In situ regeneration is a promising strategy for constructing tissue engineering heart valves (TEHVs). Currently, the decellularized heart valve (DHV) is extensively employed as a TEHV scaffold. Nevertheless, DHV exhibits limited blood compatibility and notable difficulties in endothelialization, resulting in thrombosis and graft failure. The red blood cell membrane (RBCM) exhibits excellent biocompatibility and prolonged circulation stability and is extensively applied in the camouflage of nanoparticles for drug delivery; however, there is no report on its application for large-scale modification of decellularized extracellular matrix (ECM). For the first time, we utilized a layer-by-layer assembling strategy to immobilize RBCM on the surface of DHV and construct an innovative TEHV scaffold. Our findings demonstrated that the scaffold significantly improved the hemocompatibility of DHV by effectively preventing plasma protein adsorption, activated platelet adhesion, and erythrocyte aggregation, and induced macrophage polarization toward the M2 phenotype in vitro. Moreover, RBCM modification significantly enhanced the mechanical properties and enzymatic stability of DHV. The rat models of subcutaneous embedding and abdominal aorta implantation showed that the scaffold regulated the polarization of macrophages into the anti-inflammatory and pro-modeling M2 phenotype and promoted endothelialization and ECM remodeling in the early stage without thrombosis and calcification. The novel TEHV exhibits excellent performance and can overcome the limitations of commonly used clinical prostheses.
Valvular heart disease (VHD) is a significant domain within the realm of cardiovascular diseases predominately caused by degenerative diseases, rheumatic heart disease, or congenital heart diseases. VHD poses a substantial threat to millions of people and is the primary cause of mortality among patients [1]. Heart valve (HV) replacement is the only effective treatment hitherto for patients with unrepairable VHD [2]. However, the available HV prostheses, encompassing mechanical and bioprosthetic heart valves (BHVs), face numerous shortcomings. Mechanical HVs are susceptible to thrombosis and embolic events. Therefore, patients undergoing mechanical valve replacement need lifelong anticoagulation therapy, which increases the risk of hemorrhage or thromboembolism. The hemocompatibility of BHV is favorable and patients undergoing BHV replacement need anticoagulants for only 3-6 months. However, BHV is prone to calcification and structural valve degradation, which limits the durability of BHV to approximately 15 years. Furthermore, the absence of valve growth necessitates multiple surgical procedures for pediatric patients. Research on VHD focuses on new valvular prostheses, including polymer valves and mechanical valves with favorable anticoagulation properties and in situ tissue engineering heart valves (TEHVs) [3]. In situ TEHV that can repair, remodel, and regenerate by controlling the host response to implanted biomaterials holds significant promise for treating VHD.
Decellularized heart valves (DHVs) are derived from either homogenous or xenogeneic HVs using suitable decellularization techniques. DHV is the most commonly used scaffold for TEHV construction, preserving the native structure and preventing the immune response. Commercial DHV xenografts have been used in pre-clinical and clinical trials, including CorMatrix and Matrix P [4], [5], [6]. These prostheses exhibited limited hemocompatibility and a lack of remodeling and recellularization. Moreover, the DHV scaffolds can elicit a robust immune response, resulting in leaflet thickening, calcification, and graft failure [4], [7], [8]. Furthermore, the DHV scaffold with suboptimal recellularization will be exposed to plasma proteins; hence, activating platelet (PLT) adhesion and resulting in thrombosis [9], [10]. Therefore, modifications improving hemocompatibility and accelerating recellularization after implantation can promote in situ regeneration of DHV scaffolds and extend their durability [11].
The red blood cell membrane (RBCM) exhibits exceptional hemocompatibility and anti-biofouling properties due to the hydrophilic head of the phospholipid bilayer and the steric hindrance induced by surface glycoproteins [12]. These advantageous properties, coupled with its abundance and accessibility, made RBCM the primary option for cell membrane coating. Additionally, RBCM-coated nanoparticles have been extensively used in drug delivery to prevent immune recognition and ensure prolonged stability in blood circulation [13]. RBCM coating is an auspicious candidate for surface modification, attenuating the immune response, and accelerating recellularization, but has been used only for decellularized extracellular matrix (dECM).
The extracellular matrix (ECM), the main component of DHV, is negatively charged and electrically repulsive to RBCM [14], undermining the stability of RBCM modification. Therefore, introducing other substances to co-modify DHV with RBCM and construct a stable coating can improve the function of RBCM. Polyethyleneimine (PEI) is a strong and positively charged cationic polymer that is widely used in layer-by-layer (LbL) self-assembly. However, abundant amino groups of PEI lead to a certain degree of cytotoxicity; thus, the application of PEI in biomaterials necessitates the introduction of other group modifications [15]. The catechol group inspired by mussel is a common group introduced to enhance adhesion to various substrates by covalent and noncovalent interactions. However, the stability of catechol functionalized PEI (PEI-C) combined with RBCM coating on dECM is still uncertain.
In the present study, we applied the LbL self-assembly approach to develop a novel TEHV known as DHV-PEI-C-RBCM, which employed strongly positive charged PEI-C to immobilize RBCM on the surface of DHV as an intermediary layer. The mechanical properties, hemocompatibility, and cytocompatibility of the TEHV were assessed in vitro. In addition, its histocompatibility and resistance against calcification and immune response were measured using the small animal subcutaneous embedding model and abdominal aorta implantation model. Our modification strategy can enhance the applicability of erythrocyte membranes in macroscopic-scale modification of biogenic materials, augment the hemocompatibility of TEHV, shift macrophage polarization toward the anti-inflammatory and pro-healing M2 phenotype, and facilitate in situ regeneration of valves.
2. Materials and methods
2.1. Materials
All chemicals were purchased from Sigma-Aldrich (USA), and all cell culture reagents were purchased from Gibco (USA) unless otherwise stated. PEI (MW = 25 000 kDa) and 3-(3,4-dihydroxyphenyl) propionic acid were purchased from MeilunBio Co., Ltd. (China). Cell counting kit-8 (CCK-8) and Calcein-AM/propidium iodide (PI) cell viability/cytotoxicity assay kit were purchased from Beyotime Co., Ltd. (China). Primary antibodies and fluorescent secondary antibodies were purchased from Abcam (UK).
2.2. Preparation of PEI-C and RBCM
2.2.1. Preparation of PEI-C
PEI-C was prepared as described previously [16]. In brief, 3.00 g of PEI was dissolved in 300 mL of phosphate-buffered saline (PBS), and the solution pH was adjusted to 5.5 using 1 mol∙L−1 HCl. Then, 1.52 g of 3-(3,4-dihydroxyphenyl) propionic acid and 2.71 g of 1-(3-dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride (EDC) were added to the solution, and the pH was adjusted to 5.5 using 1 mol∙L−1 NaOH. The reaction mixture was stirred using a magnetic stirrer at room temperature for 2 h. Extensive dialysis was performed to eliminate unreacted reagents and byproducts.
2.2.2. Derivation of RBCM
Whole blood was obtained from healthy male Sprague Dawley (SD) rats (body weight 150-200 g) via cardiopuncture and collected in acid-citrate-dextrose (ACD) solution-containing tubes (1:9, v/v), centrifuged at 1600 r∙min−1 for 10 min at 4 °C, and washed three times with equal volume of PBS to obtain red blood cells (RBCs). Washed RBCs were resuspended in 0.25 × PBS solution for 30 min, and then centrifuged at 12 000 r∙min−1 for 5 min at 4 °C to obtain RBC ghost. RBCM was harvested after repeated cycles of freezing and thawing.
2.3. Preparation of DHVs and glutaraldehyde-cross-linked heart valves (GLUT-HVs)
DHV was prepared as previously described [17]. Briefly, porcine aortic valves harvested from a local slaughterhouse were rinsed in cold heparin saline. Valves were subsequently soaked in M3 system (2% w/v 3-[(3-cholanidopropyl)dimethylammonio]-1-propane sulfonate (CHAPS) and 2 mmol∙L−1 tributyl phosphate (TnBP) in 40 mmol∙L−1 Tris-HCl) at 110 r∙min−1, 37 °C for 24 h; and in the M4 system (2% w/v CHAPS, 2% w/v sulfobetaine 3-10, 1% w/v amidosulfobetaine-14, and 2 mmol∙L−1 TnBP in 40 mmol∙L−1 Tris-HCl) at 110 r∙min−1, 37 °C for another 24 h. After sufficient rinse in PBS, valves were treated with a nuclease system for 24 h to obtain DHVs, which were stored at 4 °C. GLUT-HV was prepared from native HV cross-linked with 0.625% w/v glutaraldehyde solution at 110 r∙min−1, 37 °C for 24 h in dark.
2.4. Formation of PEI-C and RBCM coating on DHV
DHV-PEI-C and DHV-PEI-C-RBCM were prepared by coating PEI-C and RBCM on DHV via LbL self-assembly. The clean DHV substrates were immersed in PEI-C solution (10 mg∙L−1) at 110 r∙min−1, 37 °C for 24 h to obtain DHV-PEI-C. After sufficiently rinsing in PBS several times, DHV-PEI-C was submerged in RBCM suspension to obtain RBCM-coated DHV-PEI-C (DHV-PEI-C-RBCM). RBCM suspension was previously treated with a bath sonicator at room temperature for 24 h. Subsequently, DHV-PEI-C-RBCM was adequately washed with 1× PBS.
2.5. Atomic force microscope (AFM) scanning and roughness analysis
To detect the surface roughness of DHV, DHV-PEI-C, and DHV-PEI-C-RBCM, the AFM measurements were conducted using multimode 8 scanning probe microscope (Bruker Corporation, USA) with a Nanoscope V controller under ambient conditions (25 °C, relative humidity of 25%). Nanoscope software was used to analyze the roughness.
2.6. Hemocompatibility test
2.6.1. PLT absorption test and lactate dehydrogenase (LDH) assay
Whole blood was obtained as described before. Platelet-rich plasma (PRP) was prepared by centrifuging at 1600 r∙min−1 for 10 min. Different valve samples were incubated with PRP at 37 °C for 1 h, and washed with PBS. LDH assay (Beyotime Co., Ltd.) was used to quantify the adhered PLTs. After fixing with 2.5% w/v glutaraldehyde solution and subsequent lyophilization, the samples were visualized by scanning electron microscope (SEM).
2.6.2. Ex vivo cervical arteriovenous (A-V) shunt experiment
DHV, DHV-PEI-C, DHV-PEI-C-RBCM, and GLUT-HV samples were cut into rectangles (10 mm × 7 mm) and enclosed into the inner surface of a polyvinyl chloride (PVC) catheter. Japanese white rabbits were anesthetized by 1% w/v pentobarbital sodium solution. After isolating the left carotid artery and right jugular vein, ex vivo circulation was built through an intubation tube and catheters allowing the blood flow to contact the samples. After contacting the blood flow for 90 min, the catheters were separated and the samples were washed in PBS. Images were captured to observe thrombi on the samples. Then, the samples were fixed with 2.5% w/v Glut solution before visualization by SEM.
2.6.3. Hemolysis ratio assessment
To measure the hemolysis ratio of samples, pure RBC was obtained from citrated blood samples by centrifugation at 4000 r∙min−1 for 10 min. All samples (n = 6) were cut into plates with a diameter of 7.9 mm and incubated with RBC suspension (RBC 200 μL in normal saline 1 mL) at 37 °C for 1 h. The positive and negative controls were prepared by adding 200 μL RBC to 1 mL of diluted water and normal saline, respectively. After incubation, all tubes were centrifugated at 4000 r∙min−1 for 10 min, and then images were taken. Using a microplate reader, 200 μL of supernatant was collected to detect optical density (OD) at 540 nm. The hemolysis ratio was calculated as follows:
2.6.4. Plasma proteins absorption test
DHV, DHV-PEI-C, DHV-PEI-C-RBCM, and GLUT-HV were cut into plates with a diagram of 6.35 mm, washed in PBS, and placed in a 96-well plate. All samples were incubated with 200 μL of fluorescein isothiocyanate-labeled bovine serum albumin (FITC-BSA) or fibrinogen (FITC-Fg) at a concentration of 1 mg∙mL−1 for 1 h. The fluorescence images were acquired using a fluorescence microscope (AXIO observer 7; Zeiss, Germany). ImageJ software was used to identify the average fluorescence intensity.
2.7. Rat subcutaneous embedding model
The details of model construction and histological staining are provided in Appendix A. Rats were euthanized at 7, 14, or 28 days after the surgery. The implants were dissected and fixed. ECM changes were measured by hematoxylin and eosin (H&E), Masson, and Elastica van Gieson (EVG) staining. Calcification was evaluated by Von Kossa staining. The in vivo relative degradation was calculated by:where d0 and d28 refer to the thickness of the sample before subcutaneous embedding and 28 days after the surgery.
2.8. Rat abdominal aorta implantation model
The rat abdominal aorta implantation model was used to investigate biomechanical properties and biocompatibility under hemodynamic conditions. To construct the model, DHV, DHV–PEI-C–RBCM, and GLUT-HV were sheared into rectangles (5 mm × 5 mm) and then stitched into tubules with the ventricular surface facing inward. SD rats (male, 180 g) were anesthetized by inhaling 3% isoflurane, and the abdominal aorta was exposed through laparotomy. Then, 1 cm of the abdominal aorta was dissected, blocked with an atraumatic artery clip, and bluntly severed. In 30 min, the valvular tubular was implanted into the abdominal aorta via end-to-end anastomosis with approximately ten sutures (9-0 prolene; Ethicon, USA). We subsequently removed the artery clip, compressed for hemostasis, and closed the abdomen. Hindlimb movement was assessed every two days after the surgery, and the patency of the graft was measured four weeks after the surgery using Doppler ultrasound. Rats were sacrificed at 7, 14, and 28 days after implantation to assess ECM changes, calcification, inflammation, cellularization, and remodeling. The calcium (Ca) contents of all samples were detected using an inductively coupled plasma optical emission spectrometer (ICP-OES). In brief, all samples were lyophilized and the initial weight (m0, mg) was measured. After digesting with nitric acid and hydrogen peroxide, the solution was filled with ultrapure water to reach 25 mL volume. Then, the Ca concentration (Cx, μg∙mL−1) was measured using ICP-OES, and the Ca content (CCa, μg∙mg−1) was calculated as:
2.9. Statistical analysis
All quantitative data are expressed as mean ± standard error of the mean (SEM). Statistical analyses were conducted using GraphPad Prism or Origin software. Significant differences were assessed using one-way analysis of variance (ANOVA) followed by post hoc Tukey’s multiple comparisons. p < 0.05 was considered significant (not significant (ns), p > 0.05; *p < 0.05, **p < 0.01, ***p < 0.001).
3. Results
3.1. Fabrication of PEI-C and RBCM coating on DHV
We used 3-(3,4-dihydroxyphenyl) propionic acid to introduce the catechol group to PEI and synthesize a positively charged mussel-inspired polymer PEI-C under EDC catalysis. RBCM was prepared by hypotonic lysis from erythrocytes. The construction of DHV and the LbL assembly of DHV-PEI-C and DHV-PEI-C-RBCM are shown in Fig. 1.
The dynamic light scattering (DLS) detection revealed that the size of RBCM had an unimodal distribution with a mean particle size of 164.1 nm, and the distribution was consistent with transmission electron microscope (TEM) images (Figs. 2(a) and (b)). In addition, Fourier transform infrared spectroscopy (FTIR) analysis indicated the peaks observed at 1231 cm−1 for P=O stretching vibration and 1054 cm−1 for P-O-C stretching vibration, which are abundant in the phospholipid bilayer. These findings suggest the successful extraction of RBCM (Fig. S1(a) in Appendix A) [18]. The H nuclear magnetic resonance (HNMR) spectroscopy of PEI-C indicated that PEI-C had the proton of the benzene ring of 3-(3,4-dihydroxyphenyl) propionic acid at 6.5-7.0 parts per million (ppm), indicating successful synthesis (Fig. S1(b) in Appendix A). As shown in Fig. 2(c), PEI-C had a positive value of +74.1 mV and RBCM had a negative value of −52.3 mV, suggesting that PEI-C and RBCM were oppositely charged, had stable systems without aggregation, and could assemble on the surface of DHV [19]. Co-culture with macrophages supported the immunomodulatory function of RBCM. Immunofluorescence staining revealed that the expression of cluster of differentiation 206 (CD206) increased after co-culturing with RBCM (Fig. S1(c) in Appendix A). This finding suggests that the erythrocyte membrane can promote the differentiation of macrophages into the M2 phenotype, which is known for its immunomodulatory properties.
The general view of the valve leaflet changed from translucent silk white to brown after successively cross-linking in PEI-C solution and RBCM suspension. This change was dominantly derived from the color of PEI-C (Fig. 2(d)). SEM images showed that the wavy, branching, and curved fibers on the surface of DHV were flattened after PEI-C cross-linking. Granular structure particulates bound to the surface after RBCM coating, indicating the successful modification of PEI-C and RBCM. As the sign of cell membrane, the phosphorus element energy-dispersive spectrum (EDS) dot scanning of DHV-PEI-C-RBCM was significantly more incremental than that of DHV and DHV-PEI-C (Figs. 2(d) and (e)). To label RBCM on the surface, we used 1,1′-dioctadecyl-3,3,3′,3′-tetramethylindocarbocyanine perchlorate (DiI), a lipophilic tracer that emits red fluorescence after being combined with phospholipid. The fluorescence intensity of DHV-PEI-C-RBCM was markedly higher than that of DHV and DHV-PEI-C. The P element was accurately quantified using ICP-OES by Thermo Fisher Scientific iCAP PRO (Fig. 2(f)). The phosphorus content of DHV-PEI-C-RBCM (0.149% ± 0.036%) was significantly higher than that of DHV (0.057% ± 0.014%) and DHV-PEI-C (0.053% ± 0.020%), indicating that phospholipids were present on the surface of DHV-PEI-C-RBCM. These results confirmed the coating of RBCM on the surface of DHV.
X-ray photoelectron spectroscopy (XPS) analysis was conducted to investigate element characteristics on the surface of the three types of valves. DHVs undergoing decellularization were rich in proteins, like collagen and elastin, whereas phospholipids were the major component of the cell membrane [10], [20]. Thus, DHV-PEI-C-RBCM had a relatively higher proportion of P elements compared to DHV and DHV-PEI-C. As shown in Figs. 2(g) and (h), the phosphorus 2p (P2p) peak of DHV-PEI-C-RBCM at 133.5 eV was significantly higher than that of DHV and DHV-PEI-C. Due to the lack of -COOR in PEI-C, the high-resolution scanning of C1s spectra revealed the absence of the -COOR group in DHV-PEI-C and the recurrence of the -COOR group in DHV-PEI-C-RBCM, suggesting the successful modification of PEI-C and RBCM (Fig. S1(d) in Appendix A). The stability of RBCM coating is another issue worth considering. As shown in Figs. S1(e) and (f) in Appendix A, the fluorescent intensity of DHV-PEI-C-RBCM was significantly stronger than that of DHV-RBCM. There was still a significant fluorescence signal on the surface of DHV-PEI-C-RBCM, suggesting that PEI-C stabilized RBCM coating. These findings prove the success of the modification strategy.
3.2. Characterization of the DHV-PEI-C-RBCM
3.2.1. Surface roughness and hydrophilia
Atomic force microscopy (AFM) was conducted to observe the surface morphology of the scaffolds. As shown in Figs. 3(a) and (b), three-dimensional (3D) scanning of AFM indicated that modification of RBCM significantly increased the Ra value, an indicator of roughness, mainly due to the granular structure of particulates on the surface. Water contact angle (WCA) measurement showed that PEI-C modification decreased wettability. After subsequent modification of RBCM, the DHV-PEI-C-RBCM scaffold was significantly more hydrophilic, even compared to DHV. This phenomenon was due to the hydrophilic head of phospholipids (Fig. 3(c)). The rougher surface and better hydrophilicity may provide better cytocompatibility for DHV-PEI-C-RBCM, promoting cell adhesion and proliferation on the valve surface.
3.2.2. Mechanical properties
Adequate mechanical strength is critical in the construction of TEHV. To investigate the mechanical properties of the three scaffolds and compare them with those of bioprosthetic valves, natural porcine GLUT-HV, DHV, DHV-PEI-C, and DHV-PEI-C-RBCM samples were assessed by the uniaxial tensile test. The macroscopic graphs revealed that the sagging DHV leaflet became stiffer after PEI-C coating, while there was no significant alteration after RBCM coating (Fig. S2 in Appendix A). The stress-strain curves of the four samples are displayed in Fig. 3(d). The slope of curves hinged on the tangent modulus of samples. Steeper curves were observed in DHV-PEI-C, DHV-PEI-C-RBCM, and GLUT-HV compared to DHV, indicating better mechanical properties (Fig. 3(d)). The statistical analysis of ultimate tensile strength, ultimate strain, and Young’s modulus also confirmed the results (Figs. 3(e)-(g)). In contrast, there was no statistically significant difference between DHV-PEI-C, DHV-PEI-C-RBCM, and GLUT-HV. The result of the uniaxial tensile test indicated that PEI-C modification enhanced the mechanical properties of DHV, which was comparable to bioprosthetic valves in clinical practice.
3.2.3. Enzymatic hydrolysis stability
As the main component of ECM, the stability of collagen and elastin determines the durability of the HV. Collagenase and elastase solutions were respectively used to decompose the scaffolds in vitro. Then, dry weight loss was calculated to investigate the degradation rate. As expected, the degradation rate of DHV after collagenase treatment (77.612% ± 9.125%) was significantly higher than that of DHV-PEI-C, DHV-PEI-C-RBCM, and GLUT-HV (Fig. 3(h)). Therefore, there was a significant difference in the degradation rate between DHV-PEI-C-RBCM and DHV after elastase treatment (Fig. 3(i)). These results indicated that PEI-C and RBCM coating stabilized dECM, prevented degradation by proteases, and effectively extended durability.
3.3. The cytotoxicity and immunomodulatory properties of DHV-PEI-C-RBCM
3.3.1. Cytotoxicity
After implantation, the scaffolds interact with diverse types of cells, such as endothelial cells (ECs) in the circulating blood. In this study, human umbilical vein endothelial cells (HUVECs) were seeded on the surface of the scaffolds to investigate cytotoxicity. For cell adhesion, the fluorescence images taken at 2 and 4 h exhibited a gradual increase in cell population on the surface of DHV, DHV-PEI-C, and DHV-PEI-C-RBCM, while a significant decrease in cell population on the surface of GLUT-HV (Fig. 4(a)). The results of cell proliferation experiment revealed that an incomplete EC layer was observed in DHV, DHV-PEI-C, and DHV-PEI-C-RBCM groups after 3 days. After 7 days, a complete EC layer formed in these groups with no sign of dead cells. In contrast, no viable cells were observed in the GLUT-HV group at day 3 and day 7 (Fig. 4(b)). CCK-8 assay was used to measure the viability of surface cells on the scaffold. The cell proliferation curve validated the above conclusions (Fig. 4(c)). Furthermore, HUVECs were cultured with scaffold extracts. GLUT-HV extracts exhibited a significant inhibitory effect on cell proliferation, compared with DHV, DHV-PEI-C, and DHV-PEI-C-RBCM (Fig. 4(d)). These findings indicate that DHV-PEI-C-RBCM possesses favorable cytocompatibility and can achieve recellularization and in situ regeneration, suggesting its potential for regenerative medicine.
3.3.2. The immunomodulatory property of DHV-PEI-C-RBCM
We used RAW 264.7 cells to investigate the immune microenvironment and macrophage polarization. The immunofluorescence staining of inducible nitric oxide synthase (iNOS; M1 phenotype marker) and CD206 (M2 phenotype marker) demonstrated that compared with DHV, differentiation into the M2 phenotype significantly increased after RBCM modification (Fig. 4(e)). Inflammatory cytokine chip testing in the culture supernatant revealed that compared with DHV, the levels of pro-inflammatory cytokines, including interleukin-1β (IL-1β), IL-2, IL-6, and tumor necrosis factor-α (TNF-α), significantly decreased after RBCM coating (Fig. 4(f)). Additionally, there was a notable increase in the concentrations of anti-inflammatory cytokines, such as IL-4 and IL-10. These anti-inflammatory cytokines promote macrophage polarization toward the M2 phenotype. These results demonstrate that DHV-PEI-C-RBCM can shift macrophage polarization toward the M2 phenotype.
3.4. Hemocompatibility of DHV-PEI-C-RBCM
3.4.1. Thrombosis
The anti-thrombotic capacity of DHV-PEI-C-RBCM was assessed under dynamic blood flow and static situations. Fig. 5(a) exhibits the schematic representation of the ex vivo cervical A-V shunt model. The gross view of the model and the close images of the samples are shown in Fig. S3(a) in Appendix A. DHV and DHV-PEI-C catheter samples showed marked blockade after 2 h of circulation (Fig. S3(b) in Appendix A). Furthermore, an obvious thrombus was observed on the surface of DHV, while fewer blood components were absorbed on DHV-PEI-C. As expected, almost no thrombus was found on DHV-PEI-C-RBCM and GLUT-HV (Fig. 5(b)). SEM images indicated that many erythrocytes gathered on the surface of DHV, while few erythrocytes were observed on DHV-PEI-C-RBCM and GLUT-HV (Fig. 5(c)). In vitro experiments also confirmed this conclusion (Fig. S4 in Appendix A), indicating that PEI-C-RBCM modification prevented thrombosis compared with DHV.
3.4.2. PLT adhesion
In vitro PLT adhesion test was performed using PRP. A great number of activated PLTs adhered to the surface of DHV, with extending fused dendritic pseudopodia (Fig. 5(d)). Slighter non-activated discoidal PLTs were observed on DHV-PEI-C. After further modification of RBCM, PLTs were barely observed on DHV-PEI-C-RBCM. A similar result was observed in the GLUT-HV group. Quantitative detection of absorbed PLTs by the LDH kit confirmed the results (Fig. 5(e)). As expected, a higher OD value at 490 nm was measured in the DHV group, whereas the DHV-PEI-C group showed a significantly lower value, indicating more adhesion of PLTs in the DHV group. Furthermore, there were significantly fewer PLTs in the DHV-PEI-C-RBCM group, compared with the DHV-PEI-C group. Similar conclusions were obtained in the GLUT-HV group. The reduced number of PLTs on the surface of DHV-PEI-C-RBCM indicates the anti-biofouling properties of RBCM, suggesting the favorable blood compatibility of DHV-PEI-C-RBCM [12].
3.4.3. Hemolysis rate assessment
Images of the hemolysis test are shown in Fig. 5(f). The hemolysis rates of DHV, GLUT-HV, DHV-PEI-C, and DHV-PEI-C-RBCM were 0.363% ± 0.158%, 0.313% ± 0.198%, 0.365% ± 0.095%, and 0.343% ± 0.052%, respectively, and there was no statistical difference between groups (Fig. 5(g)). All hemolysis rates were significantly lower than 5%, which met the requirements of blood contact products for clinical application. These findings indicated that neither PEI-C nor RBCM modifications induced hemolysis.
3.4.4. Plasma proteins absorption assay
As the components of plasma, BSA with negative charges and Fg with positive charges were labeled with fluorescein isothiocyanate and incubated with samples to assess their ability to resist the absorption of plasma proteins. For FITC-BSA, elevated fluorescence was observed in the DHV group, indicating large absorption of both types of proteins on dECM scaffolds. The fluorescence intensity was significantly reduced after the coating of PEI-C on the DHV substrate. Further reduction with a statistical difference was observed in the DHV-PEI-C-RBCM group, compared to the GLUT-HV group (Figs. 5(h) and (i)). The same tendency was observed in FITC-Fg-treated samples (Figs. 5(j) and (k)). These results prove the good hemocompatibility of DHV-PEI-C-RBCM.
3.5. In vivo assessment of DHV-PEI-C-RBCM grafts after subcutaneous embedding
The results confirmed the outstanding cytocompatibility, immunomodulatory, and hemocompatibility properties of DHV-PEI-C-RBCM. Thus, after implantation, the remodeling process of DHV-PEI-C-RBCM was compared to those of DHV and GLUT-HV groups.
3.5.1. In vivo degradation and calcification assessment
The rat subcutaneous implantation model was applied to investigate in vivo histocompatibility of grafts. The representative histological images (H&E, Masson, EVG, and Von Kossa staining) and the relative degradation analysis of subcutaneous embedding are shown in Figs. 6(a) and (b) and Figs. S5-S7 in Appendix A. There were no cell nuclei in DHV and DHV-PEI-C-RBCM, indicating the favorable effect of decellularization. As expected, cells gradually infiltrated the DHV and DHV-PEI-C-RBCM grafts after embedding for 7, 14, and 28 days, but there was no infiltration in GLUT-HV at any time point (Fig. 6(a)). Masson and EVG staining showed obvious degradation of collagen and elastin in DHV, but not in DHV-PEI-C-RBCM and GLUT-HV (Figs. S5 and S6). Thickness measurement before and after 28 days of embedding also indicated that the degradation rates of DHV-PEI-C-RBCM and GLUT were significantly lower than that of DHV (Fig. S7). Von Kossa staining was conducted to investigate the calcification of the grafts. The results indicated that calcified nodules were only observed in the GLUT-HV group, and the calcification degree increased over time (Fig. 6(b)). These results confirmed the biotoxicity and calcification of glutaraldehyde cross-linked valves and were consistent with clinical results [1].
3.5.2. Assessment of immune responses
The infiltrated immune cells were identified by immunofluorescence staining of CD3 (T cell marker), CD68 (macrophage marker), iNOS (M1 phenotype marker), and CD163 (M2 phenotype marker). In general, all samples exhibited certain immune responses.
The results of T cell and macrophage infiltration after implantation is shown in Fig. 6(c). Generally, there was a gradual increase in the proportion of macrophages in both DHV and DHV-PEI-C-RBCM groups. At 7 days, a higher abundance of T cells was noted in DHV grafts, whereas the DHV-PEI-C-RBCM group exhibited a predominance of macrophages. By day 14, an increased abundance of macrophages was observed in the DHV-PEI-C-RBCM group with a low abundance of T cells. After 28 days, a high number of macrophages were seen in the DHV group along with a few T cells. Due to the cytotoxicity of GLUT-HV, immune cells only infiltrated into the fibrous tissue surrounding the grafts, but not the graft itself.
M1 and M2 phenotypes of macrophages were identified by immunofluorescence staining (Fig. 6(d)). M1 macrophages were observed in DHV grafts 7 days after implantation, whereas M2 macrophages were barely seen. The number of M1 macrophages gradually decreased at 14 and 28 days after implantation, and the number of M2 macrophages peaked at 14 days, suggesting that the acute inflammation subsided and the remodeling process was initiated. As for DHV-PEI-C-RBCM, the acute immune response was suppressed and the remodeling occurred after 7 days, showing its anti-inflammatory properties. M1 and M2 phenotypes failed to infiltrate into GLUT-HV grafts. Consistent with in vitro results, these findings illustrate the enhanced adaptability of DHV-PEI-C-RBCM to the immune response elicited by DHV.
3.6. In vivo performance of the scaffolds under hemodynamic environment
3.6.1. In vivo degradation and calcification assessment
The rat abdominal aorta implantation with an end-to-end anastomosis model was established to assess the performance of grafts in the presence of arterial blood flow [21]. Fig. 7(a) depicts the construction of the rat model of abdominal aorta implantation. DHV had a tendency to dilate over time, and an aneurysm was formed after 28 days. On the contrary, there was no obvious dilatation in the other two groups, and the grafts maintained their vascular shape. Doppler ultrasound indicated that all grafts were still open after 28 days; however, the lumen of the DHV was dilated. In addition, calcified stenosis was observed in the GLUT-HV group, resulting in increased blood flow velocity (Fig. 7(b), Fig. S8 in Appendix A).
Histological examination confirmed similar results in abdominal aorta xenografts and embedding grafts (Fig. 7(b), Figs. S9-S12 in Appendix A). To more precisely characterize calcification, we conducted Ca quantification. The results showed that the Ca content gradually increased over time. The Ca content of DHV and DHV-PEI-C-RBCM were similar but significantly lower than that of the GLUT-HV group (Fig. 7(c)).
3.6.2. Endothelialization and remodeling
The abdominal aorta implantation specimens were collected at 7, 14, and 28 days to investigate endothelialization and remodeling. Discontinuous CD31+ ECs were observed on the internal surface of the DHV-PEI-C-RBCM graft, while no obvious ECs were observed in the DHV group (Fig. 7(d)). After 28 days of implantation, there was a completed monolayer of ECs on DHV-PEI-C-RBCM, and some ECs were also found on DHV. As for the GLUT-HV group, no ECs were found on the inner surface at all three time points due to the cytotoxicity of glutaraldehyde.
The reproduction of autologous collagen represents in situ remodeling of ECM. Immunofluorescence staining exhibited that collagen I level gradually increased in DHV and DHV-PEI-C-RBCM over time (Fig. 7(e)), but collagen I level in DHV-PEI-C-RBCM was significantly more than that in DHV, indicating greater remodeling in DHV-PEI-C-RBCM. As expected, due to the cytotoxicity of glutaraldehyde, there was no cell infiltration and no collagen deposition in GLUT-HV at all time points. These findings illustrate that DHV-PEI-C-RBCM can facilitate endothelialization and ECM remodeling in vivo, underscoring its favorable regenerative efficacy.
3.6.3. The immunomodulatory property of DHV-PEI-C-RBCM
Immunofluorescence staining of CD3 (T cell marker), CD68 (macrophage marker), iNOS (M1 phenotype macrophage marker), and CD163 (M2 phenotype macrophage marker) was conducted to identify immune cells infiltrating into abdominal aorta grafts. Fig. 8(a) exhibits T cell and macrophage infiltration. Similar to the results of embedding, the abundance of macrophages increased at 14 days and decreased at 28 days after the surgery in DHV-PEI-C-RBCM. At day 14, the abundance of macrophages increased in DHV-PEI-C-RBCM, whereas almost no T cells were observed in DHV-PEI-C-RBCM. A high abundance of macrophages was observed in the DHV group after 28 days. In addition, a few T cells were also observed in this group. Due to its cytotoxicity, there was no immune cell infiltration in the GLUT-HV group at any of three time points.
M1 and M2 macrophages were identified by immunofluorescence staining (Fig. 8(b)). At 7 days post-implantation, M1 macrophages were observed in DHV grafts, while M2 macrophages were rarely detected. Compared to the DHV group, a markedly earlier disappearance of acute immune response and initiation of remodeling was evident in the DHV-PEI-C-RBCM group after 7 days, highlighting its potential anti-inflammatory properties. Similar to the staining results of T cell and macrophage infiltration, both M1 and M2 macrophages failed to infiltrate into GLUT-HV grafts. Semi-quantitative analysis of macrophages/T cells and M1/M2 ratios indicated that the DHV-PEI-C-RBCM scaffold showed earlier macrophage infiltration and enhanced macrophage polarization toward the M2 phenotype (Figs. 8(c) and (d)).
4. Discussion
The repairing, remodeling, and regenerative properties of TEHV make it an ideal candidate for addressing the limitations of current valve prostheses [22], [23]. In this study, we innovatively constructed a TEHV by modifying PEI-C and RBCM on DHV and subsequently measured the properties of this TEHV. We found that: ① The PEI-C layer immobilized RBCM on the surface of DHV and enhanced its mechanical properties and resistance to enzymatic degradation; ② stable RBCM coating significantly improved hemocompatibility by reducing the absorption of plasma proteins, attenuating the adhesion and activation of PLTs, and preventing erythrocyte deposition; ③ the DHV-PEI-C-RBCM scaffold promoted cell adhesion and proliferation and macrophage polarization; ④ the DHV-PEI-C-RBCM scaffold regulated macrophage differentiation into the M2 phenotype and accelerated endothelialization and remodeling after implantation.
Due to the negative charge of both RBCM and dECM, the challenge of modification instability due to electrical rejection limits the practical application of RBCM. Mussel-inspired materials exhibited extensive application in bio interfaces due to their strong toughness, wet adhesiveness, and self-healing properties [24], [25], [26], [27]. To improve the stability of RBCM coating, positively charged PEI-C was introduced into the modification system. Serving as a mussel-inspired polymer, PEI-C with catechol group was tightly bound to the surface of DHV as an intermediate coating for more modification [27]. The catechol group plays a significant role in enhancing the coating stability of large-scale RBCM-modified metal materials [28]. However, further studies are needed to assess the stability of biogenic materials modified by RBCM. The immobilization of RBCM coating on the surface of DHV is primarily achieved through the interactions between positive and negative charges and the interactions between the benzoquinone group of PEI-C and nucleophilic functional groups on RBCM, such as the amine group, via Schiff base/Michael addition reaction [29], [30], [31]. The subsequent morphological and elementary analysis of DHV-PEI-C-RBCM provided further evidence for the presence of the cell membrane, demonstrating the robust adhesive capability of PEI-C. Moreover, catechol group modifications have been reported to significantly enhance the mechanical properties of the substrates [16], [32], [33]. The aortic valve is exposed to a high pressure of over 100 mmHg (1 mmHg = 0.133kPa) under the physiological hemodynamic condition, thus adequate mechanical strength is necessary to design an ideal valve prosthesis [3]. In the absence of adequate mechanical strength, the valve frequently suffers from accelerated deterioration, resulting in graft dysfunction [34]. Although glutaraldehyde crosslinking is commonly used to improve mechanical properties, the cytotoxicity and subsequent calcification of the valve limit its durability. The result of the uniaxial tension test showed that PEI-C significantly enhanced the mechanical properties of the valve mainly because of the strong covalent bonds between the catechol group of PEI-C and the amine group of DHV [35].
The hemocompatibility of valve prostheses is crucial for their clinical application. Previous studies have employed numerous modification strategies to mitigate the biofouling of blood constituents and enhance the hemocompatibility of valves [36], [37], [38], [39]. In the present study, we modified dECM with an RBCM coating to enhance its hemocompatibility. Compared to alternative modification strategies, RBCM can be easily obtained and prevents interference from organelle membranes. The static WCA analysis indicated an increase in the hydrophilicity of DHV-PEI-C-RBCM compared to DHV. Furthermore, in vitro incubation with plasma proteins, PRP, and whole blood and in vivo abdominal aorta implantation, and A-V shunt experiment demonstrated that DHV-PEI-C-RBCM exhibited excellent hemocompatibility. Enhanced blood compatibility of RBCM coating can be attributed to the hydrophilic phosphoric acid head of the phospholipid bilayer and the steric hindrance effect of glycoproteins [12]. The anti-biofouling properties of cell membranes have been empirically validated. Large-scale modification of PLT membrane coating was reported to enhance the hemocompatibility of blood-contacting materials and prevent thrombosis [40]. The exceptional hydrophilicity of DHV-PEI-C-RBCM enabled the formation of a resilient hydrated layer in the aqueous environment, consequently augmenting its effectiveness in preventing biofouling. To the best of our knowledge, for the first time, this study incorporated RBCM directly into the surface of DHV. Our findings demonstrated encouraging potential for diverse clinical applications.
The host immune response is an inevitable process after implantation of biomaterials, playing a pivotal role in the reconstruction of tissue microenvironment and in-situ regeneration [41], [42]. Macrophages induce a remarkably swift immune response to implanted biomaterials [43]. Previous studies have demonstrated the efficacy of utilizing nanoparticle-coated cell membranes to target macrophages and modulate their behavior, showing promise for treating pneumonia [44]. Regulating macrophage differentiation toward the M2 phenotype represents a promising strategy for promoting regeneration and treating VHD. CD47, known as “don’t eat me” signal, is a specific 5-transmembrane receptor identified on RBCM, which has been reported to prevent macrophage phagocytosis and modulate the immune response [45], [46], [47]. Recent studies have demonstrated that cell membrane-coated nanoparticles with anti-CD47 antibodies facilitate their cellular uptake and promote macrophage-mediated phagocytosis, thereby enhancing nanoparticle targeting [48]. Additionally, a recent study on RBCM coating attempted to directly coat RBCM onto the implanted electrode, which showed satisfactory biocompatibility, attenuated the acute immune response, and prolonged the lifespan of the electrode [49]. We used the subcutaneous embedding model and the abdominal aorta transplantation model to assess the infiltration of macrophages and T cells at various time points. Interestingly, our findings revealed an increased abundance of infiltrating macrophages and a decreased abundance of T cells in the DHV-PEI-C-RBCM group. This observation suggests that T cell-mediated immune rejection was mitigated in this study [50], [51]. Macrophage phenotype analysis at different time points revealed an earlier infiltration of M2 macrophages in DHV-PEI-C-RBCM, indicating a greater anti-inflammatory and tissue remodeling response. This finding was consistent with our in vitro findings involving macrophages. In summary, the RBCM coating facilitated by PEI-C effectively promoted the polarization of macrophages toward the anti-inflammatory M2 phenotype, thereby promoting endothelialization and tissue remodeling.
Endothelialization and ECM remodeling are the hallmarks of in situ regeneration. Our study revealed that the roughness of the RBCM layer significantly increased in AFM, suggesting its potential role in enhancing cell adhesion and proliferation [52]. Furthermore, in vitro experiments with HUVECs on scaffold surfaces confirmed the cytocompatibility of RBCM. Our findings from the abdominal aorta implantation model indicated that the DHV-PEI-C-RBCM graft notably improved endothelialization and collagen deposition, exhibited adaptive degradation, and prevented calcification compared to DHV and GLUT-HV. These properties can be attributed to the favorable cytocompatibility of RBCM, which promotes EC adhesion and shifts macrophage polarization toward the M2 phenotype, thereby facilitating tissue remodeling.
In general, our modification strategy utilizing DHV-based PEI-C-mediated RBCM coating has achieved the stable modification of RBCM on dECM, showing the promising potential of TEHV for future clinical applications. However, there are limitations to the current study. Future studies should investigate molecular mechanisms by which macrophages promote re-endothelialization and matrix remodeling, and assess the function of TEHV in large animal orthotopic HV implantation models.
5. Conclusions
In conclusion, the present study developed a new strategy to use PEI-C, immobilize RBCM coating on dECM, and construct a novel TEHV. Our novel TEHV had favorable mechanical properties, hemocompatibility, and histocompatibility, induced macrophage polarization toward the M2 phenotype, and promoted endothelialization and adaptive ECM remodeling after implantation, thereby enhancing durability compared to bioprosthetic HVs cross-linked with glutaraldehyde. Generally, our strategy modified RBCM at the macroscopic level and overcame the shortcomings of mechanical and bioprosthetic valves.
Acknowledgments
This work was supported by the National Key Research and Development Program of China (2021YFA1101900 and 2023YFB3810100), the National Natural Science Foundation of China (82270381 and 81930052), and the Major Science and Technology Special Plan Project of Yunnan Province (202302AA310045).
Compliance with ethics guidelines
Yuqi Liu, Pengning Fan, Yin Xu, Junwei Zhang, Li Xu, Jinsheng Li, Shijie Wang, Fei Li, Si Chen, Jiawei Shi, Weihua Qiao, and Nianguo Dong declare that they have no conflict of interest or financial conflicts to disclose.
D.Messika-Zeitoun, H.Baumgartner, I.G.Burwash, A.Vahanian, J.Bax, P.Pibarot, et al. Unmet needs in valvular heart disease. Eur Heart J, 44 (2023), pp. 1862-1873.
[2]
C.M.Otto, R.A.Nishimura, R.O.Bonow, B.A.Carabello, J.R. ErwinIII, F.Gentile, et al. 2020 ACC/AHA guideline for the management of patients with valvular heart disease: a report of the American College of Cardiology/American Heart Association Joint Committee on Clinical Practice Guidelines. Circulation, 143 (5) (2021), pp. e72-e227.
[3]
Y.Snyder, S.Jana. Strategies for development of synthetic heart valve tissue engineering scaffolds. Prog Mater Sci, 139 (2023), Article 101173.
[4]
M.A.Padalino, B.Castaldi, M.Fedrigo, M.Gallo, F.Zucchetta, V.L.Vida, et al. Porcine intestinal submucosa (CorMatrix) for semilunar valve repair in children: a word of caution after midterm results. Semin Thorac Cardiovasc Surg, 28 (2) (2016), pp. 436-445.
[5]
M.Hofmann, M.O.Schmiady, B.E.Burkhardt, H.H.Dave, M.Hubler, O.Kretschmar, et al. Congenital aortic valve repair using CorMatrix®: a histologic evaluation. Xenotransplantation, 24 (6) (2017), p. e12341.
[6]
D.C. van derValk, A.Fomina, M.Uiterwijk, C.R.Hooijmans, A.Akiva, J.Kluin, et al. Calcification in pulmonary heart valve tissue engineering: a systematic review and meta-analysis of large-animal studies. JACC Basic Transl Sci, 8 (5) (2023), pp. 572-591.
[7]
N.Poulis, P.Breitenstein, S.Hofstede, S.P.Hoerstrup, M.Y.Emmert, E.S.Fioretta. Multiscale analysis of human tissue engineered matrices for next generation heart valve applications. Acta Biomater, 158 (2023), pp. 101-114.
[8]
A.H.Zaidi, M.Nathan, S.Emani, C.Baird, N.P.Del, K.Gauvreau, et al. Preliminary experience with porcine intestinal submucosa (CorMatrix) for valve reconstruction in congenital heart disease: histologic evaluation of explanted valves. J Thorac Cardiovasc Surg, 148 (5) (2014), pp. 2216-2225.e1.
[9]
X.Ye, X.Hu, H.Wang, J.Liu, Q.Zhao. Polyelectrolyte multilayer film on decellularized porcine aortic valve can reduce the adhesion of blood cells without affecting the growth of human circulating progenitor cells. Acta Biomater, 8 (3) (2012), pp. 1057-1067.
[10]
Y.Snyder, S.Jana. Strategies for development of decellularized heart valve scaffolds for tissue engineering. Biomaterials, 288 (2022), Article 121675.
[11]
J.Gao, X.Yu, X.Wang, Y.He, J.Ding. Biomaterial-related cell microenvironment in tissue engineering and regenerative medicine. Engineering, 13 (2022), pp. 31-45.
[12]
M.He, K.Gao, L.Zhou, Z.Jiao, M.Wu, J.Cao, et al. Zwitterionic materials for antifouling membrane surface construction. Acta Biomater, 40 (2016), pp. 142-152.
[13]
P.Nguyen, M.K.Jayasinghe, A.H.Le, B.Peng, M.T.N.Le. Advances in drug delivery systems based on red blood cells and their membrane-derived nanoparticles. ACS Nano, 17 (6) (2023), pp. 5187-5210.
[14]
M.Yin, J.Shen, G.O.Pflugfelder, K.Mullen. A fluorescent core-shell dendritic macromolecule specifically stains the extracellular matrix. J Am Chem Soc, 130 (25) (2008), pp. 7806-7807.
[15]
C.Zhao, B.Zhou. Polyethyleneimine-based drug delivery systems for cancer theranostics. J Funct Biomater, 14 (1) (2022), p. 12.
[16]
H.Lee, Y.Lee, A.R.Statz, J.Rho, T.G.Park, P.B.Messersmith. Substrate-independent layer-by-layer assembly by using mussel-adhesive-inspired polymers. Adv Mater, 20 (9) (2008), pp. 1619-1623.
[17]
W.H.Qiao, P.Liu, D.Hu, S.M.Al, X.M.Zhou, N.G.Dong. Sequential hydrophile and lipophile solubilization as an efficient method for decellularization of porcine aortic valve leaflets: structure, mechanical property and biocompatibility study. J Tissue Eng Regen Med, 12 (2) (2018), pp. e828-e840.
[18]
C.Ma, Y.Xie, X.Huang, L.Zhang, M.D.Julian, L.Zou, et al. Encapsulation of (-)-epigallocatechin gallate (EGCG) within phospholipid-based nanovesicles using W/O emulsion-transfer methods: masking bitterness and delaying release of EGCG. Food Chem, 437 (2024), Article 137913.
[19]
A.V.Kroll, R.H.Fang, Y.Jiang, J.Zhou, X.Wei, C.L.Yu, et al. Nanoparticulate delivery of cancer cell membrane elicits multiantigenic antitumor immunity. Adv Mater, 29 (47) (2017), Article 1703969.
[20]
Y.Zhang, C.Zhang, Y.Li, L.Zhou, N.Dan, J.Min, et al. Evolution of biomimetic ECM scaffolds from decellularized tissue matrix for tissue engineering: a comprehensive review. Int J Biol Macromol, 246 (2023), Article 125672.
[21]
Z.Wang, Y.Lu, K.Qin, Y.Wu, Y.Tian, J.Wang, et al. Enzyme-functionalized vascular grafts catalyze in-situ release of nitric oxide from exogenous no prodrug. J Control Release, 210 (2015), pp. 179-188.
[22]
Z.Cui, B.Yang, R.Li. Application of biomaterials in cardiac repair and regeneration. Engineering, 2 (1) (2016), pp. 141-148.
[23]
R.D.Weisel. Tissue engineering to restore cardiac function. Engineering, 13 (2022), pp. 13-17.
[24]
C.Zhang, B.Wu, Y.Zhou, F.Zhou, W.Liu, Z.Wang. Mussel-inspired hydrogels: from design principles to promising applications. Chem Soc Rev, 49 (11) (2020), pp. 3605-3637.
[25]
A.Rising, M.J.Harrington. Biological materials processing: time-tested tricks for sustainable fiber fabrication. Chem Rev, 123 (5) (2023), pp. 2155-2199.
[26]
A.Taghizadeh, M.Taghizadeh, M.K.Yazdi, P.Zarrintaj, J.D.Ramsey, F.Seidi, et al. Mussel-inspired biomaterials: from chemistry to clinic. Bioeng Transl Med, 7 (3) (2022), p. e10385.
[27]
H.Lee, S.M.Dellatore, W.M.Miller, P.B.Messersmith. Mussel-inspired surface chemistry for multifunctional coatings. Science, 318 (5849) (2007), pp. 426-430.
[28]
Z.Dong, X.Ke, S.Tang, S.Wu, W.Wu, X.Chen, et al. A stable cell membrane-based coating with antibiofouling and macrophage immunoregulatory properties for implants at the macroscopic level. Chem Mater, 33 (20) (2021), pp. 7994-8006.
[29]
M.J.Lavoie, B.L.Ostaszewski, A.Weihofen, M.G.Schlossmacher, D.J.Selkoe. Dopamine covalently modifies and functionally inactivates parkin. Nat Med, 11 (11) (2005), pp. 1214-1221.
[30]
X.Zhang, Z.Li, P.Yang, G.Duan, X.Liu, Z.Gu, et al. Polyphenol scaffolds in tissue engineering. Mater Horiz, 8 (1) (2021), pp. 145-167.
[31]
J.Li, A.D.Celiz, J.Yang, Q.Yang, I.Wamala, W.Whyte, et al. Tough adhesives for diverse wet surfaces. Science, 357 (6349) (2017), pp. 378-381.
[32]
P.Podsiadlo, Z.Liu, D.Paterson, P.B.Messersmith, N.A.Kotov. Fusion of seashell nacre and marine bioadhesive analogs: high-strength nanocompoisite by layer-by-layer assembly of clay and L-3,4-dihydroxyphenylaianine polymer. Adv Mater, 19 (7) (2007), pp. 949-955.
[33]
L.Pan, H.Li, T.Hu, S.Wen, Y.Zhou, L.Jiang, et al. Design and assembly of chain-oriented-crystalline multilayered composite with largely improved mechanical strength. Compos Sci Technol, 238 (2023), Article 110031.
[34]
Y.Yang, K.Wang, X.Gu, K.W.Leong. Biophysical regulation of cell behavior-cross talk between substrate stiffness and nanotopography. Engineering, 3 (1) (2017), pp. 36-54.
[35]
Y.Zhou, G.Yan, S.Wen, W.Y.Yim, Z.Wang, X.Chen, et al. Nitric oxide generation and endothelial progenitor cells recruitment for improving hemocompatibility and accelerating endothelialization of tissue engineering heart valve. Adv Funct Mater, 33 (9) (2023), Article 2211267.
[36]
M.Hu, S.Shi, X.Peng, X.Pu, X.Yu. A synergistic strategy of dual-crosslinking and loading intelligent nanogels for enhancing anti-coagulation, pro-endothelialization and anti-calcification properties in bioprosthetic heart valves. Acta Biomater, 171 (2023), pp. 466-481.
[37]
P.Lancellotti, A.Aqil, L.Musumeci, N.Jacques, B.Ditkowski, M.Debuisson, et al. Bioactive surface coating for preventing mechanical heart valve thrombosis. J Thromb Haemost, 21 (9) (2023), pp. 2485-2498.
[38]
X.Huang, C.Zheng, K.Ding, S.Zhang, Y.Lei, Q.Wei, et al. Dual-crosslinked bioprosthetic heart valves prepared by glutaraldehyde crosslinked pericardium and poly-2-hydroxyethyl methacrylate exhibited improved antithrombogenicity and anticalcification properties. Acta Biomater, 154 (2022), pp. 244-258.
[39]
M.Hu, X.Peng, S.Shi, C.Wan, C.Cheng, X.Yu. Dialdehyde xanthan gum and curcumin synergistically crosslinked bioprosthetic valve leaflets with anti-thrombotic, anti-inflammatory and anti-calcification properties. Carbohydr Polym, 310 (2023), Article 120724.
[40]
Z.Zhou, R.Luo, L.Chen, C.Hu, C.Chen, M.F.Maitz, et al. Dressing blood-contacting devices with platelet membrane enables large-scale multifunctional biointerfacing. Matter, 5 (7) (2022), pp. 2334-2351.
[41]
S.Franz, S.Rammelt, D.Scharnweber, J.C.Simon. Immune responses to implants—a review of the implications for the design of immunomodulatory biomaterials. Biomaterials, 32 (28) (2011), pp. 6692-6709.
[42]
Y.Zhao, Y.Wang, J.Gong, L.Yang, C.Niu, X.Ni, et al. Chitosan degradation products facilitate peripheral nerve regeneration by improving macrophage-constructed microenvironments. Biomaterials, 134 (2017), pp. 64-77.
[43]
J.M.Anderson, A.Rodriguez, D.T.Chang. Foreign body reaction to biomaterials. Semin Immunol, 20 (2) (2008), pp. 86-100.
[44]
Y.Li, X.Liu, Z.Cui, Y.Zheng, H.Jiang, Y.Zhang, et al. Inflammation and microbiota regulation potentiate pneumonia therapy by biomimetic bacteria and macrophage membrane nanosystem. Research, 6 (2023), p. 0096.
[45]
T.Matozaki, Y.Murata, H.Okazawa, H.Ohnishi. Functions and molecular mechanisms of the CD47-SIRPα signalling pathway. Trends Cell Biol, 19 (2) (2009), pp. 72-80.
[46]
Y.Liu, Y.Wang, Y.Yang, L.Weng, Q.Wu, J.Zhang, et al. Emerging phagocytosis checkpoints in cancer immunotherapy. Signal Transduct Target Ther, 8 (1) (2023), p. 104.
[47]
P.A.Oldenborg, A.Zheleznyak, Y.F.Fang, C.F.Lagenaur, H.D.Gresham, F.P.Lindberg. Role of CD47 as a marker of self on red blood cells. Science, 288 (5473) (2000), pp. 2051-2054.
[48]
L.Chen, Z.Zhou, C.Hu, M.F.Maitz, L.Yang, R.Luo, et al. Platelet membrane-coated nanocarriers targeting plaques to deliver anti-CD 47 antibody for atherosclerotic therapy. Research, 2022 (2022), Article 9845459.
[49]
B.Wang, P.Yang, Y.Ding, H.Qi, Q.Gao, C.Zhang. Improvement of the biocompatibility and potential stability of chronically implanted electrodes incorporating coating cell membranes. ACS Appl Mater Interfaces, 11 (9) (2019), pp. 8807-8817.
[50]
L.Chung, D.J. MaestasJr, F.Housseau, J.H.Elisseeff. Key players in the immune response to biomaterial scaffolds for regenerative medicine. Adv Drug Deliv Rev, 114 (2017), pp. 184-192.
[51]
E.Mariani, G.Lisignoli, R.M.Borzi, L.Pulsatelli. Biomaterials: foreign bodies or tuners for the immune response?. Int J Mol Sci, 20 (3) (2019), p. 636.
[52]
H.Chen, X.Huang, M.Zhang, F.Damanik, M.B.Baker, A.Leferink, et al. Tailoring surface nanoroughness of electrospun scaffolds for skeletal tissue engineering. Acta Biomater, 59 (2017), pp. 82-93.
PDF
(8692KB)
