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e u r o p e a n j o u r n a l o f p h a r m a c e u t i c a l s c i e n c e s 3 6 ( 2 0 0 9 ) 493–501

avai lab le at www.sc iencedi rec t .com

journa l homepage: www.e lsev ier .com/ locate /e jps

anoscale mechanical measurement determination of thelass transition temperature of poly(lactic acid)/everolimusoated stents in air and dissolution media

ing Wua, Lothar Kleinerb, Fuh-Wei Tangb, Syed Hossainyb, Martyn C. Daviesa,live J. Robertsa,∗

Laboratory of Biophysics and Surface Analysis, School of Pharmacy, The University of Nottingham, Nottingham NG7 2RD, UKAbbott Vascular, CA, USA

r t i c l e i n f o

rticle history:

eceived 12 August 2008

eceived in revised form

1 November 2008

ccepted 1 December 2008

ublished on line 6 December 2008

a b s t r a c t

Localized atomic force microscopy (AFM) force analysis on poly(lactic acid) (PLA) and

poly(lactic acid)/everolimus coated stents has been performed under ambient conditions.

Similar Young’s modulus were derived from both PLA and PLA/everolimus stent surface,

namely 2.25 ± 0.46 and 2.04 ± 0.39 GPa, respectively, indicating that the drug, everolimus

does not significantly effect the mechanical properties of PLA up to a 1:1 (w/w) drug load-

ing. Temperature controlled force measurements on PLA only coated stents in air and in a

1% Triton surfactant solution allowed the glass transition temperature (Tg) of the polymer

eywords:

FM

olylactic acid

ano-indentation

to be determined. A significant drop of the Young’s modulus in solution was observed at

36 ◦C, suggests that in vivo the Tg of the polymer is below body temperature. The possible

consequences on drug release and the mechanisms by which this may occur are considered.

© 2008 Elsevier B.V. All rights reserved.

Glass transition temperature

1. Introduction

Young’s modulus (E) is used to characterize the elasticresponse of a material to an applied stress. It is an importantproperty to consider for many pharmaceutical processes andapplications. For example with powders, the Young’s modu-lus has an effect on the efficiency of size reduction processsuch as milling and micronisation (Taylor et al., 2004; Zugneret al., 2006; Kwan et al., 2004). It also partially determines thedifficulty of compaction procedures as many pharmaceuti-cal powders are dry pressed into tablet form (Bassam et al.,

990; Roberts and Rowe, 1996; Narayan and Hancock, 2003). Ofarticular relevance here, biocompatible polymers are becom-

ng more widely used in pharmaceutical applications, such asrug delivery (Liu et al., 2005) and tissue engineering (Kong

∗ Corresponding author. Fax: +44 1158467969.E-mail address: clive.roberts@nottingham.ac.uk (C.J. Roberts).

928-0987/$ – see front matter © 2008 Elsevier B.V. All rights reserved.oi:10.1016/j.ejps.2008.12.001

et al., 2006). A reasonable mechanical strength is one of thebasic requirements for successful application, especially invivo (Verma et al., 2008), since such biopolymers are oftenrequired to remain intact and functional within the humanbody for an extended period of time.

Atomic force microscopy (AFM) is the most widely usedscanning probe microscope, producing not only high resolu-tion images of surface morphology but also the ability to applya localized force and record it effectively at the nanoscale(Raghavan et al., 2000; Du et al., 2001). Compared with conven-tional macroscale measurements or so-called nano-indenters

used to determine E, the AFM provides a much lower loadingforce and more importantly, a much higher spatial resolution.This makes it suitable for the characterization of the poly-meric materials such as hydrogels, composites, thin films, etc.

u t

494 e u r o p e a n j o u r n a l o f p h a r m a c e

(Shulha et al., 2004; Clifford and Seah, 2005; Kurokawa et al.,2002; Tsukruk et al., 2002) and importantly the ability to exe-cute in situ measurements. For example, recently, temperaturecontrolled AFM force distance curves have been used to char-acterize the glass transition behaviour of a polymer (Cappellaet al., 2005).

Here we report the measurement of E using AFM onpoly(lactic acid) (PLA) coated stents and PLA/everolimus drugeluting stents (DES). Stents are implantable medical devicesused to treat the narrowing of the coronary artery blood vessel(Barend et al., 2004; Burt and Hunter, 2006; Acharya and Park,2006). Everolimus is an anti-proliferation drug used to over-come restenosis after stent implantation (Waksman, 2002).

2. Materials and methods

2.1. Materials

PLA/everolimus stents with a drug/polymer weight ratio of1:1 and 1:3 and PLA only coated stents were provided byAbbott Vascular (Santa Clara, CA). A reference Young’s modu-lus value for PLA of 1.3–2.8 GPa was provided by Abbott fromin-house macroscale measurements. Tempfix adhesive waspurchased from Agar (Stansted, UK). Trition XL-80N was pur-chased from Sigma (Poole, UK). Tap300 AFM cantilevers withan average spring constant (�) of 40 N/m were purchased fromBudget sensors (Sofia, Bulgaria), DNP cantilevers were pur-chased from Veeco (Santa Barbara, CA) and NSG-20 cantileverswith an average spring constant, � ∼ 48 N/m, were purchasedfrom NT-MDT (Moscow, Russia). A TGT1 tip radius calibrationdiffraction grating was purchased from NT-MDT. Silicon wafersubstrates were purchased from Rockwood (Riddings, UK).

2.2. Experimental methods

The stents were cut, open, flattened and mounted onto sili-con with tempfix. AFM air imaging and force data at ambientconditions of 25 ◦C and 30% relative humidity (RH) wererecorded using a D3100 AFM (Veeco) with Tap300 tappingmode cantilevers. Adhesion measurements were performedunder ambient conditions also using a D3100 AFM, but withDNP cantilevers. Temperature controlled force measurementsin air and liquid were acquired using a Nanoscope IIIa Multi-mode AFM instrument (Veeco) using Tap300 cantilevers andNSG-20 cantilevers, respectively.

In a typical experiment under ambient conditions thespring constant of the cantilever was firstly calibrated usingthe Sader method (Sader et al., 1999; Sader, 1998), whichreduced the errors associated with cantilevers composition(Gibson et al., 2005) to between 5 and 10% (Burnham et al.,2003). Secondly, several tens to several hundreds force curveswere then recorded for individual samples. This was followedby tip self-imaging to determine tip radius using the calibra-tion grating and the calibration of the cantilever sensitivity ona hard silicon surface (Perkins et al., 2007). Each single force

curve was laterally separated by several hundred nanome-ters to ensure that the measurements were independent andwere recorded at the same experimental parameter setting.Tapping mode was used to re-image this position to assess if

i c a l s c i e n c e s 3 6 ( 2 0 0 9 ) 493–501

the surface had responded elastically or if plastic deformationwith permanent indents had taken place. Finally, the tip wasre-imaged with calibration grating in light tapping mode toconfirm the tip had not blunted significantly and to determinean average tip radius during the experiment. AFM measure-ments under ambient conditions were recorded in air on thePLA only and PLA/everolimus coated stents. By applying anappropriate loading force, indentations were able to be con-trolled either within or exceeding the coatings elastic limit.

Experiments with temperature control followed the sameprocedure with the addition that in air, force curves were cap-tured at 25, 50, 60, 70, 80 and 90 ◦C and in liquid at 28 and 36 ◦C.1% Triton XL-80 N was used as the liquid media since it wasused as the release media for the in vitro accelerated releasestudy. Mechanical properties were studied with temperaturecontrol on PLA only coated stents in air and an aqueous dis-solution media environment to provide an insight in to thebehaviour of such coating in vivo.

The Hertz model was used to process the elastic AFM forcedata to derive the Young’s modulus (E) from the force datarecorded within elastic limit. The model describes the elasticdeformation of two homogeneous surfaces under an appliedload and has often used to model AFM data (Perkins et al.,2007; Plassard et al., 2004; Sneddon, 1965; Briscoe et al., 1994).Since the Hertz model is valid only for elastic deformation,any contribution from non-elastic deformation has to be min-imized. To examine the suitability of the Hertz model, forcedata was plotted against sample indentation. Ideally, the forceand indentation would have a linear relationship for a purelyelastic deformation within the entire curve. However, in somecases only parts of the curve followed such a linear relation-ship, these regions were selected as the elastic region and apartial fit of the data made (Perkins et al., 2007).

For the inelastic case where a permanent indent wasformed, the Oliver–Pharr method is an appropriate model tocalculate the Young’s modulus from AFM force–displacementunloading curves (Sneddon, 1965; Oliver and Pharr, 1992).

3. Results and discussion

Fig. 1 shows AFM topography and phase images of (a) a PLAcoated stent, (b) PLA/everolimus coated with a drug polymerweight ratio 1/1 stent and (c) PLA/everolimus coated witha drug polymer weight ratio 1/3 stent. Under ambient con-ditions, no obvious phase separation was observed on thedrug loaded polymer at a 1:1 loading suggesting an intimatemixture of drug and polymer, while a strong phase contrastwas seen at the 1:3 loading. The visualization of the phaseseparation in the sample is due to the oscillating AFM tiplosing different amount of energy to the different materi-als surfaces as a result of changes in materials propertiessuch as stiffness and adhesion (Chen et al., 1998, 2002). Pre-vious work has identified these bright (e.g. nominally higher)areas in the topography image which correspond to the darkareas in the phase image, as being related to drug rich region

(Ranade et al., 2004). Two typical types of force indentationcurves for the PLA only coated stents are shown in Fig. 2.For the data in Fig. 2(a) the maximum loading force was160 nN. Both loading and unloading curves showed a lin-

e u r o p e a n j o u r n a l o f p h a r m a c e u t i c a l s c i e n c e s 3 6 ( 2 0 0 9 ) 493–501 495

Fig. 1 – AFM topography and phase images of PLA/everolimus stents, (a) PLA coated, (b) PLA/everolimus drug polymert r

weight ratio d/p = 1/1, (c) PLA/everolimus drug polymer weigh

40◦ for topography and phase images, respectively.

ear relationship between the force and the indentation withonly a small amount of hysteresis, indicating that only elas-tic deformation occurred. Further AFM air imaging after theforce measurement confirmed this, since no pits on the sur-face were observed (Fig. 2(c)). Such measurements were takenon several different locations and the Young’s modulus was

derived with the Hertz model to be 2.25 ± 0.46 GPa whichis within the reference value range of 1.3–2.8 GPa for PLA.Increasing the maximum applied force to 530 nN, yieldedthe second type of force curve where an obvious difference

atio d/p = 1/3. All images X-range 2 �m, Z-range 10 nm and

between the loading and unloading force indentation curvescould be seen (Fig. 2(b)). Neither of the curves displayed a linearrelationship between force and indentation which suggestedthat plastic deformation had occurred. Again further AFM airimaging by tapping mode was recorded at the same position.Permanent indents were observed on the surface (Fig. 2(d)),

confirming that the applied forces here exceeded the yieldpoint. The initial part of the unloading curves was used tocalculate the Young’s modulus using the Oliver–Pharr model.An average value of 2.04 ± 0.39 GPa was determined, which is

496 e u r o p e a n j o u r n a l o f p h a r m a c e u t i c a l s c i e n c e s 3 6 ( 2 0 0 9 ) 493–501

Fig. 2 – Typical force indentation curves derived from raw force distance curves on the PLA only coated stents. (a)encu

Nano-indentation with elastic deformation only, (b) nano-ind3 �m × 3 �m, Z-range 25 nm after acquiring the force distance

consistent with the modulus determined using only elasticdeformation.

The processed data calculated by either model can also bedisplayed as histogram distribution (Fig. 3). As the measure-ment was performed on the homogeneous pure PLA surface,a mono-model distribution should be expected. Fig. 3 showsthe histogram distributions of the data derived from the Hertzmodel (a) and the Oliver–Pharr method (b). In both casesthe histogram distribution approximately fits a mono-modeldistribution suggesting that the measurements with bothapproaches are reliable.

Since an excessive adhesion force can influence the mea-surement of the Young’s modulus of soft polymers (Gupta etal., 2007), the measurement of adhesion was also performedhere. A mean value of 19.8 ± 0.8 nN for PLA was determined.This is negligible compared to the force applied in the mea-surement of the Young’s modulus, 160 and 530 nN, indicatinga limited effect of the adhesion in this case.

Measurements performed on the drug loadedPLA/everolimus stent with drug/polymer ratio 1:1 alsoshowed a homogeneous surface. The average Young’s mod-ulus calculated with the Hertz model and the Oliver–Pharr

tation with the permanent indent, AFM Imagesrves without (c) and with (d) permanent indents.

method from several different locations was 1.97 ± 0.39 and1.87 ± 0.70 GPa, respectively. These results are again close toeach other, further confirming the reliability of the measure-ment with both methods. Although the Young’s modulus ofthe PLA/everolimus coating is marginally smaller than thepure PLA coating, the results are still within the referencerange of the Young’s modulus value of PLA and are notsignificantly different.

AFM data in Fig. 4 again shows that the drug polymerweight ratio 1:3 coating displays a phase separated surfacemorphology. This data also shows the same area after forcemeasurement without (a) or with (b) permanent indents. Theadvantage of having permanent indents on the surface is thatthe exact locations of where the force curves were recordedcould be seen. This is helpful for the analysis of the hetero-geneous phase separated surfaces especially when the phaseseparation is on the nanoscale, since there is always someunavoidable system drift of the AFM scanner or the sample,

for example a small drift could be observed between Fig. 4(aand b).

By choosing six different positions in either dark or brightareas in Fig. 4(b), the average value of Young’s modulus

e u r o p e a n j o u r n a l o f p h a r m a c e u t i c a l s c i e n c e s 3 6 ( 2 0 0 9 ) 493–501 497

ea

Fig. 3 – The histogram distribution of the Young’s modulus mdeformation only (a) and plastic deformation (b).

derived was 2.16 and 2.14 GPa, respectively, indicating a sim-ilar Young’s modulus of the dark matrix and the brightdomains. The overall mean values of the Young’s modulusfrom all measurements were 1.95 ± 0.32 and 2.38 ± 0.53 GPausing the Hertz model and Oliver–Pharr method, respectively.Histogram distributions of the processed data are displayed inFig. 5. Consistent with the lack of a Young’s modulus differencebetween the matrix and domains a mono-model distributionwas again observed. This implies that either the drug/polymerregions have a similar Young’s modulus to the polymer or thatthe drug rich regions (bright domains) are too thin on thispolymer dominated system to be distinguished. If the drugrich regions were very thin (<50 nm) then the substrate matrixbeneath would possibly have influence on the force measure-ment (Soifer et al., 2004).

Due to the absence of the difference of Young’s modulusbetween domains and matrix, it is likely that adhesion dif-ference of the AFM probe on the domain and matrix areas

is the main reason for the existence of the observed phasecontrast. The histogram distributions of the data of the max-imum adhesion force on PLA/everolimus coated stents withhomogeneous (d/p weight ratio 1/1) or phase separated sur-

Fig. 4 – AFM 5 �m × 5 �m tapping mode air images of PLA/everoacquiring the force distance curves with (b) and without (a) the

surement on the PLA only coated stent by elastic

face (d/p weight ratio 1/3) are shown in Fig. 6. A biphasicdistribution was observed on the phase separated surfacewith the centre of the peaks at 23 and 24 nN, while a mono-model distribution was observed on the homogeneous surfacewith the centre of the peak at 26.4 nN. This suggests that theadhesion differences do exist between domains and matrixleading to the visualization of the phase separation in AFMphase images, and that the drug rich areas are more adhe-sive than the polymer alone, since with the increase of thedrug loading, adhesion force increased. The dark region in thephase image corresponded with a greater energy loss of thecantilever (Chen et al., 2002) and is consistent with a higheradhesion.

Since a material’s mechanical properties change througha glass transition, temperature controlled AFM force mea-surements offer a means to locally study this phenomenaat surfaces. Several studies of determination of Tg throughforce displacement curves have been performed by AFM

(Cappella et al., 2005; Bliznyuk et al., 2002; Hinz et al., 2004).Compared to differential scanning calorimetry (DSC) or scan-ning thermal microscopy (SThM), extensively used thermalmethods for the bulk or localized Tg measurement, the E

limus stents with a drug polymer weight ration 1/3 afterpermanent indents at the same location, Z-range 15 nm.

498 e u r o p e a n j o u r n a l o f p h a r m a c e u t i c a l s c i e n c e s 3 6 ( 2 0 0 9 ) 493–501

Fig. 5 – The histogram distribution of the Young’s modulus measurement on the PLA/everolimus stents with a drugpolymer weight ration 1/3 by (a) elastic deformation only and (b) plastic deformation.

uon

observed at the elevated temperature (Fig. 7). The polymer

Fig. 6 – The histogram distribution of measurement of maximpolymer weight ratio (a) 1/1 and (b) 1/3 at the ambient conditi

measurement with AFM is taken at a steady equilibratedtemperature with an effective zero heating rate. Previously,heating rate has been noted as an important factor effect-ing the quantification of Tg in DSC or SThM, with in general

a higher value being estimated for a given Tg with increas-ing heating rate (Mathot, 1994). Here the initial parts ofthe loading curves on PLA stents were used for the mod-ulus calculation because of the significant increase of the

Fig. 7 – Young’s modulus versus temperature curve on PLA onlystandard deviation were shown in the curve.

m adhesion force on the PLA/everolimus stents with a drugs of 25 ◦C and 30% relative humidity.

adhesion force due to softening of the polymer with theincrease of temperature (Tsui et al., 2000; Cappella and Stark,2006). A significant drop of the Young’s modulus has been

lost around 80% of its stiffness at 90 ◦C compared to ambienttemperature. The onset of Tg of PLA in air derived from thismodulus versus temperature approach was approximately55 ◦C.

coated stents. The average value (30 data points) and

e u r o p e a n j o u r n a l o f p h a r m a c e u t i c

Fa

uctfds2ti0ra5wgtpmtT(liTsi2mtwwc

ig. 8 – The AFM force distance curves on PLA coated stentst (a) 28 ◦C and (b) 36 ◦C in 1% Triton XL-80N.

We have extended this approach to measure Young’s mod-lus in release media. Two representative force–displacementurves recorded at 28 and 36 ◦C are shown in Fig. 8. At bothemperatures the loading and unloading curves followed dif-erent paths, suggesting that plastic deformation occurreduring the measurements. As there was very limited adhe-ion in aqueous conditions (no capillary force (Leite et al.,003)), the initial part of the unloading curves were fitted byhe Oliver–Pharr approach. The Young’s modulus of PLA coat-ng at the temperatures of 28 and 36 ◦C was 1.73 ± 0.78 and.78 ± 0.37 GPa, respectively. The modulus acquired in liquid atoom temperature is quite close to the value in air. However, inir the polymer did not start to lose its stiffness until around5 ◦C, while in liquid a significant drop was found by 36 ◦Chere the modulus was similar to that at 80 ◦C in air. Hence the

lass transition temperature of the polymer was lowered byhe presence of the aqueous environment to below body tem-erature. On returning the temperature to 28 ◦C, the Young’sodulus recovered to 1.64 ± 0.83 GPa (curve not shown). Both

he presence of surfactants similar to Triton XL-80N, such asriton X100 (Bouissou et al., 2006a) and an uptake of waterBlasi et al., 2005) can cause polymer plasticization and hence aower Tg. The data here directly confirms this assertion, show-ng that in vivo the PLA is likely to be above its glass transition.his could influence the nature and degree of drug release,ince above Tg the polymer would expand and its chain mobil-ty would increase (Yamamoto et al., 2002; Steendam et al.,001). This change could lead to an increase in drug per-eability in the coating (Fujimori et al., 2005) and facilitate

he diffusion of both the releasing drug and the infiltratingater (Bouissou et al., 2006b; Friess and Schlapp, 2002). Suchater uptake would further plasticize the PLA and potentially

ause macro erosion at the end for the bioerodable polymer

a l s c i e n c e s 3 6 ( 2 0 0 9 ) 493–501 499

system. The drug release would hence enter the phase ofdegradation based release from the initial phase of diffusionbased release, which would greatly increase the drug releaserate.

4. Conclusions

AFM force–displacement curves have been used to calculatethe Young’s modulus on PLA stents and PLA/everolimus drugeluting stents (1:1 and 1:3 (w/w), drug to polymer ratio). Withcareful control of the applied force, the surface deformationcould be successfully controlled within or exceeding the elas-tic limit and the Young’s modulus was able to be calculatedfrom either data type by using the Hertz model or Oliver–Pharrmethod. The PLA/everolimus coated stent showed a simi-lar modulus to the PLA only coated stent, indicating thateverolimus does not significantly effect the mechanical prop-erties of PLA up to a 1:1 (w/w) drug loading. The visualizationof the phase contrast is due to the difference in the adhesionforce of PLA and everolimus. Through the Young’s modu-lus versus temperature curve in air, the onset Tg of the PLAwas determined to be around 55 ◦C. Young’s modulus mea-surement in dissolution media on the PLA only coated stentrevealed that the Tg of PLA in vivo is likely to be lower thanbody temperature. This will greatly affect the polymer struc-ture, potentially increasing the diffusion of the releasing drugand the uptake water, and hence lead to an accelerated drugrelease.

Acknowledgements

The authors would like to thank Prof. X. Chen from LBSA fordiscussion and programming the Young’s modulus algorithmsand thank Abbott Vascular and the University of Nottinghamfor providing Ming Wu with a studentship.

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