hydrogels and natural protein fibers for layered heart valve constructs

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Bioinspired engineering of poly(ethylene glycol) hydrogels and natural protein fibers for layered heart valve constructs Qian Li, Yun Bai, Tao Jin, Shuo Wang, Wei Cui, Ilinca Stanciulescu, Rui Yang, Hemin Nie, Linshan Wang, and Xing Zhang ACS Appl. Mater. Interfaces, Just Accepted Manuscript • Publication Date (Web): 27 Apr 2017 Downloaded from http://pubs.acs.org on May 1, 2017

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ACS Applied Materials & Interfaces

Bioinspired engineering of poly(ethylene glycol) hydrogels and natural protein fibers for layered heart valve constructs Qian Li†,‡,#, Yun Bai†,#, Tao Jin¶, Shuo Wang￥, Wei Cui†, Ilinca Stanciulescu¶, Rui Yang†,§, Hemin Nie￥,*, Linshan Wang#,*, Xing Zhang†,§,* †

Shenyang National Laboratory for Materials Science, Institute of Metal Research,

Chinese Academy of Sciences, Shenyang, Liaoning 110016, China ‡

Department of Chemistry, Northeastern University, Shenyang, Liaoning 110004,

China ¶

Department of Civil and Environmental Engineering, Rice University, Houston,

Texas 77005, USA ￥

Department of Biomedical Engineering, College of Biology, Hunan University,

Changsha, Hunan 410082, China §

School of Materials Science, University of Science and Technology of China,

Hefei, Anhui, 230026, China ABSTRACT: Layered constructs from poly(ethylene glycol) (PEG) hydrogels and chicken eggshell membranes (ESMs) are fabricated, which can be further crosslinked by glutaraldehyde (GA) to form GA–PEG–ESM composites. Our results indicate that ESMs composed of protein fibrous networks show elastic moduli ~3.3-5.0 MPa and elongation percentages ~47-56%, close to human heart valve leaflets. Finite element simulations reveal obvious stress concentration on a partial number of fibers in the GA-crosslinked ESM (GA–ESM) samples, which can be alleviated by efficient stress distribution among multiple layers of ESMs embedded in PEG hydrogels.

Moreover,

the polymeric networks of PEG hydrogels can prevent mineral deposition and enzyme degradation of protein fibers from incorporated ESMs. The fibrous structures of ESMs retain in the GA–PEG–ESM samples after subcutaneous implantation for four weeks, while those from ESM and GA–ESM samples show early degradation to certain extent, suggesting the prevention of enzymatic degradation of protein fibers by the polymeric network of PEG hydrogels in vivo.

Thus, these GA–PEG–ESM

layered constructs show heterogenic structures and mechanical properties comparable to heart valve leaflets, as well as improved functions to prevent progressive calcification and enzymatic degeneration, which are likely used for artificial heart valves. KEYWORDS: heart valves; poly(ethylene glycol) hydrogels; protein fibers; layered structures; calcification; enzymatic degradation

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1. INTRODUCTION Heart valve disease with major symptoms of stenosis and regurgitation is prevalent worldwide. Surgical replacement of diseased heart valves at the end-stages has been widely performed with mechanical valves (MVs) or bioprosthetic heart valves (BHVs). All these current devices have significant limitations with risks of further morbidity and mortality. For example, MVs may cause hemorrhage and thromboembolism, and require anticoagulation for the lifetime of the patients.1,2 BHVs show better hemodynamic behavior due to the composition and structural similarity to native heart valves when compared to MVs, however, they do show limited durability because of calcification and progressive degeneration.3,4 Thus, polymeric heart valve (PHV) prostheses combining the advantages of MVs and BHVs with long-term durability and no necessity for permanent anticoagulation are of great interest5 and also show potential applications in advanced transcatheter devices.6, 7 The heterogenic composition and orientation of extracellular matrices (ECMs) play important roles in heart valve functions. For example, aortic valve leaflets comprise of the extracellular matrices (ECMs) in three distinct layers: fibrosa, spongiosa and ventricularis.8 The fibrosa layer towards the outer surface is mainly composed of dense type I collagen fibers oriented along the circumferential direction, bears tensile loads and associates to mechanical stiffness for the valve leaflet.9 The spongiosa layer in the middle consists of proteoglycans and a small amount of collagen, which facilitates the movement of the leaflet. The ventricularis layer towards the inner surface consists of aligned elastin fibers interspersed with short collagen fibers, responsible for extension in diastole and recoil during systole, and


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associates to the elasticity property of the leaflet.

The layered macrostructures,

together with tailored microstructures of each layer, i.e., oriented collagen bundles in the fibrosa layer, provide anisotropic mechanical behavior10,

11

to withstand high

circumferential stress and extend radially, and allow ~3 billion open and close cycles of the leaflet in an average adult life.12, 13 Thus, it is crucial to mimic these structural features in PHVs in order to obtain mechanical behavior similar to that of heart valve leaflets. Recently, a variety of PHVs have been developed to imitate the structural heterogeneity and versatile functions of the leaflets. Poly (ethylene glycol) (PEG) hydrogels were fabricated into tri-layered structures with layer-specific formulations to mimic the macroscopic layered structures of heart valve leaflets.14, 15 Moreover, tri-layered constructs by assembling microfabricated poly (glycerol sebacate) (PGS) and fibrous PGS/poly(caprolactone) (PCL) electrospun sheets were developed with tunable anisotropic mechanical properties similar to the mechanical characteristics of native heart valves.16, 17 These tri-layered constructs can also support the growth of valvular interstitial cells (VICs) and mesenchymal stem cells (MSCs) within the three-dimensional structure and promote the deposition of heart valve extracellular matrix (ECM). Circumferentially oriented PCL nanofibrous substrates mimicking the morphology of a fibrosa layer of an aortic valve leaflet were fabricated to direct the orientation of VICs and their deposited collagens on the substrate. Moreover, hyaluronan hydrogels were employed as a biomimetic spongiosa layer of tissue engineered heart valve scaffolds. These hydrogels showed the mechanical properties



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approximate to those of the valve spongiosa layer, and healthy phenotype of valve cells was maintained in these scaffolds.18,

19

These work focused on either the

structural-mechanical properties of the scaffolds or on cell functions when cultured in the scaffolds in vitro. However, other essential functions such as anti-calcification and prevention of progressive degeneration remain poorly understood for these PHVs. The eggshell membrane (ESM) between the egg shell and the egg white consists of two distinct layers of highly interwoven fibrous meshwork (thickness ~70-100 µm) deposited during the passage through the white isthmus.20, 21 The thin inner membrane and thick outer membrane not only serve as structural support, but also resist microorganism invasion. The fibers in the ESM are mainly composed of a collagen-rich core (mainly types I, V and X collagens) and a glycoprotein-rich mantel,22 similar to the ECM compositions of valve leaflets, allowing usage as artificial heart valve scaffolds.

Moreover, considering the tunable mechanical

properties and biological functions,23-25 PEG-based hydrogels have been explored as heart valve scaffolds.15, 26 PEG hydrogels are biocompatible and non-immunogenic, and show good non-fouling or low-fouling properties.27-32

In this study, PEG–ESM

and GA–PEG–ESM (glutaraldehyde crosslinked) constructs with layered structures are fabricated from PEG hydrogels and ESMs. Microstructures, mechanical properties and biocompatibility of these samples are investigated. In addition, the influence of PEG hydrogels on the anti-calcification and enzymatic degradation of the GA–PEG– ESM composites is evaluated.


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2. EXPERIMENTAL SECTION 2.1. PEGDA synthesis. Photocrosslinkable PEG-diacrylate (PEGDA) was synthesized with minor modification of previously described methods.26, 33 Briefly, PEG powder (M.W. = 3.35 kDa, Sigma–Aldrich, St. Louis, MO, USA) was dissolved in anhydrous dichloromethane (DCM; Sigma–Aldrich) in a round bottom flask, which reacted with 3 M excess acryloyl chloride (Sigma–Aldrich) with addition of 1 M excess trimethylamine (TEA; Sigma–Aldrich) as the catalyst. The reaction was performed under argon environment overnight, and the resulting PEGDA was purified and lyophilized, and finally stored at -20°C for further use. The proton nuclear magnetic resonance spectrum (1H-NMR) of the as-obtained PEGDA was performed using a Bruker 400 MHz spectrometer (Bruker Corporation, Swiss) to verify the percentage of acrylation (Figure S1). 2.2. Preparation of eggshell membranes (ESMs). Fresh chicken eggs were procured from a local supermarket. The ESMs were carefully peeled off from eggshells and soaked in a 8 M acetic acid solution for 12 hours in a shaker at 4°C to dissolve the calcium carbonate particles, which were then rinsed three times thoroughly with ultrapure water (A 10, Millipore, Billerica, MA, USA). ESM samples with thickness ~70-90 µm were further cut off in different shapes along both hemispherical and/or latitudinal directions and stored in purified water for further experiments. ESM samples were crosslinked with a glutaraldehyde (GA) solution (0.5 wt.%, pH = 8.0) for six hours in a shaker at 4°C, and were further rinsed with a sodium bisulfite solution (40 wt.%) for 24 hours in a shaker at 4°C to remove the



excess GA, followed by washing with ultrapure water for three times. The GA crosslinked ESM samples are designated as GA–ESM in the text. 2.3. Preparation of PEG–ESM composites. A pre-polymer solution was prepared by mixing 1.5% (v/v) triethanolamine (TEOA, Sigma–Aldrich), 10 µM Eosin Y solution (Sigma–Aldrich) and 0.375% (v/v) N-vinylpyrrolidone (NVP, Sigma–Aldrich) into a 20 wt.% PEGDA solution. The pre-polymer PEGDA solution was then photo crosslinked to form a hydrogel using white light (Shenzhen Yangao Photoelectric Technology, Shenzhen, China). The ESMs were immersed into the pre-polymer solution, which was then vacuumed for 30 minutes to allow the infiltration of the PEGDA solution into the fibrous network of ESMs. Polydimethylsiloxane (PDMS) molds with different thickness (0.4-0.6 mm) were prepared by mixing a base agent and a curing agent at a weight ratio of 10:1, and then cured at 80°C for two hours in an oven.

The ESMs soaked with the PEGDA solution

were aligned in a PDMS mold to obtain relatively uniform spacing between different layers and sealed between two pieces of glass slides pretreated with sigmacote® (Sigma–Aldrich), and then exposed to white light for about two minutes to form the PEG–ESM composite (Figure S2). The PEG–ESM composite was designated as PEG–ESM in the text. PEG–ESM constructs with a number of ESMs (i.e. 1, 2, 4 and 6 pieces) and a total thickness of 400 µm or 600 µm, mimicking layered-structure of native heart valve leaflets, were fabricated. The PEG–ESM construct was further chemically crosslinked using a GA solution (0.5 wt.%, pH = 8.0) for six hours in a shaker at 4°C, which was further rinsed with a sodium bisulfite solution (40 wt.%) for


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24 hours in a shaker at 4°C to remove the excess GA, followed by washing with ultrapure water for three times. The GA crosslinked PEG–ESM constructs are designated as GA–PEG–ESM in the text. 2.4. Mechanical properties of samples. The cleaned ESM, PEG hydrogel and PEG–ESM samples before and after GA treatment were cut into 4 mm × 25 mm strips for uniaxial tensile tests by a Bose Enduratec ELF 3200 system (Eden Prairie, MN, USA), with the crosshead speed of 0.1 mm/s and the gauge length of 10 mm. Six specimens were tested for each sample. The loading force and the displacement were recorded, and divided by the un-deformed cross-sectional area and gauge length, respectively, to obtain nominal stress and strain (or stretch ratio) values. All samples were immersed in distilled water during the tests in order to avoid dehydration. The elastic modulus was determined as the slope of the stress–strain curve using the 10%– 20% strain section by the linear regression tool in Excel software.34 2.5. Finite element simulation of the mechanical property. Finite element (FE) simulations were performed to numerically investigate mechanical properties of ESM and PEG–ESM samples with or without GA treatment. The numerical investigations were based on the computational framework for fibrous biomaterials with randomly distributed fiber network structure.35,

36

The PEG hydrogel was described by the

modified neo-Hookean constitutive law (equation (1)), ଵ

ଵ

ψ = ψ௩௢௟ ሺ‫ܬ‬ሻ + ψௗ௘௩ ሺ‫ܫ‬ଵ ሻ = ‫ ܭ‬ሺ‫ܬ‬ଶ − 1 − 2݈݊‫ܬ‬ሻ + ‫ܩ‬ሺ‫ܫ‬ଵ − 3ሻ ସ ଶ

(1)

where K = E (1-2υ) /3 is the bulk modulus, G = E (1+υ) /2 is the shear modulus, J = det F is the determinant of the deformation gradient F, and Ī1 = J-2/3 tr(FTF) is the



modified tensor invariant. The hydrogel was treated as a nearly incompressible material with the Poisson’s ratio υ = 0.499. The elastic moduli of PEG hydrogels with and without GA treatment were calibrated by fitting the load-displacement relationship obtained from the FE results with the experimental data. The ESM component was modeled as an isotropic fiber network, the spatial geometry of which was generated by extending the random walk algorithm into three dimensions (3D) as previously reported.36 The fiber volume fraction µ = 70% and the mean fiber diameter d = 2 µm were chosen according the experimental observations. Each fiber was assumed to be piece-wise linear and composed by multiple segments with the same length lseg = 0.2 mm. Each fiber segment in the ESM network was modeled as a truss element with the following strain energy function (equation (2)), ߰௙ ሺߣሻ =

ଵ ݇ ‫ܮ‬ଶ ሺߣ ଶఒ್ ௙ ଴

− 1ሻଶ

(2)

where kf is the fiber axial elastic modulus, b is a dimensionless parameter controlling the fiber nonlinearity, λ is the stretch ratio and L0 is the fiber original length. Based on a variational approach, the element residual and stiffness matrix contributed by individual fiber segments can be derived. Regarding the material parameters of an individual fiber segment contained in the ESM component, the fiber axial elastic modulus was chosen kf = 16 × 10−3 Nmm−1 for ESM samples and kf = 20 × 10−3 Nmm−1 for GA–ESM samples. The dimensionless parameter of the fiber nonlinearity was b = 1.3 for both ESM and GA–ESM samples. The PEG–ESM sample was modeled with two different material components,


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the PEG hydrogel as the ground matrix and the ESM as the embedded fiber network. The total element residual and stiffness matrix were the contributions from these two material components. Although an ESM network was typically composed of a large number of fiber segments, due to the isoparametric relationship inside the isoparametric finite element, the displacements of the two ends of individual fiber segments were mapped to the element nodal displacement via the element interpolation functions. Therefore, the total number of degrees of freedom (DOFs) of the discretized system only depended on the applied FE mesh density, regardless of the number of embedded fiber segments. For detailed information about the adopted embedded fiber approach and the derivation of the FE formulation, see the reference.37 2.6. In vitro calcification of experimental samples. In vitro calcification of ESM, GA–ESM, GA–PEG–ESM samples were performed following previous studies38, 39 with minor modifications. Aortic valve leaflets were dissected from fresh porcine hearts procured from a local market and thoroughly rinsed with a phosphate buffered saline (PBS, pH = 7.4) solution. The cleaned leaflets were further crosslinked with a 0.5 wt.% GA solution for 24 hours, followed by another 0.3 wt.% GA solution for 24 hours. The leaflets were further rinsed with a sodium bisulfite solution (40 wt.%) for 24 hours in a shaker at 4°C to remove the excess GA, followed by washing with PBS for three times. These GA-crosslinked aortic valve leaflets representing prosthetic heart valve substitutes (hereafter designated as PHV leaflets) were also used for the calcification study as control.

The calcification solution with 3.87 mM



CaCl2 (Sinopharm Chemical Reagent, China) and 2.32 mM K2HPO4 (Sinopharm Chemical Reagent), was prepared by dissolving the chemicals in a HCl–Tris buffer solution (50 mM, pH = 7.4, Sinopharm Chemical Reagent) to obtain the Ca/P molar ratio of 1.67. ESM, GA–ESM, GA–PEG–ESM and PHV leaflets were soaked in the calcification solution for 1, 2, 4 weeks in a shaker at 37°C with the calcification solution refreshed every day. 2.7. In vitro enzymatic degradation of experimental samples. A type I collagenase solution (20 u/ml) was prepared by dissolving type I collagenase (~201.00 u/mg, Invitrogen, California, USA) in a PBS solution. A pepsin solution (50 u/ml) was prepared by mixing pepsin powder (800 µ/mg, Sinopharm Chemical Reagent) in a 0.5 M acetic acid solution.

The ESM, GA–ESM and GA–PEG–ESM

samples were separately soaked in the type I collagenase solution and pepsin solution at 37°C for two and four weeks. The microstructures and mechanical properties of these samples before and after enzyme treatment were investigated. 2.8. Subcutaneous implantation of samples. Sprague Dawley (SD) rats (male, six weeks old) were injected with anesthetic (10 wt.% chloral hydrate, Sinopharm Chemical Reagent, China) according to their weight (1 ml anesthetic solution/100 g). Hair from the dorsal area was clipped and the skin scrubbed in a routine fashion with an 0.5 wt.% iodophor solution (Sinopharm Chemical Reagent). Two 1.5 cm long linear incisions were separately made in the middle-left and middle-right dorsal areas. Three types of samples, ESM (four pieces of ESMs), GA–ESM (four pieces of ESMs) and GA–PEG–ESM (12 samples for each type, total 36 samples), were then


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subcutaneous embedded, with two different types of samples assigned to the left and right implantation sites for each rat. The experimental animals were anesthetized by injection of 1 ml 2 wt.% phenobarbital sodium solution (Kingyork Inc., Tianjing, China) post-surgery for two or four weeks, and the implants with adjacent skin and subcutaneous tissues (n = 6 for each group) were collected and fixed in 4 wt.% paraformaldehyde in PBS. The explants were further dehydrated and followed by embedding in paraffin, which were then cut into 10 µm thin sections for hematoxylin and eosin (H&E) staining.

All animal experiments were performed with approval by

the Animal Ethics Committee of Hunan University. 2.9. Morphology and composition analyses. ESM, GA–ESM, PEG–ESM, GA– PEG–ESM samples, as well as the implants from in vivo experiments, were dehydrated in a series of alcohol solutions at different volumetric concentrations (50%, 60%, 70%, 80%, 90%) for 15 minutes each time, followed by dehydration in pure alcohol (100%) for three times. The samples were coated with a gold film using a sputter coating instrument (Cressington, Watford, England), and

then used for

microstructure analyses by a field emission scanning electron microscope (SEM, Zeiss Supra 35, Germany), at an accelerating voltage of 20 kV.

The element

analyses for these samples were conducted using the energy disperse spectroscopy (EDS, Oxford INCAX 7582, UK).

The crystal phases for calcified granules on the

samples from the calcification study were identified by powder X-ray diffraction (XRD), which were operated at 40 kV and 100 mA at a 2θ range of 20°-70° with a 0.02° step size using a Rigaku D/max 2400 diffractometer (Rigaku Corporation,



Tokyo, Japan) with monochromated Cu Kα radiation (Kα1 = 1.5418 Å). 2.10. Statistical analysis. All test data were expressed as mean ± standard deviation (S.D.). Statistical significance was analyzed by one-way analysis of variance (ANOVA) with post hoc Tukey testing. The p value less than 0.05 was considered to be statistically significant.

3. RESULTS 3.1. Structure and composition of ESMs. All ESM samples for the experiments in this study were cut out along the latitudinal and hemispherical directions as shown in Figure 1A. The original ESM consists of a network of randomly aligned fibers with an average diameter of 2.01 ± 0.69 µm on the outer thick layer (Figure 1B), and 0.96 ± 0.26 µm on the inner thin layer (Figure 1C), based on total 60 measurements from six samples under an optical microscope.

The fibers comprised of a

collagen-rich core and a glycoprotein-rich mantel (Figure 1D,E) with slightly different compositions, which further formed trunks and branches. The main composition of the fiber core in the outer membrane was type I collagen, whereas type I and type V collagens were found in the fiber core in the inner layer.41 A small amount of calcium carbonate particles were identified on the fibers by the EDS result (Figure 1F), which formed during the initial calcification stage of eggshell formation.40 These calcium carbonate particles were completely dissolved by treatment with 8 M acetic acid for 12 hours (Figure 1G-I) without disruption of the


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original fibrous structures.

Figure 1. A photo of an original ESM (A), and SEM micrographs of the fibrous networks in the outer layer membrane (B) and the inner layer membrane (C), and the core–mantel structures of ESM fibers indicated by dashed ellipses (D, E), as well as the fibrous networks in the out layer membrane after acetic acid treatment (G, H). EDS results (F from the dashed circle spot of B, I from the dashed circle spot of H) of the outer layers of ESMs were shown to confirm the complete removal of CaCO3 granules by acetic acid treatment. 3.2. Mechanical properties of ESM, PEG–ESM and GA–PEG–ESM samples. The representative stress-strain curve of the original ESMs included a linear regime and a non-linear regime (Figure 2A), similar to those from other protein fibrous materials, such as collagen and keratin.42

The non-linear regime can be divided into

the “toe” and “heel” areas (Figure 2B), correlated to different intrinsic structures. In the “toe” regime, there was no substantial increase of stress with increase of strain,



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due to the re-organization of unstretched fibrous network.43-45 In the “heel” regime, there was a substantial increase of stress with the increase of strain, where the crosslinked telopeptide terminals, even the triple helical structures, of proteins can be stretched, with the rearrangement of molecules in the liquid environment.46 The ESM sample performed as a Hookean elastic material at the higher level of strain,34 resulting in a transition of the stress-strain curve (Figure 2A,B).

In this range, the

load mainly affected the entropic elasticity of the ESM network, and covalent bonds in the protein backbone played an important role in major stress bearing.

On the

other hand, a single linear relationship between the stress and strain was shown for PEG hydrogels, indicating the good homogeneity and elasticity of the hydrogels. The elastic moduli of original ESM, ESM after acetic acid treatment and GA– ESM (glutaraldehyde crosslinked ESM) samples, calculated from the strain range of 10-20% in the linear regime, were shown in Figure 2C.

The average elastic

modulus of original ESMs along the latitudinal direction was 4.35 (± 0.62) MPa, which was slightly higher than that along the hemispherical direction (3.76 ± 0.42 MPa), but no significantly difference (p = 0.252) likely due to the random alignment of protein fibers in the networks.

Moreover, there was no apparent difference of

elastic moduli along the two directions for samples before and after acetic acid treatment, further confirming that the acidic treatment has no obvious effect on the structures of the fibrous networks.

However, the elastic moduli of GA–ESM

samples significantly increase either in the latitudinal direction (6.91 ± 0.88 MPa) or in the hemispherical direction (5.63 ± 0.59 MPa) compared to those from original


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ESM samples.

In addition, the elastic moduli of PEG hydrogels significantly

increased after GA treatment (0.24 ± 0.06 MPa) compared to those of the original PEG hydrogels (0.15 ± 0.02 MPa), as shown in Figure 2D, suggesting the chemical crosslinking of the polymer network by GA.

Figure 2. Representative stress-strain curves for the original ESM and PEG hydrogel under tension tests in water (A), and the enlargement of the nonlinear regime (B) of the stress-strain curve in (A) marked with red color. The elastic moduli of original ESM, ESM after acetic acid treatment (“Acetic Acid–ESM”) and GA–ESM samples with the loading along the latitudinal and hemispherical directions (C), and elastic moduli of PEG hydrogels before (“PEG”) and after (“GA–PEG”) GA treatment (D). *,# p < 0.001 compared with the GA–ESM group in either latitudinal or hemispherical directions. $ p < 0.001 compared between the PEG and GA–PEG samples.



PEG–ESM constructs with different layers of ESMs were prepared, which showed an approximately linear increase of modulus by increasing the volumetric percentage of ESMs in the composites (Figure 3A). There was also a significant increase of elastic modulus of the PEG–ESM composite by GA crosslinking (Figure 3B). The elastic moduli of PEG–ESM samples with a thickness of 400 µm and four layers of ESMs were about 3.18 ± 0.27 MPa, whereas those for GA–PEG–ESM samples were approximately 3.73 ± 0.34 MPa.

GA–PEG–ESM samples with the

similar volumetric percentages of ESMs (i.e. 400 µm thickness with four layers of ESMs vs. 600 µm thickness with six layers of ESMs) showed elastic moduli close to each other (Figure 3B-D), again suggesting the great uniformity of the layered structures.

Moreover, there was no significant difference between the elastic moduli

along the latitudinal and hemispherical directions for GA–PEG–ESM samples (Figure 3D), likely due to the random alignment of protein fibers in the ESM networks. Elastic moduli of GA–PEG–ESM composites with a thickness of 400 µm and four layers of ESMs and with a thickness of 600 µm and six layers of ESMs were in the range of 3.73 ± 0.34 MPa and 3.83 ± 0.51 MPa, respectively, close to those of aortic valve leaflets (~3-15 MPa along the circumferential direction and ~1-2 MPa in the radial direction) as previously reported.26


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Figure 3. Elastic moduli of PEG–ESM samples (thickness ~400 µm) comprising 0, 1, 2, 3, and 4 layers of ESMs (A) and the corresponding GA–PEG–ESM samples (B) after GA treatment, and GA–PEG–ESM samples (thickness ~600 µm) comprising 0, 2, 4, 6 layers of ESMs (C). Comparison of elastic moduli of GA–PEG–ESM samples with a thickness of 400 µm and 4 layers of ESMs and those with a thickness of 600 µm and 6 layers of ESMs along the latitudinal and hemispherical directions (D).

3.3. Finite element analysis of stress distribution and fiber mechanics. The macroscopic mechanical behavior of hydrogels and protein fibers was largely influenced by the microscopic parameters, such as crosslinking density, fiber density and fiber chain length.35,47,48 FE simulations were employed to analyze the stress



distribution in the fibrous structures of the PEG hydrogel and composite samples at the microscopic level. Figure S3 shows an ESM fiber network generated by the random walk algorithm. In the FE simulations of ESM and PEG–ESM samples under uniaxial tension, the clamped-clamped boundary conditions were applied at the two ends of the sample to mimic the experimental setup. By comparing the load-displacement relationship obtained from the FE simulations and the experiments as shown in Figure S4, the elastic moduli for PEG hydrogels with and without GA treatment were calibrated as 0.35 MPa and 0.23 MPa, respectively. The FE simulation results for ESM, GA–ESM, PEG–ESM and GA–PEG–ESM samples under uniaxial tension were shown in Figure 4. The load-displacement relationships obtained from the FE simulation results and the experimental data showed good agreement (Figure 4A,C). At the same deformation level, the GA–ESM sample showed larger tensile stresses than the ESM sample, as shown in Figure 4A, indicating there was an significant increase of elastic modulus after GA crosslinking, consistent with previous results (Figure 2C). In addition, the histograms of the stress distribution in the fiber network of ESM and GA–ESM samples were obtained. The major stress concentration was around 0.50 MPa for the ESM sample (data not shown), whereas that was around 1.78 MPa for the GA–ESM sample (Figure 4B). Compared to the ESM samples, significant stress concentrations (marked by the red circle in Figure 4B) occurred inside the fiber network of the GA–ESM sample, suggesting that certain fibers undertook larger tensile stresses. These results showed that the chemical crosslinking by GA largely influenced the fiber mechanics at the


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microscopic level. The increase of fiber stress by increase of the stretch ratio in the GA–PEG–ESM sample was clearly shown in Figure 4E. The major stress concentration was around 1.78 MPa for the GA–ESM sample at the stretch ratio λ = 1.3 (Figure 4B), whereas that was around 0.78 MPa for the GA–PEG–ESM sample at the stretch ratio λ = 1.5 (Figure 4D). The alleviation of stress concentration in the GA–PEG–ESM sample was due to the fact that the stresses undertaken by the single layer ESM in the GA– ESM sample were now distributed among the four layers of ESMs embedded in the PEG hydrogels. Furthermore, the detailed stress distribution among the fibers in the four-layer ESMs of the GA–PEG–ESM sample were obtained (Figure 4F). The fibers located at the two ends of the sample exhibited the largest stresses (around 1 MPa, Figure 4F) due to the imposed clamped-clamped boundary conditions. Fiber segments oriented perpendicularly to the stretch direction exhibited compressive stresses due to the contraction of the cross section when the sample underwent stretch. Thus, fiber stress analysis in detail based on the FE simulations can provide important quantitative information about the mechanical response of fibers at the microscopic level.



Figure 4. FE simulation results for ESM, GA–ESM, PEG–ESM and GA–PEG–ESM samples under uniaxial tension: (A,C) comparison of the load-displacement relationships obtained from the FE simulations and experimental data, (B) histograms of the fiber stress distribution in the GA–ESM sample when the stretch ratio λ = 1.3, (D) histogram of the fiber stress distribution in the GA–PEG–ESM sample when the stretch ratio λ = 1.5, (E) evolvement of the fiber stress distribution with respect to the sample stretch ratio λ, (F) stress distribution among the four ESM layers embedded in the GA–PEG–ESM samples (the distance between the ESM layers are exaggerated


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for visualization purpose).

3.4. In vitro calcification and enzyme degradation. SEM images showed deposition of calcified granules on the surfaces of ESM, GA–ESM, BHV leaflet samples (Figure 5A,B,D) after being soaked in the calcification media for two weeks. The granules were shown to be calcium phosphate based on calcium (Ca), phosphorus (P) and oxygen (O) peaks from EDS results (Figure 5I,J,L), which were further confirmed to be hydroxyapatite (JCPDF #09-0432) by the corresponding XRD patterns (Figure S5).

Moreover, there were larger amounts of hydroxyapatite

granules formed on the surfaces of these samples after being soaked in the calcification media for four weeks (Figure 5E,F,H), compared to those soaked for only two weeks.

These results showed a series of events including the initial mineral

deposition, growth and aggregate formation mimicking the calcification consequences in vivo.

On the other hand, there was no calcium phosphate deposition occurred on

the surfaces of GA–PEG–ESM samples after soaked in the calcification media for two or four weeks (Figure 5C,G,K), suggesting that the polymeric network of PEG hydrogels efficiently prevents calcification on embedded protein fibers.



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Figure 5. SEM images of ESM (A, E), GA–ESM (B, F), GA–PEG–ESM

(C, G)

and BHV samples as the control group (D, H) after immersion in an calcification solution for 2 weeks (A-D) or 4 weeks (E-H), and the corresponding EDS results (I-L) for (A-D), respectively. The ESM, GA–ESM and GA–PEG–ESM samples were treated by type I collagenase and pepsin. Figure 6 showed the influence of enzymatic degradation on the mechanical properties of different samples. The elastic moduli of ESMs along the latitudinal direction significantly decreased from 4.35 ± 0.62 MPa to 3.00 ± 0.60 MPa and 1.64 ± 0.15 MPa after treatment with type I collagenase and pepsin for two weeks, respectively, which further decreased to 2.67 ± 0.51 MPa and 1.39 ± 0.16 MPa, respectively, after treatment for four weeks (Figure 6A).

Moreover, the elastic

moduli of ESMs after pepsin treatment were much lower than those after type I collagenase (p < 0.01) both for two weeks and for four weeks, suggesting the higher


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efficiency of protein degradation by the pepsin treatment than that of type I collagenase treatment. The elastic moduli of GA–ESM samples also significantly decreased after either type I collagenase or pepsin treatment for two or four weeks (Figure 6B, p < 0.001). However, there was no significant difference in the elastic moduli of GA–ESM samples between the pepsin and type I collagenase treatment, indicating the reduced efficiency of pepsin on protein degradation after GA crosslinking.

On the other hand, there was no significant difference in the elastic

moduli for GA–PEG–ESM samples after type I collagenase or pepsin treatment for either two or four weeks (Figure 6C, p = 1.00). These results showed that the polymeric network of the PEG hydrogels can effectively prevent the enzymatic degradation of the incorporated protein fibers from ESMs.

Figure 6. Elastic moduli for (A) ESM, (B) GA–ESM, (C) GA–PEG–ESM (400 um thickness with four layers of ESMs) samples before and after treatment with type I ACS Paragon Plus Environment


collagenase and pepsin for two and four weeks. *p < 0.001 compared with the “No treatment” group.

#

p < 0.01 compared with the Collagenase I group.

3.5. In vivo degradation and biocompatibility. HE staining of histological sections of the explanted ESM, GA–ESM and GA–PEG–ESM samples were shown in Figure 7. No obvious edema, hyperemia or tissue necrosis for the experimental rats was observed during implantation for two and four weeks, and few inflammatory cells were observed in adjacent tissues around the implants, indicating the good biocompatibility of these implant materials in vivo.

De novo tissues were found to

grow into the spaces between ESM (Figure 7A-C) or GA–ESM (Figure 7D-F) samples. The fiber structures of the ESM were preserved after two weeks implantation (Figure 7A), but were obvious deteriorated after four weeks implantation (Figure 7B,C), suggesting the degradation of the protein fibers after long-term implantation. Similar results were found for the GA–ESM group (Figure 7D-F).

On the other hand, a tight interface between the PEG hydrogel and adjacent

host tissue formed after subcutaneous implantation of GA–PEG–ESM samples (Figure 7G-I), and no newly formed tissues were found inside of these samples, showing the good biocompatibility of the composite materials and the blocking of cell penetration by the PEG hydrogels. Moreover, the protein fibers in GA–PEG–ESM samples remained uniform without structural disruption after implantation for two weeks (Figure 7G) or four weeks (Figure 7H,I), indicating the efficient prevention of protein degradation by the polymeric network of PEG hydrogels in vivo.


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Figure 7. HE staining images for ESM (A-C), GA-ESM (D-F) and GA–PEG–ESM (G-I) samples after subcutaneous implantation or two weeks (A, D, G) and four weeks (B,C,E,F,H,I). Obvious disruption of fibrous structures in ESMs were marked out by black squares (B,C). SEM images and corresponding EDS results from the explant sections were shown in Figure 8. The EDS mapping of sulphur (S) element was only from ESM protein fibrous networks (Figure 8I,L), while mapping for carbon (C, Figure 8G,J) and oxygen (O, Figure 8H,K) elements can be either from ESM protein fibrous networks or from the remaining PEG layers as well as embedded paraffin. Moreover, the EDS results from selected areas (Figure S6) showed the characteristic peaks of C, O and S elements from the explant sections after implantation for two weeks.

SEM

images showed that there was disruption of protein fibers in the networks for ESM



(Figure 8D) and GA–ESM (Figure 8E) samples after implantation for four weeks, while those for GA–PEG–ESM samples (Figure 8F) maintained original uniform fibrous structures.

These results again demonstrated that the GA–PEG–ESM

samples showed good biocompatibility, and the PEG hydrogels in the samples can prevent the pass through of cells and enzyme molecules, thus resulting in improved functions against progressive degradation in vivo.

Figure 8. SEM images and EDS mapping for ESM (A,D,G,J), GA-ESM (B,E,H,K) and GA–PEG–ESM (C,F,I,L) samples after subcutaneous implantation for two weeks (A-C, G-I) and four weeks (B-F,J-L). The EDS mapping for C (G,J), O (H,K) and S


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(I,L) elements were performed. ESMs in the samples (A-F, I, L) are marked by yellow dot lines.

4. DISCUSSION Protein fibers from the bacterial cellulose networks,49 spider silks50 and collagen-based systems51 have been of great interest for biological applications due to their unique physical and mechanical properties.

ESMs, one of the most abundant

natural biopolymers, have also been widely studied, for example, as biological dressing for burning wounds,52,53 substrates for cell growth,54 and therapeutic dietary supplements for relief of pain associated with joint and connective tissue disorders.55 The fibrous meshes of ESMs are structurally and morphologically similar to heart valve leaflets. In this study, chicken ESMs are used to fabricate novel PEG–ESM composites as polymeric heart valve (PHV) prostheses, considering that ESMs comprise of protein fibrous structures and have mechanical properties close to those of heart valve leaflets.34 The microstructures and mechanical properties play an important role in the functions of heart valve leaflets, which can undergo opening and closing for about 3 billion cycles during an adult life time.6 For example, collagen fibers in the fibrosa layer of aortic valve leaflet are the major stress-bearing components, providing the strength to maintain coaptation during diastole. The elastin fibers in the ventricularis layer extend in diastole and recoil during systole. Chicken ESMs mainly comprise collagen fibers, which form interwoven networks. There is slight difference of



chemical composition, morphology and thickness among ESM samples, resulting in different mechanical properties. The elastic moduli of original ESM samples are in the range of 3.3-5.0 MPa, close to those of aortic valve leaflet leaflets (~3-15 MPa in the circumferential direction and ~1-2 MPa in the radial direction).26 The fracture strain of ESM samples in water is in the range ~47-56%, higher than those of valve leaflets (~10.9 ± 0.6% in the circumferential direction and 19.8 ± 1.5% in the radial direction),56 allowing ESMs to withstand the extension during cardiac cycles. Thus, ESMs can serve as artificial fibrosa and ventricularis layers for PHVs. The middle spongiosa layer of the aortic valve leaflet shows a foam-like structure and binds a large amount of water, which absorbs energy during compression, and facilitates the arrangement of collagen fibrils in the fibrosa and elastin in the ventricularis during the cardiac cycles. The spongiosa layer provides proper shearing between the layers of the leaflet without compromising the leaflet’s overall structural or biological integrity when mechanical stimuli are applied.57 PEG hydrogels with a large amount of water (~80 wt.%), like the spongiosa layer of the leaflet, can absorb energy during compression and facilitate arrangement of protein fibers in the ESMs.19 In this study, PEG–ESM composites with layered structures mimicking the structures of aortic valve leaflets were prepared, making use of the good biocompatibility of PEG hydrogels.26 PEG–ESM composites at different thickness ~400-600 µm close to that of human aortic valve leaflet ~300-700 µm, can be prepared by implementing different layers of ESMs, showing the increased moduli with increase of the volumetric percentage of ESMs.


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Moreover, GA crosslinking has been widely used to improve the stability of protein fibers for BHVs58 as well as for other biological tissues.59 Degradation of ESM protein fibers by pepsin can be reduced to some extent (Figure 6A,B), suggesting improved stability of ESMs with GA crosslinking. In addition, elastic moduli of both ESM (Figure 2, 4) and PEG–ESM samples (Figure 3, 5) were significantly increased after GA crosslinking. The elastic moduli of GA–ESM samples were ~5.0-7.8 MPa, whereas those for GA–PEG–ESM samples were ~3.73 ± 0.34 MPa (400 µm with four layers of ESMs) and ~3.83 ± 0.51 MPa (600 µm with six layers of ESMs), close to those from aortic valve leaflets. Therefore, GA crosslinking was an efficient method to improve the stability of ESM protein fibers.

However,

FE simulation showed significant stress concentration on a partial number of fibers for GA–ESM samples (Figure 4B). The stress concentration can be alleviated in the GA–PEG–ESM layered constructs due to the efficient stress distribution among multiple layers of ESMs embedded in PEG hydrogels (Figure 4F), which is beneficial to cyclic stretch when used as artificial heart valve leaflets. Thus, GA– PEG–ESM layered constructs taking advantage of proper mechanical strength from fibrous ESMs and compression cushion capability from PEG hydrogels, show microstructures and mechanical properties similar to human aortic valve leaflets, superior to hydrogels14,15 or polymeric fiber scaffolds16,17 alone as PHV scaffolds. Calcification is one of the major contributors to the failure of artificial heart valves, especially BHVs, when exposed long-term to the hemodynamic environment. The early calcification can form on the substrates of protein fibers from heart valves



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or blood vessels as micro-calcification.60,61 A previous study20 showed that type X collagen identified from both membrane layers of ESMs can inhibit mineralization and likely functions as boundaries to prevent mineral deposition on the egg white and yolk during the short-term production of eggshells in the hen oviduct.

However,

ESM or GA–ESM samples were still prone to calcification (Figure 5A,B,E,F) after long-term exposure to calcification media, same as BHVs (Figure 5D,H), due to their similar compositions of protein fibers (i.e. collagens).

The endothelial cell layers

and basement membranes on the heart valve leaflets or blood vessels covering the blood-contacting surfaces can function as compact barriers to prevent calcium deposition on the underneath protein fibers.

In this study, PEG hydrogels with

well-known non-fouling or low-fouling properties28-30 were employed as the blocking barrier for embedded ESM protein fibers. The GA–PEG–ESM samples showed improved anti-calcification capability compared to ESMs and BHVs (Figure 5), likely due to the efficient reduction of ion transportation through the polymeric mesh, thus preventing calcium phosphate nucleation on the embedded protein fibers. Progressive degeneration is another major failure mode besides calcification for BHVs. Even though GA crosslinking can stabilize parts of the tissue, the major ECM components are lost from the BHVs due to enzymatic degradation,62,63 leading to compromised mechanical function and degenerative tears of BHVs.64

ESMs can be

chemically crosslinked by GA to improve stability of protein fibers, but still underwent considerably enzymatic degradation (Figure 6A,B). However, GA–PEG– ESM composites showed significantly improved functions against enzymatic


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degradation, likely due to the inhibition of enzyme molecules to pass through by the polymeric networks of PEG hydrogels. Thus, the GA–PEG–ESM layered constructs not only showed suitable mechanical properties close to human heart valve leaflets, but also exhibited versatile functions to prevent calcification and progressive degeneration, which can be used for PHV scaffolds. 5. CONCLUSIONS The compositional and spatial heterogeneity of native heart valves play important roles in their unique mechanical properties and biological functions (i.e. anti-calcification). Thus, it is critical to mimic these features in artificial heart valve leaflets to obtain the equivalent mechanical and biological functions.

In this study,

layered GA–PEG–ESM constructs were fabricated, consisting of protein fibers from natural ESMs to mimic the fibrous networks in the fibrosa and ventricularis layers for stress bearing, as well as PEG hydrogels functioning as the spongiosa layer for compression cushion. These GA–PEG–ESM layered constructs showed mechanical properties, i.e., elastic modulus and elongation percentage, close to those of human aortic valve leaflets.

Finite element simulations revealed the efficient stress

distribution on protein fibers among multiple layers of ESMs when embedded in the PEG hydrogels, which is beneficial to cyclic stretch when used as artificial heart valves. Furthermore, the presence of PEG hydrogels in GA–PEG–ESM composites improved the resistance to progressive calcification and enzymatic degradation of the embedded protein fibers, likely due to prevention of large-size hydrated ions and molecules to pass through by the polymeric networks of the hydrogels.


Thus, this


study lays down a basis for fabrication of novel layered composite materials to mimic the heterogenic structures, mechanical properties and biological functions of heart valve leaflets.

ASSOCIATED CONTENT

Supporting Information

Including Figure S1 for 1H NMR spectrum of as-prepared 3.4 kDa PEGDA, Figure S2 for schematic diagrams for preparation of the PEG–ESM composite, Figure S3 for 3D protein fiber network of the ESM generated by the random walk algorithm, Figure S4 for comparison of FE simulation results with experimental data of PEG hydrogel samples, Figure S5 for the XRD patterns of ESM , GA–ESM and BHV leaflet samples after soaked in a calcification solution, Figure S6 for SEM images and EDS results from ESM, GA–ESM and GA–PEG–ESM samples after subcutaneous implantation.

AUTHOR INFORMATION

Corresponding Authors

*E-mail: [email protected] (Xing Zhang), [email protected] (Linshan Wang), [email protected] (Hemin Nie).

Author Contributions

#

These authors contributed equally to this work.


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Notes

The authors declare no competing financial interest.

ACKNOWLEDGEMENTS

This work was supported by National Natural Science Foundation of China (31300788, 31670981) and the Hundred-Talent Program from Chinese Academy of Sciences. We thank Chen Liu and Chang Liu for assistance with preparation of experimental samples. REFERENCES (1) Cannegieter, S. C.; Rosendaal, F. R.; Briët, E. Thromboembolic and Bleeding Complications in Patients with Mechanical Heart Valve Prostheses. Circulation 1994, 89 (2), 635-641. (2) Sheriff, J.; Claiborne, T. E.; Tran, P. L.; Kothadia, R.; George, S.; Kato, Y. P.; Pinchuk, L.; Slepian, M. J.; Bluestein, D. Physical Characterization and Platelet Interactions under Shear Flows of a Novel Thermoset Polyisobutylene-based Co-polymer. ACS Appl. Mater. Interfaces 2015, 7 (39), 22058-22066. (3) Schoen, F. J.; Levy, R. J. Calcification Tissue Heart Valve Substitutes: Progress toward Understanding Prevention. Ann. Thorac. Surg. 2015, 79 (3), 1072-1080. (4) Zilla, P.; Brink, J.; Human, P.; Bezuidenhout, D. Prosthetic Heart Valves: Catering for the Few. Biomaterials 2008, 29 (4), 385-406. (5) Sachweh, J. S.; Daebritz, S. H. Novel "Biomechanical" Polymeric Valve



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hydrogels and natural protein fibers for layered heart valve constructs

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