nanoparticles for targeted drug delivery to bone

Journal of Controlled Release 122 (2007) 24 – 30 www.elsevier.com/locate/jconrel

Design of surface-modified poly(D,L-lactide-co-glycolide) nanoparticles for targeted drug delivery to bone Sung-Wook Choi, Jung-Hyun Kim ⁎ Nanosphere Process and Technology Laboratory, Department of Chemical Engineering, Yonsei University, 134 Shinchon-dong, Sudaemoon-ku, Seoul 120-749, South Korea Received 12 February 2007; accepted 7 June 2007 Available online 14 June 2007

Abstract Poly(D,L-lactide-co-glycolide) (PLGA) nanoparticles, modified with both alendronate and polyethylene glycol (PEG), were prepared by dialysis method without additional surfactant to evaluate the potency of the bone-targeted drug delivery. Alendronate, a targeting moiety that has a strong affinity for bone, was conjugated to PLGA polymer via carbodiimide chemistry. Monomethoxy PEG(mPEG)–PLGA block copolymers with different molecular weights of mPEG (Mn 550, 750, and 2000) were synthesized and used for a hydrophilic layer on the surface of the nanoparticles to avoid reticuloendothelial system (RES). The surface-modified PLGA nanoparticles with various ratios of alendronate and mPEG densities on their surface were evaluated by adsorption study onto hydroxyapatite (HA). It was confirmed that alendronate-modified nanoparticles had a strong and specific adsorption to HA. The amount of nanoparticles absorbed onto HA tended to be smaller when the content of alendronate was decreased and the large block length of mPEG was found to reduce the potency of alendronate. © 2007 Elsevier B.V. All rights reserved. Keywords: Bone-targeted drug delivery; Dialysis method; PLGA nanoparticles; Alendronate; mPEG–PLGA block copolymer

1. Introduction Targeted drug carriers with high affinity for specific organs, tissues, and cells were introduced in 1906; this concept has been gaining much attention recently [1]. Generally, there have been two major issues: site specificity for desired organ and the duration time of the circulation of the drug carrier. Targeted drug delivery is the most promising way to reduce side effects of a specific drug. For example, in the treatment of osteoporosis with estrogen, the distribution of estrogen to other tissues except the bone can cause several side effects such as intrauterine hemorrhage and occasionally both endometrial and breast cancer [2,3]. To deliver nanoparticles to a desired site, site specificity, in terms of ‘active targeting’, is based on the affinities between targeting moiety and desired organ. The attachment of a specific moiety onto the surfaces of nanoparticles can improve the targeting efficiency. Recently, Wang et al. reviewed the polymeric conjugate system for bone-targeted ⁎ Corresponding author. Tel.: +82 2 2123 7633; fax: +82 2 312 0305. E-mail address: [email protected] (J.-H. Kim). 0168-3659/$ - see front matter © 2007 Elsevier B.V. All rights reserved. doi:10.1016/j.jconrel.2007.06.003

drug delivery [4]. Many researchers [5–7] have also shown specific and efficient cellular uptake of particles that had been modified with a functional ligand with a high affinity for target cells. The second requirement of targeted drug delivery is a long blood circulation time of the nanoparticles. To avoid the reticuloendothelial system (RES), a hydrophilic surface and small particle size under 100 nm are the most often mentioned criteria. The originally hydrophobic particles, after intravenous administration, will become coated by blood components (opsonins) and rapidly taken up by RES. Therefore, nanoparticle surfaces should be modified with hydrophilic components such as PEG [8]. The goal of surface modification is to make the particles unrecognizable by the RES and guide it to the desired site. Particle size is also a crucial factor for prolonged circulation time in the blood stream. Generally the smaller nanoparticle with more hydrophilic surface shows less RES uptake. PEG–PLGA block copolymer was chosen as the micellar drug carrier because both PEG and PLGA are most often used for the drug delivery system and also approved by FDA (Food and Drug Administration). The main application of

S.-W. Choi, J.-H. Kim / Journal of Controlled Release 122 (2007) 24–30

PLGA is focused on the drug delivery system because of controllable degradation period (about 1∼6 months), while polylactic acid (PLA) has long degradation period over 12 months and crystalline polyglycolic acid (PGA) is often used for the suture application. For the successful bone-targeted drug delivery, targeting moieties with affinity to bone should be carefully taken into consideration. Alizarin Red S, tetracycline, calcein, and bisphosphonate are known to have strong affinities for hydroxyapatite (HA), a major inorganic component of hard tissues such as bone and teeth. Among those materials, bisphosphonate has been used clinically in the treatment of osteoporosis and therefore can be used as a targeting moiety of nanoparticles. In this article, alendronate (AL), a type of bisphosphonate, was chosen as the targeting moiety because of easy conjugation system between AL and PLGA through carbodiimide chemistry. We modified the surfaces of the nanoparticles with AL and PEG hydrophilic components to evaluate the potency of their targeted delivery to bone. The present results will be helpful in designing the nano-structured particles using block copolymers for the targeted drug delivery systems. 2. Materials and methods 2.1. Materials Poly(D,L-lactide-co-glycolide) (PLGA) (D,L-lactide 75: glycolide 25) with an average molecular weight of 10,000 was purchased from Wako Pure Chemicals (Japan). Distilled water from a Milli-Q water purification system (Direct-Q 3, USABedford, MD) with a resistance of 18 M Ωcm− 1 was used. Glycolide and D,L-lactide were purchased from Boehringer Ingelheim (Germany); both were recrystallized from ethyl acetate and dried under vacuum prior to use. Stannous octoacte, 17β-estradiol (estrogen), and monomethoxy PEG (mPEG) with various Mn of 550, 750, and 2000 were purchased from Sigma Chemical Company. Dicyclohexyl carbodiimide (DCC), N-hydroxyl succinimide (NHS), pyrene, and hydroxyapatite (HA) were purchased from Sigma. Dimethylformamide (DMF, purity N 99.0%), Dimethyl sulfoxide (DMSO, purity N99.0%) and diethyl ether (purity N99.0%) were purchased from Junsei Chemicals (Japan). Alendronate was supplied from M/S Drugs & Intermediate Ltd. (India). 2.2. Synthesis of mPEG–PLGA block copolymer mPEGs were dehydrated under vacuum at 70 °C for 12 h and were used without further purification. D,L-Lactide (3.94 g) and glycolide (1.06 g) (molar ratio 75:25) were put into a glass ampoule with a various content of mPEGs. Stannous octoate (dissolved in toluene) was added at a 0.05 (w/w) %. The ampoule was evacuated by a vacuum pump at 25 °C for 4 h and then sealed using torch. The ampoule was heated in an oil bath at 130°C for 24 h. The resulting polymers were purified by dissolution in acetone and then precipitated in water. The purified polymers were dried at 25 °C for 12 h in a vacuum oven.

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2.3. Conjugation of AL to PLGA Two grams of PLGA were dissolved in 20 mL acetone and activated by 100 mg DCC and 60 mg NHS overnight at room temperature. The by-product of the activation, dicyclohexylurea, was removed using a syringe filter with 0.45 μm pore size. The NHS-activated PLGA was precipitated in cold diethyl ether and dried at 25 °C for 4 h in a vacuum oven. Two hundred milligrams activated PLGA and 3 mg AL were dissolved in 20 mL of a mixture of 19 mL DMSO and 1 mL water, and then stirred for 24 h at room temperature. The conjugated polymer was precipitated in cold diethyl ether and water sequentially, and then dried at 25 °C for 4 h in a vacuum oven. 2.4. Characterization of mPEG–PLGA block copolymer and AL–PLGA conjugate The 1H NMR spectra were recorded in CDCl3 on Inova 500 MHz (Varian, USA). The mPEG–PLGA block copolymers were characterized by FT-IR (TENSOR27, BRUKER, The Netherlands), gel permeation chromatography (GPC, Waters Breeze System, Waters Co., USA). The GPC column consisted of a series of μStyragel® columns (HR5, HR4, HR1, HR5E) and tetrahydrofuran (THF) was used as an eluent at a flow rate of 1 mL/min and 1 × 103 Pa. The AL–PLGA conjugate was characterized by GPC equipped with a UV/VIS spectrophotometer (UV-1601, Japan) as the detector. PLGA and the AL–PLGA conjugate were dissolved in THF at the same concentration. For the GPC profiles, the absorbance of the eluted polymer solution was measured by UV/VIS spectrophotometer at 268 nm for PLGA and 295 nm for AL–PLGA conjugate. The micellar formation of mPEG– PLGA block copolymer was confirmed by fluorescence technique by using pyrene as a probe [9]. 2.5. Preparation and characterization of nanoparticles The nanoparticles modified with mPEG and AL were prepared by the dialysis method [10–12]. In brief, 100 mg polymer with various weight ratios (AL–PLGA conjugate and mPEG–PLGA block copolymer) and 2 mg estrogen were dissolved into 20 mL DMF. Pyrene was used as a dye to measure the concentration of nanoparticle suspended in water by UV/VIS spectrophotometer. The solution was dialyzed for 12 h in a dialysis tube (MWCO 6000∼8000) in 3 L of PBS, which was exchanged every 2 h. The final suspension was filtrated through a syringe filter with a 0.45 μm pore size and dialyzed extensively to remove unloaded estrogen before use. To determine the loading amount, freeze-dried nanoparticles were dissolved in 4 mL DMF. The estrogen concentration was measured by the UV/VIS Spectrophotometer at 285 nm. The amount of estrogen loaded onto the nanoparticles was calculated by a standard curve obtained from estrogen in DMF. The mean particle diameter of the nanoparticles suspended in water media was determined by dynamic light scattering method (Zeta plus, Brookhaven Inst. Co., U.S.A).

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2.6. HA affinity assay of AL-modified nanoparticles HA was suspended in PBS at a 10 mg/mL concentration. Two milliliters of the nanoparticle dispersion at 0.25% concentration were mixed with 2 mL of HA suspension and then gently shaken for 1 h at room temperature. After filtration with a syringe filter with 0.45 μm pore size, the concentration of nanoparticles remained in the water suspension was measured by UV/VIS spectrophotometer at 337 nm for the calculation of the amount bound to HA.

Table 1 The molecular weight and ratio of mPEG–PLGA block copolymer

mPEG550–PLGA mPEG750–PLGA mPEG2000–PLGA

Mw

Mn

PDI (GPC)

Weight ratio (MnmPEG/MnPLGA)

4804 5017 6321

2965 3279 4026

1.62 1.53 1.57

0.23 0.30 0.99

3.1. Characterization of mPEG–PLGA block copolymers and AL–PLGA conjugate

2.7. In vitro drug release Twenty milliliters of the nanoparticle suspension (3.2 mg/mL) was sealed in a dialysis tube, and incubated in 100 mL PBS with gentle shaking at 37°C. At predetermined time intervals, 4 mL of the solution was collected from the released media and replaced with fresh PBS. The released estrogen was analyzed by UV/VIS spectrophotometer at 280 nm. 3. Results and discussion Fig. 1 depicts the overall procedure for the preparation of surface-modified nanoparticle. AL was chosen as the targeting moiety due to the strong affinity to bone and PEG component was selected as the hydrophilic surface; both were incorporated into the design of the functional nanoparticle. The well-known dialysis method, which does not need surfactant, was selected for the preparation of nanoparticles since it was assumed that the usage of additional surfactant can reduce the functionality of AL. Both AL and mPEG were existed on the surface of nanoparticles due to the hydrophilicity of those molecules. The mPEG–PLGA conjugate plays a role as a stabilizer so that additional stabilizers are not necessary for stable nanoparticle formation.

Fig. 1. Overall schematic procedure for the preparation of surface-modified nanoparticle.

The mPEG–PLGA block copolymers were synthesized by ring opening polymerization to provide colloidal stability and hydrophilic layer to the resulting nanoparticles. The molecular weights and block ratios of the resulting block copolymers are listed in Table 1. The block length of the PLGA component in the mPEG–PLGA block copolymer was adjusted to be nearly constant for the evaluation of the mPEG effects; the weight ratios of mPEG to PLGA were varied (0.23, 0.30, and 0.99). The basic chemical structure of mPEG–PLGA copolymer is confirmed by 1H NMR (data not shown). The large peak at 3.65 ppm came from the methylene groups of mPEG and the multiplets at 5.13 and 4.78 ppm were corresponding to the lactic acid proton and the glycolic acid protons, respectively. The peak at 1.55 ppm was attributed to the methyl groups of lactic acid repeat units. The infrared spectra of the mPEG–PLGA copolymer, mPEG, and PLGA are shown in Fig. 2. The major peaks assigned to the structure of mPEG–PLGA were 2900– 3000 cm− 1 (C–H stretching of CH3), 1760 cm− 1(ester C = O stretching), and 1080 cm− 1 (O–CH2 stretching). It was confirmed that the mPEG–PLGA block copolymers were successfully synthesized. With the formation of micelles, the hydrophobic pyrene can be partitioned into the core of micelles, resulting in the increase in the absorbance of the dye. As seen in Fig. 3, the abrupt increase in absorbance was observed over 0.025 wt.%

Fig. 2. FT-IR spectra of PLGA, mPEG, and mPEG–PLGA block copolymer.


Fig. 3. UV–vis spectra of mPEG2000–PLGA block copolymer aqueous solution containing pyrene at 25 °C.

concentration. The critical micellar concentration (CMC) of mPEG2000–PLGA micelles was observed to be approximately 0.0255 wt.%. This result also suggests that the hydrophobic estrogen may be located in the inner hydrophobic core of the mPEG–PLGA micelles. AL was selected as the targeting moiety among the various bisphosphonates available because it had a primary amine group that was easily conjugated to PLGA via carbodiimide chemistry. The carbodiimide chemistry is well-known and often used for the amide bond reaction between a carboxyl group and a primary (or secondary) amine group [13,14]. The uncapped PLGA polymer has two reactive groups, a hydroxyl group and a carboxylic acid group. The carboxyl group of PLGA was activated with NHS in the presence of DCC. After the addition of AL, NHS was replaced by AL to form the AL–PLGA conjugated polymer. The resulting amide bond is not cleaved while the ester bond is degraded via hydrolysis. Fig. 4 shows the GPC profiles of both PLGA and AL–PLGA conjugate. It was found that the AL–PLGA conjugates were eluted a little earlier than PLGA since there was a slight difference in the molecular weight of PLGA and the AL–PLGA conjugate. From the GPC profile of the AL–PLGA conjugate, it was confirmed that the conjugation reaction was successfully conducted since the unconjugated PLGA was not detected at the 295 nm wave length [15].

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polymer. Most of the conjugate systems do not show evidence of a drug release mechanism. Only a few studies have shown a hydrolysable ester bond as a release mechanism [18]. The cleavage of the drug from the polymer backbone is assumed to be the primary release mechanism, which is not easily controlled through simple conjugation. The hydrodynamic volume of the conjugates is also relatively large since it is based on the linear water-soluble polymers. Molecular weights of backbone polymer are hard to be optimized for the prodrug system. Wang et al. reported that the molecular weights of the backbone polymers should be sufficiently large to provide adequate circulation time during which to extravasate. However, high molecular weight polymer may possibly be retained in both the kidney and bone marrow (and also the liver and spleen) [4]. The range of molecular weights should be carefully considered in polymeric prodrug systems. These drawbacks are the major limitations of the conjugate prodrug system for its clinical application. In this article, a particulate system using surfacemodified nanoparticles was proposed as a candidate for bonetargeted drug delivery. AL was chosen as the targeting moiety and the PEG component was selected as the hydrophilic surface; both were incorporated into the design of the functional nanoparticle. As previously mentioned, AL–PLGA conjugates and mPEG– PLGA block copolymers were successfully prepared. The wellknown dialysis method, which does not need surfactant, was selected for the preparation of nanoparticles since it was assumed that the usage of additional surfactant can reduce the functionality of AL. Both AL and mPEG existed on the surface of nanoparticles due to the hydrophilicity of those molecules. The mPEG–PLGA conjugate plays a role as a stabilizer so that additional stabilizers are not necessary for stable nanoparticle formation. The series of mPEG–PLGA block copolymers with different molecular weights of mPEG (550, 750, and 2000) were

3.2. Preparation and characterization of nanoparticles Much attention has been devoted to the bone-targeted drug delivery and also many studies have focused on prodrugs using bisphosphonate-drug conjugates [16–24]. This system has many advantages such as improvement in the water solubility of the hydrophobic drug, prolonged circulation time, minimized metabolism of the loaded drug, etc. [25]. However, the prodrug system also has some weak points in terms of its release mechanism and the molecular weight selection of the backbone

Fig. 4. GPC–UV profiles of the alendronate–PLGA conjugate. (PLGA: 268 nm; alendronate–PLGA conjugate: 295 nm).

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Table 2 The characteristics of nanoparticles with respect to mPEG–PLGA block copolymer Particle size mean ± SD (nm) a

mPEG550–PLGA mPEG750–PLGA mPEG2000–PLGA a

Estrogen loaded

Estrogen unloaded

43.5 ± 2.3 46.6 ± 3.6 57.3 ± 2.2

42.8 ± 3.8 45.5 ± 4.5 54.6 ± 4.3

Estrogen loading efficiency mean ± SD (%) 9.5 ± 0.27 10.1 ± 0.38 9.7 ± 0.32

Determined by dynamic light scattering measurement.

synthesized. Table 2 shows the characteristics of nanoparticles prepared with block copolymers with different mPEG lengths. The average particle size was slightly increased with respect to the increase in mPEG molecular weight, due in part to the hairy structure of the mPEG component as it extends into the water phase. Riley and Heald reported on the hydrodynamic radius (Rh) of nanoparticles prepared by block copolymers. They found that Rh of micelle was dependent on the block lengths of the copolymer and not on the polymer concentration in solvent during nanoparticle preparation when the length of the hydrophilic block was relatively larger than that of the hydrophobic block. Additionally, the hydrophilic block was the more dominant factor than was the hydrophobic block for the determination of the Rh [26,27]. Therefore, the diameter of the hydrophobic domain is assumed to be nearly the same since the molecular weight of the PLGA block is constant at approximately Mn 2300. The similar sizes of the hydrophobic domains are also the reason for nearly identical loading efficiency of drug although the average particle size is varied from 43 to 57 nm. It is known that the vasculature in bone have pores of approximately 80∼100 nm [24]. The hydrodynamic sizes of nanoparticles should be less than at least 80 nm to extravasate and be localized in bone after i.v. administration. Particles larger than the local bone vasculature pores are retained in the marrow, which cause potential toxicity [4]. The sizes of resulting nanoparticles are under 80 nm (as measured by DLS), which is assumed to be the size used in bone-targeted drug delivery based on particles.

Fig. 5. The binding of modified PLGA nanoparticles to HA (n = 3, nanoparticles modified with AL 80% and mPEG 20%).

nanoparticles was increased with an increase in the content of AL–PLGA conjugate, while nanoparticles with only mPEG have a non-specific adsorption to HA since the nanoparticle with much AL–PLGA content has more binding site to HA. Fig. 5 shows the binding capacity of AL-modified nanoparticles to HA with respect to Mn of mPEG. While the adsorbed amount is nearly the same in both mPEG500–PLGA and mPEG750– PLGA, it was less than 50% when mPEG2000–PLGA was used. This result was due to the much longer block length of the mPEG that interfered with the potency of AL binding to HA. The long chain length of PEG is practically helpful in the reduction of a RES response. However, its length should be optimized so as not to weaken the potency of the targeting moiety.

3.3. Affinity study to HA Recent research has reported the long-term adverse effects of hormone replacement therapy, which are caused by a lack of bone specificity [28,29]. The affinity of a targeting moiety to a specific site is a major point that should be considered. It was reported that both Alizarin Red S and bisphosphonate had greater affinities than did tetracycline and calcein [30]. Among these molecules, AL is considered to be the most suitable moiety for targeting to bone because of its present clinical use. The HA adsorption assay was performed to evaluate the affinity of the surface-modified nanoparticle to bone. HA is used as a model of bone since it is the major component of bone. It is reported that the binding mechanism of bisphosphonate to HA is simple adsorption [31]. The binding capacity of

Fig. 6. Release profiles of estrogen from nanoparticles with different mPEG block lengths (n = 3, nanoparticles modified with AL 80% and mPEG 20%).


3.4. In vitro drug release In our particle design, AL serves mainly as a targeting moiety to guide the nanoparticles to bone and estrogen is used as a drug, while both are clinically used as drugs to treat osteoporosis. The amide linkage between AL and PLGA is not cleavable; AL is fixed on the particle surface and is not released from the nanoparticles. If the bisphosphonate is conjugated to PLGA through an ester bond, it can serve as not only the drug but also the targeting moiety. Fig. 6 shows the release profiles of estrogen from the nanoparticles with different mPEG block lengths. The initial burst in the release of estrogen was observed because the estrogen was not chemically conjugated to the polymer but physically entrapped inside the core of the nanoparticles. It was shown that there was not much difference in their release behaviors with respect to the length of mPEG block. However, the release profile of estrogen from mPEG2000–PLGA nanoparticles was found to be slightly faster than those of the others. It seems that mPEG2000–PLGA nanoparticles with large hydrodynamic diameter accelerate the estrogen release due to the increase in PEG chain hydration and molecular mobility. 4. Conclusion For bone-targeted drug delivery, functional nanoparticles for therapeutic treatment of bone disease were designed by surface modifications with alendronate and monomethoxy polyethylene glycol as the targeting moiety and the hydrophilic surface, respectively. HA affinity assay suggested that the alendronatemodified PLGA nanoparticles had the potential to the targeted drug delivery system to bone. A detailed in vivo and kinetic release study would be required to optimize the structural design of the nanoparticles. This research is likely to be helpful in the design of functional nanoparticles for the site-specific drug delivery in the treatment of bone diseases. Acknowledgements This work was supported by Ministry of Commerce, Industry and Energy (MOCIE) through the project of NGNT (No. 10024135-2005-11) and Seoul Research and Business Development Program(10816). References [1] T.M. Fahmy, P.M. Fong, A. Goyal, M. Saltzman, Targeted for drug delivery, Materials Today 8 (8) (2005) 18–26. [2] G.A. Colditz, S.E. Hankinson, D.J. Hunter, W.C. Willett, J.E. Manson, M.J. Stampfer, C. Hennekens, B. Rosner, F.E. Speizer, The use of estrogens and progestins and the risk of breast cancer in postmenopausal women, N. Engl. J. Med. 332 (1995) 1589–1593. [3] L.G. Raisz, The osteoporosis revolution, Ann. Intern. Med. 126 (1997) 458–462. [4] D. Wang, S.C. Miller, P. Kopečková, J. Kopeček, Bone-targeting macromolecular therapeutics, Adv. Drug Deliv. Rev. 57 (2005) 1049–1076. [5] C.S. Cho, K.Y. Cho, I.K. Park, S.H. Kim, T. Sasagawa, M. Uchiyama, T. Akaike, Receptor-mediated delivery of all trans-retinoic acid to hepatocyte using poly(L-lactic acid) nanoparticles coated with galactose-carrying polystyrene, J. Control. Release 77 (2001) 7–15.

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