Journal of Controlled Release 128 (2008) 224–232

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Journal of Controlled Release

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Inhalable large porous microspheres of low molecular weight heparin: In vitro and invivo evaluation

Amit Rawat a, Quamrul H. Majumder b, Fakhrul Ahsan a,⁎a Department of Pharmaceutical Sciences, School of Pharmacy, Texas Tech University, Health Sciences Center, 1300 Coulter Drive, Amarillo, TX-79106, United Statesb Barr Pharmaceutical Inc. 223 Quaker Road, Pomona, NY 10970, United States

⁎ Corresponding author. Tel.: +1 806 356 4015x335; fE-mail address: fakhrul.ahsan@ttuhsc.edu (F. Ahsan)

0168-3659/$ – see front matter © 2008 Elsevier B.V. Aldoi:10.1016/j.jconrel.2008.03.013

A B S T R A C T

A R T I C L E I N F O

Article history:

This study tests the feasibili Received 28 December 2007Accepted 13 March 2008Available online 21 March 2008

Keywords:Large porous particles

MicrospheresLow molecular weight heparinPulmonary delivery

ty of large porous particles as long-acting carriers for pulmonary delivery of lowmolecular weight heparin (LMWH). Microspheres were prepared with a biodegradable polymer, poly(lactic-co-glycolic acid) (PLGA), by a double-emulsion–solvent-evaporation technique. The drug entrapmentefficiencies of the microspheres were increased by modifying them with three different additi-ves\polyethyleneimine (PEI), Span 60 and stearylamine. The resulting microspheres were evaluated formorphology, size, zeta potential, density, in vitro drug-release properties, cytotoxicity, and for pulmonaryabsorption in vivo. Scanning electron microscopic examination suggests that the porosity of the particlesincreased with the increase in aqueous volume fraction. The amount of aqueous volume fraction and the typeof core-modifying agent added to the aqueous interior had varying degrees of effect on the size, density andaerodynamic diameter of the particles. When PEI was incorporated in the internal aqueous phase, theentrapment efficiency was increased from 16.22±1.32% to 54.82±2.79%. The amount of drug released in theinitial burst phase and the release-rate constant for the core-modified microspheres were greater than thosefor the plain microspheres. After pulmonary administration, the half-life of the drug from the PEI- andstearylamine-modified microspheres was increased by 5- to 6-fold compared to the drug entrapped in plainmicrospheres. The viability of Calu-3 cells was not adversely affected when incubated with the microspheres.Overall, the data presented here suggest that the newly developed porous microspheres of LMWH have thepotential to be used in a form deliverable by dry-powder inhaler as an alternative to multiple parenteraladministrations of LMWH.

© 2008 Elsevier B.V. All rights reserved.

1. Introduction

The pulmonary route has recently emerged as a viable alternativeto the needle-based route of administration for an expanding array ofbiotechnology-derived drugs with superior therapeutic activity.However, the vast majority of currently available pulmonary drugdelivery systems, including recently approved inhaled insulin—Exubera, which was later withdrawn from the market—are designedas immediate-release formulations to produce local and systemiceffects for a short period of time [1]. Little has been done in thedevelopment of viable, inhaled formulations of therapeutic agents,especially biopharmaceuticals that can be administered via the lungsto release drugs for a prolonged period. In fact, inhaled long-actingformulations remain elusive because of the lack of efficient deliverydevices and optimal drug carriers. The factors that have been majorbarriers to the development of long-acting pulmonary formulationsare (i) suboptimal size and shape of the drug substance or the

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encapsulating carriers within the respirable fraction, (ii) shortresidence time of the inhaled drug particles or formulations in therespiratory tract, and (iii) poor loading of drugs into the particulatecarriers frequently used to prepare inhaled formulations [2–4].

The recent advances in the dry-powder inhalation (DPI) technol-ogy have addressed some of the limitations associated with inhaledformulations, including unwanted loss of drug due to oropharyngealdeposition [5,6]. However, the limitations associated with poordeposition of particles larger than 5 μm have yet to be overcome. Ina seminal paper, Edwards et al. first proposed that particles with massdensities b0.4 g/cm3 and geometric diameter N5 μm could be used toease respirability as well as to enhance residence time in the lungs [3].Indeed, the report of Edwards et al. spurred significant growth instudies on PLGA-based large porous particles for delivery of bio-pharmaceuticals. However, suboptimal pore size of the microspheresand poor encapsulation efficiency of many drugs remain major impe-diments to the widespread use of large porous PLGA microspheres ascarriers for inhaled long-acting formulations [7]. In order to achievesustained release properties and increase drug load, attempts havebeen made to modify particulate carriers by using polymer blends [8],

Table 1Composition and entrapment efficiency of the microspheric formulations

Formulations IAP:OP:EAP(v/v/v)

IAP composition(w/v)

Entrapment efficiency of LMWH(%)

PM-1 0.25:5:25 – 17.87±1.79PM-2 0.5:5:25 – 16.22±1.32PM-3 1:5:25 – 9.39±1.21PM-SP-1 0.5:5:25 1.25% Span 60 36.7±2.08PM-SP-2 0.5:5:25 2.5% Span 60 47.34±1.98PM-SA-1 0.5:5:25 1.25% SA 40.2±2.77PM-SA-2 0.5:5:25 2.5% SA 54.44±2.02PM-PEI-1 0.5:5:25 1.25% PEI 43.5±1.32PM-PEI-2 0.5:5:25 2.5% PEI 54.82±2.79

SP: Span 60; SA: Stearylamine; PEI: Polyethyleneimine.

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porosity-altering or surface-active agents [9], and by means of thecyclodextrin-mediated double-entrapment technique [10–13].

Over the past few years, the large porous particle technology hasbeen used for a number of biopharmaceuticals and conventional the-rapeutic agents, including insulin, testosterone, estradiol, deslorelin,tobramycin, ciprofloxacin and para-aminosalicylic acid [3,13–17].However, currently it is not known if this technology can be usedfor pulmonary delivery of hydrophilic macromolecules carrying anegative surface charge. Lowmolecular weight heparins (LMWHs) arenegatively charged oligosaccharides used in the treatment of deepvein thrombosis and pulmonary embolism [18]. Because of their re-latively highmolecular weight, negative surface charge and short half-lives, LMWHs are administered by subcutaneous injection multipletimes a day. Recent studies in our laboratory showed that pulmonaryadministration of LMWH is feasible [19,20]. But the proposed formu-lation of the drug with absorption enhancers and dendrimers failed toproduce a prolonged release of the drug. In this study, we propose thatlong-acting pulmonary formulations of LMWH can address the limita-tions imposed by the short duration of action of the drug.

Very recently, PLGA-based microspheres and nanoparticles havebeen used for oral and nasal delivery of unfractionated heparin andLMWH, although not for pulmonary delivery of LMWH. Moreover, thedrug load and release profiles of the proposed formulations were notthat encouraging [21–24]. We believe that the poor entrapment effi-ciency of the formulations was because of the excessive hydrophilicityof the drug, and that the poor release profiles were the result of theuse of nonporous particulate carriers. Thus, it is reasonable to assumethat the entrapment efficiency of the carriers can be increased bycomplexing the drug with cationic polymers or surfactants, and that arefined controlled release can be achieved by usingmicrospheres withoptimal porosity. This study, therefore, tests the hypothesis that core-modified PLGA-based large porous microspheres can enhance theentrapment efficiency of LMWH and facilitate release of the drug foran extended period after pulmonary administration.

2. Materials and methods

2.1. Materials

Poly(DL-lactide-co-glycolide) (50:50) PLGA (inherent viscosity0.15–0.25 dl/g; weight average molecular weight=10.6 kDa) was pur-chased from Boehringer Ingelheim (Lactel Absorbable Polymers,Pelham, AL). LMWH (average molecular weight and anti-factor Xaactivity, 4493 Da and 61 U/mg, respectively) was obtained from Celsuslaboratories (Cincinnati, OH). Poly vinyl alcohol (PVA), Span 60 (SP),polyethyleneimine (PEI) and stearylamine (SA) were purchased fromSigma (Sigma-Aldrich Inc., St. Louis, MO).

2.2. Preparation of LMWH-loaded porous microspheres

The LMWH-loaded PLGA microspheres were prepared by thewater-in-oil-in-water (w/o/w) emulsion and evaporation method [3].Briefly, an aqueous solution of LMWH (internal aqueous phase, IAP)was first emulsified in 5.0 ml of dichloromethane (organic phase, OP)containing (0.25 g) PLGA polymer by homogenization (Ultra-TurrexT25 basic, IKA,Wilmington, DE) at 15,000 rpm for 3min. The volume ofthe IAP was varied from 0.25 to 1ml to obtain microspheres of varyingporosity (Table 1). To prepare core-modified microspheres, three dif-ferent core-modifying agents\Span 60, stearylamine and polyethyle-neimine\were added to the IAP. As both Span 60 and stearylamine arehydrophobic in nature, they were dispersed in water by heating at85 °C for 15min. The concentrations of the core-modifying agents usedwere 1.25 and 2.5%. The resultingwater-in-oil (w/o) emulsionwas thenpoured into 25ml of 1.0% w/v PVA aqueous solution (external aqueousphase, EAP) and emulsified by homogenization at 8000 rpm for 5 min.The w/o/w emulsion thus obtained, the secondary emulsion, was

stirred overnight at room temperature for evaporation of dichlor-omethane. The polymeric particles were then washed thrice andlyophilized to get free-flowing powder. A blank microsphere formula-tion without LMWH was also prepared according to the sameprocedure. Each batch was prepared in triplicate.

2.3. Particle characterization

Microspheres were characterized for their morphology, size, zetapotential, tapped density, and aerodynamic diameter. The morphol-ogy of the formulations was studied under a scanning electron mic-roscope. The samples for SEMwere prepared on a conductive, double-sided adhesive tape and then sputter-coated with gold under argon(Emitech K550X, Kent, UK). The microphotographs were taken using aHitachi S-3400N (Freehold, NJ) scanning electron microscope. Themean volume-based diameter and size distribution of the particleswere determined by aMicrotrac® S3500 (North Largo, FL) particle sizeanalyzer after dispersing the freeze-dried microspheres in a 0.2% w/vaqueous solution of PVA. The polydispersity indices of all formulationswere calculated as the ratio of volume-averaged mean particle size tonumber-averaged mean particle size [25]. The zeta potential wasmeasured by Zeta potential analyzer V. 3.40 (Brookhaven InstrumentsCorp. Holtsville, NY) after dispersing the formulations in PBS buffer.The density of the particles was estimated from the tapped density asdescribed previously [4,12]. An aliquot of 100 mg microspheres wastransferred to a 10 (±0.05) ml graduated cylinder and the initialvolume was recorded. Tapped density of the particles (ρ) was cal-culated as the ratio between sample weight (g) and the volume (ml)occupied after 200 tappings. The aerodynamic diameter of the dry-powder formulations was determined in an eight-stage Andersencascade impactor (Westech Instruments Inc., Marietta, Georgia). Theformulations were fired 5–6 times into the cascade impactor at a flowrate of 28.3 L/min. The deposited formulation at each stage wascollected by rinsing the impactor and the samples were analyzed forLMWH content. The studies were performed in triplicate. In additionto actual aerodynamic diameter, the theoretical mass mean aero-dynamic diameter (MMAD) of the particles was also calculated usingthe following equation [26]:

MMADt ¼ d q=q0Xð Þ1=2

Where d is the geometric mean diameter obtained from particle sizeanalysis, ρ is the tapped density, ρ0 is the reference density of 1 g/cm3

and X is the shape factor, which is 1 for a sphere.

2.4. Encapsulation efficiency

The amount of heparin entrapped within microspheres was deter-mined using an azure A colorimetric assay by measuring the amountof unentrapped drug in the external aqueous solution recovered aftercentrifugation and washing of the microspheres [27]. Typically,

226 A. Rawat et al. / Journal of Controlled Release 128 (2008) 224–232

aliquots (50 µl) of aqueous samples were reacted with 150 µl of theazure A solution (0.01 mg/ml) and assayed in triplicate at 595 nm. Thedrug entrapment efficiency was expressed as the percentage of calcu-lated heparin load in the microspheres with the actual amount ofheparin added during the preparation.

2.5. In vitro release experiments

Freeze-dried microspheres (50 mg) were suspended in 10 ml ofPBS buffer (pH 7.4) in a flask containing 0.1% Tween 80 and incubatedin a water bath at 37±1 °C under gentle magnetic stirring (200 rpm).At various time intervals, 200 µl samples were withdrawn and centri-fuged for 15 min at 16,000 ×g. The supernatant was removed andassayed for heparin content according to a previously described colo-rimetric method [27].

2.6. In vivo studies

Male Sprague–Dawley rats (Charles River Laboratories, Charlotte,NC)weighing between 250 and 350 g were used for the in vivo absorptionexperiments (n=5–6). Absorption was monitored by measuring plasmaanti-factor Xa activity (4). Prior to the experiment, the animals wereanesthetized by an intramuscular injection of an anesthetic cocktailcontaining xylazine (20 mg/kg) and ketamine (100 mg/kg). Anesthesiawas maintained with additional intramuscular injections of anestheticmixture as needed throughout the experiments. Freeze-dried micro-spheres containing LMWHwere administered intratracheally (50 U/kg)by using a specially designed dry-powder insufflator for aerosol inha-lation in small animals (PennCentury, Philadelphia). The total amount ofpolymeric formulation administered was about 5 mg per animal carry-ing a 50 U/kg dose of LMWH. For subcutaneous administration, 50 U/kgLMWH in normal saline was administered as a single 100 μl injectionunder the back skin. Blood sampleswere collected from the tip of the tailat 0, 0.5, 1, 2, 4, 6, 8, 12 and 24 h in citrated microcentrifuge tubes andplaced on ice. Subsequently, the plasmawas separated by centrifugation(1600 ×g for 5min) and stored at −20 °C until further analysis. All animalstudies were conducted in accordance with the NIH Guide for the CareandUseof Laboratory Animals under an approved protocol (AM-02004).

2.7. Anti-factor Xa activity assay

Anti-factor Xa activity present in blood samples was determinedby colorimetric assay using a Chromogenix Coatest Heparin Kit®(Diapharma Group Inc., West Chester, OH). The assay was performedaccording to the protocol supplied by the manufacturer. Standardcurves for determination of anti-factor Xa activity in plasma or salinewere prepared by diluting the drug with pooled plasma or saline.Plasma obtained from untreated rats was used as a negative control toaccount for the effect of rat endogenous anti-factor Xa that otherwisecould give a false positive increase in anti-factor Xa activity.

2.8. Cytotoxicity studies

For the cytotoxicity studies, Calu-3 cells were seeded at a density of50,000 cells/well in flat-bottom, 96-well micro-titer tissue cultureplates. Immediately prior to the start of the experiment, the mediumwas removed from the wells and the cells were washed with normalsaline. Subsequently, the cells were incubated with 20 μl of formu-lations for 4 h, separately. The test samples contained 0.25 mg/ml,0.025 mg/ml or 0.0025 mg/ml of plain core-modifying surfactants ormicrospheres containing equivalent amounts of core-modifyingagents. Drug-loaded or blank microspheres with no core-modifyingagents were also tested for cytotoxicity. Sodium dodecyl sulfate (SDS)was used at 0.1% as a positive cytotoxic control. Cell viability wasmeasured by the MTT assay as previously described [28]. Briefly, after4 h, MTT (5 mg/ml) solutionwas added to each well and the cells were

incubated at 37°C for 4 h. Next, the solution in each well was removedand acidified isopropyl alcohol (100 μl of 1% v/v, concentrated hydro-chloric acid in isopropyl alcohol) was added. Finally, the plates wereincubated at 37 °C for 1 h and absorbance was measured on a micro-titer-plate reader (TECAN U.S. Inc, Research Triangle Park, NC) at570 nm. Each assay was performed on eight samples and cell viabilitywas expressed as the percentage of MTT released by cells exposed tocore-modifying agents and microspheric formulation or SDS com-pared to cells incubated with saline alone.

2.9. Data analysis

Standard noncompartmental analysis (Kinetica®, Version 4.0,Innaphase Corp., Philadelphia, PA) was performed to calculate the areaunder the plasma concentration versus time curve (AUC0→24 h), relativeand absolute bioavailability. Pharmacokinetic parameters of differentformulations were compared by ANOVA. When the differences in themeans were significant, post-hoc pair wise comparisons were con-ducted using Newman–Keuls multiple comparison (GraphPad Prism,version 3.03, GraphPad Software, SanDiego, CA). Differences in p-valuesless than 0.05 were considered statistically significant.

3. Results and discussion

3.1. Particle characterization

In this set of experiments we have used scanning electronmicroscopy to investigate how particle morphology was affected byvarying the fraction of the IAP volume and by the presence of additivesin the internal phase (Fig. 1 and Table 1). The photomicrographs ofunmodified (in absence of core-modifying agents) microspheres werespherical in shape and the pore size of the particles increased withincreasing volume of the IAP from 0.25 ml to 1 ml. Also, empty andfragmented cores were observed when the IAP volume was ≥1 ml(data not shown). This observation agrees with previously publishedstudies showing that polymeric microspheres prepared by the emul-sion solvent–evaporation technique become porous owing to thepresence of the IAP and that the degree of porosity increases as thefraction of the aqueous phase volume increases [29].

Incorporation of Span 60, stearylamine or PEI into the IAP affectedsurface morphology, porosity and particle–particle interaction of themicrospheres. Span 60 and stearylamine produced particles withrough surfaces containing sporadic pores (Fig.1A and B). Particles withsporadic pores may result from the tendency of Span 60 and steary-lamine to migrate to the oil–water interface and the polymeric phasedue to their lipophilic character. However, no significant differenceswere observed between the morphologies of the Span 60- and steary-lamine-modified particles despite the two surfactants being electro-statically different. Uniform sized particles with rough surfaces havepreviously been observed when neutral or cationic surfactants wereincorporated in the internal aqueous or OP of the primary emulsion[10,11,30]. The rough surface of the particles was caused by shrinkageof the polymer as water was being removed during drying. Further-more, the morphology of particles prepared using either of the surfac-tants was very similar, perhaps because the hydrophilic–lipophilicbalance (HLB) values for two surfactants are very close\the HLB ofSpan 60 is 4.7 and that for stearylamine is 7.9 [31].

When PEI, a polycation, was present in the IAP, the microspheresweremore porous compared to those formulatedwithout PEI, i.e., PM-2(Fig. 1C). A similar concentration-dependent increase in porosity andpore diameter was observed when PEI was complexed with oligonu-cleotides in the IAP to prepare PLGA-basedmicrospheres of the complex[32]. LMWH, like oligonucleotides, is a negatively charged molecule, sothat PEI formed an electrostatic complex with LMWH, helping to retainmore of the drug in themicrospheres. As a result, the space occupied bythe dispersed drugmay have increased alongwith the amount of bound

Fig. 2. Volume-based mean diameter (A), the data in parentheses show thepolydispersity index, PI, and Zeta potential (B) of LMWH-loaded PLGA microspheres.Data represent mean±standard error of the mean (n=3).

Fig. 1. SEM of LMWH-loaded PLGA microspheres (A) PM-SP-2, (B) PM-SA-2 and (C) PM-PEI-2. See Table 1 for composition of the different formulations.

227A. Rawat et al. / Journal of Controlled Release 128 (2008) 224–232

water. Alternatively, differences between the osmotic pressure of theinternal and external aqueous phases may have played a role in theincrease in porosity. In fact, it has been reported that increased osmoticpressure in the aqueous phase renders the particlesmore porous [9,33].As PEI is an osmotically active polycation, its presence in the IAP maylead to an increased influx of water, resulting in larger aqueous dropletsduring emulsification. Removal of water during lyophilization leavesvoid space in the particles, making them more porous and with rela-tively larger-diameter pores.

3.2. Particle size, zeta potential, tapped density and aerodynamicdiameter

Aerodynamic diameter and particle density play an important role insuccessful deposition of inhaled particles in the lungs. As discussed inthe Introduction, for traditional dry-powder inhalers, a geometric dia-meter of 1–5 μm is required for efficient deposition of particles in therespiratory tract. Recently, it has been proposed that lighter particles

with a mass density b0.4 g/ml and geometric diameter N5 μm could beused to enhance the residence time in the lungs and provide a prolongedrelease effect. To investigate if the particles meet the above criteria forpulmonary delivery, the volume-based diameter, tapped density, andaerodynamic diameter of the microspheres were determined (Figs. 2and 3). The data presented in Fig. 2A show that the particles weremoderately polydispersed, with a polydispersity index ranging from1.58 to 2.31. The mean volume-based diameter of the microspheressuggests that the size of the microspheres increases as the aqueousphase volume fraction increases; the highest size increasewas observedwith an IAP toOP ratio of 1:5. At a ratio of 2:1 a significant increase in thesize of the microspheres occurred due to the formation of a relativelyunstable secondary emulsion, as observed in the SEM photographs.

Incorporation of Span 60 in the IAP did not affect the particle size ofthe microspheres compared to those without Span 60. When PEI orstearylamine was incorporated in the aqueous phase, an increase inparticle size was observed, although no further increase was observedwhen the concentration of the agentswas increased from 1.25% to 2.5%(Fig. 2A). The basis for the increase in particle size could be that theseagents may leak out to the external aqueous phase and remain on thesurface of the negatively charged PLGA. Such a relocation of cationic

Fig. 3. Tapped density (A) and Theoretical (MMADt)/Actual mass median aerodynamic(B) diameter (MMADa) of LMWH-loaded PLGA microspheres. Data represent mean±standard error of the mean (n=3).

228 A. Rawat et al. / Journal of Controlled Release 128 (2008) 224–232

core-modifying agents onto the surface of microspheres can increasethe particle–particle interactions during the secondary emulsificationand this may lead to an increase in particle size. Furthermore, com-plexation of PEI or stearylaminewith the negatively charged drugmayhave contributed to the increase in size of the particles.

In addition to particle size, zeta potential and tapped density of theparticles were also determined. The tapped density provides importantinformation about the flowability of the particles from the inhaler device,theporosityof theparticles, theparticle sizedistribution, the truedensity,and interparticulate forces\cohesive and adhesive. The tapped density isalso affected by surface properties such aswettability and surface charge.The data presented in Fig. 2B show that plain microspheres werenegatively charged but the particles became positively charged uponaddition of cationic core-modifying agents, suggesting that the negativesurface charge of the drug and PLGApolymerwas partially neutralized bythe cationic core-modifying agents. The tapped densities decreased withthe increase in the internal aqueous volume fraction (Fig. 3A). When theinternal aqueous volume was increased from 0.25 ml to 1 ml, a twofolddecrease in the tapped density was observed, suggesting that theparticles became lighter because of the increase in porosity of themicrospheres. These data agree with the porous nature of the micro-spheres as observed in SEM photomicrographs. However, when Span 60

was used as a core-modifying agent, a reduction in tapped density wasobserved compared to plain microspheres and the magnitude ofreduction was dependent on the concentration of Span 60 used. Thepresence of Span 60may havemade the particles more hydrophobic andincreased the interparticle repulsive forces. As a result, the interparticulardistances increased and the particles became fluffier and lightercompared to the plain microspheres. However, when stearylamine orPEIwas added to theaqueous internal core, an increase in thedensitywasobserved compared tomicrosphereswithout stearylamineor PEI, and thetapped density increased slightly with the increase in concentration ofthe additives. No difference between the tapped densities of the PEI- andstearylamine-based microspheres was observed. Increased tappeddensity suggests a compact and dense internal structure of the particles.The increase in bulk density may be the result of reduced interparticulardistances because of the hydrophilic surface of the microspheresconferred by PEI or stearylamine. It is also likely that the presence ofcationic core-modifying agents rendered the microparticles' internalenvironment hydrophilic, causing the particles to retain residual waterand making them heavier than the plain microspheres, as observed byPistel and Kissel [33].

Themassmedian diameter (MMAD), a parameter important for thedeposition and distribution of particles in the respiratory tract, wasdetermined. The theoretical mass median diameter (MMADt) wasdetermined from the volume-based diameter and tapped density, andthe actual mass median diameter (MMADa) was calculated using aneight-stage Anderson cascade impactor, which is a widely used me-thod to characterize the aerodynamic particle size distribution emittedfrom therapeutic inhalation devices [34]. In terms of size distribution,the MMADt and MMADa data presented in Fig. 3B show a similarpattern. For example, for plain microspheres, an increase in MMADtand MMADa was observed when the internal aqueous phase volumewas increased. Similarly, for core-modified microspheres, the MMADaof PM-PEI-2 was greater than those of the other two core-modifiedmicrospheres. Importantly the MMADt and MMADa of the formula-tions were b6 µm, which is below the maximum value for aerosolizedparticles for efficient deposition in the lungs, as described by Edwardset al. and Van Campen and Venthoye [35]. On the whole, the datapresented in Figs. 2 and 3 suggest that the presence of additives in theIAP plays an important role in modifying the properties of particles.Moreover, the aerodynamic diameter of the particles was determinedto be within the respirable range, indicating their suitability for admi-nistering as a dry powder into the respiratory tract.

3.3. Entrapment efficiency

To prepare microspheres containing a single dose of the drug thatcan be administered in a single insufflation to the lungs, 50 mg LMWHwas encapsulated in 250 mg of PLGA microspheres. As indicatedabove, a fixed amount of the drug was dissolved in varying aqueousphase volume fractions in order to obtain microspheres of differentporosity. The data in Table 1 suggest that entrapment efficiency de-creases with increasing water volume in the internal phase. When theIAP volume was increased from 0.25 ml to 1 ml, the entrapmentefficiency was reduced from 17.87±1.79% to 9.39±1.21%. However, theentrapment efficiency of the formulations containing 0.5 ml of inter-nal phase was close to that obtained with 0.25 ml of internal phase.The reduction in entrapment efficiency observed with the increase inIAP likely occurred because of the formation of particles with largerpores, allowing the encapsulated drug to leach out through the poresduring the washing process. The large amount of IAP can also lead todestabilization of the primary emulsion and compromise the stabilityof the final particles. In fact, because of such instability of the primaryemulsion, a large number of empty cores were observed in micro-spheres prepared with 1 ml of IAP.

To enhance the entrapment efficiency and achieve extended-release properties, the drug was complexed with a polycation, PEI, or

Fig. 4. In vitro performances of the formulations. Data represent mean±standard errorof the mean. (A) Release profiles of LMWH from plain microspheres (n=3) (B) Releaseprofiles of LMWH from core-modified particles (n=3).

Table 2The microsphere surface-associated LMWH, initial burst effect and release rate ofLMWH

Formulation Surface-associated LMWH(%±SEM)

Burst release(%±SEM)

Release constant(%±SEM)

PM-1 2.0±0.5 3.33±0.69 0.259±0.012PM-2 4.6±0.6 5.79±0.75 0.232±0.018PM-3 8.2±1.2 10.36±1.21 0.102±0.01PM-SP-2 7.5±1.4 10.3±1.23 0.687±0.021PM-SA-2 13.1±1.5 21.3±1.85 1.121±0.036PM-PEI-2 17.7±2.1 25.01±1.62 1.247±0.030

SEM: Standard error of the mean.

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cationic surfactant, stearylamine, in the IAP. The effect of cationicagents on increasing the entrapment efficiency of anionic LMWH wasfurther compared with the effect of the nonionic surfactant, Span 60.For the preparation of core-modified microspheres, we chose to use0.5 ml IAP because plain microspheres prepared with this amount ofinternal phase volume showed an optimal in vitro release pattern.Incorporation of each of the three additives into the internal phasesignificantly increased the entrapment of LMWH in the microspheres(Table 1). When Span 60, stearylamine or PEI was used at a concen-tration of 1.25% or 2.5%, the LMWH entrapment efficiency wasincreased compared to that of particles prepared without any ofthese additives. The increase in entrapment efficiency was dependenton the concentration of the additives used in the internal phase. Forexample, when 1.25% PEI was used, a threefold increase in LMWHentrapment efficiency was observed (43.5±1.32%) compared toparticles prepared in the absence of PEI (16.22±1.32%). When theconcentration of PEI was increased to 2.5% of the IAP, a further increasein entrapment efficiency was observed.

The higher entrapment efficiency of LMWH in the microsphereswas because of the enhanced stability of the primary emulsion due tothe presence of any one of the three core-modifying agents: Span 60,stearylamine or PEI. However, as these additives are chemically diffe-rent, theymayhave increased the stability of the primaryemulsion andthe entrapment efficiency bydifferentmechanisms. Indeed, it has beenpreviously shown that a critical parameter for high drug-entrapment

efficiency in the final particle is the homogeneous dispersion of theaqueous phase in the OP of the emulsion [36]. The entrapment effi-ciency in the stearylamine-based formulationwas higher compared tothe Span 60-based formulation because stearylamine acts as both acomplexing and emulsifying agent. Stearylamine's positive surfacecharge may have helped the drug to complex with the surfactant andincrease the total amountof the drug in the emulsion. Furthermore, theincrease in entrapment efficiency was the highest with PEI, perhapsbecause a different mechanism is involved. It is hypothesized thatpositively charged PEI forms an electrostatic complex with the nega-tively charged LMWH. This hypothesis agrees with a previous studythat showed greater entrapment efficiency of heparin in PLGA micro-spheres prepared with Eudragit®, a quaternary ammonium group-carrying polymer, as compared to microspheres without Eudragit®[22]. We believe that PEI-LMWH complex formed in the IAP becomesrelatively hydrophobic compared to the drug alone. The resultingcomplex may migrate to the water–oil interface and facilitate theformation of a stable primary emulsion. A similar increase in entrap-ment efficiency was observed when PEI was complexed with oligo-nucleotides in the internal phase. Overall, the ability of the three agentsto increase the entrapment efficiency can be ranked as PEINstear-ylamineNSpan 60.

3.4. In vitro release profiles

The amount of heparin released from all formulations was es-timated and the release kinetics calculated. The amount of LMWHreleased at the zero time point was considered as surface-associateddrug, and that released during the first 30 min was considered as theinitial burst release. The in vitro release profile presented in Fig. 4 andTable 2 suggests a biphasic release pattern for all formulations: aninitial burst release followed by near zero-order release kinetics. Thephenomenon of burst release occurs because of the release of heparinadsorbed on the surface of the particles and development of a highconcentration gradient across the particle surface. For plain PLGA mi-crospheres containing no core-modifying agents, the amount of drugreleased during the initial burst phasewas related to the porosity of theparticles. As the porosity of the formulations increased, the amount ofdrug released also increased (Fig. 4A). As such, the amounts of drugreleased from formulations prepared with 0.25, 0.5 and 1 ml IAP were3.33±0.69, 5.79±0.75, and 10.36±1.21%, respectively (Table 2).Because of the presence of larger pores in the latter two formulations(PM-2 and PM-3), a larger amount of medium is likely to come incontactwith the particles and facilitate the release of the drug from themicrospheres.

For core-modified microspheres, the amount of drug released inthe initial burst phase was dependent on both the entrapment effi-ciency and the amount of drug associated with the surface of theparticles. Because more drug was available in the core-modifiedmicrospheres, more drug was released. For example, because theentrapment efficiency of the PEI-based formulationwas greater thanthat of the Span 60-modified microspheres, the amount of drug

Table 3Pharmacokinetic parameters of LMWH-loaded PLGA microspheric formulations

Formulations Pharmacokinetic parameters

Cmax (U/ml) Tmax (h) AUC0–24 (U h/ml) AUMC0–24 T1/2 (h) MRT (h) Frelative (%)

Plain LMWH subcutaneous 0.33±0.03 4.4±0.31 3.98±0.22 36.6±1.72 4.35±0.33 7.79±0.60 –

Plain LMWH pulmonary 0.20±0.01 2.2±0.44 1.58±0.12 9.63±0.77 4.28±0.31 6.63±0.47 39.7±3.0PM-2 0.07±0.04 0.8±0.4 1.05±0.11 10.57±2.04 12.13±0.93 18.43±1.42 26.38±2.8PM-SP-2 0.13±0.03 1.7±0.3 1.98±0.29 21.22±1.86 23.75±2.97 34.33±4.29 49.74±7.31PM-SA-2 0.20±0.04 6.9±0.89 3.64±0.36 43.14±3.7 25.28±2.81 39.09±4.34 91.45±9.01PM-PEI-2 0.21±0.03 7.6±0.70 3.37±0.40 40.8±2.38 20.13±1.68 32.39±2.70 84.67±10.11

Data represent mean±SEM (n=5–6).

Fig. 5. (A) In vivo performances of the formulations. Changes in plasma anti-factor Xaactivity after pulmonary administration of LMWH (50 U/kg) in plain and core-modifiedPLGA microspheres (n=5–6). (B) Effects of plain core-modifying agents and micro-spheres containing core-modifying agents on the viability of Calu-3 cells. The testsamples contained 0.25 mg/ml, 0.025 mg/ml or 0.0025 mg/ml of plain core-modifyingsurfactants or microspheres containing an equivalent amount of core-modifying agents.Data represent the mean±standard error of the mean (n=8).

230 A. Rawat et al. / Journal of Controlled Release 128 (2008) 224–232

released both at the initial burst and during the constant-releasephase was also higher for the PEI-modified microspheres (Fig. 4B). Inaddition to the entrapment efficiency, the amount of drug releasedwas also affected by the core-modifying agent used in the IAP. Thepresence of either of the surfactants, Span 60 or stearylamine,produced an increase in the amount of drug released during theinitial and burst phases. However, the amount of drug released fromthe Span-modified microspheres was higher than for the stearyla-mine-modified microspheres (Pb0.05) (Fig. 4B). These slight differ-ences could be because stearylamine is more hydrophilic than Span60, as suggested by their HLB values [37]. In agreement with thesedata, it was previously shown that surfactants can affect the releaseof entrapped heparin from PLGA microspheres [23]. Span 40 wasshown to decrease the release of heparin from PLGA particles;however, both Span 80 and 85 have been shown to increase therelease considerably [23]. Similarly, for PEI, which is more hydro-philic than either Span 60 or stearylamine, the rate and extent ofrelease of the drug from PEI-based formulations was higher than thatfrom the two other formulations. Hydrophilic PEI facilitates interac-tion of the aqueous medium with PLGA, which results in fasterrelease, as observed by De Rosa et al. [32,38]. Furthermore, an excessof PEI-LMWH complex may have adsorbed onto the surface of themicrospheres and been released soon after the particles came incontact with water.

Overall, in terms of the amount of drug released in the burst phase,the formulations can be ranked in the following order: PM-PEI-2NPM-SA-2NPM-SP-2. Furthermore, the rate of release was independent ofthe entrapment efficiency, suggesting a constant rate of release afterthe burst phase (Table 2). The data also suggest that core-modifiedmicrospheres can release LMWH for 24 h, and that the amount of drugreleased during 24 h is greater than that released from previouslyreported PLGA microspheres of LMWH [21,23,24]. A larger amount ofdrugwas released from this formulation because the PLGAusedwas oflow molecular weight, and low molecular weight PLGA undergoesfaster degradation in dissolution medium at 37±1 °C due to its re-latively low glass transition temperature of 28–37 °C [39]. However,caution should be exercised in correlating the in vitro release datawith the release of the drug in the lungs because the composition ofthe medium used for in vitro release studies is significantly differentfrom that of the lung fluid.

3.5. Pulmonary absorption of LMWH entrapped in microspheres

Pulmonary absorption of the encapsulated drug was studied in arodent model and the amount of drug absorbed was monitored bymeasuring plasma anti-factor Xa activity. To confirm that pulmonaryadministration of blank PLGA microspheres has no effect on anti-factor Xa activity, microspheres without LMWHwas administered andanti-factor Xa levels weremeasured. No change in anti-factor Xa levelswas observed, indicating that PLGA had no effect on endogenous anti-factor Xa levels (data not shown). When plain LMWH was adminis-tered via the pulmonary route, a rapid onset of absorption was ob-served, with a maximum increase in blood anti-factor Xa activity

(Cmax=0.20±0.01 U/ml) at 2.2±0.44 h (Tmax) and a drug half-life of4.28±0.31 h (Table 3). Administration of plain PLGA microspherescontaining an equivalent dose of the drug resulted in a considerable

231A. Rawat et al. / Journal of Controlled Release 128 (2008) 224–232

prolongation of anti-factor Xa activity (T1/2=12.13±0.93 h) in the blood,although themaximum increase in anti-factor Xa activitywas below thelevel required for an anti-thrombotic effect in a rodent model [40].

In agreement with in vitro release data presented in Fig. 4A and B,the levels of anti-factor Xa produced by the core-modified micro-spheres and the time required to achieve the Cmax was dependent onthe entrapment efficiency of the formulations (Fig. 5A). The Span 60-based formulation, PM-SP-2, showed an initial increase in anti-factorXa activity in the blood and produced a maximum increase(Cmax=0.13±0.03 U/ml at 1.7±0.3 h). Although the Cmax produced bythis formulation was lower than that produced by plain pulmonaryLMWH, the blood concentration of anti-factor Xa activity was at asteady state for 24 h. The stearylamine- and PEI-based microspheresalso showed a slower absorption phase compared to plain LMWH.However, the anti-factor Xa levels produced by these two formula-tions were similar to that produced by plain LMWH administeredby the pulmonary route. For both formulations, the time requiredto achieve maximum anti-factor Xa levels was longer than the Tmax

for plain LWMH. The stearylamine-based formulation reached itsCmax (0.20±0.04 U/ml) at 6.9±0.89 h, and the PEI-based formulationreached its Cmax (0.21±0.03 U/ml) at 7.6±0.7 h. The eliminationphase of both formulations was longer than those for subcutaneousand pulmonary administration of plain LMWH. Furthermore, a highconcentration of anti-factor Xa in the blood was observed for 24 h,underscoring the prolonged release properties of the formulations. Thedatapresented inTable3 also clearly demonstrate that thehalf-life of thedrug released from the core-modifiedmicrosphereswas increased by 5-to 6-fold compared to subcutaneous LMWH. Similar to the half-life, therelative bioavailabilities of the stearylamine- and PEI-based formula-tions were increased by 2-fold compared to plain LMWH administeredby the pulmonary route.

Several factorsmay have contributed to the increase in half-life andbioavailability of the formulations. First, slower release of the drugfrom the formulations should increase the drug's half-life, and a highermagnitude of drug release should enhance its bioavailability. Second,prolongation of the drug's effect can be explained by the fact that theparticles were too big to be phagocytosed by alveolar macrophages [3]so that the drug remained in the lung for a longer period of time. Inaddition to particle size, the rate of release of a drug from the micro-spheres may also depend on the molecular weight of the PLGA usedand the porosity of the microspheres. As discussed above, the bio-degradability of PLGA microspheres depends on the molecular weightand glass transition temperature of the polymer. The PLGA used in thisformulation has a glass transition temperature of 28–37 °C, and hencecan easily degrade at physiological temperature and release the drugfrom the formulations for absorption via the respiratory epithelium[39]. However, the fate of PLGA after release of the drug is unknown. Itis likely that some residue of PLGA remains in the lung after the drughas been released. Further investigations are required to determine thefate of PLGA after pulmonary ingestion. The porosity of the micro-spheres also plays amajor role in enhancing the rate of biodegradationby exposing more surface area of the polymer matrix to the lung fluid[41]. The presence of hydrophilic molecules such as PEI and surfactantmay also facilitate degradation of polymers by increasing the wettabi-lity and reducing the contact angle of the polymer with the phy-siological fluid, as suggested previously [42,43]. Another importantfactor that may contribute to the enhanced bioavailability is theproperty of the core-modifying agent to increase the permeability ofthe respiratory epithelium. It is hypothesized that core-modifyingagents used to prepare microspheres may separate out from the mi-crospheres and act as absorption promoters [37].

3.6. MTT cytotoxicity study

To evaluate the cytotoxicity of the core-modifying agents to Calu-3cells, the viability of the cells was assessed with the MTT assay. MTT, a

tetrazolium salt, is cleaved by mitochondrial dehydrogenase in livingcells to form a measurable, dark blue product called formazan. Da-maged or dead cells have reduced dehydrogenase activity and dimi-nished formazan production. Results of the MTT test showed highlevels of viability of Calu-3 cells after incubation with saline; incontrast, only 21% of Calu-3 cells were viable following treatment with0.1% SDS (Pb0.05) (Fig. 5B). A concentration-dependent increase incytotoxicity was observed for both plain core-modifying agents andmicrospheres containing core-modifying agents. Of the three core-modifying agents tested, the effect of PEI in cell death was morepronounced than that of the other two agents. However, when PEI wasused as a core-modifying agent to prepare LMWH-entrapped micro-spheres, a reduction in cytotoxicity was observed compared to plainPEI. This reduced cytotoxicity was because of partial neutralization ofthe positive surface charge of PEI in the presence of negatively chargedLMWH. Overall, the cytotoxicity produced by all core-modified mi-crosphereswas comparable to that produced by blank or plain LMWH-loaded microspheres and that of the saline control.

Taken together, the entrapment efficiency of LMWH in PLGAmicrospheres can be increased by incorporating Span 60, stearylamineand PEI in the IAP. In terms of particle size, porosity, aerodynamicdiameter, in vitro release behavior and cytotoxicity, the core-modifiedmicrosphereswere optimal for pulmonary administration. The drugwasreleased, both in vitro and in vivo, for 24 h and themaximum increase inanti-factor Xa levels was observed for the PEI- and stearylamine-modifiedmicrospheres. The half-life of the drug from the core-modifiedmicrospheres was considerably higher than that of plain LMWH admi-nistered via the pulmonary route. The data presented here demonstratethat PLGA-based large porous microspheres could be a viable option fordelivering LMWHvia the pulmonary route for a prolonged release effect.

Acknowledgment

This work was supported by NIH grant R 15 HL7713302 (FA).

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