the complexity demonstrated in paclitaxel-loaded glycopolymer

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Correlation between drug loading content and biological activity: the complexity demonstrated in paclitaxel-loaded glycopolymer micelle system Cheng Cao, Jiacheng Zhao, Mingxia Lu, Christopher J. Garvey, and Martina H. Stenzel Biomacromolecules, Just Accepted Manuscript • DOI: 10.1021/acs.biomac.8b01707 • Publication Date (Web): 15 Feb 2019 Downloaded from http://pubs.acs.org on February 21, 2019

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Correlation between drug loading content and biological activity: the complexity demonstrated in paclitaxelloaded glycopolymer micelle system Cheng Cao, a,b Jiacheng Zhao, a Mingxia Lu,a Christopher J. Garvey,*,b and Martina H. Stenzel*,a aCentre

for Advanced Macromolecular Design, School of Chemistry, The University of New South

Wales, Sydney, Australia. bAustralia

Nuclear Science and Technology Organisation, Lucas Heights, Australia.

ABSTRACT Drug delivery carriers are now widely established as they can increase the therapeutic efficiency of drugs. In general, the aim in this field is to create effective carriers that have large amounts of drugs loaded in order to minimize drug carrier material that needs to be disposed of. However, there was little attention so far in the literature on the effect of the amount of loaded drugs on the biological activity. In this paper, we are trying to answer the question how the drug loading content will affect the in vitro activ-ity. Here, we use two methods to load paclitaxel (PTX) into micelles based on the glycopolymer, poly(1-O-methacryloyl-β-D-fructopyranose) -block-poly(methyl methacylate) (Poly(1-O-MAFru)35-bPMMA145). In the one-step method, the drug is loaded into the particles during the self-assembly ACS Paragon Plus Environment

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process. However, the size of nanoparticle increased with the PTX content from 26 nm to 50 nm triggering enhanced cellular uptake by MCF-7 and MDA-MB-231, which was caused by changes in diameter size and not by changes in drug concentration. To keep the nanoparticle size constant, preformed micelles were loaded with PTX (two-step process). The increasing amount of loaded drug led to decreased cellular uptake and reduced cytotoxicity by the cancer cell lines. Small angle neutron scattering (SANS) and small angle X-ray scattering (SAXS), supported by Transmission Electron Microscopy (TEM) and Dynamic light scattering (DLS) exposed the PTX location in the shell. This caused shrinkage of the shell and lower levels of shell hydration, resulting in the lower cellular uptake and lower cytotoxicity. Upon release of PTX, the shell regained its original level of hydration. We could show that since drug loading causes morphology changes, either in the shell or the size, it is impossible to separate the parameters that will influence the biological activity. Although the same phenomenon may not apply to every drug delivery system, it needs to be considered that except for the well-known parameters that affect cell uptake – size, shape, surface chemistry, type of nanoparticle and presence of bioactive groups – the amount of loaded drugs might change the physico-chemical parameters of the nanoparticle, thus the in vitro and potentially the in vivo outcomes.

INTRODUCTION Over the years we have gained in-depth understanding on the correlation between physico-chemical properties of nanocarriers and their biological function.1-4 The size of the nanoparticle, the shape, the type of nanoparticle and the surface chemistry, which encompasses the charge and the presence of functional groups, are typically regarded as the deciding parameters that influence the fate of the internalized nanoparticles in the cell and in the body. Among the nanocarriers studied, polymeric micelles have attracted significant attention as drug delivery system due to their versatility, high drug loading efficiency and long-circulating time.5-6 Moreover, the core- and shell forming polymer block can be easily adjusted to maximize polymer-drug interaction, but also to create self-assembled ACS Paragon Plus Environment

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aggregates of different size and shapes. Drug encapsulation strategies range from chemical conjugation (covalent bonding and coordination bonding) to noncovalent bonding (electrostatic complex formation, hydrophobic interaction and hydrogen bonding).7-10 The most common pathway for drug loading is the physical entrapment of hydrophobic drugs into the hydrophobic core of the micelle, which is influenced by the crystallinity of the core,11 the drug-polymer compatibility,12-16 the volume of the core,17 the glass transition temperature18 and other interactions such as hydrogen bonding. Theoretical studies such as the one by Kumar and Prud’homme are an important contribution to this topic as they highlight that there is a thermodynamic upper limit for the amount of drug that can be loaded into a micelle when the drug is physically encapsulated. The model encompasses the entropic term, which is a function of the hydrophobic block, the enthalpic term, which is predominantly dictated by the Flory-Huggins interaction parameter , and the pressure-volume work, which is derived from the interfacial tension between core, shell and solvent.19 Comparison of the theoretical model with experimental results using PEG-b-PLA polymer and PTX20 revealed a very good agreement between both highlighting the importance of block lengths to achieve high drug loading capacities. However, the hydrophobicity may not always be the determining factor as H-bonding can overcome disparate polarities as demonstrated using poly(2-butyl-2-oxazoline).21 Normally, the increased drug solubilisation with increasing hydrophobic block has been experimentally confirmed in various studies with the maximum amount being strongly dependent on the interaction between polymer and drug.17, 22 It is noticeable that the focus in this type of work is to increase the drug loading capacity as the polymer, which can be potentially toxic, needs to be disposed of at the end of the treatment and should therefore be minimized. Large amounts of polymers can render a drug carrier clinically irrelevant as the amount of polymer required to deliver a therapeutic amount of drug exceeds what is practically feasible.23 As there are limits to how much drugs a micelle can entrap, chemical conjugation strategies are often more successful in increasing the drug loading content.24-25 If chemical conjugation is not possible, other forces can be employed to increase drug loading such as H-bonding26 or donor–acceptor coordination.27 In all approaches the aims seems to maximize drug loading content. However, there are some cases that 3 ACS Paragon Plus Environment


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show that nanoparticles with lower drug loading outperform micelle with high loading.28-30 Although unexpected, the outcomes can sometimes be explained by analyzing the drug carrier in detail. For example, the location of the drug in the micelle – core or shell – can affect the activity of the drug as the drug is more accessible and readily released.31 The presence of the drug at various locations in the micelle may also affect the zeta potential, thus the cellular uptake.32 To make matters more complicated, lessons learned from one drug and one drug carrier can often not be translated to another drug as different drugs can influence the physico-chemical behavior differently, which will subsequently affect the fate in vivo.33 These examples show that drug-polymer interaction should not be oversimplified. One of the assumptions that the hydrophobic drug is always located in the hydrophobic core of the micelle has shown to be wrong as strong hydrogen bonding can overcome polarity disparities.30, 34-35 Very often, the micelle morphology such as the aggregation number, core and shell radius influence the drug loading, but drug loading will also influence the micelle morphology. It therefore seems an almost impossible task to answer the question if there are any advantages in term of biological activity to have a higher payload (apart from the obvious point that excess polymer materials needs to be avoided for the drug carrier to have any clinical significance). In the following work, we demonstrate how the drug loading procedure and the amount of drug influences the morphology of the micelle, which subsequently affects the biological activity of the drug carrier. The main target was to learn how the drug loading content can affect the in vitro results of fructose based glycopolymers, which are known to have a high affinity to GLUT5 overexpressing cells such as breast cancer cells.36-37 In an attempt to decouple the effect of drug loading from the selfassembly process, we are evaluating two drug loading strategies of polymeric micelles: simultaneous drug encapsulation and self-assembly (one step) and loading of pre-existing micelles (two step). The chosen polymer, Poly(1-O-MAFru)36-b-PMMA145, which is known undergo self-assembly into micelles,37-38 was loaded with PTX. With the help of small angle X-ray scattering (SAXS) and small ACS Paragon Plus Environment

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angle neutron scattering (SANS), the drug location in different drug loading system can be identified. The drug loaded micelles where then studies in regards to their ability to interact with cells and cause cell death. Structural insight gained from the scattering studies was then correlated to the biological activity. Experimental Materials D-fructose (99%, Aldrich), dichloromethane (DCM; anhydrous, >99.8%, Aldrich), methacrylic anhydride (94%, Aldrich), 4-dimethylaminopyridine (DMAP; 94%, Aldrich), sulfuric acid (95%-98%, Ajax Finechem), acetone (HPLC grade, Ajax Finechem), tetrahydrofuran (THF; 99%, Ajax Finechem), fluorescein O-methacrylate (97%, Aldrich) and N, N-dimethylformamide (DMF; 99%, Ajax Finechem) were used as received. 1,4-dioxane (99%, Ajax Finechem) and pyridine (99%, Ajax Finechem) were purified by reduced-pressure distillation. Methyl methacrylate (>99%, Aldrich) was passed over basic aluminium oxide to remove the inhibitors. 2,2-azobis(isobutyronitrile) (AIBN; 98%, Fluka) was recrystallized from methanol for purification. Synthesis of Poly(1-O-MAFru)36-b-PMMA145 The synthesis of 1-O-methacryloyl-2,3:4,5-di-O-isopropylidene-β-D-fructopyranose followed a similar procedure as reported previously.38 The synthesis of macro RAFT agent was described as follows: in a Schlenk tube, 1-O-methacryloyl-2,3:4,5-di-O-isopropylidene-β-D-fructopyranose (1 g, 3.05 mmol), AIBN (1.33 mg, 8.1 × 10-3 mmol) and CPADB (17 mg, 0.061 mmol) were dissolved in 1,4-dioxane (4.3 mL). Then the solution was degassed by three freeze-pump-thaw cycles. The polymerization was carried out at 70 ℃ and stopped at 7 h by cooling the solution using ice-cold water. The polymer solution was then poured into a large excess of n-hexane for precipitation. The viscous polymer was dried under vacuum for 24 h before being used as macro RAFT agent for the following chain extension. The chain extension was carried out like this: macro RAFT agent (200 mg, 2 × 10-2 mmol), methyl ACS Paragon Plus Environment

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methacrylate (516 mg, 5.16 mmol), AIBN (0.64mg, 4× 10-3 mmol) and fluorescein O-methacrylate (14 mg, 0.037mmol) were dissolved in 1,4-dioxane (3 mL) at first. The solution was degassed by three freeze-pump-thaw cycles and reacted at 70 ℃ for 18 h. A large excess of n-hexane was used for precipitation. The viscous polymer was dried under vacuum for 24 h. The deprotection of the block copolymers was carried out under acidic conditions. The polymer (80 mg) was added into 1.59 mL of TFA/H2O (9:1 v/v) in a vial with stirring at room temperature for 30 minutes. After reaction, the polymer solution was dialyzed against deionized water for two days (MWCO 3500). The deprotected polymer was then lyophilized.

One step method self-assembly of Poly(1-O-MAFru)36-b-PMMA145 The preparation of spherical micelles was described as follows: The deprotected polymer (4 mg) of three batches with different amount of PTX (0 mg/mL, 0.25 mg/mL and 0.5 mg/mL) were separately dissolved in DMF (0.2 mL) firstly. Then 1.8 mL MQ water was added to the polymer solutions with different amount of PTX using a syringe pump with a rate of 0.2 mL/h respectively. The self-assembly process was carried out at 25 ºC with stirring rate of 1000 rpm. A glass cooling tube was used to control the temperature of solution precisely and tempering medium was pumped around the double walled vessel. The stirring rate was controlled by heat plate. The micelle solution was dialyzed against deionized water to remove DMF after self-assembly. Unloaded paclitaxel was removed by centrifugation at 3000 rpm for 5 minutes.

Two step self-assembly of Poly(1-O-MAFru)36-b-PMMA145 Size-controlled micelles were prepared in advance before drug loading, which was carried by the similar procedures to one-step self-assembly of Poly(1-O-MAFru)36-b-PMMA145 without PTX. Then the micelle solution (20 mL; 2mg/mL) was separated into five batches. Paclitaxel was dissolved in DMSO ACS Paragon Plus Environment

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with a concentration of 1 mg/ml. Then different volume (10 L, 20 L, 40 L, 80L and 160 L) of paclitaxel solution was added dropwise into micelle solution over a period of 3 hours. The drug loaded micelles were dialysis against MQ water to remove DMSO. Free paclitaxel was removed by centrifugation at 3000 rpm for 5 minutes.

IC50 Determination via Sulforhodamine B Assay Human breast cancer MCF-7 and MDA-MB-231cells are from Cell bank Australia. They were seeded in 96-well plates (4000 cells per well for MCF-7 Cells and 8000 cells per well for MDA-MB-231 cells) with culture medium Dulbecco’s modified Eagle’s medium (DMEM) supplemented with 2.2 g/L NaHCO3, 10% (v/v) foetal bovine serum (FBS), 100 U/mL penicillin and 100 μg/mL streptomycin in the incubator (5% CO2/ 95% air atmosphere at 37 °C) for 24 h. The sample solution to be tested was sterilized via UV irradiation (20 min) before the solution was serially halved via dilution in sterile MilliQ water. The micellar solutions were then loaded into the plate at 100 μL per well. After incubation for 72 h, cells were treated with trichloroacetic acid 10% w/v (TCA) and incubated at 4 °C for 40 min, and then washed five times with Milli-Q water to get rid of the TCA solution. TCA- fixed cells were stained for 30 min with 0.4% (w/v) sulforhodamine B (SRB) dissolved in 1% acetic acid. SRB outside the cells was removed by washing the plates with 1% acetic acid. The plates were left to air-dry overnight followed by the addition of 100 μL of 10 mM Tris buffer per well to dissolve the dye in the cells. The absorbance at 570 nm of each well was measured using a microtiter plate reader scanning spectrophotometer to calculate cell viability [cell viability (%) = (test - blank) / (control - blank) × 100]. The data are mean  SD (n = 4), and the IC50 values are mean  SD (n = 3).

Cellular uptake

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Cellular uptake was carried out using flow cytometry. After culturing sufficient amount of cells using medium Dulbecco’s modified Eagle’s medium (DMEM) including fetal bovine serum (FBS, Bovogen Biologicals) and plasmocin at 37 °C in a humidified atmosphere including 5% CO2, MCF-7 human breast cancer cells were seeded in the 6-well plates at a density of 3 × 105 cells per well with DMEM cell culture medium at 37 °C with 5% CO2 for 2 days prior to nanoparticle treatment. For the treatment, micelles (0.6 mg in 3 mL DMEM) were incubated with cells for 8 hours. The cells were initially washed 4 times with cold PBS to remove excess micelles before they were detached from the plates using tryspin/EDTA. The cell suspensions were then centrifuged and resuspended in cold Hank’s buffer. The fluorescence intensity of the cells from the suspensions (20000 cell events) was measured indicative of cellular uptake on a BD FACSCanto TM II Analyser using an excitation laser wavelength of 480 nm and a band-pass filter of 530/30 nm for emission spectra. Raw data was analysed using FlowJo® software and the results are shown using median fluorescence intensity (MFI) averaged from 3 wells in different plates for each sample. Drug release study The micelles were prepared by the similar procedures as described under the two-step process. Paclitaxel (1 mg/mL, 240 L) in DMSO was used for drug loading. The release study was carried out at 37 oC in the pH 7.4 phosphate buffer (20mM). The micelles (2mg/mL) carrying of PTX was transferred into 10mL two different dialysis bags (MWCO 3500 kDa) in 1L MQ water at 37 oC. The polymer solution inside the bag was sampled every 2 h. The drug concentration in the nanoparticle was determined using the HPLC. After all the PTX was released, the empty micelles were collected and analyzed using TEM (JEOL1400). Analysis Size Exclusion Chromatography (SEC)


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The molecular weight and polydispersity of synthesized polymers were analyzed via size exclusion chromatography (SEC). A Shimadzu modular system comprising a SIL-10AD auto-injector, DGU-12A degasser, LC-10AT pump, CTO-10A column oven and a RID-10A refractive index detector was used. A 5.0-m bead-size guard column (50 × 7.8 mm) followed by four 300 × 7.8 mm linear columns (500, 103, 104, and 105 Å pore size, 5 μm particle size) were employed for analysis. N, N-Dimethylacetamide [DMAc; HPLC grade, 0.05% w/v 2,6-di-butyl-4-methylphenol (BHT) and 0.03% w/v LiBr] with a flow rate of 1 mL/min at 50 C was used as mobile phase. 50 μL of polymer solution with a concentration of 2 mg/mL in DMAc was used for every injection. The calibration was performed using commercially available narrow-polydispersity PMMA standards (0.5-1000 kDa, Polymer Laboratories). Dynamic Light Scattering (DLS) The hydrodynamic diameter Dh was determined using a Malvern Zetaplus particle size analyzer (laser, angle = 173°) at a copolymer concentration of 1 mg/mL. Samples were prepared in deionized water and sonicated for 30 min prior to the measurements. The polydispersity index (PDI) as a fitting parameter in the cumulants analysis of autocorrelation function shows the DLS data’s quality. Zeta potential is measured to indicate the surface charge of the micelles. High Performance Liquid Chromatography (HPLC) The percent of paclitaxel incorporated during nanoparticle preparation was determined via highperformance liquid chromatography (HPLC). HPLC system is a Shimadzu modular system consisting of a SPD-20A UV-Vis detector, DGU-20A3 degasser, LC-20AD pump (1 mL/min) and a Cole-Palmer single compact HPLC column heater (30 °C). The column was a Phenomenex Viva C18 (250 × 4.6 mm) with a pore size of 5 μm. The injection volume of each sample was 25 μL. The 50:50 – 90:10 mixtures of acetonitrile (HPLC grade) and Milli-Q water was used as the mobile phase. A range of paclitaxel solutions with different, and known concentrations was prepared in the same solvent and used to prepare standard curve. The elution time of paclitaxel was 12.1 min. PTX-loaded micelles with a volume of 25 ACS Paragon Plus Environment

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μL were injected, followed by the HPLC analysis using the standard curve described above (Figure S1). The encapsulated PTX was measured with spectrophotometer. The maximum wavelength of PTX is 220 nm, the drug loading efficiency (DLE) (%) were calculated by:

DLE(wt%) =

the amount of PTX in micelles the amount of PTX in feed

× 100%

(Equation 1)

The drug loading capacity (DLC) (%) were calculated by:

DLC(wt%) =

the amount of PTX in micelles the amount of polymer

× 100%

(Equation 2)

UV-Vis absorbance of the polymer samples from HPLC regarding the different concentration of PTX is used to describe the change of absorbance during the drug release.

Transmission Electron Microscopy (TEM) The TEM micrographs were obtained using a JEOL1400 transmission electron microscope comprising of a dispersive X-ray analyzer and a Gatan CCD facilitating the acquisition of digital images. The measurement was conducted at an accelerating voltage of 80 kV. The samples were prepared by casting the micellar solution (1 mg/mL) onto a copper grid. The grids were dried by air and then negatively stained with uranyl acetate. Fluorescence Spectroscopy Fluorescence measurements were carried on a CARY Eclipse fluorescence spectrophotometer at room temperature. The excitation of measurements was 480 nm for the experiments. The fluorescent wavelength was collect at 480 – 610 nm and excitation slit is 2.5 and 5. ACS Paragon Plus Environment

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Small-angle X-ray scattering and small-angle neutron scattering SAXS measurements were made at the Australian Synchrotron’s SAXS/WAXS beamline from samples in 1.5 mm quartz capillaries (Charles Supper, Natick, USA). Isotropic scattering patterns were collected on a Pilatus 1 M detector (Dectris, Baden-Daettwil, Switzerland) using a wavelength of 1.033 Å (20 KeV) and a sample to detector distance of 7.2 m. The measurement geometry was used to convert the counts per detector pixel into the radially averaged intensity versus the scattering vector on an absolute intensity scale after subtraction of background due to solvent filled capillary and normalizing to the scattering of water using the beamline software Scatterbrain. The scattering vector, q, is defined by q=

4π θ λ sin (2)

(Equation 3)

and /2 is the scattering angle. Small angle neutron scattering (SANS) measurements were made on the fixed wavelength reactor based pin-hole instrument Quokka (ANSTO, Lucas Heights, Australia). For each sample measurements were made at three different samples to detector distances: 1.3 m, 10 m and 20 m, where the counting times were 30 min, 1 h and 4 h respectively. The first two configurations were measured using 5 Å neutrons and the final at 8 Å. In both cases the  / was 10%. The isotropic counts per pixel on the area detector was converted a continuous scattering curve using the by radial averaged 2-dimensional data after first correcting for the detector sensitivity, subtracting the dark field counts and the background signal due to the solvent filled empty cell and finally normalising to the incident flux of neutrons on the sample. Small angle scattering curves were obtained at several different concentrations of the polymers (SAXS: 0.5 mg/mL, 1 mg/mL, 2 mg/mL; SANS: 0.5 mg/mL, 1 mg/mL, 2 mg/mL). When normalized to concentration of particles, both SANS and SAXS curves were independent of concentration. We can therefore assume that all scattering curves were obtained under dilute conditions, and that the scattering curve is dominated by the shape and internal structure of the particles. For SANS measurements, the


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contrast between the shell and solvent was enhanced by using D2O as the solvent. The structure model for the NPs described was fitted separately to the reduced SAXS and SANS data by using the program SASview39 and specific macros written for IgorPro.40

Figure 1. Synthesis of Poly(1-O-MAFru)36-b-PMMA145 and the subsequent simultaneous self-assembly of the polymer (2 mg/mL) in presence of different amounts of PTX: 0 mg/mL, 0.25 mg/mL and 0.50 mg/mL by TEM. Scale bar = 100 nm. Results and discussion One step: Simultaneous self-assembly and drug encapsulation and the impact on biological activity Self-assembly of block copolymer to form polymeric micelles is widely used to improve the bioavailability and solubility of hydrophobic anti-cancer drugs such as paclitaxel.41-42 Most commonly these micelle are loaded with hydrophobic drugs using the solvent switch technique, where the amphiphilic polymer and drug are dissolved in a common solvent, followed by water addition, which triggers self-assembly and simultaneous drug encapsulation. The block copolymer, Poly(1-O-MAFru)36b-PMMA145 (Figure 1 and S2), and PTX were dissolved in DMF, followed by adding of MQ water (Figure 1). Dialysis and centrifugation were used to remove excess DMF and free PTX. The drug ACS Paragon Plus Environment

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loading amount of the micelles was determined by HPLC. Drug loading efficiency (DLE) and drug loading capacity (DLC) are listed in Table 1. TEM analysis shows that there are no morphological changes and independent of the amount of PTX spherical nanoparticles were obtained. However, DLS revealed that the diameter increased from 25.9 nm to 49.1 nm with increasing amount of PTX in the solution. This is not unexpected as it has been hypothesized that hydrophobic drugs are located in the hydrophobic core resulting in increased the particle radius by changing the self-assembly, thus in increased micelle diameters (Table 1). This has indeed been frequently observed in the literature although there are cases where the presence of drugs can lead to a shrinkage of micelles.43 In the following, we will coin the nanoparticles obtained micelles as we have a clear core-shell structure. It should be pointed out here that these are unlikely micelles that are in equilibrium as the high Tg of PMMA will prevent any dynamic exchange.

Figure 2. Flow cytometry on MDA-MB-231 and MCF-7 cell lines. The polymer concentrations were set to 0.2 mg/mL, and the concentration of loaded drug ranges from 0 to 0.015 mg mL-1 (see Table S1). Data are presented as mean  SD (n=3); P 0.01, P 0.001, P 0.0001 (One-Way Anova followed by Turkey’s multiple comparison test). Changes in micellar sizes prevent the direct study of the effect of drug loading on the biological activity as alteration in micelle size will directly influence the cellular uptake.44-46 In general, nanoparticles of approximately 50 nm in diameter were reported to display maximum cellular uptake with lower and larger sizes leading to decreased endocytosis.44, 46-47 This was indeed observed with our current system ACS Paragon Plus Environment

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where the micelles, loaded with the highest amount of PTX and therefore biggest size, led to the highest cell internalization. In this experiment, the three micelles (No drug, low drug and high drug samples) (Table 1), which were labelled using fluorescein O-methacrylate (fluorescein O-methacrylate< 1wt%), were incubated with MDA-MB-231 and MCF-7 cell lines, followed by analysis using flow cytometry. Prior to the analysis, the fluorescent intensity was measured to account for any intensity fluctuation (Table S1 and 2). The enhanced cellular uptake of the micelles with the highest amount of PTX was observed in both cell lines, MDA-MB-231 and MCF-7 cell lines (Figure 2). The enhanced uptake then coincided with a higher toxicity as more drug was delivered into the cells (Table S3).

Figure 3. Size controlled self-assembly of Poly(1-O-MAFru)36-b-PMMA145 and morphologies by self-assembly of Poly(1-O-MAFru)36-b-PMMA145 (2 mg/mL) in existence of different amount of PTX: 10 L, 20 L, 40L, 80L and 160 L, [PTX] = 1 mg/mL by TEM. Scale bar =100 nm.

This example highlights the struggle when attempting to study the effect of drug loading on the biological performance. Contribution factors to the higher toxicity of the high drug loading micelle can range from the enhanced higher cellular thanks to the larger size, to the higher micelle stability in the presence of drugs or maybe the higher localized drug concentration as less micelles are required to package the same amount of drugs. Therefore, discussions around the effect of the amount of drug in micelles on biological activity are challenged by all these competing influence that will not allow for a ACS Paragon Plus Environment

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clear conclusion to be drawn. The simultaneous self-assembly and drug loading is therefore not suitable to study the drug loading effect in this system. Two step: Self-assembly followed by PTX loading and the impact on biological activities To avoid that the drug influences the self-assembly process, empty micelles were prepared first, followed by PTX diffusion into the micelles (Figure 3). Therefore, the block copolymer, Poly(1-OMAFru)36-b-PMMA145, was dissolved in DMF, followed by the addition of MQ. After DMF removal by dialysis, the micelle solution was mixed with different amounts of PTX dissolved in the DMSO for drug loading. Free drug and DMSO were subsequently removed by dialysis and centrifugation and the spherical morphology of the resulting micelles was confirmed using TEM (Figure 3). The drug loading efficiency and the resulting drug loading capacity were calculated using HPLC data (Table 2 and Figure S1). It is noticeable that this technique leads to very low drug loading efficiency and may not be suitable from a practical point, but important for this study is that the diameters are now less affected by the amount of drug and all micelle have sizes of around 20 nm. In contrast to the one step process where the diameter of micelles increased with increasing amount of drugs (Table 1), the post-loading technique led to a slight decline of the micelle size (Table 2). The opposite influence on the micelle size depending on the type of loading is not unexpected considering that PTX can potentially interrelate with both blocks in the one-step process interfering with the self-assembly process while in the two step-process the PTX will be initially only in contact with the glycopolymer and the drug has to pass through the shell to reach the hydrophobic core, if it will reach the core at all.


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Figure 4. Flow cytometry on MDA-MB-231 cell line (a) and MCF-7 cell line (b). The polymer concentrations were set to 0.1 mg/mL, and the concentration of loaded drug ranges from 0 to 2.65 g/mL (see Table S4). Data are presented as mean  SD (n=3); P 0.05, P 0.01, P 0.001, P 0.0001 (One-Way Anova followed by Turkey’s multiple comparison test). (c), (d) IC50 curves of PTXloaded micelles in 2D MDA-MB-231 and MCF-7 cell culture models. The cytotoxicity data is mean  SD (n = 4). The IC50 value is mean  SD (n = 3).

Subsequently, the PTX-free micelles M0 and PTX-loaded micelles M1-M5 can be used to the direct comparison in regards to the drug loading amount on biological activity without having to take size factors into account. As all micelles have similar diameters, the size dependency on the cellular uptake is not the determining factor anymore. Moreover, the zeta potential is independent of the amount of PTX as all micelles have now similar surface charge, which is in contrast to the micelle loaded in the one-step procedure where drug loading led a more negative surface charge. This is caused by the different structure of the micelles (one-step drug particles: big core, small shell vs two- step drug particles: small core, big shell) (Table S8). The subsequent cell uptake measurements were carried out ACS Paragon Plus Environment

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using similar conditions and preparations (Table S4 and 5). It was expected that the cellular uptake is now comparable enabling an answer to the question if it is better to concentrate the drug in a few micelles or to deliver the drugs in micelles of low loading capacity. Unexpectedly, micelles with high concentration of drug resulted in lower uptake by MDA-MB-231 and MCF-7 cell lines (Figure 4a and b). Moreover, high PTX loading led to an overall lower cytotoxicity per PTX (Figure 4 c and d and Table S6), which can be directly correlated to a lower cellular uptake. As the size and the surface charge (Table 2) do not change dramatically, the strong dependency on the PTX amount on uptake is rather unexpected. The unusual correlations must be related to the internal structure change caused by the drug location creating cues for the cell to reject the nanoparticle. Thus, small angle scattering was used to reveal the structure information.

Figure 5. (a) Cartoon of the model core shell particle shows the radial variation of the SLD. (b) SANS data for micelles with different amount of PTX, M0 – M5 and their respective curves fits. Each curve has been offset for clarity of presentation.

Physico-chemical characterization for PTX loaded micelles Techniques such as TEM and dynamic light scattering (DLS) are commonly used to investigate nanoparticle shape and size, by direct visualisation in the case of TEM and indirectly based on ACS Paragon Plus Environment

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assumptions of the hydrodynamic behaviour of the particles in the case of DLS. Small angle scattering (SAS), small angle neutron scattering (SANS) and small angle X-ray scattering (SAXS), provide this information on a statistically relevant ensembles of particles as well as probe the internal structure of nanoparticles.48-50 In our previous work, SAS not only revealed the internal structural information (core and shell) of self-assembled micelles, but also details of a drug’s location within this structure. Here we utilize SAS to provide insight on the location of a drug, PTX, in a nanoparticle structure. The analysis of SAS data has been carried out by the model-dependent fitting39 where the absolutely scaled scattering curves were fitted with a mathematical model including spatial and compositional information. A coreshell model was used to model SAXS and SANS data.51 This model is shown in the Figure 5a. The scattering length density (SLD) and the thickness of shells as well as the radius of the micelles were first constrained with physically reasonable parameters. An initial indication of the particle’s size and spherical shape was found from TEM analysis (Figure 1 and 3 and Table 1 and 2). DLS measurements were used to constrain an upper value of the radius (hydrodynamic radius) and the particles dispersity. Compositional information of the core and the shell of the particles were determined by modelling the scattering length density (SLD). The SLD is a description of the atoms in a representative volume:

SLD =

∑ibc

i

Vm

(Equation 4)

where bci is the bound coherent scattering density of all i atoms over the molecular or monomer volume, Vm. For SAXS and SANS, the coherent scattering cross-sections are from the electronic and nuclear environments respectively.52 The estimated SLDs of PTX, D2O and Fructose were used to constrain the SLD of core and shell (Table S7).

The core of the nanoparticle is composed of amorphous PMMA with the density and chemical composition used as a starting point for the modelling of core SLD for both SAXS and SANS curves.53 ACS Paragon Plus Environment

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The SANS curves of one step method self-assembled micelles (Table 1) and size-controlled selfassembled micelles (Table 2) with different amounts of PTX were fitted with a core-shell micelle model (Figure 5a). The fitting results are interpreted using the following quantities: aggregation number, Nagg (Equation S2); SLD core, SLD shell; and volume fraction of fructose, ∅fructose (Equation S3).

The size (core radius plus the thickness of shell) of the drug loaded micelles, obtained by self-assembly in the one step method, increased by 16 nm with increasing amount of PTX (Figure S3 and Table S8) and is therefore in agreement with DLS data (Table 1). The scattering length density (SLD) declined for core and shell suggesting that PTX interfered with the self-assembly process resulting in larger micelles. More interesting is now to understand why the two-step procedure led to such unexpected biological results despite all micelle having similar sizes. Figure 5b compiles all SANS fitting curves of samples M0 - M5 for the size controlled self-assembly samples (two-step method). All the fitting parameters are shown in the Table S9. The radius of the PMMA core, Rcore, is almost independent from the amount of drug loading, though a slight increase from 6.0 (M0)- 6.5 nm (M5) can be noticed with increasing amounts of PTX. In contrast, the thickness of the fructose shell, Tshell, decreases with increasing amount of PTX from 3.7 (M0) to 2.9 (M5) nm (Figure 6).


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Figure 6. The variation of core radius and the thickness of shell with drug loading. The lines are only to guide the eye.

It is noticeable that the post-loading effect has some effect on the size of the shell. However, this does not explain the reduced cellular uptake. The shrinkage of the shell can only be explained by the dehydration of the shell triggered by the presence of the hydrophobic PTX, which seems to be unfavorable to the biological activity. It seems to be crucial to look deeper into the structure of the drug loaded micelles, by extracting more information from the SANS data.

Figure 7. Volume fraction of water (D2O) in the fructose shell with respect to the different amount of PTX according to the equation S3. ACS Paragon Plus Environment

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The SLD of the core remains more or less constant increasing only marginally from 3.50  10-6 Å-2 to 3.61  10-6 Å-2 (M1-M5) with a value close to pure PMMA, and the modelling permits the narrow range of SLD values for the core close to the PMMA. This suggests that most PTX may indeed be loaded into the shell. With increasing amount of PTX, the SLD of the shell decreased from 4.16  10-6 Å-2 (M0) to 3.78  10-6 Å-2 (M5) (Table S9), underlining the reduced hydration. The chemical composition of the shells could subsequently be calculated using 𝑆LD shell = ∅fructose ∙ SLDfructose + ∅D2O ∙ SLDD2O + ∅PTX ∙ SLDPTX (equation S3). According to the SLD change of shell in the presence of PTX, the volume fraction of water, ∅𝐷2𝑂 , in the shell can now be calculated (Figure 7 and Table S10). So far, the fraction of drug in the shell is unknown, but can be elucidate using the aggregation number Nagg (Table S10). The sample M0 without drug loading can be used to calculated the aggregation number, Nagg = 46, according to the measured 𝑅𝑐𝑜𝑟𝑒 and 𝑇𝑠ℎ𝑒𝑙𝑙 from the data fitting (Table S9 and Equation S2). As samples M1- M5 were prepared by loading different amounts of PTX into sample M0, Nagg should be similar across all samples. Employing the drug loading capacities calculated from the HPLC data (Table 2), the amount of loaded drug per particle can be calculated (Table S10 and Figure S4). Finally, the water content of shell can be obtained, which leads to the conclusion that the water content of shell decreases with the increasing amount of PTX. It seems that the main change in the micelle structure is the declining shell thickness caused by the presence of increasing amounts of PTX, which coincides with a decreasing level of hydration. This result can now be connected to the biological testing observed in Figure 4 suggesting that nature of the shell and its hydration is a defining feature that influences cellular uptake and thus cytotoxicity.


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Figure 8. (a) SAXS data for PTX release from micelles in 0 h, 5 h and 72h, and their respective curves fits: (b) Drug released samples for flow cytometry on MDA-MB-231 and MCF-7 cell lines. Data are presented as mean  SD (n=3); P 0.05, P 0.01, P 0.001, P 0.0001 (One-Way Anova followed by Turkey’s multiple comparison test).

Characterization of PTX released micelles If the PTX content in the shell is in fact detrimental to the cell uptake, the release of PTX should be able to reverse the process. Therefore, the release of PTX from the micelle (new sample by two step method), loaded with 1 mg/mL PTX (240 L), was monitored over 72 h in PBS solution (pH = 7.4). The drug loaded micelles was dialyzed against water and samples were taken at regular time intervals to determine the amount of PTX released. After 72 h, most of the drug (88 %) was released from the micelles (Figure S5). The morphology of the micelles was tested at three times points (0 h, 5 h and 72 h) using TEM and SAXS (Figure 8a). Although the apparent morphology as determined by TEM does not change after the drug release (Figure S6), SAXS analysis shows an increase of shell thickness from 4.8 nm (0h) to 5.6 nm (72h). At the same time, the core radius slightly decreased from 8.6 nm to 8.1 nm after release (Table S11). In theory, with release of PTX, the glycopolymers on the surface should regain their original level of hydration, which should directly influence the cellular uptake. Hence, the cell uptake experiment using MCF-7 and MDA-MB-231 cell lines was repeated with these three ACS Paragon Plus Environment

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micelles, which contained high (0h), medium (5h) and low amounts (72h) of PTX (Figure 8b). In both cell line, the originally low cellular uptake – caused by PTX loading and subsequent shell collapse – was reversed upon drug release confirming that it is indeed PTX, and the changes it caused, that influence the cellular uptake. In this work, we only investigate one polymer concentration, as PTX is very toxic, and we therefore cannot use higher concentrations. The system created here is theoretically an unsuitable drug delivery system as the drug loaded micelle has a lower cellular uptake than the empty micelle. It does however highlight that the influence of the drug on micelles need to be seriously considered. The remaining question might be why the hydration may play such a big role here. The cellular uptake of fructose containing polymers is facilitated by the GLUT5 receptors overexpressed on the surface. It is therefore important that fructose is freely available for interaction. In earlier solid state NMR studies it was shown that hydration can be the determining factor to enable high glycopolymer chain mobility. Dehydration led to low availability of the fructose ligand, thus reduced cellular uptake. Hydration may however by a parameter specific to glycopolymers as their mobility can influence the interaction with cell surface receptors.29, 54

Conclusion In summary, we examined the effect of the PTX loading process into Poly(1-O-MAFru)36-b-PMMA145 micelles using either the traditional one step solvent switch technique or a two step-process where the micelles were prepared first and then incubated with PTX for loading. Aim was to understand how the amount of drug in a nanoparticle will affect the biological activity as low loading means that more micelles are required to achieve the same amount of PTX concentrations. Unfortunately, this question cannot be answered easily with these types of micelles as changes to the drug loading was found to inflict changes to the morphology. The one-step process led to increasing micelle sizes with increasing amount of drugs, while the two-step process led to shell shrinkage and dehydration according to scattering experiments. Both morphology changes led to changes in activity. The one-step process led to 23 ACS Paragon Plus Environment


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increased uptake with higher drug loading. This was caused by an increase in micelle size with drug loading towards to theoretically optimum 50 nm required for efficient endocytosis. In contrast, the two step process led to a decline in cell uptake and cytotoxicity although the sizes were now independent from the amount of loaded drug. Using SANS and SAXS, we found the drug was now located in the shell, which affect the hydration of shell. We could show that different self- assembly and drug loading methods affect the biological activities, resulting from size change and drug location. The lessons learned here is that the drug loading process and its location can influence the biological activity. While in most papers only the size, type of nanoparticle, shape and surface chemistry/functionalization are discussed, we should here that the presence of drugs is a crucial and mostly underestimated parameter.

ASSOCIATED CONTENT Electronic Supplementary Information (ESI) available: [PTX release, TEM after PTX release, HPLC, calculations of aggregation number and volume fraction of water in the shell, SAXS and SANS fitting parameter].

AUTHOR INFORMATION Corresponding Author * E-mail: [email protected] [email protected];

ACKNOWLEDGMENT MHS and CJG like to thank the Australian Research Council ARC for funding DP160101172. We also like to thank the Electron Microscopy Unit of the UNSW Mark Wainwright Analytical Centre for their help. We acknowledge the support of the Australian Centre for Neutron Scattering, Australian Nuclear ACS Paragon Plus Environment

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Science and Technology Organization, in providing the neutron research used in this work. This research was also undertaken on the (SAXS/WAXS) beamline at the Australian Synchrotron, part of ANSTO. REFERENCES 1. Sykes, E. A.; Dai, Q.; Sarsons, C. D.; Chen, J.; Rocheleau, J. V.; Hwang, D. M.; Zheng, G.; Cramb, D. T.; Rinker, K. D.; Chan, W. C., Tailoring nanoparticle designs to target cancer based on tumor pathophysiology. Proc. Natl Acad. Sci. USA 2016, 201521265. 2. Walkey, C. D.; Olsen, J. B.; Song, F.; Liu, R.; Guo, H.; Olsen, D. W. H.; Cohen, Y.; Emili, A.; Chan, W. C. W., Protein Corona Fingerprinting Predicts the Cellular Interaction of Gold and Silver Nanoparticles. ACS Nano 2014, 8, 2439-2455. 3. Sykes, E. A.; Chen, J.; Zheng, G.; Chan, W. C. W., Investigating the Impact of Nanoparticle Size on Active and Passive Tumor Targeting Efficiency. ACS Nano 2014, 8, 5696-5706. 4. Müllner, M.; Dodds, S. J.; Nguyen, T.-H.; Senyschyn, D.; Porter, C. J. H.; Boyd, B. J.; Caruso, F., Size and Rigidity of Cylindrical Polymer Brushes Dictate Long Circulating Properties In Vivo. ACS Nano 2015, 9, 1294-1304. 5. Cabral, H.; Miyata, K.; Osada, K.; Kataoka, K., Block Copolymer Micelles in Nanomedicine Applications. Chemical Reviews 2018, 118, 6844-6892. 6. Song, Z.; Han, Z.; Lv, S.; Chen, C.; Chen, L.; Yin, L.; Cheng, J., Synthetic polypeptides: from polymer design to supramolecular assembly and biomedical application. Chem. Soc. Rev. 2017, 46, 6570-6599. 7. Kim, J. O.; Kabanov, A. V.; Bronich, T. K., Polymer micelles with cross-linked polyanion core for delivery of a cationic drug doxorubicin. J. Control. Release 2009, 138, 197-204. 8. Bae, Y.; Fukushima, S.; Harada, A.; Kataoka, K., Design of environment‐sensitive supramolecular assemblies for intracellular drug delivery: Polymeric micelles that are responsive to intracellular pH change. Angew. Chem. Inter. Ed. 2003, 115, 4788-4791. 9. Yang, C.; Ebrahim Attia, A. B.; Tan, J. P. K.; Ke, X.; Gao, S.; Hedrick, J. L.; Yang, Y.-Y., The role of non-covalent interactions in anticancer drug loading and kinetic stability of polymeric micelles. Biomaterials 2012, 33, 2971-2979. 10. Xie, Z.; Guan, H.; Chen, X.; Lu, C.; Chen, L.; Hu, X.; Shi, Q.; Jing, X., A novel polymer– paclitaxel conjugate based on amphiphilic triblock copolymer. J. Control. Release 2007, 117, 210-216. 11. Gou, J.; Feng, S.; Xu, H.; Fang, G.; Chao, Y.; Zhang, Y.; Xu, H.; Tang, X., Decreased Core Crystallinity Facilitated Drug Loading in Polymeric Micelles without Affecting Their Biological Performances. Biomacromolecules 2015, 16, 2920-2929. 12. Kakde, D.; Taresco, V.; Bansal, K. K.; Magennis, E. P.; Howdle, S. M.; Mantovani, G.; Irvine, D. J.; Alexander, C., Amphiphilic block copolymers from a renewable ε-decalactone monomer: prediction and characterization of micellar core effects on drug encapsulation and release. J. Mater. Chem. B 2016, 4, 7119-7129. 13. Qian, F.; Huang, J.; Hussain, M. A., Drug–Polymer Solubility and Miscibility: Stability Consideration and Practical Challenges in Amorphous Solid Dispersion Development. J. Pharm. Sci. 2010, 99, 2941-2947. 14. Zhao, Y.; Fay, F.; Hak, S.; Manuel Perez-Aguilar, J.; Sanchez-Gaytan, B. L.; Goode, B.; Duivenvoorden, R.; de Lange Davies, C.; Bjørkøy, A.; Weinstein, H.; Fayad, Z. A.; Pérez-Medina, C.; Mulder, W. J. M., Augmenting drug–carrier compatibility improves tumour nanotherapy efficacy. Nat. Commun. 2016, 7, 11221. 15. Liu, J.; Xiao, Y.; Allen, C., Polymer–drug compatibility: A guide to the development of delivery systems for the anticancer agent, ellipticine. J. Pharm. Sci. 2004, 93, 132-143. ACS Paragon Plus Environment

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Table 1. Size of self-assembled drug loaded Poly(1-O-MAFru)36-b-PMMA145 micelles at different PTX feed amounts.

Micelles

[PTX]/ mg/mL

Drug loading efficiency/ %

Drug loading capacity/ %

Dh / nm (DLS)

PDI

Zeta Potential / mV

No drug

0

0

0

24.7  1.2

0.112  0.023

-9.9  2.7

Low drug

0.25

29.0

3.6

27.8  3.8

0.213  0.035

-22.3  3.4

High drug

0.50

29.2

7.3

45.6  2.6

0.178  0.047

-21.2  1.5

Table 2. Size of self-assembled drug loaded Poly(1-O-MAFru)36-b-PMMA145 micelles with the different amount of PTX after loading of a micelle solution prepared using a polymer concentration of 2 mg/mL.

Micelles

Actual Drug Drug PTX solution amount of loading loading added (L) loaded drug efficiency/ capacity/ % % (g/mL)

Dh / nm (DLS)

PDI

Zeta Potential / mV

M0

0

0

0

0

24.7  1.2

0.112  0.023

-9.9  2.7

M1

10

2.63

26.30

0.13

21.2  1.4

0.116  0.045

-11.3  2.4

M2

20

6.00

30.00

0.30

20.5  2.1

0.154  0.042

-12.2  5.3

M3

40

19.76

49.40

0.99

19.2  1.4

0.131  0.034

-13.1 3.2

M4

80

20.71

25.90

1.04

17.5  1.6

0.142  0.012

-14.0  5.5

M5

160

26.56

16.60

1.33

19.4  2.1

0.127  0.024

-11.1  4.6


28

the complexity demonstrated in paclitaxel-loaded glycopolymer

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