Acta Biomaterialia 10 (2014) 1194–1205

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Acta Biomaterialia

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Poly(ethylene glycol) methacrylate hydrolyzable microspheresfor transient vascular embolization

1742-7061/$ - see front matter � 2013 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.http://dx.doi.org/10.1016/j.actbio.2013.11.028

⇑ Corresponding author at: Université Paris-Sud, Institut Galien Paris-Sud, LabExLERMIT, Faculté de Pharmacie, 5 rue J.B. Clément, 92296 Châtenay-Malabry, France.

E-mail address: laurence.moine@u-psud.fr (L. Moine).

Stéphanie Louguet a, Valentin Verret a,b, Laurent Bédouet a, Emeline Servais a, Florentina Pascale b,Michel Wassef c, Denis Labarre d,e, Alexandre Laurent c,f, Laurence Moine d,e,⇑a Occlugel S.A.S., 12 Rue Charles de Gaulle, 78350 Jouy en Josas, Franceb Archimmed S.A.R.L., 12 Rue Charles de Gaulle, 78350 Jouy en Josas, Francec AP-HP hôpital Lariboisière, Department of Pathology, University of Paris 7 – Denis Diderot, Faculty of Medicine, 2 rue Ambroise Paré, 75010 Paris, Franced Université Paris-Sud, Institut Galien Paris-Sud, LabEx LERMIT, Faculté de Pharmacie, 5 rue J.B. Clément, 92296 Châtenay-Malabry, Francee CNRS UMR 8612, Institut Galien Paris-Sud, LabEx LERMIT, 5 rue J.B. Clément, 92296 Châtenay-Malabry, Francef ‘‘Laboratoire Matières et Systèmes Complexes’’, CNRS 7057, University of Paris 7, Bâtiment Condorcet, 10 rue Alice Domon et Léonie Duquet, 75205 Paris Cedex 13, France

a r t i c l e i n f o a b s t r a c t

Article history:Received 25 June 2013Received in revised form 6 November 2013Accepted 27 November 2013Available online 7 December 2013

Keywords:EmbolizationMicrosphereDegradableHydrolyzable crosslinkerPEG-based hydrogel

Poly(ethylene glycol) methacrylate (PEGMA) hydrolyzable microspheres intended for biomedical applica-tions were readily prepared from poly(lactide-co-glycolide) (PLGA)–poly(ethylene glycol) (PEG)–PLGAcrosslinker and PEGMA as a monomer using a suspension polymerization process. Additional co-mono-mers, methacrylic acid and 2-methylene-1,3-dioxepane (MDO), were incorporated into the initial formu-lation to improve the properties of the microspheres. All synthesized microspheres were spherical inshape, calibrated in the 300–500 lm range, swelled in phosphate-buffered saline (PBS) and easily inject-able through a microcatheter. Hydrolytic degradation experiments performed in PBS at 37 �C showed thatall of the formulations tested were totally degraded in less than 2 days. The resulting degradation prod-ucts were a mixture of low-molecular-weight compounds (PEG, lactic and glycolic acids) and water-sol-uble polymethacrylate chains having molecular weights below the threshold for renal filtration of50 kg mol�1 for the microspheres containing MDO. Both the microspheres and the degradation productswere determined to exhibit minimal cytotoxicity against L929 fibroblasts. Additionally, in vivo implanta-tion in a subcutaneous rabbit model supported the in vitro results of a rapid degradation rate of micro-spheres and provided only a mild and transient inflammatory reaction comparable to that of the controlgroup.

� 2013 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

1. Introduction

Embolization, i.e. occlusion of arteries feeding organs with in-jected biomaterials, is a technique used for the treatment of vari-ous pathologies such as tumors, arterio-venous shunts andhemorrhages. It is commonly achieved with particles, which canbe either degradable or non-degradable. The degradable particlesare preferred when a durable vessel occlusion is not required, forexample in uterine fibroids that are very sensitive to ischemia, inbleedings where the blood leak can be stopped if the flow isblocked for a few hours, or in some tumors, specially malignantones, when embolization has to be repeated in the same feederarteries.

Non-degradable particles have evolved considerably in the lastfew decades. The irregular poly(vinylalcohol) (PVA) particles of the

1970s, which were poorly calibrated, aggregated proximally inclusters and gave a chronic inflammatory response [1–3], havebeen replaced by smooth, biocompatible and calibrated micro-spheres (MS). They drastically changed the conditions of emboliza-tion, since the radiologist could adapt the size of the MS to the sizeof the vessels to allow precise targeting and accurate devasculari-zation. These MS have since been chemically grafted, to be loadedionically with cationic drugs, and have thereby become actual drugdelivery systems for local chemotherapy in tumor embolization[4,5].

By comparison, the evolution of the degradable particles hasbeen very slow. Porcine sponge particles (GSP) were used at thevery beginning of embolization in the 1970s and are still com-monly used today. The enzymatic degradation of GSP is highly var-iable over time, from 3 weeks to 4 months [6–10], and isaccompanied by a chronic inflammatory response and a vesselremodelling process [6,9], so that functional recanalization cannotbe assured [11,12].

S. Louguet et al. / Acta Biomaterialia 10 (2014) 1194–1205 1195

It is only very recently that two types of degradable MS have be-come available for embolization. However, their severe limitationshamper their competitiveness with non-degradable MS. Starch MS(Embocept™, Pharmacept) have been proposed to enhance drugdiffusion in tumoral tissues when coadministered with chemother-apeutic drugs [13–15]. However, their small diameter (<100 lm)limits their use to small vessels, and they produce very transientvascular occlusions (�1 h) that cannot lead to tumor necrosis. Thisis a major drawback for the treatment of ischemia-responding tu-mors, such as uterine fibroids, which also require large MS(>500 lm). Collagen-coated poly(lactide-co-glycolide) (PLGA) MS(Occlusin™500, IMBiotechnologies) are also available, but in onlyone small size (150–212 lm). Further, their in vivo degradation re-quires several months and is accompanied by an inflammatory andfibrotic reaction [16].

We have therefore conceived a degradable microsphere thatencompasses the useful properties of non-degradable MS. First,we aimed to make a degradable microsphere which temporary oc-cludes the vessel between a few hours and a few days, which is suf-ficient to achieve an ischemia of tumors (in the case of fibroids), avessel repair (in the case of bleeding) or the delivery of a drug totumors (in the case of chemoembolization). Second, this biomate-rial should, after its degradation, be eliminated quickly, before theonset of a chronic inflammatory response and vessel wall remodel-ling [l2]. Third, the MS must be calibrated in several size ranges,between 100 and 1000 lm. Fourth, the MS should be easily sus-pended in physiological solutions and contrast medium. Fifth, thematerial must be soft and elastic enough to be injected via micro-catheters of diameters smaller than the MS, and to recover theirsize and shape upon exit [17]. Sixth, the MS must be loadable withdrugs for transient chemoembolization procedures.

To fulfill the above-listed requirements, we have successfullydeveloped a tunable and degradable MS that combines poly(ethyl-ene glycol) methacrylate (PEGMA) and a hydrolyzable crosslinker(PLGA–poly(ethylene glycol) (PEG)–PLGA). PEGMA was selectedas the monomer because of its amphiphilic nature, which resultsfrom its water-soluble PEG side chain and its hydrophobic methac-rylate group. The amphiphilic nature of PEGMA makes possiblemixing it directly with the hydrolyzable crosslinker in hydrocarbonphase to perform the polymerization by direct oil in water suspen-sion process for making microspheres. The hydrolyzable cross-linker was designed to achieve rapid degradation time. Ananionic co-monomer, methacrylic acid (MA) was combined to thePEGMA monomer and the crosslinker to adjust the water uptakeand to allow the ionic loading of drugs. A large part of the strategywas devoted to the biocompatibility of the material. First, the ini-tial components were chosen to generate water-soluble macromol-ecules after hydrolysis in order to avoid any accumulation in theorganism, which might lead to inflammatory reactions [18]. Sec-ond, additional ester linkages from a cyclic ketene acetal monomer,2-methylene-1,3-dioxepane (MDO), were incorporated into thehydrogel backbone to shorten the size of the polymer degradationproducts (below 50 kg mol�1), facilitating their renal elimination[19,20].

We report here the feasibility of preparing suitable degradablemicrospheres for embolization with the selected components(hydrolyzable crosslinker, PEGMA, MA and MDO). The impact ofeach component on the degradation time and on the size of thedegradation products was evaluated. The hydrolytic degradationof the different formulations was characterized and the effect ofdegradation products on the biocompatibility was studied by twodifferent cytotoxic tests. Subcutaneous implantations were thenperformed in rabbits to assess the in vivo degradation of MS andthe local inflammatory reaction.

2. Material and methods

2.1. Materials

PEGMA (Mn = 300 g mol�1), stannous octanoate, triethylamine,methacrylic anhydride, 88% hydrolyzed PVA, glycolide, 1-hexa-nethiol and polyethylene glycol (Mw = 600 g mol�1) (PEG) werepurchased from Aldrich (St Quentin Fallavier, France). 2,20-Azobis-isobutyronitrile (AIBN), used as a polymerization initiator, andtrimethylsilyldiazomethane (2 M in hexane; TMS) were obtainedfrom Acros Organic (Geel, Belgium). D,L-Lactide was obtained fromBiovalley (Marne la Vallée, France). MDO was synthesized accord-ing to the procedure published in Ref. [21]. Doxorubicin hydrochlo-ride (Adriblastin) was purchased from Pfizer (France). Embosphere(non-degradable gelatin microsphere; GMS) was purchased fromBiosphere Medical (Roissy, France). DC-beadsTM microsphereswere purchased from Biocompatibles (Farnham, UK). All reagentswere used without any further purification. Analytical grade sol-vents were supplied by Carlo Erba (Val de Rueil, France).

2.2. Instrumentation

2.2.1. Nuclear magnetic resonance (NMR)Products were analyzed by 1H NMR spectroscopy using a Bruker

DPX300 FT NMR spectrometer, with the solvent peak as areference.

2.2.2. Fourier transform infrared spectroscopy (FTIR)FTIR was used to characterized the chemical structure of poly-

mers and MS. Spectra were recorded on a Spectrum One Perkin El-mer infrared spectrometer equipped with a diamond attenuatedtotal reflectance module in a range of 500–4000 cm�1. One hun-dred interferograms were averaged per spectrum at a resolutionof 2 cm�1.

2.2.3. Size exclusion chromatography (SEC)The molecular weight and molecular weight distribution of

polymers were measured by SEC at 30 �C on a system equippedwith a guard column and two GMH HRm columns (Viscotek) withtetrahydrofuran (THF) as the eluent at 1 ml min�1. A differentialrefractometer and a double SEC detector (model 270 Dual, Visco-tek) with RALS and viscosimeter in series were used to analyzesamples. The data obtained were treated with OmniSEC software(Viscotek).

2.3. PLGA–PEG–PLGA-hydrolyzable crosslinker synthesis

The crosslinker synthesis is a two-step reaction comprisingpolyester block polymerization initiated by hydroxyl-terminatedPEG chains and end polymer chain modification with polymeriz-able double bonds.

PEG (10.07 g, 16.8 mmol), DL-lactide (7.26 g, 50.4 mmol), glyco-lide (5.86 g, 50.5 mmol) and stannous 2-ethylhexanoate (128 mg,0.3 mmol) were added to a dry Schlenk tube and subjected to sev-eral vacuum–argon cycles. The Schlenk tube was then heated at115 �C for 20 h under argon atmosphere with continuous magneticstirring. After the reaction had been completed, the polymer wascooled and then dissolved in 40 ml of dichloromethane and precip-itated twice, first in a 1 l equivolumic mixture of diethyl ether andpetroleum ether and then in petroleum ether, to remove any tracesof unreacted monomer. Purified polymer was dried under a vac-uum at room temperature and characterized by 1H NMR. The finalproduct is a translucent gel (yield 95%).

1196 S. Louguet et al. / Acta Biomaterialia 10 (2014) 1194–1205

1H NMR (CDCl3) d (ppm): 5.06–5.35 (m, CH, LA), 4.56–4.93 (m,CH2, GA), 4.2–4.45 (m, CO–O–CH2 of PEG, CH or CH2 of the last LAor GA units), 3.36–3.81 (m, CH2 of PEG), 1.37–1.71 (m, CH3, LA).The peak corresponding to the CH2 of ethylene glycolide was takenas a reference for the calculation of the polymerization degree ofboth lactide and glycolide units.

PLGA–PEG–PLGA copolymer (3 g, 2.2 mmol) was added to aSchlenk tube and dissolved in 30 ml of degassed ethyl acetate.The reaction mixture was cooled at 0 �C in a glass bath and, after5 min of gentle stirring, triethylamine (6 eq.) was added dropwiseunder an argon flow. Methacrylic anhydride (6 eq.) was then addeddropwise, also under the argon flow. The final solution was stirredfor 1 h at 0 �C, then heated at 80 �C for 8 h. The mixture was precip-itated three times in petroleum ether to remove excess methacrylicanhydride and triethylamine. Purified polymer was dried under avacuum at room temperature and characterized by 1H NMR. The fi-nal product is a light yellow gel (yield 95%).

1H NMR (CDCl3) d (ppm): 6.13–6.29 ppm (m, CH@), 5.55–5.7 ppm (m, CH@), 5.06–5.35 ppm (m, CH, LA), 4.56–4.93 ppm(m, CH2, GA), 4.31 (m, CO–O–CH2, PEG), 3.36–3.81 ppm (m, PEG),1.8–2.05 ppm (m, CH3 methacrylate), 1.37–1.71 ppm (m, CH3,LA). The functionalization degree was calculated using the peakcorresponding to the CH2 of ethylene glycolide as a reference andthe peak of the CH of the methacrylic group.

2.4. Linear polymers and gels synthesis

In a typical linear polymer synthesis, PEGMA (2 g, 6.6 mmol),MDO (85 mg, 0.7 mmol) and 1-hexanethiol (28 ll, 0.2 mmol) wereadded to a vial and dissolved in 2 g of toluene. The vial was heatedat 80 �C for 8 h under gentle magnetic stirring. After the reactionhad been completed, 10 ml of dichloromethane was added to thevial and the reactive medium was precipitated in petroleum etherto remove any traces of unreacted monomer. Purified polymer wasdried under a vacuum at room temperature, then characterized by1H NMR and SEC. The final product was a slightly white gel (yield98%).

1H NMR (CDCl3) d (ppm): 3.8–4.3 (m, CO–O–CH2, MDO andPEGMA), 3.5–3.8 (m, CH2, PEGMA), 3.3–3.5 (s, CH3, PEGMA) and0.7–2.2 (m, CH2–CO and CH2, MDO and CH2–C(CH3)–, PEGMA).The MDO content was calculated using the peak correspondingto the CH3 of PEGMA as a reference and the peak of CO–O–CH2

of the MDO and PEGMA blocks.For gels synthesis, the PLGA–PEG–PLGA crosslinker was added

to the reactive medium at 5 mol.%. After the reaction had beencompleted, gels were washed with acetone and water to removeany traces of unreactive monomer.

2.5. Microspheres preparation

The MS process parameters (i.e. the oil-to-water ratio, themonomer mass concentration of the organic phase, the reactiontemperature and the stirring speed) were selected to obtain MSwith a diameter between 300 and 500 lm and kept constant alongthis study.

In a typical experiment, an aqueous solution (111 ml) containingPVA (0.5 wt.%) and NaCl (3 wt.%) was added to a 0.25 l reactor. Thedispersed phase, containing PEGMA (5.38 g, 18 mmol), PLGA–PEG–PLGA crosslinker (1.816 g, 1.2 mmol), MA (206 mg, 2.4 mmol), MDO(273 mg, 2.4 mmol), 1-hexanethiol (76 ll, 0.54 mmol) and AIBN(42 mg, 0.25 mmol) solubilized in 6.15 g of toluene, was degassedby nitrogen bubbling for 15 min. The solution was added to theaqueous phase at 80 �C and stirred at 180 rpm by using an Inoximpeller for 8 h. The MS were collected by filtration on a 100 lmsieve and washed extensively with acetone and water. They werethen sieved with decreasing sizes of sieve (800, 500, 300 and

100 lm). All MS were freeze dried immediately after purification.Dry MS pellets were then sterilized by beta-irradiation (Ionisos,France) and stored at �20 �C until use. The mass yield of the MSsynthesis is 78%.

2.6. Hydration kinetics

Dry pellets of sterilized MS (300–500 lm) were loaded onto a glassslide using a plastic tip and pictures were taken with microscope (LeitzDIAPLAN) (magnification� 2.5). The dry pellets were then hydratedby adding saline solution (Versol, Laboratoire Aguettant) and, after5 min, pictures of the MS were taken with a microscope. The wet MSwere transferred to a 15 ml tube with 1 ml of saline and incubatedat room temperature with occasional shaking. After 30 min and 3 hof incubation, MS were removed from the tube to a glass slide and pic-tures were taken. The same experiment of MS swelling was performedin a 1/1 mixture of water with non-ionic contrast medium (Omni-paque, GE). The diameters of MS (521 dry MS, 1614 MS swollen in sal-ine and 1627 MS swollen in saline/Omnipaque mixture, respectively)were determined using the ImageJ software.

2.7. Swelling evaluation

In triplicate, 1 ml of wet MS was placed in pre-weighted 15 mlpolypropylene vials before freeze-drying. The dry MS weight wasdetermined. MS were then incubated for 10 min in distilled waterand the supernatant and interstitial water between the MS wereremoved by aspiration with a pipette tip. The tube was weightedagain and the amount of absorbed water calculated. The same pro-cedure was repeated with phosphate buffer (0.1 M phosphate-buf-fered saline (PBS) with 0.9% NaCl and 0.02% NaN3, pH 7.4). Thesolvent taken up by MS was calculated according to the formula(Ws �Wd)/Wd, where Ws was the weight of the swollen MS andWd was the weight of the dry MS.

2.8. MS injectability

To test their mechanical compliance, MS were dispersed in asolution composed of 50 vol.% saline solution (Versol, LaboratoireAguettant) and 50 vol.% contrast medium (Omnipaque, GE), theninjected via a microcatheter (EV3, Echelon 10, inner diameter430 lm) in a vial.

2.9. Doxorubicin loading on MS

Aliquots of 0.1 ml of MS, either degradable MS or DC-beadsTM

(Biocompatibles, Farnham, UK), were incubated with 3.5 mg oflyophilized doxorubicin (Adriblastin, Pfizer) reconstituted in water(2.5 mg ml�1). Next, 50 ll of sodium hydrogen carbonate 1.4%(Fresenius Kabi, France) was added to the tubes containing degrad-able MS to slightly elevate the pH of the solution. Tubes were thenincubated for 1 h at room temperature under agitation.

The doxorubicin remaining in the supernatant after 1 h wasquantified by optical density reading at 492 nm in a 96-well micro-plate reader. The loaded dose was calculated by subtracting the fi-nal amount of doxorubicin from the initial amount. The loadingefficiency was calculated by the following equation:

Loading efficiency¼Doxorubicin in feed�doxorubicin in supernatantDoxorubicin in feed

�100

2.10. In vitro degradation

In triplicate, wet MS were placed in pre-weighted 15 mlpolypropylene tubes before freeze-drying. Dry MS weights weredetermined and phosphate buffer was added to make a

S. Louguet et al. / Acta Biomaterialia 10 (2014) 1194–1205 1197

concentration of 10 mg ml�1 MS. The tubes were then shaken hor-izontally at 200 rpm at 37 �C in an orbital shaker oven (IKA,KS4000i control). The samples were taken out after 0, 6, 15, 24and 48 h. After MS decantation, the pH values of the degradationmedia were collected (Mettler Toledo, type M8220). The mediawere then removed and the MS were rinsed (3 � 10 ml) withdeionized water before freeze-drying. Tubes were re-weighed todetermine the weight loss using the following equation: weightloss (%) = (W0 �Wt)/W0 � 100, where W0 and Wt are the dryweight of the sample before and after degradation, respectively.The reported weight loss was the average of three samples.

2.11. Linear polymers, gels and MS degradation products

The hydrolytic degradation of linear polymers, gels or MS wasperformed as follows: 150 mg of linear polymers, 150 mg of gelsor 1 ml of wet MS sediment was added to a 15 ml Falcon tubeand made up to 10 ml with 0.1 M NaOH solution. The solutionswere mechanically stirred for 48 h at 37 �C in an incubator(176 rpm) and dialyzed against deionized water for 48 h to removesalt and low-molecular-weight degradation products (dialysismembrane molecular weight cut off: 100–500 Da).

1H NMR analysis of the MS degradation products was per-formed in dimethylsulfoxide-d6 on lyophilized MS degradationproducts (see the figure in the Supplementary data).

To solubilize the degradation products in the SEC solvent (THF),they were chemically modified through the methylation of carbox-ylic groups using trimethylsilydiazomethane. For this, 10 ml ofdialysate was added to a 100 ml round-bottom flask togwther with40 ml of THF. TMS solution (roughly 1 ml) was added dropwise tothe solution, which turned yellow. The solution was stirred over-night at 20 �C. After the completion of the reaction, any excessTHF, hexane and TMS was removed by rotary evaporation. Finally,the methylated polymer was lyophilized prior to its solubilizationin the SEC solvent for analysis.

2.12. Cytotoxicity of microsphere extracts

Cytotoxicity of MS was analyzed using extracts of MS preparedin cell culture medium. Briefly, mouse fibroblasts (L929) cultureswere maintained in high-glucose Dulbecco’s modified Eagle’s med-ium (DMEM) with 10% fetal bovine serum, 2 mM L-glutamine,50 lg ml�1 streptomycin and 50 units ml�1 penicillin in a CO2

incubator at 37 �C. L929 cells harvesting was performed using tryp-sin–ethylenediaminetetraacetic acid (Lonza) and subculturesstarted in 96-well plates (NUNC) at densities of 5 � 103 -cells well�1. MS extracts were prepared in sterile tubes by adding500 ll of MS pellets in DMEM and making up the volume to 3 mlwith cell culture medium without serum. Samples were incubatedat 37 �C under agitation until the MS had completely degraded.Chirurgical glove fragments (latex) and Embosphere were used asthe positive and negative control of cytotoxicity, respectively.The day after cell seeding, bovine serum was added to the MS ex-tracts and the pH was adjusted to pH �7 with 20 mM sodiumhydrogen carbonate (Fresenius Kabi, France) before non-confluentfibroblasts were added (6–8 wells per condition). Extracts obtainedfrom the chirurgical gloves and GMS were also added to mousefibroblasts. After 72 h of culture (37 �C, 5% CO2), the medium wasremoved and the cells were washed with 100 ll of PBS, beforethe addition of 100 ll of bicinchoninic acid solution (BCA proteinreagent, Sigma) containing 0.08 wt./vol.% CuSO4 and 0.05% TritonX-100. After incubation (1 h at 37 �C), absorbance was measuredat 570 nm and the amount of proteins was obtained by extrapola-tion from the standard curve using bovine serum albumin.

2.13. Contact toxicity

In 24-wells plates, L929 cells were seeded at 50,000 cells well�1

in cell culture medium (see above). The day after seeding, the med-ium was removed and replaced with 0.9 ml of culture mediumcontaining 100 mM HEPES. Then 100 ll of sterilized microspheresediment was added (in duplicate) and the cells were co-culturedwith MS for 1 week (37 �C, 5% CO2). Pieces of latex gloves andnon-degradable GMS were used as the positive and negative con-trol of cytotoxicity, respectively. At the end of culture, pH of cul-ture medium was estimated using pH indicator paper andpresence of non-degradable MS was checked. The amount of totalcell proteins was determined by a BCA assay. Briefly, the medium isremoved and the cells are washed with PBS, then lysed with 300 llof PBS containing 0.05% Triton X-100 and 1 unit of DNAse I for 1 hat 37 �C. The proteins are determined in triplicate using 20 ll ofcell lysate.

2.14. Subcutaneous implantation in rabbit

Experimental procedures were performed at the Center of Re-search in Interventional Radiology (Cr2i/APHP/National Instituteof Agronomic Research INRA, Jouy-en-Josas, France). The studyprotocol was approved by the Institutional Animal Care and UseCommittee of the Center and was conducted according to Euro-pean Community Rules of Animal Care.

Different degradable MS (PEG–PLGA, PEG–PLGA–MA and PEG–PLGA–MA–MDO) were implanted into the back skin of whiteNew Zealand rabbits (n = 2 per group). Briefly, anesthesia was in-duced by 5% isoflurane/95% oxygen through a mask. The implanta-tion sites were shaved and tattooed. A 0.6 ml volume ofmicrospheres swelled in saline was injected through a 14G needlein six 4 cm long parallel rails per rabbit to avoid a cluster effect.Rabbits were then sacrificed on day 2 or 7. In addition, four rabbitswere injected with saline only (sham) as controls and sacrificed onday 2 or 7.

At sacrifice, the skin was explanted according to the tattooedmarks, attached to a plate and fixed in 10% neutral buffered forma-lin for at least 24 h. For each rabbit, a 4 mm thick tissue slice wascut transversally from each of the six implant rails. The sampleswere dehydrated in a series of alcohols, set in xylene and embed-ded in paraffin. Sections 3–4 lm thick were cut and mounted ona microscope slide. After rehydration in a series of alcohols, sec-tions were stained with hematein–eosin–saffron.

Tissue reaction and material degradation were assessed by his-tology. The frequency of observation of amounts of (almost de-graded) MS in the histological slides was calculated.

The main organs were harvested at sacrifice and checked by avet and a pathologist for abnormalities that would suggest sys-temic toxicity.

3. Results

3.1. Formulation and preparation of hydrolyzable microspheres

A hydrolyzable crosslinker, PLGA–PEG–PLGA dimethacrylate,was prepared via an established protocol [22]. Precursor was firstsynthesized by ring-opening polymerization of lactide and glyco-lide using Mw 600 PEG as the initiator. The composition of the oli-gomer obtained was found to be similar to the correspondingmonomer feed ratio. The hydroxy-terminated precursor was thenallowed to react with methacrylic anhydride to form crosslinkablepolyesters. The complete conversion of the terminal OH groups bythe methacrylate group was confirmed by 1H NMR.

1198 S. Louguet et al. / Acta Biomaterialia 10 (2014) 1194–1205

A series of microspheres were then prepared by suspensionpolymerization using 5 mol.% hydrolyzable crosslinker and PEGMAas the main monomer (Fig. 1). Co-monomers such as MA and MDOwere incorporated to improve the degradation characteristics ofthe microspheres: MA allows drug loading and improves water up-take when ionized; MDO decreases the molecular weight of theresidual polymer chains after hydrolysis.

The cyclic monomer MDO is able to (co)polymerize through afree radical mechanism either by a ring-opening reaction to forman ester in the main chain or by a vinylic addition reaction at thedouble bond with retention of the ring structure [23]. Prior to itsincorporation into the microspheres, the MDO was copolymerizedwith PEGMA at different feed ratios, using AIBN at 80 �C as the ini-tiator. In all cases, the molar ratio of PEGMA/MDO, as determinedby 1H NMR, was higher in the linear copolymers than in the initialfeed (Table 1). No peak around 100–110 ppm, corresponding to theacetal carbon of the ring-retained structure of MDO, was evidencedin the 13C NMR spectra of the copolymers. Thus, copolymerizationof MDO with PEGMA in this condition led to quantitative ringopening of MDO.

To determine the influence of the introduction of MDO on thereduction of the molar mass, gels of MDO/PEGMA containing5 mol.% hydrolyzable crosslinker with increasing proportions ofMDO in the monomers mixtures were synthesized in toluene at80 �C using AIBN (Table 1). The number-average molecular weightsof the remaining polymer chains after degradation of the gels un-der alkaline conditions rapidly decreased with increasing molarfeed ratio of MDO to PEGMA. At 10 mol.% MDO in the monomersmixture, the molecular weight (Mn = 24,000 g mol�1) was far be-low the threshold of 50,000 g mol�1. Above 20 mol.% MDO, it wasnot possible to determine the molecular weight precisely, as theywere similar to or below the permeation limit of the columns.

Microspheres were prepared from different monomer compo-sitions (Table 2). Whatever the formulations used, the resultingMS were all spherical in shape (Fig. 2), with a diameter mainlyin the range 300–500 lm. The FTIR spectra of the hydrolyzablecrosslinker and the PEG–PLGA MS are shown in Fig. 3. Thecharacteristic absorption bands of C = CH2 stretch (1634 and810 cm�1) from the crosslinker disappeared in the PEG–PLGAmicrosphere due to their consumption during the suspensionpolymerization. Meanwhile, the absorption peak of the saturated

Fig. 1. Reaction scheme for the synth

ester groups (1736 cm�1) appeared as a shoulder in the IR spec-trum of the MS. A strong absorbance was seen at 1755 cm�1 inthe crosslinker, indicating the presence of ester stretch in PLAand PGA units. This absorbance remained after PEG-PLGA forma-tion of the MS, indicating the good incorporation of the hydrolyz-able crosslinker. The peak at 1726 cm�1 in the MS spectrum wasattributed to the carbonyl absorption of the saturated estergroups of PEGMA.

3.2. Swelling behaviour

The microsphere swelling ratio was measured in water (pH�6) and PBS (pH 7.4) (Table 2). In both media, the presence of10 mol.% MA increased the swelling ratio to approximately 20%in water and 50% in PBS. In PBS at pH 7.4, the carboxylic groupsof the MA units were more dissociated than in water, resultingin swelling of the MS. Incorporation of MDO into the polymericnetwork induced a slight decrease in the swelling ratio. Despiteits hydrophobic structure, the low level of MDO incorporation(around 7 mol.% according to NMR) did not modify the swellingproperties of the MS.

After the lyophilization and sterilization steps, the MS regainedtheir initial swelling state rapidly (see the figure in the Supplemen-tary data). The diameter of dry MS (259 ± 17 lm) increased to410 ± 26 and 407 ± 23 lm after 5 min of incubation in saline orin the saline:/Omnipaque mixture, respectively. After 3 h at roomtemperature the diameter of the MS in saline did not change(405 ± 27 lm). In the saline/contrast medium mixture, the diame-ter increased slightly (429 ± 28 lm at 30 min) and did not changeup to 3 h (431 ± 29 lm).

3.3. Injectability

All the MS were easily injected through a echelon™ 10 micro-catheter (inner diameter 430 lm). The force needed for manualinjection was similar to that needed for injection of 300–500 lmEmbosphere

�under the same conditions. No aggregation or cathe-

ter blockage was found. The MS recovered their spherical shape atthe exit of the catheter without residual deformation.

esis of PEG–PLGA microspheres.

Table 1Characterization of linear polymers and crosslinked gels containing increasing proportions of MDO.

MDO:PEGMA (mol.%) Crosslinker (mol.%) PEGMA:MDO (1H NMR) Mn (g mol�1) Mn after hydrolysis (g mol�1)a Mw/Mn

Linear 10:90 – 7:93 24 000 14,000 1.6Linear 20:80 – 13:87 28 100 8000 1.7Linear 30:70 – 20:80 29 200 6700 1.6

Gel 0:95 5 ND – 61,000 1.2Gel 10:85 5 ND – 24,000 1.3Gel 20:75 5 ND – <5000 NDGel 30:65 5 ND – <5000 NDGel 40:55 5 ND – <5000 ND

a Molar masses were determined after hydrolysis in 1 M NaOH on residual polymer chains.

Table 2Composition and swelling ratio of degradable microspheres.

MS PEGMA (mol.%) MA (mol.%) MDO (mol.%) Swelling ratio in water (g g�1) Swelling ratio in PBS (g g�1)

PEG–PLGA 95 0 0 4.0 ± 0.2 3.5 ± 0.1PEG–PLGA–MA 85 10 0 5.0 ± 0.2 5.4 ± 0.1PEG–PLGA–MDO 85 0 10 4.2 ± 0.1 3.4 ± 0.1PEG–PLGA–MA–MDO 75 10 10 4.5 ± 0.1 5.1 ± 0.1

Fig. 2. Optical and SEM micrographs of PEG–PLGA–MA degradable microspheres (scale bar for optical microscopy = 200 lm).

Fig. 3. FTIR spectra of (a) PLGA–PEG–PLGA crosslinker, (b) PEGMA polymer and (c) PEG–PLGA degradable microspheres.

S. Louguet et al. / Acta Biomaterialia 10 (2014) 1194–1205 1199

1200 S. Louguet et al. / Acta Biomaterialia 10 (2014) 1194–1205

3.4. Loading of doxorubicin

Loading of the anticancer drug doxorubicin into the degradableMS was carried out by simply mixing the drug in the aqueous MSsuspension at pH 7 (Fig. 4). A high loading efficiency was achievedin less than 1 h for degradable MS containing carboxylic acidgroups. The efficiency was as high as that for the gold standardDCbeadsTM, which contain sulfonic groups. A much lower loadingefficiency was obtained with MS free of carboxylic groups (�40vs. 92% for PEG–PLGA and PEG–PLGA–MA, respectively).

Fig. 5. Degradation profiles of sterile hydrolyzable MS in PBS at 37 �C. Errors barscorrespond to mean ± 1 SD for n = 3.

3.5. Hydrolytic degradation

All MS formulations used were totally degraded within 2 days(100% weight loss; Fig. 5). The degradation rate was higher whenMA was present in the polymer backbone. After 24 h of incubationin PBS, the presence of MA increased the MS weight loss signifi-cantly, to 16.5% (p = 0.0027). The presence of ionized MA increasedthe diffusion of water into the polymer network and thus sped upthe crosslinker hydrolysis and MS degradation. The incorporationof around 7 mol.% MDO units in the MS containing the MA unitsdid not change their degradation profile. Similarly to the swellingbehavior, the relatively low incorporation of this hydrophobicmonomer was compensated by the presence of MA and did notmodify the degradation profile.

The pH drop during degradation experiments for each MS sam-ple was limited, reaching a value of 7–7.1 (see the table in the Sup-plementary data), suggesting that small amounts of acid werereleased from MS during the crosslinker hydrolysis reaction. Thisresult was consistent with the low content of lactic acid and gly-colic acid initially introduced to the MS (2.5 mol.%).

The 1H NMR spectrum of PEG–PLGA MS degradation productsshowed the COO–CH2, PEG, CH2 and CH3 signals of the polymeth-acrylate chains containing mainly PEG as the lateral group (seethe figure in the Supplementary data). The signal of free lactic acidformed after cleavage of the ester bond was observed at 1.1 ppm(the doublet of the CH3 group). The signal of glycolic acid andthe second signal of the lactic acid (CH group) were hidden bythe peak at 4.1 ppm, corresponding to the methylene group ofPEGMA.

To gain further insight into the nature of the polymethacrylatechains, the molecular weights of the polymer fraction of the differ-ent degraded MS were determined by SEC. The number-averagemolecular weight (Mn) of the polymethacrylate chains of MS

0

20

40

60

80

100

Dox

orub

icin

load

ing

(%)

PEG-PLGA PEG-PLGA-MA

PEG-PLGA-MA-MDO

DC Bead®

40±3%N=3

95±2%N=391±4%

N=590±2%

N=4

Fig. 4. Doxorubicin loading efficiency after 1 h, for degradable MS containing, ornot containing, carboxylic acid groups, expressed as a percentage of the initial dose(100% = 35 mg ml�1 of MS), and comparison to DCbeads™.

without MDO was around 50 kg mol�1 (Table 3). The addition ofMA to the MS had no influence on the Mn and polydispersity indexvalues. However, the Mn of the degradation products of MS con-taining MDO was significantly decreased (25–30 kg mol�1).

3.6. In vitro cytotoxicity

3.6.1. Contact cytotoxicityIn this experimental setting, cells grow up to confluence while

MS gradually release their degradation products, as would happenafter implantation. Culture with fragments of latex gloves triggeredcell death while co-culture with non-degradable GMS did not af-fect cell proliferation after 1 week (Fig. 6). A control experimentperformed with cell medium of 50 mM 2-(N-morpholino)ethane-sulfonic acid (MES; pH �6.4) showed that acidity reduced cell pro-liferation. The addition of degradable MS to adherent and sub-confluent cells did not induce cell death. At the end of culture,the MS were fully degraded and the cells had reached confluenceand formed a continuous monolayer with all degradable materials.A protein assay indicated that, compared to cells incubated withGMS, there was only a marginally significant difference betweenPEG–PLGA and PEG–PLGA–MA. The presence of MA in the MS com-position had no effect on cell proliferation for MS prepared with orwithout MDO.

3.6.2. Toxicity of MS degradation productsAfter complete degradation of the MS in culture medium, the

resulting degradation products were added to sub-confluent cells.Extracts obtained from latex gloves were cytotoxic, leading to celldeath, while no cytotoxic compounds were released from GMSafter 3 days (Fig. 7). At the end of culture with extracts obtainedfrom all of the batches of degradable MS (n = 12), continuous cell

Table 3Number-average molecular weight and polydispersity index (Mw/Mn) of the poly-methacrylic chains after degradation of MS in PBS.

MS Mn (kg mol�1) Mw/Mn

PEG–PLGA 48 1.5PEG–PLGA–MA 52 1.4PEG–PLGA–MDO 32 1.4PEG–PLGA–MA–MDO 23 1.5

-AGLP-GEPSMGxetaL0

20

40

60

80

100

120

PEG-PLGA PEG-PLGA-MA MDO

PEG-PLGA-MA-MDO

pH6.4

10.36+/-5.23

83.75+/-8.37

66.42+/-

16.06

77.17+/-

11.26

84.18 +/- 2.14

84.39+/-9.06

44.81+/-

14.29

Cel

lpro

lifér

atio

n (%

of c

ontro

l)

p = 0.0683 (MW)

p = 0.6602 (MW)

*p = 0.0018 p = 0.0367

p = 0.4689 p = 0.3864

p < 0.0001

Fig. 6. Contact cytotoxicity of MS determined after culture with L929 cells for 1 week. Sub-confluent cells in 24-wells plate were cultured in duplicate with 100 ll ofsterilized MS for 7 days before total cell proteins assay. Cell medium at pH 6.4 is obtained after supplementation with 50 mM MES. Comparison with non-degradable GMSwas performed using the Mann–Whitney test. ⁄Comparison with non-degradable GMS. Experiments were conducted at least twice.

0

20

40

60

80

100

120

PEG-PLGA PEG-PLGA-MA

Latex GMS PEG-PLGA-MDO

PEG-PLGA-MA-MDO

pH6.4

21.65 +/-6.08

100.81 +/-3.44

72.28+/-3.04

72.10+/-3.46

78.07+/-1.35

74.80+/-4.02

49.14+/-

15.85

p = 0.5936 (MW)

p = 0.2418 (MW)

Cel

lpro

lifér

atio

n (%

of c

ontro

l)

p < 0.0001 p < 0.0001p = 0.0003

p = 0.0003

Fig. 7. Cytotoxicity of extracts after complete degradation of sterilized MS. Extracts neutralized with 20 mmol of sodium hydrogen carbonate were added to wells (n = 6 or 8)containing non-confluent L929 cells before determination of total cell proteins content after 3 days of culture. Comparison was done using the Mann–Whitney non-parametric test. ⁄Comparison with non-degradable GMS. Experiments were conducted at least twice.

S. Louguet et al. / Acta Biomaterialia 10 (2014) 1194–1205 1201

monolayers were visible. A quantitative protein assay revealedlower cell proliferation in the presence of MS extracts comparedto the negative cytotoxicity controls. This could be explained bythe pH values, which were still lower than that of the control med-ium (pH 7.4) despite neutralization with sodium hydrogen carbon-atehydrogen carbonate. Incorporation of either MA or MDO in theMS had no effect on cell proliferation. In line with the in vitro cyto-toxicity results, MS extracts did not reveal cell death events, but,rather, revealed a reduction in cell growth – probably due to acid-ity, which remain even after the addition of sodium hydrogencarbonate.

3.7. In vivo investigations: subcutaneous implantation in rabbit

To evaluate the in vivo degradation and local biocompatibility,MS were implanted into subcutaneous tissues of rabbits. MS were

placed in 4 cm long rails in order to maximize the tissue–materialinteractions and avoid a depot effect, which could impact the deg-radation time and cell infiltration. Rabbits were sacrificed on day 2to evaluate the tissue response during degradation and on day 7 tocheck the full degradation of the material and the complete healingof the tissue. We examined 594 different skin sections by histology.In the sham group there were only few neutrophils on day 2, and aslight scar reaction with macrophages and fibroblast-like cells onday 7.

On day 2, amounts of degradable MS were present in about one-quarter of the examined slides but absent on day 7 after full deg-radation (Table 4). The inflammatory reaction was limited to someneutrophils in tunnels on day 2, and to a few foamy macrophageswhich were located in the surrounding tissues on days 2 and 7(Fig. 8). Cell types were observed at a density similar to that insham group. No distinction in inflammation could be made

Table 4Frequency of observation of amounts of degradable MS in histological slides.

MS Frequency of observationon day 2

Frequency of observationon day 7

PEG–PLGA 27%(23/85)

0%(0/65)

PEG–PLGA–MA 25%(18/72)

2%(1/63)

PEG–PLGA–MA–MDO

17%(15/88)

0%(0/78)

Sham 0%(0/74)

0%(0/69)

1202 S. Louguet et al. / Acta Biomaterialia 10 (2014) 1194–1205

between groups with or without MA (PEG–PLGA–MA vs. PEG–PLGA) or with or without MDO (PEG–PLGA–MA–MDO vs. PEG–PLGA–MA).

Examination of the draining lymph nodes, liver, kidney, heart,lungs, spleen, brain and bone marrow of the rabbits did not revealany abnormality (necrosis, malformation, inflammation) whichcould have suggested a toxic effect of the degradation products(no difference with sham).

4. Discussion

We have developed a calibrated microsphere made of hydrolyt-ically degradable hydrogel that can be easily injected via micro-catheters to temporarily occlude vessels, then is quicklyeliminated, before the onset of a deleterious chronic inflammatoryresponse. This microsphere can be loaded with drugs for transientchemoembolization procedures. We selected PEG as the main com-ponent since it is used in many biomedical applications (wounddressing, cell encapsulation, drug delivery devices) because of itselasticity, biocompatibility and hydrophilicity, which decreasesprotein adsorption and cell adhesion [24]. To achieve this goal, fourdifferent formulations of MS, with or without an anionic co-mono-mer (MA) and the cyclic ketene acetal MDO, were produced andconsidered in terms of softness, swelling, degradation time andbiocompatibility to optimize the properties of the final materialcontaining all of the constituents.

4.1. Swelling

Swelling properties are particularly important for embolizationand chemoembolization applications since they determine theability of the material to be injected through a catheter. The differ-ent microspheres investigated in this study exhibited water uptake

Fig. 8. Histological aspects of degradable MS 2 days (D2) and 7 days (D7) after implantatiday 2, amounts of degradable MS (stars) were at an advanced stage of degradation, andreaction was mild and comparable to that with the sham injection (Scale: bar = 100 lm

values of the same order of magnitude as those of commerciallyavailable non-degradable microspheres (7.5 g g�1 for Embosphere�

measured under the same conditions as in this study). Incorpora-tion of 10 mol.% MA into the polymer matrices increased the swell-ing ratio of microspheres in PBS (�50%, Table 2). This was due tothe ionization of the carboxylic groups at pH 7.4, which led to anincrease in water uptake by electrostatic repulsion and osmoticpressure difference. In contrast, incorporation of the hydrophobicMDO (7 mol.%) had almost no effect on swelling. Therefore, theswelling properties of the MS were mainly conferred by the ioniza-tion state of MA.

4.2. Injectability

To prevent catheter clumping, a good suspension has to beachieved when dry MS are brought into contact with the aqueousphase, avoiding any residual MS aggregates. After lyophilization,the MS rapidly regained their spherical shape and recovered theirinitial swelling state prior to injection. It is important to note that,after lyophilization and sterilization, MS did not aggregate andcould be easily dispersed at a desired concentration in a matterof minutes without vigorous shaking or sonication.

During the injection test through a microcatheter, the MS didnot meet any resistance, suggesting that the absence of MS aggre-gates and sufficiently compressible behavior of our material couldbe easily achieved. After injection, the MS quickly recovered theirshape without compromising their integrity. The PEG-based mono-mer and the low concentration of hydrolyzable crosslinker create afavorable environment for water uptake while the gel maintainssufficient resistance to let it recover its shape after injectionthrough a catheter.

4.3. Doxorubicin loading

The difference in the loading of MS with or without MA indi-cates that doxorubicin, an amphipathic base (pKa = 8.2), wasloaded by electrostatic interaction with the anionic groups of theMS. The anionic groups used for the drug loading in our degradableMS were provided by carboxylic groups, unlike the sulfonate sidegroups present on the non-degradable ionically loadable DCBeadsTM. To avoid any potential cytotoxicity linked to degradationproducts bearing strong acidic groups, a weak acid group waspreferred. As shown below, the degradation products bearingcarboxylic groups did not generate significant cytotoxicity effects(no cell death in cell culture or tissue necrosis after subcutaneousimplantation in rabbit). Additionally, a high loading efficiency wasrapidly achieved with our degradable MS bearing carboxylic

on in subcutaneous tissues in rabbits, and comparison to saline injection (sham). Onwere fully degraded on day 7, leaving a small scar track (arrow). The inflammatory).

S. Louguet et al. / Acta Biomaterialia 10 (2014) 1194–1205 1203

groups (�90% in less than 1 h), which is comparable to the loadingefficiency of the gold standard.

4.4. Degradation

For all MS formulations, degradation occurred rapidly, in under2 days (Fig. 4), which is essential for temporary embolization pro-cedures. This rapid degradation can be attributed to the small sizeof the lactoyl/glycolidyl segments [25], the relatively low crosslink-ing density and the ability of PEGMA to attract water into the net-work. Similar degradation times were obtained by Sawhney et al.[22] and Lu and Anseth [25] with PEG hydrogels constituted witha sole hydrolyzable crosslinker. To achieve such a short degrada-tion time, authors have synthesized a crosslinker with a PEG hav-ing a molecular weight of at least 3000 g mol�1 in order to increasethe accessibility of water inside the polymer matrix. However,when the ratio of the DP of the PLGA to the DP of the PEG is lessthan 0.45, the PLGA–PEG–PLGA crosslinker becomes soluble inwater [22]. In our study, a smaller PEG of 600 g mol�1 was pre-ferred to allow solubilization of the resulting crosslinker in thehydrocarbon phase. In the absence of a co-monomer, hydrogelsformed by homopolymerization of our sole crosslinker led to a longdegradation time (more than 30 days in PBS at 37 �C). The additionof PEGMA as the main component provides the necessary hydro-philicity to the hydrogel to counterbalance the relative hydropho-bicity of the PLGA–PEG–PLGA crosslinker and reduces thedegradation time considerably.

The presence of MA in the polymer matrices accelerates the deg-radation (+16.5%). As seen above, the infiltration of water moleculesinto the hydrogel is promoted by the presence of anionic groups,especially in PBS at pH 7.4, and, as a consequence, there is greaterwater uptake and quicker hydrolysis of the ester bonds take place.The presence of MDO associated with MA had no impact on the deg-radation profile despite its relative hydrophobicity. Thus, by adjust-ing the ability of the network to attract more water, the degradationtimes were speeded up even after a short period of time.

According to NMR analyses, the degradation products were amixture of low-molecular-weight compounds (lactic acid, glycolicacid and Mw 600 PEG) with a number of linear polymethacrylatechains. This finding is in agreement with the expected degradationmechanism reported by other groups [26–28]; that is, degradationof the crosslinker by hydrolysis of the ester bond within the net-work. This process leads to a progressive decrease in crosslinkingdensity and an increase in swelling until the complete breakdownof the network. Degraded individual repeat units from the degrad-able segments, such as lactic acid or glycolic acid, are metabolizedby the organism are thus considered bioresorbable. However, anexcessive release of acidic degradation products in a short timeperiod can have a deleterious effect on cells in immediate contact.The influence of such products on the local pH of the medium cancause cell death, as reported by Bruining et al. [29]. Compared topure PLGA MS composed with 100 mol.% acidic units, our MS for-mulation contained only 2.5 mol.% lactic and glycolic acids. Thus,the release of acidic degradation products during MS degradationwould be restricted. This was confirmed by the in vitro degradationin PBS (low pH variation) and by the low inflammatory signal ob-served in vivo (see below).

The fate of polymethacrylate chains in the body depends mainlyon their molecular weight. To avoid any accumulation that couldbe toxic over time, the macromolecules must have a hydrodynamicdiameter smaller than the pores of the kidney glomerules (4–14 nm) to be filtered from the blood to urine. This size correspondsvery roughly to linear molecules below a threshold of 50 kg mol�1

[19,20]. With all our MS formulations, the polymethacrylate chainsresulting from the degradation possess an average molecularweight of �50 kg mol�1 (Table 3). Chen et al. [30] have shown that

linear polyacrylic acid polymers grafted with PEG of Mn = 32 and66 kg mol�1 had quite similar elimination times (t1/2 = 22.3 and25.0 h, respectively). A similar trend could be expected accordingto the chemical structure of the degradation products of our MS.Nevertheless, to ensure rapid elimination and limit any possibleinflammatory reaction, supplementary ester linkages were incor-porated within the polymethacrylate backbone to reduce the finalmolecular weight of the degradation products. The ester linkageswere introduced through radical ring-opening copolymerizationof MDO, a cyclic ketene acetal, during the suspension polymeriza-tion process. As already observed by Jin et al. [31], copolymeriza-tion of MDO with PEGMA using AIBN at 80–90 �C led toquantitative ring opening of MDO. Despite the relative hydropho-bicity of this monomer, swelling and degradation properties ofthe resulting MS were approximately equivalent to those withoutMDO. The polymethacrylate chains after hydrolysis showed a re-duced molecular weight of roughly 50% compared to the one with-out MDO. This result confirms the good incorporation of the ring-opened MDO form into the polymer backbone, with a random dis-tribution. For this range of molecular weight, numerous studieshave established that polymers such as PEG, PVA and hydroxypro-pylmethacrylamide were rapidly cleared from the blood and elim-inated by renal filtration after intravenous injection [32,33,19,34].With our MS, the polymethacrylate chains had a molecular weightbelow 30 kg mol�1, so it can be assumed they will be rapidlycleared in the urine after their degradation.

4.5. Biocompatibility

Degradable embolization MS should not release toxic productsor produce adverse reactions upon implantation. Here, the com-plete degradation of the MS allowed us to investigate the cytotox-icity of all of the degradation products, since the whole MS masswas converted into water-soluble degradation products. As the re-lease of protons during hydrolysis of the crosslinker jeopardizesthe cell viability [29], the acidity created in cell co-culture mediumduring MS degradation was buffered with HEPES, while addition ofsodium hydrogen carbonate raised the pH of MS extracts to around7. The pH drop observed in the MS extracts could be raised close toneutrality with only a small amount of sodium hydrogen carbonate(20 mmol), which corroborated the minimal release of acidic deg-radation products during MS degradation. Our results, obtained ata high concentration (10–20 mg ml�1), did not show any signs ofcell lysis, suggesting that the polymethacrylate chains containingMA groups did not disrupt cell membranes [35,36]. The reductionin cell proliferation observed in our experiments is consistent withthe cytotoxicity thresholds reported in the literature [36], and indi-cates good biocompatibility.

The in vivo degradation of the MS in fact appeared to be rela-tively close to that determined in the in vitro setting in PBS. Twodays after their implantation, the MS were at an advanced stageof degradation, and were fully degraded after 7 days, with verylimited scar tracks. Generally, the inflammatory lesions inducedby embolic materials are closely related to their ability to degradeover a long period [6–9,37–40]. It is therefore not surprising that,for all implanted types of MS, the inflammatory reaction was mildand comparable to that of the sham group, and did not persist afterthe material was degraded. The cell types, mainly neutrophils after2 days and macrophages after 1 week, corresponded to the classi-cal acute and progression phases of the foreign body reaction[41]. No lymphocytes or foreign body giant cells were seen forany product at any time. The inflammatory reaction of the degrad-able MS is far less pronounced than those of many other materialsused for embolization. By comparison, the long-standing enzy-matic degradation of agents such as gelatin sponge is accompaniedby a chronic inflammatory response [6–9,42] that can provoke

1204 S. Louguet et al. / Acta Biomaterialia 10 (2014) 1194–1205

arthritis in patients, and compromise the full recovery of the arte-rial lumen despite the material degradation [11,12].

5. Conclusion

To conclude, PEGMA hydrolyzable MS fulfill the requirements ofa degradable MS for embolization as defined in the introduction.They have similar properties (calibrated, swellable, injectable andloadable) as the non-degradable MS currently used today. A rapiddegradation time (less than 1 week) in a subcutaneous site can beobtained, thanks to both the hydrophilic nature of PEGMA and thechemical structure of the hydrolyzable crosslinker, without notice-able toxicity. The combination of a quick degradation time and alow inflammatory profile allow the assumption that the PEGMAhydrolyzable MS will lead to recanalization with a complete recov-ery of the initial lumen. A complementary study in a validatedmodel of embolization has recently demonstrated that this PEG-based hydrolyzable MS allowed controlled devascularization underthe same conditions of use as non-degradable embolics and, moreinterestingly, led to a full angiographic recanalization [43]. There-fore, this novel degradable MS, not loaded with drugs, can reason-ably be proposed for transient arterial embolization. The MS needto be evaluated in vivo after drug loading to prove their potentialas a chemoembolization agent.

Disclosure

This work was funded by the company Occlugel. V.V., L.B., S.L.,E.S., M.W., L.M., D.L. and A.L. are co-inventors of a patent on resorb-able microspheres co-owned by Public Institutions (CNRS, APHP,Universities VII & XI) and by OccluGel Company.

Appendix A. Figures with essential colour discrimination

Certain figures in this article, particularly Figs. 3,5 and 8 are dif-ficult to interpret in black and white. The full colour images can befound in the on-line version, at http://dx.doi.org/10.1016/j.actbio.2013.08.014.

Appendix B. Supplementary data

Supplementary data associated with this article can be found, inthe online version, at http://dx.doi.org/10.1016/j.actbio.2013.11.028.

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