Block Copolymer Modified Surfaces for Conjugation of

Dec 24, 2012 - Polymer brush layers based on block copolymers of poly(oligo(ethylene glycol) methacrylate) (POEGMA) and poly(glycidyl methacrylate) ...
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Block Copolymer Modified Surfaces for Conjugation of Biomacromolecules with Control of Quantity and Activity Xin Li, Mengmeng Wang, Lei Wang, Xiujuan Shi, Yajun Xu, Bo Song, and Hong Chen* Jiangsu Key Laboratory of Advanced Functional Polymer Design and Application, Department of Polymer Science and Engineering, College of Chemistry, Chemical Engineering and Materials Science, Soochow University, 199 Ren’ai Road, Suzhou, 215123, P. R. China ABSTRACT: Polymer brush layers based on block copolymers of poly(oligo(ethylene glycol) methacrylate) (POEGMA) and poly(glycidyl methacrylate) (PGMA) were formed on silicon wafers by activators generated by electron transfer atom transfer radical polymerization (AGET ATRP). Different types of biomolecule can be conjugated to these brush layers by reaction of PGMA epoxide groups with amino groups in the biomolecule, while POEGMA, which resists nonspecific protein adsorption, provides an antifouling environment. Surfaces were characterized by water contact angle, ellipsometry, and Fourier transform infrared spectroscopy (FTIR) to confirm the modification reactions. Phase segregation of the copolymer blocks in the layers was observed by AFM. The effect of surface properties on protein conjugation was investigated using radiolabeling methods. It was shown that surfaces with POEGMA layers were protein resistant, while the quantity of protein conjugated to the diblock copolymer modified surfaces increased with increasing PGMA layer thickness. The activity of lysozyme conjugated on the surface could also be controlled by varying the thickness of the copolymer layer. When biotin was conjugated to the block copolymer grafts, the surface remained resistant to nonspecific protein adsorption but showed specific binding of avidin. These properties, that is, well-controlled quantity and activity of conjugated biomolecules and specificity of interaction with target biomolecules may be exploited for the improvement of signal-to-noise ratio in sensor applications. More generally, such surfaces may be useful as biological recognition elements of high specificity for functional biomaterials.

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INTRODUCTION Surface modification is an effective method to improve the properties of biomaterials.1−5 The use of “living” polymerization techniques to tether polymer brushes on surfaces with accurate control over layer thickness, composition, and architecture is an effective method of modifying surface properties. The conjugation of biomolecules to the polymer chains can be easily achieved through functional groups, providing a convenient method for the preparation of functional biomaterials. In general, for biocompatibility, nonspecific interactions should be inhibited and specific interactions promoted.6−9 Strategies for the design of such bioactive surfaces include side chain or end group modification of antifouling polymers10−12 and random/block copolymer modification with antifouling and bioconjugating properties.13−15 Biomolecules can be easily “attached” to surfaces via noncovalent interactions, such as hydrophobic interactions,16 electrostatic interactions,17,18 and metal-ion affinity interactions.19 However, such noncovalent interactions are relatively weak. Therefore, covalent biomolecule attachment is a more attractive strategy. N-Hydroxysuccinimide (NHS) ester is the most commonly used linker to make conjugates of primary amine-containing bioactive molecules,20 since the reactions occur under mild conditions. A drawback of the NHS ester method is that under aqueous conditions rapid hydrolysis can © 2012 American Chemical Society

compete with biomolecule coupling. As an alternative, poly(glycidyl methacrylate) (PGMA) has been used for bioactive surface preparation because of the ease of conversion of epoxide groups to a variety of functional groups such as −NH2 and −COOH through a ring-opening reaction.21 Sung et al. synthesized an epoxide-containing random copolymer containing polyethylene glycol groups for prevention of nonspecific protein adsorption, epoxide groups for conjugation of biomolecules, and different groups for surface attachment with various substrates.22 Using these epoxide-containing multifunctional copolymers, various types of biomolecules could be easily anchored to different substrate surfaces. In other work by Li and co-workers, random POEGMA-co-GMA polymer brushes were incorporated into different bioassay systems and improved signal-to-noise ratios were demonstrated.23,24 However, control of the quantity and activity of conjugated biomolecules by varying each component of the functional copolymer has received little attention. With those considerations in mind, we have synthesized POEGMA-b-PGMA diblock copolymer modified bioactive surfaces using surface-initiated activators generated by electron transfer atom transfer radical polymerization (SI-AGET ATRP) Received: November 7, 2012 Revised: December 20, 2012 Published: December 24, 2012 1122

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performed to assess strength of binding. The samples with adsorbed radiolabeled protein were incubated in a 2% solution of SDS for 3 h at room temperature. The protein remaining on the surface was then measured. Assay of Lysozyme Activity. The activity of desorbed lysozyme was measured as described previously.28 The samples were incubated with lysozyme in PBS (pH 7.4) at room temperature for 3 h with the final concentration of lysozyme at 1 mg/mL. The samples were then washed in PBS and immersed in the fluorescent substrate EnzChek (Lysozyme Assay Kit, Invitrogen). The reaction was allowed to proceed at 37 °C for 0.5 h, and the products of the fluorescent substrate were measured at Ex495/Em525. Preparation of Biotinylated Surfaces. Surface immobilization of biotin hydrazide via covalent amide bonds was carried out as previously described.15 Briefly, the surfaces were immersed in an absolute ethanol solution of biotin-NH2 (1 mg/mL) for 24 h at room temperature, washed with ethanol, and dried in an argon stream. To deactivate the remaining epoxy groups, the surfaces were immersed in the ethanolic solution of (EG)2NH2 (0.1 mg/mL) for 3 h at room temperature, rinsed with ethanol, and dried in an argon stream. The resulting surfaces were incubated in a solution containing FITC-avidin (0.1 mg/mL) or FITC-lysozyme (0.1 mg/mL) for 3 h at room temperature. They were then thoroughly rinsed with the reaction solvent and dried in an argon stream. The surfaces were examined for adsorbed avidin and lysozyme by fluorescence microscopy (IX71, Olympus, Japan).

method. As an oxygen-tolerant living radical polymerization process, AGET ATRP allows precise control of the length of each block in the copolymer and thus of the copolymer layer thickness, thus giving control over the bioconjugation reactions. The immobilization properties of the resulting polymer brush surfaces and the activity of the immobilized biomolecules were systematically investigated in this work.

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EXPERIMENTAL SECTION

Materials. (100)-Oriented single crystal silicon wafers were purchased from Guangzhou Semiconductor Materials (Guangzhou, China). The as-received wafers were polished on one side and diced into square chips of 0.5 cm × 0.5 cm in size. Oligo(ethylene glycol) methacrylate (OEGMA, Mn = 475 g/mol, Aldrich) was distilled over CaH2 under vacuum prior to use. Glycidyl methacrylate (GMA, Aldrich), 2-(2-aminoethoxy) ethanol ((EG)2NH2, Aldrich), biotin hydrazide (Biotin-NH 2 , Sigma), 3-aminopropyltriethoxysilane (APTES, Aldrich), bromoisobutyryl bromide (BIBB, Aldrich), and 2,2′-bipyridine (Bpy, Aldrich) were used as received. Ascorbic acid (AscA) and copper(II) chloride (CuCl2) were obtained from China National Medicines Corporation Ltd. and used as received. Fluorescein isothiocyanate-labeled avidin (FITC-avidin) was purchased from Wuhan Boster Biological Technology, Ltd. Fibrinogen (MW = 341 kDa, pI = 5.5) was purchased from Calbiochem (La Jolla, CA), and lysozyme (MW = 14.7 kDa, pI = 12) was obtained from Sigma and labeled with FITC.25 All other solvents were purchased from Shanghai Chemical Reagent Co. and purified according to standard methods before use. Deionized (DI) water was purified using a Millipore water purification system to give a minimum resistivity of 18.2 MΩ·cm and used in all experiments. Preparation of POEGMA-b-PGMA Grafted Silicon Wafers. The pretreatment of silicon wafers for immobilization of initiator followed procedures reported in our previous work.26 SI-ATRP grafting of OEGMA was carried out in a glovebox purged with dry N2 gas. OEGMA (4.75g, 10 mmol), Bpy (25 mg, 0.16 mmol), CuCl2 (13.6 mg, 0.08 mmol), and AscA (14.1 mg, 0.08 mmol) were dissolved in a 1:1 mixture of methanol and water (20 mL), and the resulting solution was deoxygenated by purging with argon for 30 min at room temperature. The solution was added to a glass vessel containing the initiator-functionalized Si wafers. The polymerization was carried out at room temperature under nitrogen. After desired periods, the obtained POEGMA grafted silicon wafers were removed from solution, cleaned ultrasonically in water and methanol, then rinsed thoroughly, and dried under an argon flow. The SI-ATRP of GMA from grafted POEGMA was carried out the same as for OEGMA. The resulting surfaces were cleaned ultrasonically in methanol and dichloromethane and finally dried under an argon flow. Surface Characterization. Static water contact angles of the pristine and functionalized silicon surfaces were measured using the sessile drop method (C201 optical contact angle meter, Solon Information Technology Co., Ltd.). The thickness of the polymer grafts on the silicon substrate was measured using an M-88 spectroscopic ellipsometer (J.A. Woollam Co., Inc.). Fourier transform infrared spectroscopy (FTIR, Nicolet 6700) gave an indication of the variation in chemical composition of the modified silicon surfaces. Multiple (1024) scans were acquired for each sample. A new background was acquired before each scan to account for small changes in the atmospheric composition. The topography of the surfaces was studied using a Nanoscope V atomic force microscope (AFM, Bruker). To investigate the surface morphologies, the samples were immersed in DI water for 3 h and then dried under an argon flow prior to AFM examination. Protein Conjugation. Fibrinogen and lysozyme were dissolved in phosphate buffered saline (PBS, pH 7.4) and radiolabeled with 125I using the iodine monochloride (ICl) technique.27 For studies of protein adsorption from buffer, labeled and unlabeled proteins were mixed (1/9, labeled/unlabeled) to give a total concentration of 1 mg/ mL. Adsorption was allowed to proceed for 3 h under static conditions at room temperature. Elution with sodium dodecyl sulfate (SDS) was

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RESULTS AND DISCUSSION Formation of POEGMA-b-PGMA Diblock Brushes on Flat Silicon Substrates via Consecutive Surface-Initiated AGET ATRP. POEGMA has been widely used as an antifouling surface modifier and can be surface-grafted in a well controlled manner using SI-ATRP.29−32 In this work, POEGMA was used as the first block of the diblock POEGMA-PGMA grafts. It then served as macroinitiator for the subsequent SI-AGET ATRP of GMA. The synthesis procedure for the formation of POEGMAb-PGMA copolymer brushes on silicon using SI-AGET ATRP is illustrated in Scheme 1. Scheme 1. Synthesis Procedure for POEGMA-b-PGMA Copolymer Brushes on Silicon Wafer via SI-AGET ATRP

In the following discussion, Si-POEGMA10nm refers to surfaces with a POEGMA layer thickness of 10 nm, and bPGMA3nm refers to surfaces with a PGMA block thickness of 3 nm initiated from Si-POEGMA10nm. The thickness of the polymer grafts (determined by ellipsometry) as a function of polymerization time is shown in Figure 1. It is seen that at 1 h the POEGMA layer thickness was 4.7 nm, and increased in a linear manner at a rate of about 4.5 nm/h to 25.5 nm after 6 h. Good control of thickness is thus possible. The increase of grafted layer thickness slowed after 6 h; the most likely cause is the loss of terminal C−Br bonds, presumably through 1123

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Figure 2. FTIR-spectra of (a) Si-POEGMA10nm and (b) SiPOEGMA-b-PGMA13nm surfaces.

Figure 1. Dependence of thickness of grafted POEGMA layer (blue squares) and PGMA layer initiated from Si-POEGMA10 nm (red triangles) on polymerization time.

suggest that surface modification with the PGMA block polymer was successful. Surface Topography. Surface morphology was investigated by AFM following immersion in DI water for 3 h. The main focus of the work is the effect of chemical composition, therefore only phase images were examined in detail. Different surface morphologies can be achieved by varying the block length in this type of surface.18,40,41 Because of the difference in wettability of POEGMA and PGMA, the surface morphology evolved as the thickness of the PGMA layer increased. As shown in Figure 3, the homopolymer modified surfaces Si-POEGMA10nm (Figure 3a) and Si-PGMA18nm (Figure 3f) were flat and uniform, indicating uniform coverage of polymer. Figure 3b−e shows the morphologies of block copolymer modified surfaces with different PGMA thickness. When the PGMA layer was less than 10 nm in thickness (Figure 3b,c), spherical aggregate patterns were formed, and the sphere diameter increased with increasing PGMA length. Similar patterns have been observed for other block copolymer modified surfaces.18,40 It is believed that this morphology results from the wettability difference of POEGMA and PGMA: the short hydrophobic PGMA chains aggregate to form spherical domains in water. At a PGMA layer thickness of 12.8 nm (Figure 3d), larger, wormlike forms appeared probably due to aggregation of the smaller PGMA spheres, and surface coverage of PGMA increased. When the PGMA layer was much thicker than the POEGMA (Figure 3e), PGMA coverage was almost complete and the morphology was similar to that of the PGMA homopolymer. For the thickest PGMA layer (Figure 3f), the surface appeared identical to the PGMA homopolymer. A cartoon of the surface morphology showing the arrangement of the chains at several block thicknesses is shown in Figure 4 to illustrate the morphology changes for different layer thickness. Different surface morphologies resulting from varying copolymer composition are expected to affect surface properties and hence biomolecule-surface interactions. As shown in this work, surface morphology could easily be adjusted by controlling the thickness of each block via polymerization time. Protein Conjugation. As a major plasma protein, fibrinogen is often chosen to evaluate resistance to protein adsorption.15,42 In this work it was used as a model protein to investigate the effect of polymer layer thickness on protein adsorption to POEGMA homopolymer modified surfaces.

termination reactions or due to increased steric interference in chain growth for longer polymer brushes as propagating chains becoming buried in the polymer brush.30,33,34 Meanwhile, changes in the conformation of the grafts (relaxation) can also lead to a breakdown in the thickness-chain length relationship. Subsequent ATRP of the second (PGMA) block was carried out on Si-POEGMA10nm surfaces (2 h OEGMA polymerization). The POEGMA-grafted silicon wafers were exposed to GMA solution, and samples were removed at different times. The thickness of the diblock layers was determined by ellipsometry (Figure 1, red triangles). The increase in thickness due to PGMA was 3 nm at 30 min and 15.8 nm 2.5 h, with a linear chain growth rate of about 6 nm/h. As for the POEGMA, the PGMA thickness could thus be closely controlled by varying the polymerization time. It should also be emphasized that AGET ATRP does not require an oxygen-free environment. It is thus a convenient method which provides a means for control of grafted layer thickness. Surface Characterization. The wettability of the surfaces at different stages of the modification process was investigated by static water contact angle measurement. The WCA of the unmodified surface was 73°, decreasing to 40° after POEGMA modification, confirming the formation of a POEGMA layer. The WCA of the diblock polymer modified surfaces increased with PGMA polymerization time, reaching a value of 59° when the increase in layer thickness was 10 nm, and then remaining constant. The WCA of a PGMA homopolymer modified surface, Si-PGMA, also showed a contact angle of 59°, in agreement with values reported in the literature.35 These data suggest that the PGMA segments of the block copolymers dominate the surface at PGMA block thickness ≥ 10 nm. FTIR spectra of the surfaces are shown in Figure 2. For SiPOEGMA10nm, strong bands at 1728 and 1108 cm−1 assigned to carbonyl groups and C−O−C groups on POEGMA chains respectively11,36,37 are observed. The band at 2881 cm−1 is assigned to the C−H vibrations in OEGMA475 monomer. These data demonstrate that OEGMA graft polymerization was successful. The spectrum of Si-POEGMA-b-PGMA13nm surface (Figure 2b) shows strong bands at 1730 and 1149 cm−1 assigned to ester groups the same as for POEGMA. A new band at 1254 cm−1 is assigned to symmetrical stretching vibrations of epoxy groups. The band at 3000 cm−1 (C−H vibrations) is also due to epoxy groups.38,39 These results 1124

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Figure 3. AFM images of surfaces after water treatment.

Figure 4. Surface morphology at different block layer thicknesses.

Figure 5. (a) Adsorption of fibrinogen from 1 mg/mL solution in PBS; (b) conjugation and subsequent elution with 2 wt % SDS; (c) conjugation vs PGMA/POEGMA block thickness ratio. Data are means ± standard error (n = 3). 1125

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eluted from the Si−OH and Si−POEGMA surfaces by SDS, but much less (∼30%) from the block copolymer modified surfaces, suggesting that the proteins were bound via relatively strong interactions, possibly covalent bonds between lysine residues in the proteins and epoxy groups on the surfaces. Surfaces with different PGMA/POEGMA thickness ratio were used to further investigate protein conjugation. As shown in Figure 5c, fibrinogen conjugation increased with increasing PGMA/POEGMA thickness ratio, and changed only slightly at ratios near 1.0, probably because the surface morphology tended to be unchanged in that case, resulting in little influence on the conjugation of protein. These results demonstrate that protein conjugation on POEGMA-b-PGMA modified surfaces is strongly dependent on the epoxide content available, which is determined by the thickness of each layer and the surface morphology of the copolymer. Variation of the relative thickness of the layers can therefore be used to control biomolecular interactions. Assay of Lysozyme Activity. For a bioactive surface, activity as well as quantity of biomolecules on the surface needs to be considered. To investigate this aspect, the enzymatic activity of lysozyme conjugated on the surfaces was measured. Although the amount of lysozyme adsorbed on the SiPOEGMA10nm surface was low at 0.046 μg/cm2, the activity was substantial at 0.94 U/cm2 (Table 1). The activity of lysozyme conjugated on POEGMA-b-PGMA modified surfaces with a base of Si-POEGMA10nm increased with PGMA layer thickness to a maximum value of 1.78 U/cm2 on Si-bPGMA13nm surface. The amount conjugated on the SiPGMA surface was 8-fold higher than that on the SiPOEGMA10nm surface, but with a lower activity of only 0.86 U/cm2, indicating further that as a hydrophilic polymer layer POEGMA contributes to the conservation of activity of the conjugated enzyme. The specific activity of lysozyme in solution was 33.31 U/μg. Of the different surfaces, the b-PGMA3-nm showed the highest specific activity at 31.31 U/μg, that is, essentially the same as that of the enzyme in solution. The specific activity was 7.37 U/ μg for a PGMA block thickness of 13 nm. The PGMA homopolymer surface showed the lowest specific activity of only 1.81 U/μg, indicating almost total loss of activity presumably due to conformational change and/or inaccessibility of active sites. The specific activity of lysozyme conjugated on the POEGMA-b-PGMA modified surfaces was greater than that on the PGMA modified surfaces of similar layer thickness; it seems likely that the hydrophilic environment provided by POEGMA keeps the conjugated enzyme in the

Adsorption from a 1 mg/mL solution in PBS was measured at room temperature. As shown in Figure 5a, adsorption was reduced significantly on the surfaces modified with POEGMA, decreasing with increasing layer thickness to a value of 0.1 μg/ cm2 at 10 nm and then remaining constant. Similar behavior has been observed for poly(N-vinylpyrrolidone)-modified silicon with minimum adsorption at a layer thickness of 13 nm.26 These results indicate that fibrinogen adsorption is at a minimum for a 10 nm thick POEGMA layer. Therefore, surfaces based on Si-POEGMA10 nm with different PGMA thickness were used to investigate the protein conjugation properties of the block copolymer modified surfaces. Figure 5b shows fibrinogen conjugation on POEGMA-blockPGMA modified surfaces. It is seen that the surfaces modified with POEGMA were protein resistant. With the introduction of epoxy groups, conjugation increased from 0.23 to 0.76 μg/cm2 as PGMA block thickness increased from 3 to 13 nm. Lysozyme, a much smaller protein, showed similar behavior (see Table 1). Meanwhile, even lysozyme conjugation on Table 1. Activity of Lysozyme Conjugated on Surfaces at Room Temperaturea conjugated lysozyme (μg/cm2)

total activity (U/cm2)

specific activity (U/μg)

lysozyme in solution

0.200 μg

6.66 U

33.31

PO10nm b-PGMA3nm b-PGMA13nm Si-PGMA2nm

0.046 0.050 0.242 0.373

0.94 1.57 1.78 0.86

20.61 31.31 7.37 1.81

sample

a

Data are means of three samples.

POEGMA-b-PGMA13 nm surfaces was much lower than on PGMA homopolymer modified surfaces when the thickness of PGMA was only 2 nm. Clearly the POEGMA sublayer modulates the effect of the PGMA. From the AFM images in Figure 3, it is seen that the surface morphology evolves as the thickness of the layers changes. With Si-POEGMA10 nm as a base, the surface area of the PGMA aggregates increased with increasing PGMA layer thickness, and the POEGMA was gradually covered by PGMA, resulting in an increase of epoxide available to conjugate proteins and the loss of protein resistance. It can be inferred that the PGMA layer would cover the entire surface if the PGMA block was thick enough (Figure 3e), in which case the effect of the POEGMA layer would be greatly reduced. In addition, most of the protein was

Figure 6. Fluorescence microscopy images of surfaces after exposure to FITC-avidin (a−e) or FITC-lysozyme (f−j) for 3 h: (a, f) Si-PGMA50nmbiotin; (b, g) Si-POEGMA10nm; (c, h) b-PGMA3nm-biotin; (d, i) b-PGMA6nm-biotin; (e, j) b-PGMA13nm-biotin. (Scale bar in all images = 100 μm.) 1126

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native conformation.43−45 From the water contact angle and AFM data, it appears that the block copolymer modified surfaces became more and more PGMA-like as the PGMA layer thickness increased; the POEGMA layer was gradually covered by PGMA, resulting in a decrease of surface hydrophilicity and specific activity of the conjugated lysozyme. From these data, it appears that the amount and activity of enzyme on these surfaces can be regulated by varying the relative thickness of the blocks. It is possible that this conclusion could be extended to the conjugation and bioactivity of biomolecules more generally. Biotin Conjugation for Specific Binding of Avidin. It is well-known that the specific interaction between biotin and avidin is the strongest known noncovalent interaction between biomolecules. The biotin−avidin system is therefore useful for investigations of biospecificity.46,47 In this regard, we prepared biotin-conjugated surfaces and measured avidin binding using a fluorimetric assay. Lysozyme adsorption was measured as a control. Samples were incubated in PBS containing FITCavidin or FITC-lysozyme, and the surfaces were examined with fluorescence microscopy (Figure 6). Figure 6a and f shows the images for PGMA homopolymer− biotin surface. Both avidin and lysozyme fluorescence was strong, with the latter indicating nonspecific binding. The SiPOEGMA surface showed no fluorescence (Figure 6b and g), presumably due to the protein resistance of POEGMA. Resistance to FITC-lysozyme was also shown by the various block polymer surfaces (Figure 6h−j). For the corresponding FITC-biotin modified surfaces, fluorescence increased with increasing PGMA layer thickness (Figure 6c−e), with the bPGMA13nm-biotin surface showing the strongest fluorescence. This behavior resulted, presumably, from the increasing biotin content and decreasing protein resistance of the POEGMA layer as the PGMA thickness increased. This result clearly indicates that the biotin-conjugated SiPOEGMA-b-PGMA surface is resistant to nonspecific protein adsorption and, at the same time, is able to bind avidin. Moreover, the extent of avidin binding can be controlled by varying the thickness of the polymer layers. Surfaces with such properties may be used to improve signal-to-noise ratio in sensor applications.

Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS The authors thank Dr. J. Brash of McMaster University for helpful discussions. This work was supported by the National Science Fund for Distinguished Young Scholars (21125418), the National Natural Science Foundation of China (20920102035 and 20974086), and the Priority Academic Program Development of Jiangsu Higher Education Institutions (PAPD).

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CONCLUSIONS In this work, we successfully prepared POEGMA-b-PGMA diblock copolymer brush modified surfaces using SI-AGET ATRP. Good control of the thickness of the polymer grafts was demonstrated using ellipsometry. Surface morphology was shown to be dependent on block thickness and to influence surface properties. Fibrinogen and lysozyme conjugation measurements indicated that the POEGMA layer not only made the bioactive surfaces protein resistant but also provided a hydrophilic microenvironment to conserve the activity of the conjugated lysozyme. It was shown that the conjugation of bioactive molecules to these surfaces (for example, biotin with specific affinity for avidin) could be controlled by varying the thickness of the POEGMA and PGMA layers. This strategy offers a simple way to fabricate surfaces with controllable properties for immobilization of bioactive molecules, which may be useful in biomedical applications.

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REFERENCES

AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]. 1127

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dx.doi.org/10.1021/la3044472 | Langmuir 2013, 29, 1122−1128