissues after treatment by 6 different formulations. The tumor sections were stained with fluorescein-dUTP (green) for apoptosis and the nuclei were stained by DAPI (blue). Dox and (Fe) represent doxorubicin and iron-mediated micelle core crosslinking. Scale bar: 50 µm.


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2.9 MRI Measurement. The contrasting ability of crosslinked micelles was tested both in vitro and in vivo. Regarding the in vitro measurement, catechol-Fe3+coordinated micelles demonstrated T1-weighted (spin-lattice) relaxation signals with a calculated r1 relaxivity of 1.53 mM-1 s-1 at neutral condition (Figure S19), implying that these crosslinked micelles could be self-reporting nanocarriers (Figure S20). Since the tumor microenvironment is characteristic of a relatively lower pH (than 7.4) and the intracellular endosome/lysosome is also acidic, we also investigated the effect of pH on the MRI performance of coordination micelles in vitro (Figure S19). As expected, the imaging property of micelles was better at pH 5.0. Actually there have been continuous interests in developing Fe3+complexes as MRI contrast agents to replace toxic Gd3+ chelates.12 The in vivo MRI study employed the tumor-bearing mice; the micellar dose (112 µg Fe per mouse) via i.v. administration was set equivalent to that in the efficacy study. However, the images taken by 1.2 T MRI didn’t show noticeable contrast in the tumor tissue. To demonstrate the “proof-of-concept, we carried out a control study for which the crosslinked micelles was administered intratumorly at a dose of 5.6 µg Fe per mouse that corresponded to a 5% of the i.v. dose. The reason of such design was that the accumulated dose in tumors is often below 5% irrespective of delivery route.40 Encouragingly, the tumor regions exhibited enhanced contrast upon local delivery of coordination micelles (Figure 10); the signal intensity was plotted in Figure S21. These results well suggested that the Fe3+ content accumulated via EPR effect was not sufficient to generate a detectable MRI signal at the relatively low magnetic field intensity of 1.2 T. Theoretically, the r1 relaxivity of nanosystem could be enhanced by simply increasing either the Fe3+ concentration via utilizing multivalent carrier. From the viewpoint of



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particle engineering, enriching Fe3+ in micelles could be readily achieved by introducing more catechol groups in the conjugates due to the ratiometric coordination between them. In addition, the in situ release profile of Fe3+ in the tumor site would also affect the MRI performance of coordination micelles, which deserves further investigation. Due to the employment of intra-tumor administration approach, the micelles would mainly reside in the tumor site and thus no detectable signal in other healthy organs was observed.

Figure 10. In vivo images of 4T1 tumor-bearing mice after the administration of catechol-Fe3+ coordination crosslinked micelles (0 h signifies pre-injection); the analysis was performed using a T1-weighted MRI instrument (1.2 T). The circled area indicates the tumor location. The iron dose was 5.6 µg for one mouse with a body weighted of 18 g. CONCLUSION In summary, we have presented a simple bioinspired strategy to assemble enhanced stability, pH-triggered on-demand drug release, and MRI in a single micelle. The catechol-Fe3+ coordination enabled crosslinked micelles that could resist high volume dilution without losing their integrity during systemic circulation. The stability enhancement of micelles promoted the accumulation of model drug in the tumor tissue


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via EPR effect, which was in good correlation with high degree of apoptosis, and enhanced in vivo efficacy. The ratiometric coordination also endowed the micelles acidsensitive drug release, which would not induce delay of onset of pharmacological action. The presence of paramagnetic Fe3+ in the micelle core also made possible the T1weighted MRI whose sensitivity could be enhanced by further particle tailoring. The study broadened the applications of polymer micelles to multifunctional theranostic nanomedicine via a simple one-step coordination, which could be applied to a diverse range of hydrophobic drugs. 3. EXPERIMENTAL SECTION 3.1 Micelle Preparation. The preparation of micellar samples employed a typical dialysis method.25 In brief, 200 mg mPEG-S-Trityl-Cys-Dopa was dissolved in 5 mL DMF and the solution was dialyzed against excess MOPS buffer (pH 7.4) using a regenerated cellulose tube (MWCO: 1,000 Da). After 24 h, the un-crosslinked placebo micelles were collected ready for use. The production of iron crosslinked micelles utilized a similar procedure as described above; the only difference was that dialysis medium (MOPS buffer, pH 7.4) was supplemented with nitrilotriacetic acid (1.2 mM) and Fe3+ (0.4 mM) to enable catechol-iron coordinated crosslinking. After 24 h, the samples were further dialyzed against MOPS buffer (pH 7.4) to remove excess free Fe3+. Dox was selected as the model drug. Before drug loading, hydrophilic Dox·HCl and excess TEA were put together to generate the hydrophobic version of Dox. Then the hydrophobic Dox were loaded in both non-crosslinked and iron crosslinked micelles. Dox (20 mg) was placed together with 200 mg mPEG-S-Trityl-Cys-Dopa conjugate within the



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dialysis tube. The subsequent loading process is analogous to that of producing placebo micelles. 3.2 Micelle Characterization. To assess the iron content in the crosslinked micelles, both placebo and Dox-loaded lyophilized micelles (7 mg) was dissolved in 5 mL hydrochloric acid (37%, w/w); the samples were maintained at ambient temperature for 1 h followed by dilution with water. Then the iron content was analyzed by a 180-80 polarized Zeeman atomic absorption spectrophotometer (Hitachi High-Technologies Co. Ltd., Shanghai, China) (n = 3). The wavelength was fixed at 248.7 nm with a slit of 0.2 nm. The molar ratio of Dopa and Fe3+ was calculated to analyze the coordination relation between catechol and Fe3+ within both micelles. The hydrodynamic diameter and zeta potential of four kinds of nanocarriers were analyzed by a Zetasizer Nano ZS (Malvern Instrument Ltd., Malvern, UK) at 25°C (n = 3). The samples were diluted in PBS buffer (pH 7.4) prior to dynamic light scattering (DLS) measurement. For zeta potential analysis, the samples were dispersed in 1 mM KCl aqueous solution. The core size and morphology of micelles were analyzed by JEM-100CX II transmission electron microscope (JEOL Ltd., Tokyo, Japan). 3.3 Micelle Stability Assessment. The CMC of un-crosslinked mPEG-S-Trityl-CysDopa micelle was determined using pyrene as a fluorescence probe.52 The concentration of micellar aqueous solution varied from 400 ng/mL to 400 µg/mL and the probe concentration was set at 0.5 µM. The fluorescence spectra were recorded using a FLS980 fluorescence spectrometer (Edinburgh Instruments Ltd., Livingston, UK) with the excitation wavelength of 333 nm. The emission spectra were recorded from 350 nm to 450 nm and spectral slit width of 5 nm. The ratio of sample band intensity at 384 nm and


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373 nm was drawn against the logarithm of the sample concentration and the flexion point of the curve indicated the CMC value. In addition, a simple dilution method was also employed to evaluate the physical stability of micelles. Placebo micelles with or without iron crosslinking (1 mg/mL) dispersed in PBS buffer (pH 7.4) were diluted 1,000 times using the same buffer to reach a concentration far below the CMC. Alternatively, the samples were diluted with organic DMF (5 times). The temperature was controlled at 25oC. Then DLS was utilized to get the hydrodynamic size of different diluted micellar samples. 3.4 In Vitro Drug Release. The Dox loading in micelles were determined using HPLC. A fixed amount of micelles were dissolved in DMF and then drug content was quantified (n = 3). The separation employed a Dionex 3000 HPLC system coupled with a Phenomenex Gemini C18 column (250 mm × 4.6 mm, 5 µm) and an ultraviolet detector (detection wavelength at 500 nm) at 30oC. The injection volume is 20 µL and the mobile phase is the mixture of water and acetonitrile (1:1, v/v) with a constant flow rate at 1 mL/min. The in vitro drug release utilized a vertical Franz type diffusion cell.29 The donor and receiver compartments were separated by the regenerated cellulose membrane (1,000 Da). The release medium was the citric acid/disodium hydrogen phosphate buffer at different pH (5.0, 7.4) containing 5% SDS to maintain sink conditions. Dox-loaded micelles (ca. 20 mg) dispersed in aqueous solutions (with or without iron crosslinking) were supplemented to the donor section (n = 3). The temperature was maintained at 37oC and the drug content in the receiver compartment was determined by HPLC at predetermined time points. The release curve was plotted with the percentage of accumulative Dox across the membrane against time.



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3.5 Cell Line and Animal. Murine breast cancer cells (4T1) that were originally from American Type Culture Collection (ATCC) were cultured in RPMI 1640 medium (Gibco, NY, USA). The medium was supplemented with 10% fetal bovine serum (FBS) (BioInd, CT, USA), 100 U/mL penicillin and 0.1 mg/mL streptomycin (Gibco, NY, USA) at 5% CO2 and 37oC. BALB/c mice (female, 6 weeks, 18-22 g) were from Vital River Laboratory Animal Technology Co. Ltd. (Beijing, China). All animal experiments were carried out following the policies of the Animal Ethics Committee of Tianjin University. 3.6 In Vitro Cytotoxicity. 4T1 cells were in 96-well plates at a density of 4 × 103 cells/well in 100 µL and incubated overnight. Dox-loaded micelles (with or without crosslinking) were added to each well and the drug concentration located with the range of 0-4 µg/mL; two types of placebo micelles were used as negative control (n = 6). After 24 h, the cells were washed with PBS and then treated with the fresh medium containing 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT, 0.5 mg/mL) for additional 4 h. Afterwards, the MTT medium was replaced with 100 µL DMSO and the plates were gently shaken for 20 min at ambient temperature to dissolve the formazan crystals. Absorbance at 490 nm was measured using a SynergyTM 4 hybrid microplate reader (BioTek Instruments Inc., VT, USA). Cell viability was expressed as the percentage of viable cell relative to the untreated control cells and the half maximal inhibitory concentration (IC50) was calculated accordingly. 3.7 Cellular uptake. To examine the intracellular uptake of the Dox-loaded micelles, 4T1 cells were seeded in 35-mm plates at a density of 1×106 cells/well and maintained at 37°C for 24 h. The medium was then replaced with 1 ml of fresh medium containing one of the following samples: free Dox·HCl, Dox micelle, and Dox micelle (Fe), respectively.


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The drug content was identical at 10 µg/mL. At pre-designated time points (2 h and 4 h), the cells were washed with PBS in triplicate, followed by paraformaldehyde (4%, w/v) fixation for 20 min and DAPI (5 µg/mL) staining for 10 min. Subsequently, the cells were washed with PBS three times before fluorescence imaging using a Leica TCS SP8 confocal laser scanning microscope (Leica Microsystems Inc., IL, USA). 3.8

Biodistribution

and

pharmacokinetics.

BALB/c

mice

were

inoculated

subcutaneously with 4T1 cells in RPMI 1640 medium (50 µL of 2 × 107 cells/mL). Tumors were allowed to grow for one week to reach a size of ca. 50-100 mm3; then the mice were subdivided into three groups. Three types of samples (free Dox·HCl, Dox micelles, and Dox micelles (Fe)) were injected intravenously (i.v.) via the tail vein, respectively. The drug dose was kept at 5 mg/Kg. At 2 h, 6 h, and 24 h post administration, the animals were sacrificed; the major organs including heart, liver, spleen, lung, kidney and tumor were collected for ex vivo biodistribution study. The drug content of each sample was determined by a Cri Maestro in vivo imaging system (Cambridge Research & Instrumentation, Inc., MA, USA) with a green filter set (i.e. excitation filter at 523 nm and emission filter at 560 nm longpass). The mean fluorescence of each organs was obtained via background subtraction (n = 3). To assess the in vivo tissue distribution of DOX formulations quantitatively, the 4T1 tumor-bearing mice were intravenously injected with free DOX, DOX micelles, DOX micelles (Fe) at the dose of 5 mg DOX/kg (n = 3), respectively. The animal handing and drug quantification method was from previously published report with minor modification.53-55 In Brief, mice were sacrificed at 2 h and 6 h post administration, and the major tissues were excised, weighed, homogenized in 0.5 ml of 5% glucose solution. Then, 100 µL of



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sample were supplemented with 100 µL of 5% glucose solution, 100 µL of 5% ammonia aqueous solution and 3.5 mL of ethyl acetate, followed by ultrasonic extraction for 5 min. After centrifugation at 3,500 rpm for 5 min, the supernatant was collected and dried under a stream of nitrogen at 35oC. The dried sample was then dissolved in the mobile phase and quantified by an Agilent 6520 Q-TOF LC/MS system coupled with a ZORBAX Extend-C18 column (150 mm × 4.6 mm, 5 µm) and an electrospray ionization source. The fragment ion peak m/z 397 was picked for DOX determination, and the content was expressed as percentage of the injected dose per gram of tissue (% ID/g tissue). For pharmacokinetic measurements, free DOX, DOX micelles, DOX micelles (Fe) were administrated at the dose of 5 mg DOX/kg mice (n = 3), respectively. The blood (20 µL/sample) was collected at pre-determined time intervals and stored in heparin-coated centrifuge tubes. Then 50 µL of 5% glucose solution and 1.6 mL of methyl alcohol were added and vortexed for 5 min, followed by centrifugation at 13,300 rpm for 5 min. The DOX concentration in the supernatant was measured by HPLC-MS. 3.9 In Vivo Efficacy. Mice bearing 4T1 tumor were randomly divided into six groups (n = 6). Six types of samples were administrated intravenously including saline, free Dox·HCl, placebo micelle, Dox micelle, placebo micelle (Fe), and Dox micelle (Fe). The Dox dose was fixed at 5 mg/kg and the mass of placebo samples was maintained as the same as the corresponding drug-loaded micelles. The dosing was given every four days for a total of twice administration. The body weight and tumor volume of all mice were continuously monitored every other day. The major and minor axes of tumor (i.e. length and width) were measured by a caliper and the tumor volume was calculated by the well-


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known equation (length × width2/2). The relative ratio of tumor size was determined by the tumor volume at a fixed time divided by the initial tumor volume. 3.10 Histological and Apoptosis Analysis. At the end of in vivo efficacy experiment, all mice were sacrificed and the pictures of excised tumors were taken and compared. The major organ, i.e. heart, liver, spleen, lung, and kidney were also harvested. The histological analysis of non-tumor organs employed the typical H&E staining method. The toxicity effect on healthy organs was tested with regard to saline, Dox micelle and Dox micelle (Fe). The tumor apoptosis was examined by the TUNEL assay in accordance with the in situ cell death detection kit from Promega Biotech Co., Ltd. (Beijing, China). The cell nuclei were labeled with DAPI. 3.11 MRI measurement. The in vitro MRI measurement of crosslinked placebo micelle differing in concentration was carried out using a 1.2 T HT/MRSI60-60KY system from Huantong Corporation (Shanghai, China). The 1/ T1 relaxation time (s−1) was plotted against the Fe3+ concentration (mM) and the r1 relativity was obtained by the calculating the slope. Water and un-crosslinked micelles were taken as control. The in vivo MRI analysis employed the same 4T1 tumor-bearing BALB/c mice; 200 µL crosslinked Dox micelles (iron concentration at 11 mM) were administered intravenously. Alternatively, 20 µL crosslinked placebo micelles with the iron concentration at 5 mM were injected intratumorly. The MRI scanning was performed at different time points (0 h, 2 h and 6 h). The spin-lattice time T1-weighted imaging parameters were listed as follows: repetition time (TR) = 100 ms, echo time (TE) = 8.8 ms, field of view =100 × 50 mm2, matrix = 512 × 256, slice thickness = 1 mm, temperature = 20oC.



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3.12 Statistical Analysis. The data were expressed as the mean ± standard deviation (SD). A statistically significant difference was determined at a minimal level of significance of 0.05 via Student’s t-test or analysis of variance (ANOVA) followed by appropriate post-hoc analysis. ASSOCIATED CONTENT Supporting Information. The synthesis and characterization of polymer conjugates and relevant intermediate products and other supplementary figures. This material is available free of charge via the Internet at http://pubs.acs.org. AUTHOR INFORMATION Corresponding Author *Email: [email protected] Author Contributions The manuscript was written through contributions of all authors. All authors have given approval to the final version of the manuscript. † K. Xin and M. Li contributed equally. Notes The authors declare no competing financial interest. ACKNOWLEDGMENT We acknowledge the funding support from National Basic Research Program of China (2015CB856500), the open fund of the State Key Laboratory of Medicinal Chemical


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Biology (Nankai University) (201503007), National Natural Science Foundation of China (2134068; 81220108015), and Tianjin Research Program of Application Foundation and Advanced Technology (14JCZDJC38400). ABBREVIATIONS MRI, magnetic resonance imaging; Dox, doxorubicin; mPEG, methoxyl poly(ethylene glycol); Dopa, dopamine; MW, molecular weight; MOPS, 3-morpholinopropanesulfonic acid; TEM, transmission electron microscopy; CMC, critical micelle concentration; DMF, dimethyl formamide, HPLC, high performance liquid chromatography; SDS, sodium dodecyl sulfonate; NIR, near infrared; EPR, enhanced permeability and retention; H&E, hematoxylin and eosin; DLS, dynamic light scattering; MTT, 3-(4,5-dimethylthiazol-2yl)-2,5-diphenyltetrazolium bromide; TUNEL, terminal deoxynucleotidyl transferase dUTP nick end labeling; ANOVA, analysis of variance. REFERENCES (1) Farokhzad, O. C.; Langer, R. Impact of Nanotechnology on Drug Delivery. ACS Nano 2009, 3, 16-20. (2) Tibbitt, M. W.; Dahlman, J. E.; Langer, R. Emerging Frontiers in Drug Delivery. J. Am. Chem. Soc. 2016, 138, 704-717. (3) Mitragotri, S.; Lahann, J. Physical Approaches to Biomaterial Design. Nat. Mater. 2009, 8, 15-23.



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Graphical Abstract.


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Bioinspired Coordination Micelles Integrating High Stability, Triggered

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