International Journal of Pharmaceutics 475 (2014) 547–557

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International Journal of Pharmaceutics journal homepage: www.elsevier.com/locate/ijpharm

Poly aspartic acid peptide-linked PLGA based nanoscale particles: Potential for bone-targeting drug delivery applications Tao Jiang a,b,c , Xiaohua Yu a,b,e, Erica J. Carbone a,b,c, Clarke Nelson a,b , Ho Man Kan a,b,e, Kevin W.-H. Lo a,b,c,d, * a

Institute for Regenerative Engineering, University of Connecticut Health Center, School of Medicine, Farmington, CT 06030, United States The Raymond and Beverly Sackler Center for Biomedical, Biological, Physical and Engineering Sciences, University of Connecticut Health Center, School of Medicine, Farmington, CT 06030, United States c Department of Medicine, Division of Endocrinology, University of Connecticut Health Center, School of Medicine, Farmington, CT 06030, United States d Department of Biomedical Engineering, University of Connecticut, School of Engineering, Storrs, CT 06268, United States e Department of Orthopedic Surgery, University of Connecticut Health Center, School of Medicine, Farmington, CT 06030, United States b

A R T I C L E I N F O

A B S T R A C T

Article history: Received 2 July 2014 Received in revised form 24 August 2014 Accepted 27 August 2014 Available online 4 September 2014

Delivering drugs specifically to bone tissue is very challenging due to the architecture and structure of bone tissue. Poly(lactic-co-glycolic acid) (PLGA)-based nanoparticles (NPs) hold great promise for the delivery of therapeutics to bone tissue. The goal of the present research was to formulate a PLGA-based NP drug delivery system for bone tissue exclusively. Since poly-aspartic acids (poly-Asp) peptide sequence has been shown to bind to hydroxyapatite (HA), and has been suggested as a molecular tool for bone-targeting applications, we fabricated PLGA-based NPs linked with poly-Asp peptide sequence. Nanoparticles made of methoxy – poly(ethylene glycol) (PEG)-PLGA and maleimide-PEG-PLGA were prepared using a water-in-oil-in-water double emulsion and solvent evaporation method. Fluorescein isothiocyanate (FITC)-tagged poly-Asp peptide was conjugated to the surface of the nanoparticles via the alkylation reaction between the sulfhydryl groups at the N-terminal of the peptide and the CQC double bond of maleimide at one end of the polymer chain to form thioether bonds. The conjugation of FITCtagged poly-Asp peptide to PLGA NPs was confirmed by NMR analysis and fluorescent microscopy. The developed nanoparticle system is highly aqueous dispersible with an average particle size of 80 nm. In vitro binding analyses demonstrated that FITC-poly-Asp NPs were able to bind to HA gel as well as to mineralized matrices produced by human mesenchymal stem cells and mouse bone marrow stromal cells. Using a confocal microscopy technique, an ex vivo binding study of mouse major organ ground sections revealed that the FITC-poly-Asp NPs were able to bind specifically to the bone tissue. In addition, proliferation studies indicated that our FITC-poly-Asp NPs did not induce cytotoxicity to human osteoblast-like MG63 cell lines. Altogether, these promising results indicated that this nanoscale targeting system was able to bind to bone tissue specifically and might have a great potential for bone disease therapy in clinical applications. ã 2014 Elsevier B.V. All rights reserved.

Keywords: Nanoparticles Targeted drug delivery Peptides Bone diseases

1. Introduction Drugs administrated systematically for bone disorders, such as osteoporosis or osteosarcoma, are often problematic because they are rapidly cleared from the body due to the body’s excretory

* Corresponding author at: University of Connecticut Health Center, School of Medicine, Farmington, CT 06030, United States. Tel.: +1 860 679 2949; fax: +1 860 679 1553. E-mail address: [email protected] (K.W. -H. Lo). http://dx.doi.org/10.1016/j.ijpharm.2014.08.067 0378-5173/ ã 2014 Elsevier B.V. All rights reserved.

system, sometimes even before they are able to fully affect their target sites such as diseased bone. To combat this issue, drugs are generally administered in high dosages and/or frequently which can lead to detrimental systemic side effects (Rizzoli and Reginster, 2013). For instance, the osteoporosis drug Forteo1 (generic name: teriparatide) must be administered daily via injection and is associated with severe systemic side effects such as headaches, orthostatic hypertension, hypercalcaemia, and potentially tumors (Cipriani et al., 2012). It would be safer and more effective if the drug could be delivered specifically to bone tissue via a controllable, sustained release delivery system (Aoki et al., 2012). Therefore,

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there exists a strong need for developing a targeted delivery system specific to bone tissue by using a sustained release drug delivery device. Delivering drugs specifically to bone tissue is very challenging due to the architecture and structure of bone tissue. Bone possesses one of the most complex hierarchical micro- and nano-structures in the human body, and these unique characteristics call for the proper selection of drug carriers that can target specific diseased portions of the bone. Novel delivery devices must provide exposure of the drug to the complex mineralized structure of bone before being excreted by the body. Nanoscale particles are promising therapeutic delivery devices because of their relatively high drug encapsulation capacity and their small size, which makes them compatible with various administration routes including intravenous injection (Tautzenberger et al., 2012). Moreover, the ability of nanoparticles to traverse the nanostructure of bone allows for them to reach bone fracture sites and promote drug release more easily than larger particles. One of the most common biodegradable polymers used and studied for nanoscale drug delivery is the copolymer poly(lactide-coglycolide) (PLGA) (Vert et al., 1998; Buescher and Margaritis, 2007; Xiao et al., 2010). This polymer has demonstrated excellent host biocompatibility, variable physicochemical properties, and predictable degradation rates (Mooney et al., 1996; Sabir et al., 2009). Furthermore, it has been FDA-approved for a number of biomedical applications and is currently used clinically in surgical sutures and drug delivery devices (Jain, 2000; Jain et al., 1998; Lu et al., 2009). Another advantage of using PLGA-based particles is that PLGA can be easily modified and functionalized to allow the covalent attachment of biological molecules. Surface hydrolysis, aminolysis, and oxygen plasma treatment have been proven to be effective methods to attach reactive groups such as carboxyl, amine, hydroxyl, or peroxyl groups to PLGA marcromolecular chains (Croll et al., 2004; Wan et al., 2004). These modifications lead to enhanced material hydrophilicity, creation of cell recognition sites, as well as introduction of functional groups that are readily activated to covalently bind with peptides or growth factors. On the other hand, functionalities can also be introduced into PLGA by synthesizing PLGA-based block copolymers. Such block copolymers can be synthesized by copolymerization of lactide and glycolide in the presence of other monomers or polymers containing desired functional groups (Zhao et al., 2005; Lin et al., 2010). In certain cases, a second functional polymer can be linked to PLGA macromolecules through the terminal carboxyl groups (Nam et al., 2003; Cao et al., 2010). With the introduction of various functionalities, it is then possible to conjugate biological molecules, such as peptides, to the PLGA-based particles for target specificity. For example, the peptide sequence cyclo(1,12)PenITDGEATDSGC (cLABL) was conjugated to the surface of PLGA nanoparticles to target human umbilical cord vascular endothelial cells (HUVECs) which upregulated intercellular cell-adhesion molecule-1 (ICAM-1) expression (Zhang et al., 2008). Therefore, it is believed that the use of short bone-targeting peptide linked PLGA nanoparticles will provide exciting possibilities for bone therapies. Hydroxyapatite (HA) binding domains have been identified from several noncollagenous proteins, including osteocalcin (OCN) and osteopontin (OPN), and have demonstrated high affinity to bone tissue (Fujisawa and Kuboki, 1991; Oldberg et al., 1986, 1988; Gorski, 1992). Interestingly, a short peptide sequence of repetitive aspartic acid (Asp) amino acids has been shown to interact exclusively with hard tissues (bone and teeth) in vitro and in vivo (Kasugai et al., 2000; Yokogawa et al., 2001; Ouyang et al., 2009; Murphy et al., 2007; Ogawa et al., 2013). In fact, aspartic acid peptide sequence has been applied previously by several groups to target drugs to the bone tissue. For instance, it has been reported in

pre-clinical animal studies to promote bone accumulation of small molecular weight agents, such as radiogallium-labeled bone imaging agent (Ogawa et al., 2013). These observations prompted us to fabricate PLGA-based NPs linked with poly-Asp peptide sequence so that the poly-Asp peptide will promote the NPs to interact with bone tissue specifically. To facilitate the imaging of the peptide-linked NPs, fluorescein isothiocyanate (FITC) was tagged to the C-terminus of the poly-Asp peptide sequence. A series of in vitro and ex vivo binding assays were performed to show the exclusive binding affinity of these FITC-poly-Asp NPs to bone tissue. 2. Materials and methods 2.1. Reagents FITC tagged AspAspAspAspAspAspAspCys peptide sequence was synthesized by LifeTein (South Plainfield, NJ). Poly(lactide-coglycolide)-b-poly(ethylene glycol)-maleimide (maleimide-PEGPLGA) and methoxy poly(ethylene glycol)-b-poly(lactide-coglycolide)(methoxy-PEG-PLGA) were purchased from Polyscitech (West Lafayette, IN). HA gels, 10X TBS buffer (Tris buffered saline), and 10X PBS buffer (phosphate buffered saline) were purchased from Bio-Rad (Hercules, CA). 2.2. Fabrication of the peptide-linked nanoparticles Nanoparticles made of 9:1 (w/w) ratio of methoxy-PEG-PLGA and maleimide-PEG-PLGA were prepared using a water-in-oil-inwater (W/O/W) double emulsion and solvent evaporation method as described by Luo et al. (2010). In brief, approximately 108 mg of methoxy-PEG-PLGA and 12 mg of maleimide-PEG-PLGA were dissolved in 4 ml of methylene chloride. Two hundred microliters (200 ml) of 0.22 mm filtered DI water was added to the polymer solution and emulsified using a probe sonicator (Misonix Sonicator 3000, Farmingdale, NY) at 9 W for 30 s on ice. Eight milliliters (8 ml) of 1% sodium cholate solution was then added to the primary emulsion and sonicated at 15 W for 30 s on ice. The formed w/o/w emulsion was subsequently poured into 152 ml 0.5% sodium cholate solution and stirred vigorously for 2 h to allow the evaporation of methylene chloride. Nanoparticles were then washed with DI water and collected by centrifuging at 25,000 rpm with an OptimaTM LE-80 K Ultracentrifuge equipped with a SW 28 rotor (Beckman Coulter, Inc.) for 20 min. FITC-tagged AspAspAspAspAspAspAspCys (FITC-DDDDDDDC or FITC-poly-Asp) peptide sequence was conjugated to the surface of the nanoparticles via the alkylation reaction between the sulfhydryl groups at the N-terminal of the peptide and the CQC double bond of maleimide at one end of the polymer chain to form thioether bonds. Briefly, the peptide solution was mixed with the aqueous nanoparticle suspension at a molar ratio of 1.3:1 and incubated at neutral pH and room temperature overnight. The peptide conjugated nanoparticles were washed extensively with DI water and were either re-suspended in DI water or collected by centrifugation and lyophilization for future uses. Fig. 1 shows the schematic diagram for the synthesis of FITC-poly-Asp-conjugated nanoparticles. 2.3. Characterization of the peptide-linked nanoparticles The morphology of the methoxy-PEG-PLGA/maleimide-PEGPLGA nanoparticles was evaluated by transmission electron microscopy (TEM). In brief, a drop of the nanoparticle solution (10 mg/ml) was deposited onto a PELCO1 grid and then imaged using a FEI Tecnai T12 S/TEM at an accelerating voltage of 80 kV (FEI, Hillsboro, OR).

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The sizes of the nanoparticles were determined by dynamic light scattering (DLS) measurements which were conducted using DynaPro MS800 DLS instrument (Wyatt Technology, Santa Barbara, CA). The DLS data were processed using Dynamics 6.7.3 software. To confirm the successful conjugation of the peptide to the nanoparticles, non-conjugated and FITC-tagged peptide conjugated nanoparticles were visualized under a Zeiss fluorescent microscope (Carl Zeiss MicroImaging, Inc., Thornwood, NY). Furthermore, the nanoparticles were dissolved in deuterated methylene chloride supplemented with 0.1% tetramethylsilane (TMS) (Sigma–Aldrich) and 1H nuclear magnetic resonance (NMR) spectra of non-conjugated and peptide conjugated nanoparticles were recorded utilizing a 800 MHz Agilent VNMRS spectrometer equipped with a triple resonance HCN cold probe. The proton NMR spectra were collected at 25  C with a 10,000 Hz bandwidth, 16 K complex data points, 128 transients, a flip angle of 45 , and a 32-s

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recycle delay to ensure complete relaxation for proper integrations. All spectra were processed and analyzed using Mnova software from Mestrelab Research (Santiago de Compostela, Spain). 2.4. Cell cultures Human mescenchymal stem cells (hMSC) were purchased from Lonza (Walkersville, MD). The cells were maintained in mesenchymal stem cell basal medium (Lonza) supplemented with 10% MCGS, L-glutamine, and 1% antibiotics. hMSCs cultured in osteogenic medium (Lonza) served as a positive control for matrix mineralization. Human osteoblast-like MG63 cells, mouse premyoblasts C2C12, and mouse Schwann nerve SW10 cells were purchased from ATCC (Manassas, VA). They were cultured in the appropriate growth medium according to the suppliers'

Fig. 1. The schematic diagram for the synthesis of FITC-poly-Asp-conjugated nanoparticles (NPs). NPs made of 9:1 weight ratio of methoxy-PEG-PLGA and maleimide-PEGPLGA were prepared using a W/O/W double emulsion and solvent evaporation method. FITC-DDDDDDDC peptide sequence was conjugated to the surface of the NPs via the alkylation reaction between the sulfhydryl groups at the N-terminal of the peptide and the CQC double bond of maleimide at one end of the polymer chain to form thioether bonds.

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formalin at 4  C for three days, and tibias were placed in 95% ethanol for 3 days. All organs and tibias were then transferred to 30% sucrose (Sigma–Aldrich) in PBS overnight. The following day, the organs and tibias were removed from sucrose and embedded in cryosection matrix gel (Thermo) and submerged in 2-methylbutane (Fisher Scientific) at 78.5  C. Then for each organ or tibia, twenty individual 10 mm sections were taken in the coronal plane, placed on charged glass (UltraClear Microscope Slides, Denville Scientific), and mounted with 1% chitosan (Sigma–Aldrich). Once dried, the organ sections were incubated for 1 h at 100 mg/ml of FITC-peptide linked NPs in 1X TBS was added to the surface of the section and incubated for 1 h at room temp with rocking. The sections were subsequently washed three times with 1 ml of 1X TBS to remove any unbound FITC-peptide linked NPs, and coverslipped with a mixture of 50% v/v glycerol(Sigma–Aldrich)/ PBS. Fluorescence was then read using a fluorescent microscope (Axio Observer Z1, Zeiss) and accompanying software (Zen Pro Zeiss). To visualize the FITC-peptide linked NPs, an EGFP filter was utilized with excitation and emission at 488 nm and 509 nm respectively. The organ sections were then additionally visualized using a transmitted light-DIC filter. 2.7. Cytotoxicity assay Cytotoxicity studies were performed using a non-radioactive cell proliferation assay kit (MTS) (Promega, Madison, WI). Various concentrations of the FITC-peptide linked NPs were added to human osteoblast-like MG63 cells in regular growth medium at the time of cell seeding. Cells were collected at days 4 and 7 for the proliferation assays. Cells were incubated with the MTS substrate at 37  C for at least 1 h, or until significant color change was observed. Cell suspensions were then measured using a Biotek SynergyTM HT microplate reader (Winooski, VT) at 490 nm. Fig. 2. The peptide FITC-poly Asp binds directly to HA gels. (A) A decrease in the % of spectrophotometric absorbance in the supernatant corresponds with the relative amount of the FITC-peptide bound to the HA gel. (B) Picture revealed that the peptide FITC-poly Asp bound well to the HA gel (yellowish green). In contrast, FITCpoly Gly hardly bound to HA gel (white). (C) The relative level of FITC-peptides that were bound to HA gel. Error bars represent mean  SD (n = 4). (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)

instructions. The medium was replaced every 3–4 days. All cell cultures were maintained at 37  C in a humidified atmosphere with 5% CO2. 2.5. Preparation of mouse bone marrow stromal cells Bone marrow stromal cells (BMSCs) were gifted from Dr. Cato Laurencin's laboratory at the University of Connecticut Health Center. Briefly, BMSCs were flushed from bone marrow isolated from adult (>6 weeks of age) tibias and femurs of wild type mice by mechanical flushing using a 25 gauge needle. After isolation, the BMSCs were cultured in growth medium (aMEM, 10% FBS, 1% penicillin/streptomycin) on tissue culture polystyrene (TCP). After 3 days of cell culture, half of the medium was exchanged for fresh cell culture medium, and all of the media were changed at 6 days post-harvest. After 7 days in culture, the cells were harvested for calcification studies. 2.6. Fluorescence assay Tibias and other major organ tissues from mice (heart, liver, kidney, spleen, lung, GI tract, and brain) were gifted from Dr. Cato Laurencin’s laboratory at the University of Connecticut Health Center. Briefly, the organs of mice were resected and placed in 10%

2.8. HA gel interaction assays For each binding assay, 50 ml of 50% slurry of HA gels was first washed three times with 1 ml of 1X TBS for equilibration. FITCpeptide at 0.1 mM final concentration or 100 mg of the FITCpeptide linked NPs was added, and the suspension was agitated at room temperature for 1 h. The HA gels were then spun down and the absorbances of the supernatants were measured at the wavelength of the spectrophotometric absorbance peak (490 nm) of the FITC using a Biotek SynergyTM HT microplate reader (Winooski, VT). A decrease in absorbance corresponds with the relative amounts of the FITC-peptides or FITC-peptide linked NPs absorbed to the HA gels. The HA gels were subsequently washed three times with 1 ml of 1X TBS to remove any unbound FITC-peptide or FITC-peptide linked NPs. The bound FITC-peptide or FITC-peptide linked NPs was eluted with 100 ml of 10X PBS buffer and the elutants were measured using a plate reader at 490 nm. 2.9. FITC-peptide linked NPs/cell culture interaction assays SW10 and C2C12 cell lines were cultured in regular growth media for at least 7 days. hMSCs and mouse BMSCs were cultured in osteogenic media (Lonza) for 21 days in order to achieve matrix mineralization. The media were then replaced by the FITC-peptide linked NPs (100 mg/ml) in 1X TBS and incubated for 1 h at room temperature with rocking. The cells were subsequently washed three times with 1 ml of 1X TBS to remove any unbound FITC-peptide linked NPs. Cells were fixed in ice cold 100% methanol for 10 min. A mounting medium with propidium iodide (VECTASHIELD mounting medium1, Vector Laboratories, Burlingame, CA) was used for the counterstaining of DNA and the mounting of coverslips. The bound

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Fig. 3. Characterization of the FITC-poly-Asp conjugated nanoparticles. (A) A TEM micrograph of non-conjugated methoxy-PEG-PLGA/maleimide-PEG-PLGA nanoparticles (NPs) showing round morphology. Size distribution of (B) non-conjugated NPs showing an average size of 79.8 nm and (C) FITC-poly-Asp conjugated NPs showing an average size of 83.3 nm as determined by dynamic light scattering analysis. (D) Fluorescent micrograph of non-conjugated NPs showing minimal fluorescent signals. (E) Proton nuclear magnetic resonance (1H NMR) spectrum of non-conjugated methoxy-PEG-PLGA/maleimide-PEG-PLGA NPs. The maleimide group shows its characteristic peak at 6.7 ppm. (F) Fluorescent micrograph of FITC-poly-Asp conjugated NPs showing abundant green fluorescence demonstrating successful conjugation of the peptide onto the nanoparticle surface. (G) 1H NMR spectrum of FITC-poly-Asp conjugated NPs shows dramatically diminished peak at 6.7 ppm, confirming successful reaction of the thiol group of the peptide with maleimide group. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)

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images of the samples were captured using a high resolution digital scanner. 2.12. Statistical analysis Statistical analyses were performed on samples using the Student’s t-test (Microsoft Excel). The experiments were performed in at least triplicate and the level of significance was set at p < 0.05. An asterisk (*) in the figures denotes significant differences between groups. 3. Results 3.1. Poly-Asp peptide sequence binds directly to HA

Fig. 4. The FITC-poly-Asp NPs bind directly to HA gel. (A) A decrease in the % of spectrophotometric absorbance in the supernatant corresponds with the relative amount of the NPs bound to the HA gel. (B) Picture revealed that the peptide FITCpoly Asp NPs bound well to the HA gel (yellowish green). In contrast, FITC-poly-Gly NPs hardly bound to HA gel (white). (C) The relative levels of FITC-peptide linked NPs that were bound to HA gel. Error bars represent mean  SD (n = 3). (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)

FITC-peptide linked NPs were visualized using a laser confocal fluorescence microscope (Zeiss LSM 510 Meta). 2.10. Matrix mineralization assay After 21 days of culture in osteogenic media, the mineralized matrices of the cultured cells were quantified by colorimetric determination of total calcium deposited in each matrix (Calcium Liquicolor, Stanbio Laboratory, Boerne, TX). This procedure was performed as previously described (Lo et al., 2012a; 2013).

2.11. Alizarin red stain hMSCs or BMSCs were rinsed three times with calcium-fee PBS to remove any unattached cells. Following a wash with distilled, de-ionized water (DDI H2O), the samples were covered in 58.4 mM Alizarin Red solution (Sigma–Aldrich, St. Louis, MO) and incubated by shaking for 15 min at room temperature. The samples were then washed four times with DDI H2O. The

The direct interaction between poly-Asp peptide and HA was demonstrated by a “pull-down” assay using a commercially available HA gel product (Fig. 2). To facilitate the visibility of the peptides, the C-terminus of the peptides was chemically conjugated with a fluorescent probe, fluorescein isothiocyanate (FITC). FITC tagged GlyGlyGlyGlyGlyGlyCys peptide sequence (FITC-poly-Gly) was synthesized as a control. It is believed that small neutral charged amino acids, such as glycine or alanine, are the most appropriate control amino acids for the poly-Asp sequence (Genevaux et al., 2002). Fig. 2A suggests that both FITC-poly-Asp and FITC-poly-Gly peptides were absorbed to the HA gels as shown by a significant decrease in the absorbance percentage of the supernatants. It should be noted that FITC-poly-Asp had greater affinity to HA gel than FITCpoly-Gly since a greater decrease in the percentage of absorbance of the supernatant was observed, i.e., 96% vs. 13%. To confirm the direct interaction between the FITC-peptides and HA gel, the FITC-peptides bound to the HA gels were then washed extensively with 1X TBS to remove any non-specific binding. The bound FITC-peptides were subsequently eluted with 10X PBS. Fig. 2B and C clearly demonstrates that FITC-poly-Asp bound well to HA, whereas FITC-poly-Gly did not bind to HA, indicating that there is no interaction between the FITC tag and the HA gel. Taken together, these observations confirmed the direct interaction between poly-Asp sequence and HA. 3.2. Characterization of the FITC-poly-Asp linked NPs The morphology of the nanoparticles was evaluated by TEM. Fig. 3A shows a typical TEM micrograph of non-conjugated nanoparticles. The nanoparticles were found to be round in morphology. The nanoparticle size distribution was analyzed by dynamic light scattering (DLS) measurements (Fig. 3B and C). The non-conjugated nanoparticles had an average size of 79.8 nm with a polydispersity of 68.7%. In comparison, the peptide conjugated nanoparticles had an average size of 83.3 nm with a polydispersity of 71.5%. Peptide conjugation onto the polymer nanoparticle surface was verified by fluorescent microscopy. Since the peptide was tagged with the fluorescent FITC, the conjugated nanoparticles showed green fluorescence. Fig. 3D shows no visible fluorescence for the non-conjugated nanoparticles. However, the peptide conjugated nanoparticles showed abundant fluorescence (Fig. 3F), indicating the presence of the FITC-tagged peptide on the nanoparticle surface. Conjugation of the peptide onto the polymer nanoparticles was further confirmed by 1H NMR. The 1H NMR spectrum of the non-conjugated nanoparticles is shown in Fig. 3E. The characteristic peak for the hydrogens of the maleimide group appeared at 6.7 ppm. Hydrogen peaks near 5.2 ppm and 1.5 ppm were attributed to the CH group and CH3 group of lactide, respectively; while the peak near 4.8 ppm was attributed to the CH2 group of glycolide. The hydrogen peaks of poly(ethylene glycol) appeared between 3.3 ppm and 3.6 ppm. The 1H NMR spectrum of the peptide conjugated nanoparticles is shown in Fig. 3G. After peptide

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Fig. 5. The FITC-poly-Asp NPs bind to mineralized matrix in hMSC as well as mouse BMSC cultures. (A and C) After 21 days of culture of hMSC or mouse BMSC in osteogenic medium, quantitative calcium levels were determined. Calcium levels were significantly increased in cells cultured in osteogenic medium. In addition, mineralized matrix was visualized using alizarin red staining of deposited calcium. FITC-poly-Asp NPs were able to bind to the mineralized matrix produced from hMSC (B) as well as mouse BMSC (D) when they were cultured in an osteogenic medium condition. As a control experiment, FITC-poly-Asp NPs bound to neither hMSC nor BMSC cultures when they were grown in regular growth medium. Podium iodide (PI) stained nuclei (red) indicates the presence of cells. Scale bars are 20 mm for panel B and 50 mm for panel D. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)

Fig. 6. The FITC-poly-Asp NPs do not interact with C2C12 and SW10 cell cultures. Podium iodide (PI) stained nuclei (red) indicates the presence of cells. Scale bars are 50 mm. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)

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conjugation, the characteristic peak of maleimide at 6.7 ppm was still present but diminished dramatically. The area underneath each peak was further integrated and determined using Mnova software. Using the CH3 group of lactide (peak near 1.5 ppm) as an internal reference, we found that the area of the maleimide peak reduced by more than 60%, indicating that a large amount of the maleimide functional group had reacted with the peptide. 3.3. FITC-poly-Asp NPs bind to HA gel and mineralized matrix in cultured cells We next sought to determine whether FITC-poly-Asp NPs are able to bind to HA gel (Fig. 4). FITC-poly-Gly NPs were used as a control. Consistent with the data in Fig. 2, both FITC-poly-Asp NPs and FITC-poly-Gly NPs were absorbed by HA gels, as revealed by a significant decrease in the percentage of absorbance in the supernatants (Fig. 4A). It should be noted that FITC-poly-Asp NPs had greater affinity to HA gel than FITC-poly-Gly NPs as the percentage absorbance of the supernatant decreased more, i.e. 85% vs. 32%. The direct interactions between the FITC-peptide NPS and HA gel were confirmed by quantifying the FITC-peptides NPs in the HA gels. Fig. 4B and C clearly demonstrate that FITCpoly-Asp NPs bound well to HA, whereas FITC-poly-Gly NPs did not bind well to HA. We then investigated the interaction between FITC-poly-Asp NPs and mineralized matrix produced by mouse bone marrowderived stromal cells (BMSCs) and mesenchymal stem cells (hMSCs) (Fig. 5). Fig. 5A and C confirms the deposition of mineral on both hMSC and mouse BMSC after they were cultured in osteogenic media for 21 days. FITC-poly-Asp NPs were confirmed to bind to the mineralized matrix produced from hMSC (Fig. 5B) as well as from mouse BMSC cultures (Fig. 5D). In contrast, FITC-poly-Asp NPs did not bind to either hMSC or BMSC cultures when they were grown in regular growth media, indicating the binding specificity of FITC-polyAsp NPs to mineralized matrix. It is also worth noting that FITC-poly-Asp NPs did not interact with other cell cultures such as mouse premyoblast (C2C12) or mouse Schwann nerve

cells (SW10) (Fig. 6), suggesting that the binding of FITC-polyAsp NPs is specific for mineralized tissue. Taken together, these observations indicate that the FITC-poly-Asp NPs binds specifically to HA gel as well as to mineralized matrix in vitro. 3.4. FITC-poly-Asp NPs bind specifically to tibia sections ex vivo Because the FITC-poly-Asp NPs were able to bind to mineralized tissue in vitro, we hypothesized that FITC-poly-Asp NPs would also bind to bone tissue due to the high mineral content in bone. To test this hypothesis, we used mouse tibia sections for our binding assays. As demonstrated in Fig. 7A, FITC-poly-Asp NPs successfully bound to the tissue section surface by covering almost the whole section, whereas the control FITC-poly-Gly NPs failed to do so indicating the potential specificity of the poly-Asp tag. To demonstrate that the FITC-poly-Asp NPs specifically bind to bone tissue, we also evaluated the binding of our NPs to other major organ tissues including brain, heart, liver, spleen, kidney, lung, and GI tract. Fig. 7B reveals that the FITC-poly-Asp NPs interact preferably to bone tissue ex vivo. 3.5. Cellular proliferation in nanoparticles-treated osteoblast-like MG63 cells The effects of FITC-poly-Asp peptide linked nanoparticles on cell proliferation were assessed. Human osteoblast-like MG63 cells were cultured in regular growth medium with various concentrations of nanoparticleasindicatedinFig.8.Cellularproliferationwasevaluated using the MTS assay method, which is a routinely used assay for accessing in vitro cytotoxicity of a biomaterial (Malich et al., 1997). Untreated cells were used as a control. The absorbance in the figure serves as a surrogate for the number of living cells in culture (Lo et al., 2013). As shown in Fig. 8, there were no significant differences in cellular proliferation at day 4 and day 7 for cells treated with nanoparticles when compared to the untreated control group. These observations indicate that different concentrations of our nanoparticles, i.e.,1–100 mg/ml, do not induce cytotoxicity to MG63 cells.

Fig. 7. FITC-poly-Asp NPs bind to mouse tibia section. (A) FITC-poly-Asp NPs or FITC-poly-Gly NPs (100 mg/ml) in 1X TBS was added to the surface of the tibia section and incubated for 1 h at room temp with rocking. The sections were subsequently washed three times with 1 ml of 1X TBS to remove any unbound FITC-peptide linked NPs. The FITC signals were captured by differential interference contrast (DIC) microscopy. Scale bars are 100 mm. (B) FITC-poly-Asp NPs (100 mg/ml) in 1X TBS was added to the surface of the various organ sections, as indicated in the figure, and incubated for 1 h at room temp with rocking. The sections were subsequently washed three times with 1 ml of 1X TBS to remove any unbound FITC-peptide linked NPs. The FITC signals were captured by differential interference contrast (DIC) microscopy. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)

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Fig. 7. (Continued)

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Fig. 8. Effects of various FITC-poly-Asp NPs concentrations on cellular proliferations in human osteoblast-like MG63 cells. Cellular proliferation of MG63 cells was measured at (A) day 4 and (B) day 7 using a commercial cell proliferation assay kit (MTS). Error bars represent mean  SD (n = 4).

4. Discussion The objective of the present research is to design and develop a novel targeted PLGA-based drug delivery device for treatment of bone diseases. It should be noted that protein and peptide-based medications, vaccines, small molecule drugs, and nucleotides can be easily incorporated into biodegradable PLGA-based nanoscale delivery systems by a variety of methods (Chen, 2010; Gu et al., 2013). Several types of common incorporation techniques include physical adsorption, covalent bonding, and entrapment (Maia et al., 2013; Laurencin et al., 2014). Each method has inherent advantages depending on the biomolecule's chemical structure, intended construct fabrication method, and desired drug release profile (Laurencin et al., 2014; Lo et al., 2012b). Therefore, the NP system proposed here may be useful in delivering current and/or future therapeutics specifically to bone tissue. It is interesting to note that PLGA-based NPs have also been proposed as gene carriers for gene therapy treatment of various disease models (Bivas-Benita et al., 2004; Zou et al., 2009; Liang et al., 2011). Gene therapy for osteoporosis treatment has been actively investigated by a number of research groups with promising results (Baltzer et al., 2001; Egermann et al., 2005; Zhang et al., 2012); therefore, it is anticipated that the success of our NP system may also serve as a novel non-viral gene delivery vehicle for osteoporosis treatment. Although monoclonal antibodies have shown promise as tissuetargeting ligands, their potential for use as targeted drug delivery

systems has been limited by several factors. These include large molecular size, high production cost, and unwanted immune response (Lu et al., 2013). In addition, antibody-based drug delivery may result in nonspecific antibody uptake by Fc receptor-expressing normal cells, which will cause higher toxicity in bone marrow and liver (Cheng and Allen, 2010). These limitations can be overcome by using a short peptide sequence, which is smaller in molecular size, less immunogenic, inexpensive to produce, and easy to manipulate (Lu et al., 2013; Lee et al., 2007). It should also be noted that small molecule alendronate (ALN) has been proposed as a bone-targeting moiety (Chen et al., 2012; de Miguel et al., 2014). An elegant study conducted by Wang et al. demonstrated that aspartic acid peptide sequence would favorably recognize resorption sites in skeletal tissues, whereas ALN directs the delivery system to both formation and resorption sites (Wang et al., 2007). Thus, aspartic acid peptide conjugated drug delivery would be a better targeted therapeutic strategy for osteoporosis. It is also worth noting that the ability of NPs to interact with their target tissue can be improved by prolonging the circulation times (Luhmann et al., 2011; de Miguel et al., 2014). One major factor affecting circulation time is the size of the particles. For instance, it has been shown that smaller NP sizes (10–70 nm) penetrate very small capillaries, while moderate NP sizes (70–200 nm) have been reported to have extended circulation times (Luhmann et al., 2011). Thus our peptide-linked NPs with an average size of 80 nm seem to be optimal for bone-targeting application since their size should favor extended circulation within the bloodstream. Here, we would like to emphasize that binding of NPs to the tibia sections ex vivo does not represent the interaction between the NPs and the bones in vivo. A completely different set of factors would determine ability to attain bone targeting by the NPs in vivo, including alternative disposition pathways, number and stability of the poly-aspartate residues on the surface of the NPs, as well as the accessibility of the bone surface to the NPs. Nevertheless, future in vivo pharmaco kinetic and pharmacodynamic studies in small animal models are needed to address these questions. To date, the exact binding mechanism between poly-Asp peptide sequence and HA is still unclear. It has been proposed that the ionic interaction is responsible for the binding via the negatively charged poly-Asp peptide sequence and the positively charged calcium ion within the mineral component of bone at physiological pH (Murphy et al., 2007; Sarig, 2004). Additionally, in vitro, the negatively charged membrane surface of cultured cells may repel the negatively charged poly-Asp peptide (Goldenberg and Steinberg, 2010). These observations may account for the specificity of the poly-Asp peptide to bone tissue in our in vitro cell studies (Figs. 5 and 6). It should be emphasized that PLGA-based biomaterials have been approved by the FDA for a number of clinical applications, and their bioavailability and safety in humans are well-documented. In addition, our cytotoxicity study suggests that our NP system does not induce cytotoxicity to MG63 cells at the amount used in this study (Fig.8).Itisthereforeanticipatedthatourdrugdeliverysystemislikely to be translatable to clinical applications. Nevertheless, more extensive in vitro and in vivo characterizations of our NP system are needed in order to move forward to clinical studies. In conclusion, bone-targeted NPs hold great potential for clinical applications in delivering drugs to bone niches, increasing local drug concentrations, reducing off-target side effects, and lengthening the therapeutic window. In this study, it has been shown that (i) FITCpoly-Asp NPs interact specifically to mineralized matrix in vitro; (ii) FITC-poly-Asp NPs bind exclusively to tibia bone section; (iii) FITCpoly-Asp NPs do not induce cytotoxicity to cell cultures. The data presented here provide a necessary prerequisite for the use of the poly-Asp NPs for future in vivo bone-targeting studies in animals.

T. Jiang et al. / International Journal of Pharmaceutics 475 (2014) 547–557

Acknowledgements The work was partly supported by the Connecticut Stem Cell Research Foundation grant (13-SCA-UCHC-01) to Dr. Kevin Lo. We would like to thank Dr. Cato T. Laurencin, the Director of the Institute for Regenerative Engineering (IRE) and the University Professor of the University of Connecticut, for his excellent leadership. We wish to thank the Raymond and Beverly Sackler Center for Biomedical, Biological, Physical and Engineering Sciences Foundation for supporting our Institute. The authors would also like to thank Dr. Ewa Folta-Stogniew at the Keck Biotechnology Resource Laboratory of Yale University School of Medicine for her technical assistance on the dynamic light scattering experiments. Finally, we wish to thank all members of the IRE, past and present, and numerous colleagues and friends for their helpful discussions. References Aoki, K., Alles, N., Soysa, N., Ohya, K., 2012. Peptide-based delivery to bone. Adv. Drug Deliv. Rev. 64, 1220–1238. 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