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A Biosurfactant-containing TSD Strategy to Modify Bovine Pericardial Bioprosthetic Valves for Anticalcification
RESEARCH ARTICLE | Updated：2022-12-19
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    • A Biosurfactant-containing TSD Strategy to Modify Bovine Pericardial Bioprosthetic Valves for Anticalcification

    • Gao Cai-Yun

      ,  

      Wang Gang

      ,  

      Wang Lin

      ,  

      Wang Qun-Song

      ,  

      Wang Han-Cheng

      ,  

      Yu Lin

      ,  

      Liu Jian-Xiong

      ,  

      Ding Jian-Dong

      ,  
    • Chinese Journal of Polymer Science   Vol. 41, Issue 1, Pages: 51-66(2023)

    • DOI：10.1007/s10118-022-2843-9    

      CLC：

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  • Cite this article

  • Cai-Yun Gao, Gang Wang, Lin Wang, et al. A Biosurfactant-containing TSD Strategy to Modify Bovine Pericardial Bioprosthetic Valves for Anticalcification. [J]. Chinese Journal of Polymer Science 41(1):51-66(2023) DOI： 10.1007/s10118-022-2843-9.

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    Abstract

    Bioprosthetic heart valves (BHVs) are important for transcatheter valve replacement. Current commercial BHVs on the market are basically porcine or bovine pericardium (BP) crosslinked with glutaraldehyde (GA). Simply applying GA to BHVs can enhance mechanical stability, but cannot alleviate in vivo calcification. In this work, we developed a two-step decellularization (TSD) strategy to modify this biomacromolecular network, in which BP was post-treated, as the second step of decellularization, with a mild biosurfactant n-dodecyl-β-D-maltoside in a mixture of isopropanol and phosphate-buffered saline after the first step of traditional decellularization and GA cross-linking. The TSD-treated BP exhibited not only low cytotoxicity and excellent mechanical properties in vitro, but also low immune responses and significant anticalcification in vivo. After 60 days of subcutaneous implantation in the back of Wistar rats, the calcium content was, as quantified with an inductively coupled plasma optical emission spectrometer, only 1.1 μg/mg compared to 138.6 μg/mg in the control group without the post-treatment. In addition, collagen fibrils were observed with field emitting scanning electron microscopy (SEM), and the morphology and composition of the calcified sites resulting from in vivo biomineralization were studied with SEM with energy dispersive spectroscopy and also X-ray diffraction. This study proposes a facile yet effective anticalcification strategy for the modification of the bovine pericardial bioprosthetic heart valve, a natural biomacromolecular network.

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    Keywords

    Heart valve; Anticalcification; Collagen; In vivo biomineralization ; Transcatheter aortic valve implantation (TAVI); Extracellular matrix; Biomacromolecular network; Biosurfactant; Bovine pericardium

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    INTRODUCTION

    The heart valve maintains human blood circulation by achieving unidirectional blood flow. Valvular heart disease is caused by valve stenosis and regurgitation, which is a leading cause of morbidity and mortality.[

    1,2] With the accumulation of clinical experience, technical improvement and design iteration of transcatheter heart valves over the past decade, transcatheter aortic valve implantation (TAVI) has been the first treatment of heart valve disease.[3,4] Over 300,000 TAVIs have been performed worldwide, and the number is expected to increase to 850,000 by 2050.[5−7] Mechanical heart valves and bioprosthetic heart valves (BHVs) are the main artificial heart valves. Compared to mechanical valves，BHVs have been widely used in clinical surgery and interventional treatment of heart valves because of their inherent merits, such as superior hemodynamic performance, no need for lifelong anticoagulation, and strong anti-infection ability.
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    The commercial BHVs in clinics are basically porcine pericardium (PP) or bovine pericardium (BP) treated by glutaraldehyde (GA). GA can significantly improve the mechanical property of a collagen network and the resistance to enzymatic hydrolysis owing to its excellent cross-linking ability.[

    8] GA treatment is the current standard in this field to crosslink the biomacromolecular network in a bioderived material.[9,10] Nevertheless, BHVs generally exhibit unsatisfactory durability with a lifespan of only 10 years, and GA-treated BHVs suffer from cytotoxicity,[11] immune response and calcification.[12]
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    Thereinto, calcification is the major cause of bioprosthetic valve failure.[

    1] While the mechanism of calcification is still far away from conclusive, most scholars believe that the accumulation of calcium phosphate is related to the phospholipids in plasma membranes and intracellular membranes as well as some residual treatment substances.[13−16] Recently, non-glutaraldehyde cross-linking agents, such as curcumin,[17] procyanidins,[18] carbodiimide,[19] and genipin[20] have been tried to reduce calcification. In addition, some anticalcification treatments employed biofunctional coatings as a protective layer for calcium deposition to resist calcification, such as dopamine-modified alginate coating,[21] biopolymer films of chitosan,[22] hyaluronic acid coating loaded with vascular endothelial growth factor,[23,24] amphoteric ionic hydrogel coating,[25] and poly(ethylene glycol).[26] These methods, however, have limitations. For example, although non-glutaraldehyde substances can improve cytocompatibility to a certain extent, their cross-linking efficacies are typically not as good as GA, and an insufficient reactivity might influence the durability of the resultant valves.[27]
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    Herein, we developed a two-step decellularization (TSD) strategy, in which the first step of traditional decellularization and GA cross-linking was followed by the second step of treatment of n-dodecyl-β-D-maltoside (DDM) in the mixture of isopropanol (IPA) and phosphate-buffered saline (PBS). The basic procedures are schematically presented in Fig. 1. Although IPA has been widely used in the treatment of BP to remove residual lipids and serves as a storage solution,[

    28,29] it only shows limited effects in reducing calcification. It has been known that the combination of alcohol and surfactant can remove phospholipids and reduce calcification,[30−32] but these additives are generally cytotoxic to some extent. Different from other reported surfactants, DDM is a mild bio-based nonionic surfactant with low toxicity, which can effectively dissolve membrane-bound proteins while maintaining protein conformation within the solution phase and promoting refolding of proteins after denaturation.[33,34] We hypothesized that the combination of DDM and IPA could effectively remove lipids, especially residual cell phospholipids in tissue (thus as the second decellularization), and also replace the residue toxic surfactants used in the one-step decellularization such as Triton X-100 and thus reduce cytotoxicity of the resultant BHV. The biosurfactant-containing TSD strategy might not only maintain the excellent cross-linking ability of GA, but also overcome the inherent disadvantages such as cytotoxicity, immune response and calcification in one-step decellularization. The present study attempts to report this TSD approach and confirm its feasibility, and meanwhile, provides a new insight into the further development of low-toxic nonionic surfactant in bioprosthetic products. We also investigated the in vivo biomineralized structure via field emitting scanning electron microscopy (SEM), etc.
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    fig

    Fig 1  Schematic illustration of the two-step decellularization (TSD) processes of the bovine pericardial bioprosthetic valves. First, the BP was decellularized by Triton X-100, sodium deoxycholate (SDC), DNase I and RNase to get a decellularized bovine pericardium (D-BP), followed by glutaraldehyde (GA) cross-linking. Second, glutaraldehyde-treated bovine pericardium (GA-BP) was post-treated with isopropanol (IPA) and a mild biosurfactant—n-dodecyl-β-D-maltoside (DDM). The final product IPA-DDM-BP or TSD-BP was implanted in rats to assess the extent of in vivo calcification following the ISO standard for corresponding medical implants. The calcification of tissue collagen in the implanted bioprosthetic heart valves (BHVs) was also observed via field emitting SEM. The red spots refer to the residual hazardous substances after the 1st decellularization, mainly some internally phospholipids originally in cell membranes and externally toxic surfactants such as SDC and Triton X-100. These substrates are expected to be replaced by the mild biosurfactants during the post-treatment or the 2nd decellularization.

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    EXPERIMENTAL

    Materials

    Fresh BP was provided by Lifetech Scientific Co. Ltd. (Shenzhen, China). GA (25 wt%), IPA, DDM and any other surfactants were purchased from Aladdin (Shanghai, China). Collagenase II was a product of Sigma-Aldrich (US). PBS (pH 7.4), fetal bovine serum (FBS), Tris-HCl, penicillin-streptomycin-amphotericin B, 4% paraformaldehyde (PFA) solution, 4,6-diamidino-2-phenylindole (DAPI), phalloidin-rhodamine solution and any other cell culture medium were purchased from Beyotime Biotechnology (Shanghai, China). Human umbilical vein endothelial cells (HUVECs) were purchased from Shanghai Institute of Biochemistry and Cell Biology (Shanghai, China). Cell counting kit-8 (CCK-8) was from Dojondo (Japan). Alizarin red, Masson, hematoxylin-eosin staining (HE) and all antibody reagents were purchased from Servicebio (Shanghai, China).

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    Fabrication of BP Bioprosthetic Valves

    In order to fabricate decellularized bovine pericardium (D-BP), the fresh BP collected from the slaughterhouse was immediately washed with 4 °C normal saline within 4 h, and then immersed in a mixed solution containing 0.5 wt% Triton X-100 and 0.5 wt% SDC for 12 h. After being rinsed with PBS, the specimens were treated with 50 U/mL DNase I and 1000 μg/mL RNase overnight at 37 °C. The enzyme dilution was Tris buffer (10 mmol/L Tris-HCl at pH 7.5, 2.5 mmol/L MgCl2, 0.1 mmol/L CaCl2). Finally, the specimens were rinsed repeatedly with normal saline to thoroughly wash off the surface activity and enzyme residues on the BP. D-BP was stored in PBS containing penicillin (100 U/mL), streptomycin (0.1 mg/mL), and amphotericin B (0.25 μg/mL) at 4 °C.

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    As the fabrication of GA-BP was concerned, the above D-BP was spread out on a fixed plate to prevent curling and then placed in a 0.625 wt% GA solution (pH 7.4) at room temperature (25 °C) to experience cross-linking for 72 h. GA-BP was thus obtained.

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    In terms of fabrication of IPA-DDM-BP, namely TSD-BP, the GA-BP was cut into small pieces of about 10 cm × 10 cm, and post-treated in a GA solution (pH 7.4) containing 20 wt% IPA, and 0.5 wt% DDM at 37 °C for 12 h. The reagent residues were repeatedly washed out with PBS. The obtained modified valve IPA-DDM-BP or TSD-BP was stored in 0.625 wt% GA solution (pH 7.4).

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    Physicochemical Characterizations of BPs before and after Treatments

    Prior to morphology observation, specimens of D-BP, GA-BP, and TSD-BP were freeze-dried and cut into a size of 50 mm × 50 mm. Gold was sprayed under 10 mA for 90 s. The morphology of collagen fibrils was observed at an accelerating voltage of 3 kV with a scanning electron microscope (SEM, Ultra 55, Germany).

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    The cross-linking properties of BP valves were characterized in a thermal shrinkage thermometer (JY-PS-83, Yuanmore Co., Ltd, China). The BP tissue specimens of each group were cut into standard long splines of 56 mm × 3 mm (n=5), and the two ends of the splines were hung on the two hooks of the heat shrinkage apparatus. The specimens were then immersed in normal saline to heat from 20 °C to 90 °C at a rate of 2 °C/min. The test method refers to the standard QB/T 2713—Leather Physical and Mechanical Test Determination of Shrinkage Temperature.

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    Differential scanning calorimetry (DSC, TA Q2000, USA) is a usual tool to measure the thermal denaturation temperature of a biomacromolecule-containing material. The specimens were thoroughly rinsed with PBS and then freeze-dried. The dried specimens were weighed (10 mg per specimen) and sealed in aluminium crucibles. The test temperature ranged from 30 °C to 140 °C at a heating rate of 10 °C/min.

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    The contact angle test was used to characterize the hydrophilicity of the sample surface, and the static water contact angle (WCA) was measured with a contact angle meter (JC2000DM, Zhongchen). Briefly, 10 μL of deionized water was dripped on the surface of the specimen and a camera was used to take a picture. Contact angles were viewed on the monitor's screen by using SCA 20 software. Four parallel samples were taken for each specimen.

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    In order to test tensile performance, D-BP, GA-BP and TSD-BP were uniformly cut into dumbbell shapes of 57.5 mm × 12.5 mm (n=7 for each group) along the fiber direction with a standard dumbbell steel die. Specimens stored in normal saline were slightly dried with paper and then tested with a universal tensile testing machine (Instron 5966, USA) using a 100 N sensor at a tensile rate of 50 mm/min. The stress-strain curve of the specimen was recorded, and the elastic modulus and the breaking stress of the rope were measured. Data for specimens with fractures in the grips were excluded from the statistical analysis.

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    Collagen stability in fresh and cross-linked BPs was examined by collagenase treatment. The lyophilized D-BP, GA-BP and TSD-BP were cut into 1 cm × 1 cm (n=5 for each group). The initial weight of each specimen was recorded as W0. The dried valve was then soaked in PBS (pH 7.4) containing collagenase II (125 U/mL, pH 7.4) and placed in a constant temperature shaker at 37 °C with 60 r/min for 24 h. The specimen was washed with PBS and lyophilized again. After the moisture was completely removed, a specimen was weighed as Wt. The weight loss rate upon enzymatic degradation is ready to be calculated from

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    CollagenHydrolysis=W0−WtW0×100% 1

    Characterization of Cytocompatibility

    We designed two protocols for evaluating the difference of toxicity of surfactants. Experiment 1: after different groups of surfactants including sodium dodecyl sulfate (SDS), sodium deoxycholate (SDC), Triton X-100, Tween 80 and DDM were incubated in a 48-well plate for 24 h (n=4), every well was washed twice with PBS solution soaking for 5 min each time and discarded. Here we took the PBS solution as the blank control. After culturing HUVECs (EA.hy926) for 1 day and 3 days, we accessed cytotoxicity using CCK-8. Experiment 2: This experiment is to further explore the cytotoxicity followed by washing with DDM after incubation with different groups of surfactants. After different groups of surfactants were incubated in a 48-well plate for 24 h (n=4), the original surfactant solution was discarded, and 0.5% DDM was added to incubate for 24 h more. Then, every well was washed twice with PBS solution (soaking for 5 min each time) and discarded. Lastly, we performed cell seeding and cell viability test similar to Experiment 1.

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    The cytotoxicity of BP was evaluated according to ISO 10993-5:2009/GB/T 16886.5—Biological Evaluation of Medical Devices Part 5. An extraction method was used in this experiment. GA-BP and TSD-BP specimens were cut into 1 cm × 1 cm (n=4), placed in a 24-well plate, soaked in 75% ethanol solution for sterilization and washed with PBS. Subsequently, 1 mL of high-glucose DMEM containing 10% FBS, 1% penicillin-streptomycin, and 1% glutamine was added to continue the culture. The samples were incubated in a sterile CO2 incubator at 37 °C for 72 h to obtain the material extract. HUVECs (EA.hy926) was chosen to study the cytotoxicity of the samples. Cells were seeded in 96-well cell culture plates at a density of 1×104 cells per well. After culturing for 24 h, the original medium was discarded and replaced with the sample extract, and incubated for another 24 or 72 h. Add 100 μL of 10% CCK-8 solution to the well and incubate at 37 °C for 1 h. After discarding the leaching solution, 100 μL of 10% CCK-8 solution was added to each well and incubated at 37 °C for 1 h. Finally, the absorbance of each specimen at 450 nm was measured with a multifunctional microporous plate detector (BioTek, Cytation3, USA). The medium without any sample was used as the blank control of cytotoxicity.

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    In order to carry out endothelial cell adhesion tests, GA-BP and TSD-BP were cut into 1cm × 1cm (n=4) and placed in a 24-well plate. The specimens were soaked in a 75% ethanol solution for sterilization and washed with PBS. After that, human endothelial cells were seeded on the sample surface at a density of 1×104 cells per well. The culture lasted for 24 h. At room temperature, the cells were stained with 300 μL of 1 μg/mL phalloidin-rhodamine staining solution in the dark for 30 min, and 2 μg/mL of DAPI staining solution in the dark for 8 min. After staining, the adhesion of endothelial cells to the sample surface was observed with a fluorescence inverted microscope (Axio Scope A1, Carl Zeiss Optics, Germany).

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    Characterization of Hemocompatibility

    Prior to platelet adsorption tests, the whole blood of a rabbit was collected using sodium citrate anticoagulant tubes. The blood was centrifuged at 100 g for 10 min, and the supernatant was centrifuged again at 400 g for 10 min to take the transition layer to obtain platelet-rich plasma (PRP). The specimens of GA-BP (control) and TSD-BP (φ=10 mm) were placed in a 24-well plate, wetted with PBS, which was then discarded. Then, 1 mL of PRP was added and incubated at 37 °C for 1 h. Discard the PRP and rinse the sample for 3 times with PBS. Afterwards the specimens were fixed with 4% paraformaldehyde for 10 min and then rinsed for 3 times with PBS. After washed and dehydrated with a series of gradients of alcohol solutions (30%, 50%, 75%, 90%, and 100%, each for 20 min), the specimens were dried at room temperature, sprayed with gold at 10 mA for 90 s. The images of platelet adsorption were captured at an accelerating voltage of 3 kV with a scanning electron microscope (SEM, Ultra 55, Germany).

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    In order to carry out hemolysis tests, we took 8 mL of fresh rabbit anticoagulant and diluted it with 10 mL of 0.9% NaCl solution. We cut GA-BP and TSD-BP into 2 cm × 2 cm, put them into a 15 mL centrifuge tube, and added 5 mL of normal saline. After incubated at 37 °C for 30 min, each tube was added with 0.1 mL of the diluted anticoagulant and incubated for 60 min more at 37 °C. Next, we carried out centrifugation at 800 g for 5 min and took the supernatant to measure the absorbance at 540 nm with a multifunctional microplate detector (BioTek, Cytation3, USA). The hemolysis was calculated from the absorbance values. Distilled water and normal saline were used as positive and negative controls, respectively.

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    Animal Experiments

    Specimens were cut into 1 cm × 1 cm and washed thoroughly with sterile PBS solution 3 times before implantation. We selected 3-week-old SPF male Wistar rats with a weight of 45−55 g as our experimental animals. All animal usage obeyed the regulations as set forth in the Animal Welfare Act. Rats were first anesthetized, shaved, and routinely sterilized. A surgical opening of about 1.5 cm was made on the back, and there were totally four pockets on the left and right sides. One specimen was placed in each pocket, and 10 specimens were implanted in each group. Rats were euthanized by cervical dislocation and 5 implants were removed after 30 and 60 days. Half of the explanted tissue was preserved for histology tests; the other half was frozen for calcium analysis.

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    Prior to in vivo calcification evaluation, we removed the intertwined fibrous capsules and tissues from the sample, washed them with normal saline and dried at 60 °C for 24 h (n=5 for each group). Then we weighed specimens and digested them in 6 N HCl at 97 °C for 24 h. Calcium content was determined using an inductively coupled plasma optical emission spectrometer (ICP-OES, iCAP 7400, Thermo Fisher, US). The morphologies of collagen fibrils and calcified sites were observed at an accelerating voltage of 3 kV with a scanning electron microscope (SEM, Ultra 55, Germany). The elemental mapping of specimens was tested using SEM with energy dispersive spectroscopy (EDS-SEM) (n=4 for each group). The composition of the calcified sites was also characterized by X-ray diffraction (XRD, Bruker D8 diffractometer, Germany) in the range of 10°−70° and Fourier-transform infrared spectroscopy (FTIR).

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    In order to carry out in vivo immunohistochemical and histological evaluation, another part of the sample was fixed in 4% paraformaldehyde, embedded in paraffin and sectioned. The slices were stained with alizarin red, Masson, hematoxylin and eosin (HE). The HE staining was used for histopathological observation, Masson staining for the spatial distribution of collagen and elastin biomacromolecules, and alizarin red staining for visualization of mineralized calcium salts. Slice scanning was performed with Olympus OlyVIA VS200. In addition, we carried out immunohistochemical staining with cluster of differentiation 3 (CD3), F4/80, tumor necrosis factor alpha (TNF-α) and interleukin 10 (IL-10) markers to analyze inflammation and immune responses. The antibodies used for immunohistochemical detection are listed in Table S1 (in the electronic supplementary information, ESI).

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    Ethics Statement: in this study, all animal experiments were conducted following the Guide for the Care and Use of Laboratory Animals of the National Institutes of Health. We try our best to minimize the suffering of animals.

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    Statistical Analysis

    All experimental results in this study are presented as mean ± standard deviation (SD). Statistical analysis was performed with Origin 2021 (Origin lab Inc., USA). One-way ANOVA was performed to assess significant differences. In this study, “*” stands for p<0.05; “**” stands forp<0.01; and “***” stands forp<0.001; “****” stands for extremely significant difference withp<0.0001.

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    RESULTS

    Decellularization and Characterization of TSD-BP

    The extent of decellularization is a key indicator to a bioderived material. If a pericardium is not completely decellularized, residual cells or cellular debris lead to severe immunogenicity and calcification after the valve is implanted.[

    14] In this study, we employed the detergent-enzyme method to decellularize BP. As shown in Fig. 2(a), the nuclei of cells in a fresh BP were visible in HE-stained tissue slices. After decellularization, most of the cellular components were removed and no noticeable cells or their debris were observed in an optical microscope.
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    fig

    Fig 2  Some basic characterizations of BP valves. (a) Optical micrographs of HE-stained histopathological slices to demonstrate the decellularization efficacy; (b) Water contact angles of GA-BP and TSD-BP; (c) SEM images of the indicated specimens.

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    As shown in Fig. S1 (in ESI), we compared the effects of five decellularization methods. Different detergents led to distinct levels of decellularization. In the single detergent group, strip-shaped (vessel) or dot-shaped (single cell) cell residues were seen in the extracellular matrix (ECM), and the removal of cells was highly insufficient compared with that in the detergent-enzyme group. In addition, ECM was damaged to some extent in the group added with 0.5% SDS. In contrast, the detergent-enzyme group resulted in successful decellularization and preserved the ECM structure. As the water contact angle is concerned, no significant difference was found between the two groups (Fig. 2b), although TSD-BP resulted in slightly lower value than GA-BP; both of the valves preferred to be hydrophilic. SEM observations showed collagen fibrils in the ECM of BP. Compared with D-BP, GA-BP and TSD-BP exhibited denser collagen fibrils (Fig. 2c).

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    Mechanical Properties and Structural Stability

    Mechanical properties and structural stability are important for a valve device. Both denaturation temperature and thermal shrinkage temperature reflect cross-linking performance: the higher the temperature, the better the cross-linking.[

    35] Fig. 3(A) shows typical DSC curves, where the endothermic peak was considered as thermal denaturation temperature and used to assess the degree of crosslinking. The resultant denaturation temperatures were 95.3±2 °C for D-BP, 107±3 °C for GA-BP, and 108±1.5 °C for TSD-BP.
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    fig

    Fig 3  Thermodynamic stability of BP valves experiencing different treatments. (a) Typical DSC thermodynamic test curve and the denaturation temperature of the D-BP, GA-BP and TSD-BP; (b) The schematic diagram and statistical results of the thermal shrinkage temperature test up to 100 °C according to QB/T 2713-2005 and ISO 3380:2002: Leather-Physical and Mechanical Tests—Determination of Shrinkage Temperature. The temperature when the BP valve starts to change from flat to shrinking and curling during the heating process is defined as thermal shrinkage temperature. “***” Means a very significant difference with p < 0.001.

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    Thermal shrinkage temperature test can reflect the crosslinking extent of a BP valve. As depicted in Fig. 3(b), this characteristic valve was measured from the temperature at which the BP valves got to shrink and curl in PBS during the heating process. D-BP, GA-BP and TSD-BP resulted in the thermal shrinkage temperatures of 67.5, 85.3, and 84.5 °C, respectively. Compared with D-BP, TSD-BP exhibited the increase of both the denaturation temperature and heat shrinkage temperature significantly.

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    The mechanical properties of valve materials were further quantified. Typical stress-strain curves in tensile tests are shown in Fig. 4(a). The statistical results confirmed the significant effect of GA crosslinking of the BPs. The predominant component of the pericardium is type I collagen.[

    10] Collagen is a natural macromolecule and the most predominant component in ECM of mammals; three α-peptide chains with left-hand helix are self-assembled into a right-hand triple-helix superstructure.[36] Collagen molecules are hydrolysed by enzymes to peptides, and the partially denaturation of collagen under the condition such as a high temperature results in gelatin.[37]
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    fig

    Fig 4  Mechanical properties and structural stability of the BP valves. (a) Typical stress-strain curve and statistical results of elastic modulus and tensile strength of BP valves; (b) The schematic diagram of triple-helix structure of a collagen biomacromolecule, crosslinking of collagens by GA, collagen hydrolysis by enzymes, and the experimental results of the indicated BP valves experiencing collagenase degradation for 24 h. “***” Means a very significant difference with p<0.001.

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    The GA cross-linking is aimed not only to improve the mechanical properties, but also to enhance biochemical stability to resist enzymatic hydrolysis and chemical dissociation after the valve is implanted into human body. Therefore, it is necessary to perform an enzymatic hydrolysis test to assess the stability of a valve tissue. As illustrated in Fig. 4(b), D-BP exhibited poor resistance to enzymatic hydrolysis and its weight loss rate was 95.9% after 24 h degradation; in contrast, TSD-BP and GA-BP did not exhibit significant weight loss. Hence, while collagens were easily degraded by collagenase, the crosslinked collagen fibrils of TSD-BP and GA-BP remained relatively stable under the examined condition.

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    The experimental results from SEM observations, mechanical, thermal stability and anti-enzymolysis tests were consistent with each other, and the GA crosslinking of collagens could improve the heat resistance, enzymatic hydrolysis resistance and mechanical ability of the valve. The post-treatment using biosurfactants did not affect the GA cross-linking performance of the decellularized valves, and the resultant TSD-BPs exhibited excellent mechanical properties, thermodynamic stability and biological stability.

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    Cytocompatibility of the Modified Valve Materials

    Cytotoxicity test is necessary to access the possible harmful effect of a chemical relevant to a medical device or pharmaceutics. According to ISO 10993-5: 2009, the cytotoxicity of a material can be evaluated semi-quantitatively. The reactivity of cells in a leaching solution or in direct contact with the device for 24−72 h was scored on a scale: 0, none; 1, slight; 2, mild; 3, moderate; 4, severe. A biomaterial is considered of acceptable cytotoxicity with Grade 0 or 1. In this study, we used the CCK-8 method of the material leaching solution to test the cytotoxicity. The acceptance criteria for cell viability are 70% according to ISO 10993-5:2009.

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    Generally speaking, surfactants follow the order of toxicity as cationic surfactants>anionic surfactants>nonionic surfactants.[

    38−40] We compared cytotoxicities of five different surfactants. As shown in Fig. 5(a), after the surfactants of different groups were washed with PBS, the cell viabilities of the residual surfactants after 1 day followed the sequence SDS<SDC<Triton-100<Tween 80<DDM. Those for 3 days followed the sequence again. Hence, DDM always exhibited least cytotoxicity, while other residual surfactants were significantly cytotoxic even after washing with PBS twice during the treatment process of a valve or other ECM components.
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    fig

    Fig 5  Cytocompatibility of the chemicals used for the treatment of BPs and the resultant BP valves. (a) Cell viabilities in the media containing the indicated different surfactants and cytotoxicity tests of the different surfactants after DDM rinsing for 24 h (n=4). The dashed line represents the acceptance criteria for cell viability according to ISO 10993-5:2009. (b) In vitro cytotoxicity evaluation of extracts of the GA-BP (control) and TSD-BP (n=4). (c) Fluorescence micrographs of HUVECs on GA-BP and TSD-BP after incubating for 24 h. The upper and the lower are, respectively, grayscale photographs and color photographs under fluorescence excitation with phalloidin-rhodamine-marked F-actin in red and DAPI-marked nuclei in blue. “***” Means a very significant difference with p<0.001.

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    We further designed experiments followed by washing with DDM after incubation with different surfactants, which simulates the post-treatment or the second decellularization step in our TSD treatment of the valve. As a consequence, the toxicities of all surfactants were significantly reduced after the DDM treatment. The cell viabilities of the basic culture for 1 day were levelled to 80%, and those of the culture for 3 days were over 96%. These results imply that all the residues have been replaced during the second step of the biosurfactant-containing TSD strategy, namely, the DDM treatment.

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    We also checked cytotoxicity of the extract medium. As shown in Fig. 5(b), HUVECs exhibited much higher viability in the TSD-BP extracts compared to those without the second treatment with the biosurfactant. The direct contact experiment showed that the number of cells on the surface of the TSD-BP valve increased significantly, and the cells exhibited good adherent growth (Fig. 5c).

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    Both GA-BP and TSD-BP were stored in a 0.625% GA solution, and the valves were washed under the same condition prior to cytotoxicity tests. So, the lower cell viability of the group of GA-BP than that of TSD-BP in Fig. 5(b) did not arise from the presence of GA in the former group. The lower cytotoxicity of TSD-BP came from the second step of the biosurfactant-containing TSD strategy. The term “decellularization” in this study like in Fig. 1 usually indicates the 1st or traditional decellularization, because this process has removed almost all cells and most of cell debris. Nevertheless, the minor cell debris along with the remained strong surfactants used in the 1st decellularization could lead to significant cytotoxicity, as seen in Fig. 5. So, we introduced a post-treatment, or the 2nd decellularization to replace those minor yet cytotoxic surfactants, that constitutes the advantage of our TSD strategy.

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    Hemocompatibility

    To assess the hemocompatibility of the valve, the platelet adsorption assay was used to examine the coagulation properties of the specimen. TSD-BP showed slightly less adsorption to platelets than that of GA-BP (Fig. 6a). In addition, its hemolytic properties were evaluated with a semi-quantitative hemolytic test. As is seen from Fig. 6(b), the TSD-BP meets the safety standards for blood-contacting materials (hemolysis<5%).

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    fig

    Fig 6  Hemocompatibility of BP valves. (a) SEM images of BP surfaces with platelet adsorption; (b) Global views of hemolysis tests and quantitative results. Distilled water and normal saline were used as positive and negative controls, respectively. For each group, n=5. The dashed line represents 5%, the accepted lowest safety standards of hemolysis for blood-contacting materials according to ISO 10993-4:2002.

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    In vivo Calcification Evaluation

    Anticalcification is our primary goal of the modification of bioprosthetic valves. We examined in vivo calcification after implantation the BP sample in Wistar rats. As is schematically shown in Fig. 7(a), the valves were implanted and taken out after the pre-determined days. After dried, the valves experienced acid hydrolysis and ICP-OES tests. Some valves taken out after implantation for 30 and 60 days are globally viewed in Fig. 7(b). The degree of calcification can be roughly reflected from the areas of calcified plaque and curling levels of the valves. Large areas of flat, yellow plaques on the BPs were formed; in contrast, the TSD-BP exhibited the least area of calcification plaques on the dried valves along with relatively most transparent and curled appearance.

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    Fig 7  In vivo calcification evaluation of GA-BP, IPA-BP, and IPA-DDM-BP (TSD-BP). (a) Schematic representation of the site of subcutaneous implantation in a Wistar rat and the process for testing calcium content of the implants; (b) The global view of BP valves (1 cm × 1 cm) after being taken out from experimental rats on the indicated days of post-implantation and then dried; (c) Quantitative analysis of calcium content determined via ICP-OES.

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    We quantified the extent of calcium using ICP-OES, and the results are shown in Fig. 7(c). After implantation for 30 days, the content of calcification in the control group reached 90.6±3.9 μg/mg, and after 60 days, it reached 138.6±23.5 μg/mg. The group treated with IPA alone was slightly less calcified, with 45.2 μg/mg on day 30 and 93.9 μg/mg on day 60. In contrast, the final calcification level of TSD-BP was only 0.7±0.1 μg/mg on day 30 and 1.1±0.2 μg/mg on day 60. These results illustrated the effectiveness of our method which used a biosurfactant-containing TSD strategy to tackle with calcification of BHVs.

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    We also observed and semi-quantified the element distribution on the dried valves. In terms of calcium content, the EDS-SEM results in Fig. 8(a) are consistent with ICP-OES analysis in Fig. 7(c). There was a significant difference between the control group and the experimental group, and the calcification level of TSD-BP was much lower than that of GA-BP. Elemental analysis of the calcified and uncalcified samples by EDS-SEM showed that the P and Ca elements in the calcified specimens increased significantly, and the distributions of P and Ca elements basically coincide with each other. Moreover, the Ca/P stoichiometric ratio of GA-BP was 1.70 (Table S2 in ESI), which was basically consistent with the stoichiometric ratio of Ca/P (1.67) in hydroxyapatite (HAp, Ca10(PO4)6(OH)2).

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    Fig 8  Characterization of calcified specimens 60 days after implantation. (a) The upper images show the elemental mapping of GA-BP and TSD-BP by using scanning electron microscopy with energy dispersive spectroscopy (EDS-SEM). The lower images are typical EDS spectra and quantitative statistics of weight fractions of P and Ca elements of the specimens. (b) Typical XRD spectra of GA-BP and TSD-BP. (c) A typical FTIR spectrum of a calcified specimen.

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    To further confirm the calcified compounds of GA-BP samples experiencing in vivo biomineralization, we carried out XRD and FTIR measurements. From the XRD spectrum in Fig. 8(b), the diffraction peaks of the (211) and (310) crystal planes of HAp appeared in GA-BP. It can be seen from the FTIR spectrum of Fig. 8(c) that GA-BP was a complex of HAp and collagen. The 2930, 1654, 1552 and 1240 cm−1 signals were attributed to the characteristic peaks of the amide bond of collagen. The 3450 cm−1 signal was ascribed to the stretching vibration band of ―OH of Ca10(PO4)6(OH)2, and the 1037, 873, 605 and 565 cm−1 signals belonged to the characteristic absorption band of PO43−. These results confirmed that the calcifying compound was mainly composed of HAp.

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    Observations of Spatial Distribution and Orientation of Regenerated Collagen Fibrils and Calcified Aggregates in Tissues­

    The BP is composed of wavy type I collagen fibers and long elastic fibers, which consitutue the basic skeleton of a variety of connective tissues. One collagen molecule is made up of three polypeptide chains forming a superhelix approximately 1.5 nm in diameter.[

    41] These collagen molecules allign to orientated collagen fibrils, which provide the tissue with high tensile strength. In our in vivo calcified specimens, collagen fibrils coexisted with calcium nodules, as seen from SEM images in Figs. 9(a) and 9(b), We measured the fibril diameters by SEM, which read 85±19 nm. Therefore, a collagen fibril came from self-assembly of many collagen molecules.
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    Fig 9  SEM images of in vivo formed calcified sites with both calcium nodes and collagen fibrils. (a) A whole image with magnification ×10000; (b) local images with magnification ×20000 with cobblestone-like (upper row) and coral-like (lower row) morphology.

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    The microscopic morphology of the attached calcium clusters or nodules was observed in a higher magnification. We classified calcified crystal points roughlly into two types of microscopic morphologies, cobblestone-type and coral-type (Fig. 9b). In addition, the collagen fibrils were less ordered in the specimens with stronger calcification.

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    The physiology of connective tissues is highly dependent upon the organization of structural biomacromolecules, in particular, collagen.[

    42] In order to better assess the spatial distribution and orientation of regenerated collagen fibrils and calcified aggregates in tissues, histological sections were stained for GA-BP and TSD-BP after being implanted in rats for 60 days. The dye of Alizarin red preferred to be located in the calcium deposition area. As shown in Fig. 10, the basic valve in the control group was significantly calcified in vivo, and the whole valve was stained in bright red by Alizarin red, while TSD-BP had no noticeable calcified area. Masson staining indicated that TSD-BP collagen fibrils were complete and stable. In contrast, in the control group free of the biosurfactant treatment at the second stage, no clear wavy stripes were observed, and collagen fibrils were less oriented. HE staining showed that the fibrous capsule of TSD-BP had a clear boundary with the valve leaflets and there was basically no infiltration of inflammatory cells. In contrast, the fibrous capsule was thickened in the control group, and a large number of inflammatory cells gathered outside the valve.
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    fig

    Fig 10  Optical micrographs of tissue slides experiencing the indicated stainings. The slicing was made after taking out the valves from rats after subcutaneous implantation for 60 days.

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    In vivo Immunization Evaluation

    To explore the immune response of the external biomaterials, TSD-BP (experimental group) and GA-BP (control group) implanted in rats subcutaneously for 30 and 60 days were taken out for immunohistochemical staining. T cells are key cells of the adaptive immune system, and the characteristic marker of T cells is the cluster of differentiation 3 (CD3), a membrane protein. Macrophages belong to a core innate immune cell type; they functionally secrete a wide variety of cytokines and chemokines, which makes these cells powerful regulators in the host immune system. F4/80 is a marker for mature murine macrophages, which is divided into two types according to the stimulation-induced polarization, M1 and M2. M1 macrophages are pro-inflammatory and secrete inflammatory cytokines and chemokines, such as tumor necrosis factor-α (TNF-α). M2 macrophages are thought to have immunosuppressive effects and secrete anti-inflammatory factors such as interleukin 10 (IL-10), which affect tissue remodelling process of the valve in vivo.[

    43,44] Therefore, in this study, CD3, F4/80, TNF-α and IL-10 antibodies were used as markers for T cells, macrophages, M1 macrophages, and M2 macrophages, respectively. The immunohistochemical results are shown in Fig. 11. No matter 30 or 60 days after implantation, the inflammatory responses induced by TSD-BP were significantly lower than those of the control group.
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    Fig 11  Immune responses induced by GA-BP (control group) and TSD-BP (experimental group) after subcutaneous implantation in rats for 30 and 60 days. (a) Optical micrographs. Immunohistochemical staining was performed for sections towards CD3, F4/80, TNF-α and IL-10 markers. (b) Quantitative statistics of positive cells. For each specimen, three sites were randomly selected and Image J software was used for data collection. 100% means that all cells within the statistical field of view are antibody immunopositive. “**” and “***” mean significant differences with p < 0.01 and p < 0.001, respectively.

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    DISCUSSION

    Structural biomacromolecules are complementary to synthetic polymers,[

    45-48] and many biomacromolecules exist in the network of extracellular matrix (ECM).[49] Among various biomaterials developed so far,[50−57] bioderived materials, in particular, decellularized matrixes from some tissues, are unique owing to their biomechanical properties and natural bioactivities.[45] Nevertheless, a shortcoming of a bioderived material comes from its calcification preference in vivo when implanted in cardiovascular systems,[58] although calcification is expected for implants in musculoskeletal and some other systems. While cell-material interactions have been investigated from different aspects[59−65] and many approaches to improve the biocompatibility of medical materials have been suggested,[45,66−69] the anticalcification strategies for natural biomacromolecular networks are still limited and much desired to be developed.
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    Calcification is related to collagen-relevant mechanical injury in BPs, as revealed by an in vitro calcification experiment on valves after uniaxial cyclic stretching of 1 million times or failure.[

    70] The process of calcification formation in vivo is more complex, and the influence of cells and many other biological factors exacerbate calcification. Therefore, our experiments focused on in vivo calcification. We studied the spatial distribution and orientation of collagen fibrils and calcified aggregates in the calcified sample taken out of rats. The calcified collagen fibrils were less orientated. Our study has confirmed that the major components of calcified heart valves are HAp and collagen. The structural relationship between the collagen matrix and HAp remains unclear. In order to provide a theoretical support for preventing or promoting biomineralization dependent on the demand, identifying the fine structure and the formation mechanism of calcified collagen fibrils are expected. Some efforts have been made. For instance, Xu et al. proposed that intermolecular channels guide crystallographic orientation in mineralized collagen.[71] They revealed that confinement within the pores of collagen in human bone and anisotropic growth of HAp determine the orientation of HAp crystals within collagen fibrils and that HAp crystals are only uniaxially oriented relative to collagen. In the present study, the collagen fibrils were formed of diameter about 85 nm, and the inorganic clusters were formed afterwards with diverse sizes and cobblestone-like or coral-like morphology.
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    Calcification is a serious problem limiting the lifespan of a bioprosthetic valve in clinical applications. Phospholipids can bind calcium and play an essential role in the formation of calcium phosphate crystals. Combination of alcohol and surfactants has been reported to reduce calcification. In 1989, Jones et al. first used surfactants as an anticalcification treatment for bioprosthetic valves; they evaluated the impacts of four kinds of reagents (surfactants, phosphoramidate, toluidine blue and polyacrylamide) to valvular tissues for the mitral and tricuspid replacement in sheep and found that only the surfactants substantially reduced calcification.[

    72] In 2001, Shen et al. tried ethanol, ether and the surfactant Tween-80 to eliminate lipids such as free fatty acids, cholesterol and phospholipids; their results showed that the effective anticalcification treatment was the combination of ethanol and Tween-80 or the combination of ethanol, ether and Tween-80.[30] In 2017, Park et al. treated decellularized BP with 75% ethanol and 5% octanol to achieve anticalcification.[32] In 2019, Agathos et al. effectively attenuated calcification with 20% ethanol, 1.2% Tween 80 (surfactant), 1% calcitonin, 10% sodium bisulfite at 37 °C.[73] In most of previous reports, Tween-80 was chosen as the surfactant for the alcohol and surfactant combinations. However, Tween-80 is a synthetic surfactant with cytotoxicity. As previously reported, the cellular microenvironment constructed by implanted biomaterials is a central scientific issue.[74−76] For example, 3D cell-free porous scaffolds that mimic ECM promote cell adhesion and growth;[77] bioactive surface modification can improve biocompatibility of medical devices.[78,79] For acellular matrix implants, the cytotoxicity and bioactivity introduced by the matrix itself are critical. Therefore, it is particularly important to minimize cytotoxicity in the treatment of bioprosthetic valves.
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    In this study, we introduced n-dodecyl-β-D-maltoside (DDM) into the TSD strategy for the first time. Its chemical structure is shown in Fig. 1. As a mild non-ionic surfactant with a hydrophilic maltose head and a hydrophobic long-chain alkyl tail, DDM can effectively dissolve membrane-bound proteins. The primary properties of DDM are its ability to extract hydrophobins while maintaining protein conformation in the solution phase and to facilitate renaturation after the denaturation of proteins. The raw materials of DDM could be extracted from coconut oil and corn starch as reviewed in 1998,[

    33] which indirectly demonstrates the nontoxicity and biosafety of DDM as a mild glycoside biosurfactant. Alkyl polyglycosides are not toxic or harmful in acute toxicity tests based on available data.[34] In addition, the acute oral toxicity value (LD50, namely, half lethal dose) of DDM is a few grams per kilogram of body weight.[80] It is little probable for DDM to present significant toxicological risk based on the dose of DDM and the washing procedure in our experiment. Our experiments also confirmed the very low cytotoxicity of this biosurfactant (Figs. 5 and 6).
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    Our research used a natural plant-derived nonionic surfactant DDM in the post-treatment of GA-fixed valves, instead of any chemically synthesized or semi-synthetic surfactant. The experiment of subcutaneous implantation in Wistar rats proved its effectiveness in anticalcification (Figs. 7−11). While the mechanisms underlying the anticalcification efficacy remain open at the moment, we suppose that the following aspects might be taken into consideration: (1) the alcohol and surfactant can remove residue lipids, mainly phospholipids in cell membranes and cell remnants in decellularized valve tissue, preventing calcium from binding to phosphate and inhibiting the formation of calcium salts; (2) the mild surfactant can replace some residue cytotoxic surfactants in the first decellularization of the valve; (3) the biosurfactants can shadow the surface charges of the membrane, matrix or collagen fibrils; (4) the surfactants with hydrophilic groups can form a barrier to prevent the phospholipids from penetrating into valvular tissue. Some viewpoints are stimulated from the literature.[

    72] Changes in the surface charge distribution and hydrophobicity of a material might affect its protein adsorption after implantation. Preventing nonspecific adsorption of proteins is important to inhibit thrombosis, immune response, and calcification for a bioprosthetic or polymeric heart valve.[81] We would like to specifically indicate that DDM and IPA might not only remove phospholipids, but also remove residual toxic substances in the system, thereby reduce the immune response and in vivo calcification. Further studies are required to resolve the detailed mechanisms.
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    CONCLUSIONS

    In this study, we propose a “two-step decellularization” (TSD) strategy by using a biosurfactant at the second stage for the treatment of bioprosthetic valves. Experiments in vitro and in vivo confirmed that BHVs treated by TSD (IPA-DDM treatment following GA crosslinking of the decellularized matrix) exhibited high anticalcification efficacy as well as low cytotoxicity and excellent mechanical properties. In particular, after 60 days of implanting a BHV into a rat, the calcium content in the TSD-treated implant was only 1.1 μg/mg in contrast to 138.6 μg/mg without the biosurfactant-containing post-treatment. In addition, the morphology and composition of the calcified sites were also investigated. The revealed structure of the calcification site with both collagen fibrils and calcified clusters might be helpful for understanding the in vivo biomineralization containing both biomacromolecules and inorganic compounds under tissue microenvironment. This work provides a facile and effective approach to enhance the anticalcification of BHVs. Moreover, the employment of the mild biosurfactant DDM will broaden the horizons for researchers who attempt to improve the service life and safety of interventional bioprosthetic products in the fields of cardiovascular and other biomedical materials.

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