Novel Functionalized Biodegradable Polymers for Nanoparticle Drug

Biomacromolecules 2005, 6, 1885-1894

1885

Novel Functionalized Biodegradable Polymers for Nanoparticle Drug Delivery Systems Paraskevi Kallinteri,† Sean Higgins,‡ Gillian A. Hutcheon,‡ Christopher B. St. Pourc¸ ain,§ and Martin C. Garnett*,† School of Pharmacy, University of Nottingham, University Park, Nottingham, NG7 2RD United Kingdom, School of Pharmacy and Chemistry, Liverpool John Moores University, James Parsons Building, Byrom Street, Liverpool, L3 3AF United Kingdom, and School of Engineering and Applied Science, Aston University, Aston Triangle, Birmingham, B4 7ET United Kingdom

Biomacromolecules 2005.6:1885-1894. Downloaded from pubs.acs.org by HONG KONG UNIV SCIENCE TECHLGY on 08/19/18. For personal use only.

Received December 17, 2004; Revised Manuscript Received March 18, 2005

We have prepared and screened a library of novel functionalized polymers for development of nanoparticle drug delivery systems. The polymer backbone consisting of two ester-linked, nontoxic, biological monomers, glycerol and adipic acid, was prepared using a hydrolytic enzyme. The specificity of the chosen enzyme yields a linear polymer with one free pendant hydroxyl group per repeat unit, which can be further functionalized. This protocol gives control over the backbone polymer molecular weight, together with the ability to incorporate various amounts of different fatty acyl substituents. These functionalized polymers are able to self-assemble into well-defined small particles of high homogeneity with a very low toxicity. They are able to incorporate a water soluble drug, dexamethasone phosphate, with a high efficiency and drug loading which varies with the polymer specification. The above characteristics strongly suggest that these polymers could be developed into useful nanoparticulate drug delivery systems. Introduction Microparticle and nanoparticle drug delivery systems have been widely studied over the past 20 years.1 For parenteral delivery systems, it has been shown that nanosized particles and liposomes have great potential in cancer therapy due to their ability to extravasate from the leaky vasculature of tumors.2 However, only a few such delivery systems have progressed far toward clinical use.3 For liposomal systems, their stability in vivo is low.4 A number of polymeric systems have been investigated, with PLGA [poly(D,L-lactide-coglycolide)] and PLA-PEG [poly(D,L-lactic acid)-poly(ethylene glycol)] being the most widely studied because they are biodegradable, of low antigenicity, and approved for drug use. Unfortunately, the reported drug incorporation levels have been generally quite low,5-7 thus making it difficult to encapsulate sufficient drug for therapeutic efficacy. Improved drug incorporation has been seen using cyanoacrylate polymers, but these systems show some toxicity.8,9 Another disadvantage for many of the nanoparticles produced using the nanoprecipitation/solvent extraction technique used is the need for a surfactant during nanoparticle formation10-12 and sufficient removal of the surfactant is always a problem. We favor the use of polymeric nanoparticles for the advantages of stability, cost, and ease of formulation. We reasoned that drug incorporation and control of drug release, * To whom correspondence should be addressed. Phone: +44-(0)11595-15045. Fax: +44-(0)115-95-15102. E-mail: martin.garnett@ nottingham.ac.uk. † University of Nottingham. ‡ Liverpool John Moores University. § Aston University.

could be altered by the introduction of moieties into the polymers which could increase the level of interaction with the drug. This strategy would require biodegradable polymers to enable the in vivo biodegradation and subsequent removal from the body. A few functionalized polymers have previously been reported for use in drug delivery, e.g., polysaccharides,13 poly(amino acids),14 and poly(L-lysine citramide).15 However, these polymers are variously not readily degraded, have significant toxicity, or are difficult to produce. The chemical synthesis of functional polymers, particularly readily hydrolyzable polymers such as polyesters, is difficult due to the need for the synthesis of a suitable monomer, with a protected functional group, and the post-polymerization deprotection of the functional group.16 Frequently, the deprotection steps lead to the partial hydrolysis of the polymer, which is wasteful and leads to poorly characterized products.17 To overcome these problems, we can use enzymes to facilitate the polymerization procedure.18,19 An exceptional enzyme catalyzed process for the synthesis of functional polyesters has been reported.20 These polyesters have pendant hydroxyl groups that could be substituted to provide the basis to investigate our hypothesis. To illustrate this concept, the potent, water soluble, antiinflammatory steroid drug dexamethasone phosphate has been chosen. This drug has potential in cancer therapy for the relief of inflammation and swelling in brain tumors. However, when used chronically, this drug has a range of undesirable side effects, which would benefit from reduction by a suitable delivery system,21 and so is particularly appropriate. Steroids are known to interact with the acyl groups of phospholipids in membranes, so introduction of

10.1021/bm049200j CCC: $30.25 © 2005 American Chemical Society Published on Web 04/27/2005

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acyl groups into polymers seemed an ideal way to increase interaction between drug and nanoparticle core. In this paper, we therefore describe the synthesis and characterization of acyl substituted functional polymers and their utility in enhancing the incorporation of dexamethasone phosphate into a nanoparticle delivery system. Experimental Section Materials. Novozyme 435 (a lipase [9001-62-1] derived from Candida antarctica and immobilized on an acrylic macroporous resin) was donated by Novozyme Co., Denmark and stored over P2O5 at 5 °C prior to use. Divinyl adipate [4074-90-2] was purchased from Flurochem, U.K., and used as received. The acid chlorides were synthesized, by standard procedures, from the parent carboxylic acid using thionyl chloride and distilled prior to use. All other chemicals and reagents were purchased from VWR or Sigma-Aldrich and used as received, except tetrahydrofuran (THF), which was freshly distilled from sodium/benzophenone ketyl prior to use, and pyridine, which was distilled from KOH prior to use. Polymer Characterization. Gel permeation chromatography (GPC) results were obtained using a Polymer Laboratories system, employing 2 mixed bed (D) columns at 40 °C, flow rate 1 mL min-1 in THF, using an evaporative light scattering detector which was calibrated with 10 narrow polystyrene standards. MALDI-TOF data was collected on an Applied Biosystems QSTAR (Q-ToF) mass spectrometer. The samples were dissolved at a concentration of 10 mg mL-1 using a matrix of dithranol and silver trifluoracetate.20 The NMR spectra were recorded on a Bruker DPX 250 MHz spectrometer and are expressed in parts per million (δ) from internal tetramethylsilane. The relative intensities of the methyl triplet peak δ(1.0-0.8) (end of acid chains) and the methylene peaks δ(1.8-1.5) (β to carbonyls) was used to confirm the acylation reaction. Backbone Polymer Synthesis. A typical procedure for the synthesis of 12 kDa poly(glycerol-adipate) backbone was as follows. An oven dried 500 mL three-necked RB flask, equipped with center stirrer guide, and an open top condenser, to act as an outlet for the acetaldehyde produced, was charged with divinyl adipate (49.55 g, 250 mmol), glycerol (23.38 g, 250 mmol), and 20 mL anhydrous THF. This mixture was stirred for 30 min to allow reactants to warm to the water bath temperature (50 °C). The enzyme complex (1.45 g (2% monomer weight)) was then added and the mixture stirred for 24 h (200 rpm using an overhead stirrer fitted with a Teflon stirrer paddle). The enzyme complex was then removed by filtration, washing with a further 50 mL of THF. The solvents were removed by rotary evaporation (final bath temperature set at 100 °C to deactivate any free enzyme) to yield ∼45 g of the polyester as a pale yellow viscous liquid, which was used without further purification. Mw ) 12 554 Da (PDI ) 2.77). The other two backbone polymers were prepared in a similar manner, except for the time in contact with the enzyme, thus 2 kDa poly(glycerol-adipate) only 1 h enzyme contact, Mw ) 2298 Da (PDI ) 1.61) and 6 kDa poly(glycerol-adipate) only 4 h enzyme contact, Mw ) 5,777 Da (PDI ) 1.60).

Kallinteri et al.

Acylation of Backbone Polyester. The acylation of the polymer backbones was readily achieved by reaction of a percentage of the pendant hydroxyl groups, from the backbone polyester, with acid chlorides using pyridine as a catalyst and acid scavenger. The typical procedure for synthesis of 40% C18 (12 kDa) polymer was as follows. The percentage acylation was calculated using the repeat unit size of the backbone polyester (202.21 g mol-1); thus, a 50 mL three-necked RB, fitted with a condenser and dropping funnel, was charged with 12 kDa poly(glycerol-adipate) (2.11 g, 10.43 mmol) and 10 mL of THF. The mixture was warmed to reflux to dissolve the polyester, and stearoyl chloride (1.26 g, 4.16 mmol) was added. Pyridine (2 mL) was then added dropwise, producing acid fumes and a white precipitate as the reaction proceeded. The mixture was refluxed for 2 h and then poured onto 100 mL of 2 M HCl, followed by extraction three times with 50 mL of dichloromethane. The combined organics were washed twice with 100 mL of water, dried over magnesium sulfate and the solvent removed by rotary evaporation to yield 2.51 g of a white waxy solid, which was used without further purification. Nanoparticle Preparation. The nanoparticle dispersions were prepared using the interfacial deposition method.6,22 Briefly, 20 mg of polymer was dissolved in acetone (2 mL), and this solution was added dropwise to 5 mL of water under stirring to obtain surfactant free empty nanoparticles. For surfactant or drug containing particles, the water was replaced by aqueous solutions of polysorbate 20 or polysorbate 80 (1% or 4%) or of dexamethasone phosphate (1-12 mg). The dispersions were left overnight to allow the acetone to evaporate. The particles were filtered through a 1.2 µm filter (Whatman, cellulose nitrate membrane filters) to remove flocculated particles and polymers, as necessary. The excess of the drug and/or surfactant was separated, by eluting the particles through a Sepharose 4B-CL column (2.5 cm × 30 cm). The drug encapsulation efficiency was estimated indirectly by spectrophotometric measurement of the unencapsulated drug (λ ) 240 nm) eluted as a second peak off the column. Drug incorporation was expressed both as actual drug loading (% w/w) and encapsulation efficiency (%) represented by eqs 1 and 2, respectively: actual drug loading (% w/w) ) mass of drug in nanoparticles x 100/mass of nanoparticles (1) encapsulation efficiency (%) ) mass of drug in nanoparticles x 100/mass of initial drug used (2) Morphological Studies. Morphological evaluation of the nanoparticles was performed using transmission electron microscopy (TEM) (JEOL Jem 1010 Electron Microscope, Japan) following negative staining with phosphotungstic acid solution (3% w/v, adjusted to pH 4.7 with KOH). Physicochemical Characterization. The particle size was determined by dynamic light scattering, using a Malvern system 4700 instrument, with vertically polarized light supplied by an argon-ion laser (Cyonics) operated at 40 mW. All experiments were performed at a temperature of 25.0 (

Polymers for Nanoparticle Drug Delivery Systems

Figure 1. Lipase catalyzed polycondensation of glycerol and divinyl adipate to form backbone polyester, poly(glycerol-adipate), and the subsequent acylation reaction using acyl chloride at a controlled percentage.

0.10 °C at a measuring angle of 90 ° to the incident beam. The analysis mode was the CONTIN and 30 readings were performed for each sample. The ζ-potential of the nanoparticles was determined by laser doppler anemometry using a Malvern Zetasizer. Measurements were performed at 20 ( 0.10 °C, on samples appropriately diluted with the medium described each time. Five measurements of the ζ-potential for each sample were made. Cytotoxicity Studies. Cytotoxicity studies were carried out on HL-60 and HepG2 cells for 3 days using the 3-(4,5dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT) assay.23 This assay is based on the ability of live cells to convert the MTT (solution of yellow color) into blue formazan salts. The formazan salt was quantitated by the method of Gortzi et al.24 The polymer dispersions were also treated alone with the MTT solution in order to ensure that the polymers themselves did not give a positive reaction during the procedure. The cytotoxicity results were represented as % of viable cells in relation to cells not treated with the polymers (control) in each experiment. Results Backbone Polymer Synthesis and Characterization. The initial design reported for the backbone polymer was a polyester formed by the polycondensation of a functionalized diacid, divinyl adipate and glycerol with a molecular weight of 10 kDa.20 We have therefore utilized the regiospecificity of an enzyme, for primary hydroxyl groups, to produce a linear polyester with pendant hydroxyl groups at regular intervals along the polymer backbone as detailed in Figure 1. However utilizing the literature method20 of shaking monomers with enzyme, we were only able to achieve this target molecular weight when using small quantities of reactants. Simply put, as the polymerization proceeded, the viscosity of the mixture increased to a level where diffusion of reactants became difficult. Therefore, to enable us to produce large enough quantities of backbone polymer, for further acylation and nanoparticle studies, we had to develop a new polymerization procedure.

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We achieved this by making use of mechanical stirring, solvents, temperature control, removal of byproduct, and time in contact with the enzyme. After the initial screening of conditions, we have successfully prepared the polymer by mixing the molar equivalents of the two monomers, with the stabilized lipase (1% equivalent of the combined weight of the monomers), at 50 °C in THF. The formed polymer was separated from the insoluble enzyme complex by filtration and washing with solvent to ensure maximum yield, thus producing a clean polymer, which required little workup prior to further functionalization. While stirring the reaction mixture we ensured that the stirring paddle did not grind up the enzyme complex and release any of the enzyme. Similarly, as part of the polymer workup, we heated the final polymer to 100 °C, to assist with the complete removal of solvent and to hopefully denature any enzyme that might have leached out of the stabilized resin. It was necessary to ensure no lipase contamination of the final product, as in an aqueous environment, the enzyme catalyses the reverse pathway and will start to degrade our polymer. We found that essentially the same molecular weight range, by GPC, is recorded so long as we use the same temperature and percentage of enzyme. From our initial GPC results, we estimated the molecular weight in the range of 1-15 kDa (versus polystyrene), depending on time in contact with the enzyme. Thus, we have prepared polymer batches with three different molecular weights nominally termed 2, 6, and 12 kDa (1, 4, or 24 h contact with enzyme). In contrast, Kline et al.20 reported that a 4 h reaction time, at 50 °C, produces a polymer with Mw ) 1485 ( 202 Da, whereas a 24 h reaction time produced a polymer with Mw ) 10 391 ( 404 Da (again versus polystyrene). Therefore, our improved stirred system obtained a higher final molecular weight and offered greater control of the final molecular weight and polydispersity. 1H NMR studies have proved that the polymer formation begins by the first hour and that there is little monomer left after 8 h as evident by the gradual disappearance of the low field olefinic peaks (7.27, 4.87, and 4.56 ppm for the vinyl groups of the starting divinyl adipate). These polymers are soluble in organic solvents indicating that the required linear system has formed and we have no network formation. Further evidence comes from the MALDI-TOF data. Our results concur with those of Kline et al.20 in that we mainly see sets of peaks separated by the polymer repeat unit of ∼202 Da. Similarly, the NMR data implies we have a 1:1 ratio of the monomers and hence no network formation. Acylation of Backbone Polyester. As our selected drug is known to interact with lipids, we have designed our system to contain alkyl moieties, with the intention of producing a lipid like environment to entrap the drug molecules. The acylation of the polymer backbone was simply achieved by reaction of a percentage (20%, 40%, 60%, 80%, or 100%) of the pendant hydroxyl groups with various linear aliphatic acid chlorides (caprylic acid (C8) or stearic acid (C18)), which, once bound to the polymer backbone, provide a relatively hydrophobic environment for nanoparticle formation and drug incorporation. The acid chlorides were simply prepared from the parent carboxylic acid and coupled to the backbone

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Table 1. Physicochemical Properties of Nanoparticles Made by Polymers (12 kDa Polymer Length) Substituted in a Range between 20 and 80 %, with Two Acyl Chain Lengths (C8 and C18) in the Absence and in the Presence of Surfactants (Polysorbate 20 and Polysorbate 80) 1% surfactant particle component 20% C8 polysorbate 20 polysorbate 80 40% C8 polysorbate 20 polysorbate 80 60% C8 polysorbate 20 polysorbate 80 80% C8 polysorbate 20 polysorbate 80 20% C18 polysorbate 20 polysorbate 80 40% C18 polysorbate 20 polysorbate 80 60% C18 polysorbate 20 polysorbate 80 80% C18 polysorbate 20 polysorbate 80

4% surfactant

size, nm, (SD) polydispersity]

ζ-potential, mV (SD)

size, nm, (SD) [polydispersity]


189.4 (7.3) [0.042] 610.4 (23.7) [0.214] 185.2 (5.9) [0.069] 191.0 (2.0) [0.111] 194.7 (4.4) [0.108] 179.8 (2.9) [0.177] 209.9 (4.6) [0.129] 212.9 (3.7) [0.153] 173.5 (4.0) [0.166] no particles 203.9 (2.5) [0.116] 209.7 (3.2) [0.158] 171.5 (2.8) [0.178] 275.0 (17) [0.359] 217.7 (10.3) [0.319] 190.6 (5.2) [0.141] 244.5 (6.4) [0.213] 150.6 (2.1) [0.136] 237.3 (10.5) [0.211] 270.3 (6.5) [0.257] 192.8 (17.6) [0.246] 231.8 (3.6) [0.136] 275.8 (24.3) [0.300] 180.7 (8.7) [0.292]

-32.0 (0.5) -10.9 (1.3) -25.6 (3.7) -34.0 (1.4) -27.4 (1.1) -28.5 (1.5) -31.0 (1.0) -20.7 (0.3) -27.5 (4.7) no particles -28.8 (1.1) -22.5 (2.0) -30.0 (1.7) no value -26.2 (1.0) -22.5 (1.3) -18.2 (1.2) -28.2 (5.2) -24.3 (0.9) -26.9 (0.6) -23.8 (1.5) -34.8 (0.8) -25.4 (1.4) -25.7 (1.2)

189.4 (7.3) [0.042] no value 163.8 (4.2) [0.128] 191.0 (2.0) [0.111] 216.1 (11.7) [0.158] 179.1 (5.5) [0.157] 209.9 (4.6) [0.129] 194.8 (12.1) [0.220] 169.3 (2.0) [0.151] no particles 206.3 (2.3) [0.156] 222.7 (10.1) [0.217] 171.5 (2.8) [0.178] 196.7 (5.2) [0.139] 200.5 (4.1) [0.302] 190.6 (5.2) [0.141] 211.3 (4.0) [0.096] 186.0 (16.0) [0.369] 237.3 (10.5) [0.211] 259.0 (8.3) [0.260] 263.0 (20.7) [0.412] 231.8 (3.6) [0.136] 225.6 (3.5) [0.130] 199.4 (20.5) [0.258]

-32.0 (0.5) no value -22.2 (0.9) -34.0 (1.4) -26.0 (5.0) -25.7 (1.0) -31.0 (1.0) -26.3 (1.0) -24.9 (3.3) no particles -22.9 (1.4) -22.2 (0.7) -30.0 (1.7) no value -34.3 (0.7) -22.5 (1.3) -29.2 (1.3) -21.8 (1.3) -24.3 (0.9) no value -21.6 (1.2) -34.8 (0.8) -32.2 (0.6) -24.4 (2.4)

using pyridine as a catalyst and acid scavenger (Figure 1). GPC and NMR confirmed the level of acylation. Influence of Surfactant on the Formation of Nanoparticles. Table 1 presents the size, polydispersity, and ζ-potential of particles made from polymers with pendant acyl chains (C8 or C18) in a variety of percentage substitutions. The particles were produced by the interfacial-deposition method6,22 in the absence and presence of the surfactants, polysorbate 20 or polysorbate 80 at 1% or 4% (w/v) concentration in the dispersion medium. Surfactants can have a number of different influences. With many polymers, the addition of a surfactant during particle formation is necessary in order to reduce the surface tension of the polymers when the latter come into contact with the water, to promote the polymer assembly into nanoparticles. Previously25-28 1% of polysorbate 80 was used to facilitate the nanoparticle formation from a number of polymers such as PEG-hexadecylcyanoacrylate, polybutylcyanoacrylate, and PLGA, respectively. In other cases, polysorbates stabilize the colloid dispersion from aggregation and/or they can alter the particle biodistribution.29 In our studies, we therefore used 1% of surfactant as a “standard” procedure and 4% to investigate the possible effect of the surfactant on the particle size. Polysorbates 20 and 80 used in this work are polyoxyethylene sorbitan esters carrying the fatty acyl groups laurate or oleate, respectively. The presence of the surfactant did not have a consistent effect on particle size. Polysorbate 20 tended to increase the size of C18 nanoparticles. This probably happened because the hydrophobic part of polysorbate 20 consists of a relatively short acyl chain (C11H23)COO] comparable in length to the C8 chain (C7H15)COO] but

shorter than the C18 (C17H35)COO] acyl length attached to our polymers. In contrast, polysorbate 80 slightly reduced the size of the same particles, but it increased the polydispersity of the particle population for the C18 particles in a more obvious way. The decrease in particle size may be attributable to interactions of the acyl chains of the polysorbate enhancing the condensation or packing of the particles. The polydispersity increase could be due to a contribution of polysorbate micelles present during the size and the polydispersity estimation of the formulation. The surfactant addition reduced the particle ζ-potential to less negative values, as expected due to the addition of polyethylene-oxide layer of the polysorbate moving the Stern layer further from the particle surface.30 However, an increase in side chain substitution of the polymer tended to increase the particle size in the absence of the surfactant. This could be either due to the bigger space occupied in the particle core by the acyl groups or due to an increased aggregation number. However, with the present polymers, particle formation does not require surfactant addition. The surfactant addition stabilizes the particles after formation but in much lower concentrations (0.1%), when the dispersion medium is a buffer solution. All subsequent work has been carried out using either 0.1% surfactant (polysorbate 80) or no surfactant as a stabilizer. Effect of Polymer Properties on Particle Formation and Drug Loading. The effect of polymer specification on drug loading and efficiency of drug incorporation and physical characterization of nanoparticles is shown in detail in Table 2. From the data presented in Table 2, it can be seen that the particle sizes ranged between 120 and 320 nm with a very low polydispersity index with most particles falling in



Table 2. Physicochemical Characteristics and Drug Loading of Nanoparticles Made by Polymers Substituted with C8 or C18 Acyl Groups in a Range between 0 and 100 % in the Absence of Surfactanta (-) DP

(+) DP




174.3 (2.9) [0.029] 148.5 (2.76) [0.032] 186.6 (0.5) [0.06]

-14.65 (2.33) -23.45 (0.07) -19.5 (0.4)

173.15 (5.6) [0.117] 140.7 (5.52) [0.135] 122.8 (0.35) [0.103]

-27.9 (3.4) -30 (7.14) -34.2 (2.8)

40% C8 (2 kDa) 40% C8 (6 kDa) 40% C8 (12 kDa)

213.7 (0.8) [0.027] no particles 191.0 (2.0) [0.111]

-18.4 (0.7) no particles -34.0 (1.4)

164.1 (0.9) [0.075] 207.8 (0.4) [0.039] 184.4 (4.8) [0.090]

-28.8 (1.1) -29.0 (1.1) -32.0 (1.0)

80% C8 (2 kDa) 80% C8 (6 kDa) 80% C8 (12 kDa)

no particles no particles no particles

no particles no particles no particles

221.9 (1.5) [0.168] 207.2 (3.4) [0.122] 225.1 (5.5) [0.102]

-46.7 (0.5) -48.7 (0.9) -29.9 (1.0)

no particles 276.4 (3.5) [0.193]

no particles -47.5 (0.7)

254.0 (4.4) [0.221] 251.9 (4.9) [0.225]

-47.3 (1.3) -44 (1.5)

40% C18 (2 kDa) 40% C18 (6 kDa) 40% C18 (12 kDa)

no particles 232 (23.6) [0.133] 212.67 (20.6) [0.139]

no particles -32.4 (4.3) -25.9 (3.31)

244.75 (5.7) [0.217] 226.7 (1.0) [0.121] 211.0 (3.6) [0.154]

-35.25 (0.6) -46.3 -34.5 (4.3)

80% C18 (2 kDa) 80% C18 (6 kDa) 80% C18 (12 kDa)

250.1 (11.6) [0.126] 246.3 (23.3) [0.076] no particles

-32.4 (2.7) -34.7 (1.6) no particles

239.8(12.7) [0.527] 229.4 (4.4) [0.104] 256.3 (1.9) [0.135]

-35.7 (2.6) -43.3 -48.0

100% C18 (2 kDa) 100% C18 (6 kDa) 100% C18 (12 kDa)

227 (7.9) [0.213] 225.3 (0.9) [0.065] 244.4 (2.1) [0.095]

-40.7 (3.1) -39.7 (3.3)

316.1 (16.5) [0.258] 232.45 (13.4)[0.09] 224 (2.4) [0.077]

- 32.25 (3.04) -49.5

polymer type 0% (2 kDa) 0% (6 kDa) 0% (12 kDa)

100% C8 (2 kDa) 100% C8 (6 kDa) 100% C8 (12 kDa)*


a Particles were formed in the absence [(-) DP] and the presence [(+) DP] of drug, to evaluate the possible drug contribution to nanoparticle formation. Nanoparticles were made in the presence of 4 mg of dexamethasone phosphate. Asterisk (*) indicates that no results are given for the 100% C8 (12 kDa) polymer because it was not available for these studies.

Figure 2. Transmission electron micrographs of nanoparticles produced from (A) 12 kDa polymer backbone (magnification 80K) and (B) 40% C18 (12 kDa) produced in the absence of surfactant (magnification 15K). Polymer stained with phosphotungstic acid.

the size range of 160-240 nm. This indicates that, under the simple conditions of the preparative procedure, the resulting dispersions exhibit high homogeneity. The nanoparticles produced were of a good spherical shape (Figure 2) whether produced from nonacylated (Figure 2A) or acylated (Figure 2B) polymer in the absence of surfactant. Similar images were obtained from particles in the presence

of surfactant (data not shown). The negative ζ-potential of the particles arose from the free terminal carboxyl groups of the polymer (each polymer chain may terminate at either end with glycerol or adipic acid monomer). Also, the negative ζ-potential contributes to the particle stability without aggregation during storage at 4 °C, even after a month. The influence of the pendant acyl chain length (C8 or C18) and the polymer backbone molecular weight (2, 6, and 12 kDa) on the encapsulation efficiency is shown in Figure 3, parts A and B. Reproducibility determined for % drug loading with some key polymers was good (80% C18 (6 kDa), 100% C18 (6 kDa), and 40% C18 (12 kDa) had values of 7.48 [0.42], 4.16 [0.96], and 6.09 [0.52] respectively for drug loading [std dev.] with n ) 3). Consequently the data shown for drug loading studies with polymers of different specifications is the mean of two determinations. It seems that for the polymer of 2 kDa, the C8 pendant polymer exhibited a higher drug loading than the C18 polymer. For the C8 and C18 (2 kDa) polymers, optimum drug incorporation occurred when the polymer was fully substituted (Figure 3A,B). The dramatic influence of the acyl chain length on the encapsulation efficiency was very obvious with the polymer of 6 kDa. When the C8 acyl chain was attached to the polymer of 6 kDa, the encapsulation efficiency remained stable and low up to a substitution of 80% but reached the highest value when all of the hydroxyl groups of the polymer backbone had been substituted (Figure 3A). In contrast, for the C18 polymer of 6 kDa, the drug loading started to increase as the % of acylation increased

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Figure 4. Encapsulation efficiency and drug loading of 40% C18 (12 kDa) nanoparticles produced by the interfacial deposition method in the absence of surfactant using a range of initial drug concentrations.

Figure 3. Effect of % acylation and polymer backbone molecular weight on drug encapsulation efficiency for (A) C8 pendant acyl chain or (B) C18 pendant acyl chain.

and reached its highest value at 80% substitution (Figure 3B). Drug incorporation for 80% C18 (6 kDa) value was the same as that obtained using the 100% C8 (6 kDa) polymer. The drug incorporation obtained with the C8 or C18 (12 kDa) polymer was higher than that obtained with the respective 2 and 6 kDa polymers, especially at the lower percent acyl substitution of each polymer. The C8 (12 kDa) polymer reached maximum drug loading at a higher percent substitution than the C18 one. The latter reached a plateau from 40% to 80% substituted polymer backbone, whereas the former reached the same value at 80% substitution. This phenomenon suggests that the drug molecules interact more strongly with the C18 acyl chain. It was noteworthy that, for the 6 and 12 kDa polymers, the amount of incorporated drug decreased at 100% substitution. This could suggest that in the 100% substituted particles there was not enough aqueous space for the drug to be accommodated into the particles. Another possibility is that the polymer backbone does not offer any hydroxyl groups to interact with the phosphate groups of the drug so hydroxyl groups facilitate the partitioning of the drug into the amphipathic matrix of the particle. Additionally, the impact of molecular weight on the drug incorporation was very clear.

Generally, the 12 kDa polymer exhibited the highest encapsulation efficiency whatever the % substitution and acyl chain at this polymer molecular length. The only exception is the 100% C18 (6 kDa), which showed higher encapsulation efficiency than the 100% C18 (12 kDa) polymer. For most polymer substitutions, the C18 version gave the highest efficiency of drug incorporation. However, for 100% substitution, the C8 substituted polymers of 2 and 6 kDa were almost as good as the C18 polymers. The effect of drug amount on drug loading and efficiency is shown in Figure 4. This shows that particle saturation occurred after the addition of 8 mg of dexamethasone phosphate. (The values shown in Figure 4 were the mean of two repetitions (with the exception of the 6 mg sample which was from one determination.) Reproducibility was good for low initial drug concentrations, but poor when high drug concentrations were used. Drug Influence on Physicochemical Properties of Particles. The impact of drug encapsulation, in the absence of surfactant, on the particle physicochemical characteristics made by polymers carrying either C8 or C18 acyl chain over the whole range of substitution is presented in Table 2. The drug loaded C18 particles were slightly larger in comparison to the C8 particles of the same substitution and the ζ-potential was also shifted to more negative values (Table 2). The negative ζ-potential probably arose from the free terminal carboxyl groups of the polymers. Therefore, the % of the acylation or the type of the acyl group attached on the polymers does not influence the surface charge of the particles. The estimated ζ-potential of the particles made in the presence of the drug was usually more negative than for the preparations in its absence. This is probably due to the negatively charged phosphate group of the drug, some of which must be located on the particle surface. However, the ζ-potential of the particles made by lower molecular weight polymers is more negative because many more polymer molecules need to be assembled to form nanoparticles of the same size so a larger number of exposed carboxyl groups would be expected which will tend to orientate themselves on the particle surface in the hydrophilic aqueous environment. It is apparent that the presence of the dexamethasone phosphate tends to decrease slightly the size of the particles made from C8 polymer. Also for some polymers, nanoparticle



Figure 5. 125I-Deoxy-uridine release profile from 40% C18 (12 kDa) particles in water and 10% serum using diffusion cells.

assembly only takes place in the presence of added drug. These observations strongly suggest that the drug assisted the condensation of the polymers in the same or even a better way than polysorbate. This indicates that the dexamethasone phosphate, which is water soluble because of the phosphate moiety but has a hydrophobic steroid ring, interacts very strongly with the polymers. To show the quite remarkable retention of the drug by our polymeric carriers, we report the drug release profile of a radiolabeled water soluble molecule, deoxy-uridine (Figure 5) from 40% C18 (12 kDa) particles in buffer and 10% serum. The iodinated compound 125I-Deoxy-uridine was incorporated into particles in anticipation of biodistribution and pharmacokinetics studies. The drug was released slowly without the burst release phenomenon, which is always the case for the PLGA and PLA-PEG particles.31-33 Thus, 73% of the drug was released after 30 days in water, whereas the same percentage of release occurred after 8 days in 10% serum. Cytotoxicity Studies. The acceptability of any drug delivery system depends not only on the ability to incorporate and release the drug but also on the polymer component being nontoxic. Consequently, we have carried out basic cytotoxicity studies on several key polymer formulations using the MTT assay23 on nonadherent HL-60 (leukemia; Figure 6) and adherent HepG2 (hepatoma; Figure 7) cell lines. HL60 cells represent circulating blood cells, whereas HepG2 cells represent liver cells, the two cell types, which will be mostly, exposed to parenteral drug delivery formulations. As the polymer component itself is insoluble in aqueous systems, these tests have been carried out using nanoparticles made from these various polymers without added drug. The cytotoxicity of the nanoparticles is expressed as the % cell viability in comparison to control cells (incubated in the absence of particles) against particle concentration in µg mL-1. The range of polymer concentrations studied was based on potential therapeutic dosing. The usual therapeutic dose of dexamethasone phosphate is 4 mg. If the highest drug amount encapsulated in 20 mg of particles is 1.5 mg, then the total amount of polymer needed for formulation is 54 mg. If the amount of the blood volume in our body is 5 L, then the polymer concentration in the blood is 10 µg mL-1. Testing at a top dose of 1 mg mL-1, 100 times this amount, seems appropriate. The polymer backbone at chain lengths from 2 to 12 kDa without attached acyl groups showed little cytotoxicity

Figure 6. Cell (HL-60) viability assessment in the presence of particles made using (A) the unsubstituted polymer backbone and (B) 40% substituted polymers.

against either HL60 or HepG2 cells (Figures 6A and 7A), respectively, with only a slight apparent reduction of cell viability on HL60 cells at the top dose of 1 mg mL-1 particles made from 12 kDa polymer. Polymers at 40% acyl substitution were also investigated using either the C8 acyl groups on 2 kDa backbone or C18 acyl groups on 6 or 12 kDa backbones. None of these polymers showed any significant inhibition of cell viability (Figures 6B and 7B). We have also investigated the toxicity of nanoparticles coated with polysorbate 80 (0.1%) as a stabilizer (Figure 8). In this test, the three formulations, 100% C8 (6 kDa), 80% C18 (6 kDa), and 40% C18 (12 kDa), all showed a decrease in cell viability to approximately 50% at 1 mg mL-1 of polymer. However, this toxicity was entirely attributable to the polysorbate 80 coating. Discussion In an attempt to make more versatile drug carriers, a range of novel polymers were synthesized. The main repeating unit of these polymers consisted of glycerol and adipic acid in a 1:1 molar ratio. Pendant fatty acyl chains, either C8 or C18, were attached to the hydroxyl group offered by the glycerol monomer in a variety of substitutions ranging between 0 and 100%. The polymer backbone synthesis was achieved by using the enzyme lipase from Candida antarctica which catalyzed the ester bond formation.19,20 Varying the length of the reaction time gave us the ability to produce polymer

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Figure 7. Cytotoxicity of particles according to (A) molecular weight of the polymer and (B) the polymers substituted with the 40% C8 or C18 acyl groups on the HepG2 cells.

Figure 8. Effect of polysorbate 80 coated nanoparticles on the cytotoxicity to HepG2 cells. The cell viability was assessed using MTT assay on the nanoparticles made by the polymers with the highest ability for drug encapsulation, which were coated with the polysorbate 80 in order to achieve a more stable dispersion.

backbones of different molecular weight (2, 6, and 12 kDa, respectively). This work was based on the idea of providing polymers, which will make particles of different physicochemical characteristics with the ability to encapsulate a wide range of drugs and desirable characteristics for the release of the entrapped drugs. The aim of the work presented in this paper was to screen a selection of polymers with different polymer

Kallinteri et al.

backbone lengths, % of acylation, and the acyl chain length pendant on the hydroxyl groups for the ability to form particles and incorporate drugs. PLGA and PLA-PEG are the most widely used polymers for nanoparticle formation because they are biodegradable with low antigenicity. These polymers have a significant disadvantage due to poor drug loading as shown by Govender et al.6,7 Another disadvantage for many of the polymers used to produce nanoparticles using the nanoprecipitation/solvent extraction technique is the unavoidable presence of surfactant10-12 and sufficient removal of the surfactant is always a problem. Using poly(glycerol-adipic acid) based polymers, the resulting nanoparticles were of small size, ranging between 120 and 320 nm using the interfacial/deposition method. Surprisingly, particles could be formed in the absence of surfactant. Under these conditions, the particles produced were well-defined spheres as shown in the TEM images (Figure 2). Also, the polydispersity index of the particle population was extremely low, which means a homogeneous dispersion for the size. To improve the particle size of our system, the additions of polysorbate 20 and polysorbate 80 were investigated during particle formation. The result was that only polysorbate 80 decreased the size, whereas the polydispersity index was increased. This particular surfactant may also be useful, because it has been shown previously to enhance nanoparticle accumulation in the brain.25,34 Our data for the physicochemical characteristics and the influence of the surfactant on particles agree with the results of Cammas et al.,35 who used the polymers of malic acid, poly(β-malic acid neohexyl ester), with molecular weight 16 kDa, which is similar to our 12 kDa polymer. The particle formation mechanism depends on the strength of the hydrophobic forces drawing the fatty acyl chains together into the particle core and the molecular weight of the polymer backbone, so the flexibility increases as the molecular weight increases and the substitution decreases. We propose that the polymer assembly into nanoparticles in the absence of surfactant could be explained by the presence of the hydroxyl groups of the glycerol moiety increasing the polarity of the polymeric molecule. On this basis we can explain the inability of the polymers with high % substitution to form into particles. The 80% (2, 6, and 12 kDa) and 100% C8 (2 kDa) as well as the 80% C18 (12 kDa) polymer failed to form particles but instead formed an aggregate adhering to the stirred vessel. We believe that this was due to the reduction in hydroxyl groups resulting in a reduced stabilization during particle formation leading to aggregation. On the other hand, the addition of dexamethasone phosphate increased the polarity due to the phosphate group while simultaneously being involved in strong hydrophobic interactions to the acyl chains pendant on the polymer backbone enabling particle formation in the presence of the drug. However, we do see particle formation with 100% C18 (2-12 kDa). Particle formation involves many factors, and in this case, we believe that the increased lipophilicity of the C18 side chain leads to good particle formation rather than the formation of loose aggregates.


The highly negative ζ-potential of the particles is likely to be a problem if we wish to administer them in vivo. The ζ-potential presented here is of particles made in water. When those particles are coated with 0.5% polysorbate 80 dispersed in phosphate buffered saline (pH 7.4), the ζ-potential value drops to an average of (-) 5.65 mV which may be sufficiently close to neutral to allow the particles to escape uptake by the reticuloendothelial system. Furthermore, the dispersion stability of the noncoated nanoparticles produced is attributed to the electrostatic repulsion due to the ζ-potential. The stability of the coated particles is due to both the electrostatic repulsion and the steric hindrance from the surfactant layer around the particle surface. As shown in Table 1, the particle size tended to increase as the % substitution increased. Higher polymer substitution means an increased number of acyl groups inside the nanoparticles. This could result in increased particle size either through a more bulky interior or, alternatively, by an increased aggregation number. A similar effect has been seen with particles made by the water soluble polymer, alkylated poly(L-lysine citramide).11 Having investigated the polymer behavior under defined preparative conditions, we decided to use dexamethasone phosphate as a drug model for various reasons. First, dexamethasone is an antiinflammatory agent for the treatment of oedema caused by primary or secondary brain tumors given to the patients in its water soluble form, dexamethasone phosphate. However, there are many side effects, including increased glucose levels, peripheral oedema, psychiatric disorders, and Cushing’s syndrome,21 so effective delivery of this agent would be advantageous. Second, it is an appropriate model due to the steroid ring, which is lipophilic. We have studied the influence of the polymer structure on the drug loading into the nanoparticles. It seems that dexamethasone phosphate interacts well with the substituted polymers and especially with the polymers of long acyl chain. Additionally, dexamethasone phosphate facilitates the assembly of polymers that are unable to form particles. In contrast, the polymer backbone encapsulates only a minimal amount of drug, and a hydrophilic phosphate analogue (adenosine monophosphate) also failed to give good drug incorporation (data not shown). These observations show that the acyl groups on the polymer are important in enhancing drug entrapment and hence strongly suggested an interaction between the drug and the substituted polymers. However, as concluded from Figure 3A,B and Table 2, the drug loading is not only a matter of drug-polymer interaction but it depends on the mechanism of polymer assembly into nanoparticles. More specifically, the free space in the particle core for the drug accommodation is important, as shown from the drug loading data in the 40% C18 (12 kDa) polymer. A similar conclusion can be drawn from the data shown for particles made by poly(L-lysine citramide) hydrophobized with heptyl and lauryl side chains in different proportions.11 A key finding for our polymers was the improvement in drug loading obtained. For example, Eroglu et al.36 achieved an encapsulation efficiency of the dexamethasone phosphate of 12.7% into PLGA:L-PLA microspheres, whereas our 40%


C18 (12 kDa), 80% C18 (6 kDa), and 80% C18 (12 kDa) polymers incorporated the same drug in an efficiency of 39%, 37.425%, and 37.3%, respectively (Table 2). The slow drug release suggested by iodo-deoxyuridine would also be particularly useful for drug delivery. Finally, one more advantage of these novel polymers is their low cytotoxicity demonstrated with HepG2 and HL-60 cells (Figures 6, 7). Our data complement those reported by Wang et al.,37 who made a cross-linked polymer scaffold very similar to our polymer backbone, without substitution, for tissue engineering applications. If polysorbate 80 is used to cover the particle surface, the toxicity increases, but the surfactant coated particles are still less toxic than the surfactant itself (Figure 8). Also, according to Gelperina et al.,9 the presence of polysorbate 80 did not contribute any changes of toxicity in any in vivo toxicological studies of polysorbate 80 coated poly(butylcyanoacrylate) nanoparticles in rats. Conclusions In summary, we have synthesized a range of novel substituted functionalized polymers using an enzymecatalyzed synthesis. The polymers consist of nontoxic biological monomers and can be made in good yield and suitable quantities. This procedure gives both the flexibility of synthesizing a variety of polymer backbone molecular weights together with the incorporation of various amounts of different acyl substituents through subsequent modification. These polymers are able to self-assemble into welldefined particles of relatively small size and high homogeneity with an ability to entrap dexamethasone phosphate with a high efficiency. Preliminary experiments show a slow and prolonged release profile of a radio-iodinated compound even in the presence of serum. These characteristics strongly suggest that these polymers could be developed into useful nanoparticulate drug delivery systems. Acknowledgment. This work was supported by BBSRC, Grant No. GR42/13897. We thank Novozyme, Denmark, for their kind gift of the immobilized enzyme and Mr. Trevor Gray and Mr. Phil Hinson for their valuable help with the TEM images. References and Notes (1) Allemann, E.; Gurny, R.; Doelker, E. Eur. J. Pharm. Biopharm. 1993, 39, 173-191. (2) Nagayasu, A.; Uchiyama, K.; Kiwada, H. AdV. Drug DeliVery ReV. 1999, 40, 75-87. (3) Moses, M. A.; Brem, H.; Langer, R. Cancer Cell 2003, 4, 337341. (4) Barratt, G. Cell. Mol. Life Sci. 2003, 60, 21-37. (5) Leo, E.; Brina, B.; Forni, F.; Vandelli, M. A. Int. J. Pharm. 2004, 278, 133-141. (6) Govender, T.; Stolnik, S.; Garnett, M. C.; Illum, L.; Davis, S. S. J. Controlled Release 1999, 57, 171-185. (7) Govender, T.; Riley, T.; Ehtezazi, T.; Garnett, M. C.; Stolnik, S.; Illum, L.; Davis, S. S. Int. J. Pharm. 2000, 199, 95-110. (8) Vauthier, C.; Dubernet, C.; Fattal, E.; Pinto-Alphandary, H.; Couvreur, P. AdV. Drug DeliVery ReV. 2003, 55, 519-548. (9) Gelperina, S. E.; Khalansky, A. S.; Skidan, I. N.; Smirnova, Z. S.; Bobruskin, A. I.; Severin, S. E.; Turowski, B.; Zanella, F. E.; Kreuter, J. Toxicol. Lett. 2002, 126, 131-141.

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Kallinteri et al. (25) Ramge, P.; Unger, R. E.; Oltrogge, B. J.; Zenker, D.; Begley, D.; Kreuter, J.; von Briesenet, H. Eur. J. Neurosci. 2000, 12, 19311940. (26) Brigger, I.; Morizet, J.; Aubert, G.; Chacun, H.; Terrier-Lacombe, M. J.; Couvreur, P.; Vassal, G. J. Pharmacol. Exp. Therapeutics 2002, 303, 928-936. (27) Redhead, H. M.; Davis, S. S.; Illum, L. J. Controlled Release 2001, 70, 353-363. (28) Dunn, S. E.; Coombes, A. G. A.; Garnett, M. C.; Davis, S. S.; Davies, M. C.; Illum, L. J. Controlled Release 1997, 44, 65-76. (29) Tro¨ster, S. D.; Mu¨ller, U.; Kreuter, J. Int. J. Pharm. 1990, 61, 85100. (30) Attwood, D.; Florence, A. T. Surfactant Systems: their chemistry, pharmacy and biology; Chapman and Hall: London, 1983; pp 493, 573. (31) Soppimath, K. S.; Aminabhavi, T. M.; Kulkavni, A. R.; Rudzinski, W. E. J. Controlled Release 2001, 70, 1-20. (32) Vandervoort, J.; Yoncheva, K.; Ludwig, A. Chem. Pharm. Bull. 2004, 52, 1273-1279. (33) Ricci, M.; Blasi, P.; Giovagnoli, S.; Perioli, L.; Vescovi, C.; Rossi, C. Int. J. Pharm. 2004, 275, 61-72. (34) Kreuter, J. AdV. Drug DeliVery ReV. 2001, 47, 65-81. (35) Cammas, S.; Bear M. M.; Moine, L.; Escalup, R.; Ponchel, G.; Kataoka, K.; Guerin, P. Int. J. Biol. Macromol. 1992, 25, 273-282. (36) Eroglu, H.; Kas, S. H.; Oner, L.; Turkoglu, O. F.; Akalan, N.; Sargon, M. F.; Ozer, N. J. Microencapsul. 2001, 18, 603-612. (37) Wang, Y.; Ameer, G. A.; Sheppard, B. J.; Langer, R. Nat. Biotechnol. 2002, 20, 602-606.

BM049200J

Novel Functionalized Biodegradable Polymers for Nanoparticle Drug

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