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Regulating Cell Adhesion by Material Surface Properties

Xuanshu Zhong , Zongjian Liu , Xue Geng , Lin Ye , Zengguo Feng , Jianing Xi

Progress in Chemistry ›› 2022, Vol. 34 ›› Issue (5) : 1153 -1165.

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Progress in Chemistry ›› 2022, Vol. 34 ›› Issue (5) : 1153-1165 DOI: 10.7536/PC210510
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Regulating Cell Adhesion by Material Surface Properties
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Abstract

Biomaterials aim to achieve tissue regeneration and repair by regulating the interaction between materials and cells. The adhesion determines whether cells could perform the expected biological functions absolutely. Therefore, it is critical for biomaterials to regulate cell adhesion by surface physical and chemical properties, which have aroused more and more attention recently. The physical modification on materials surface usually includes the regulation of surface roughness, morphology, modulus, and porous properties in order to build a suitable environment for cell adhesion. On the other hand, the chemical modifications such as surface charge and hydrophobicity regulation, covalently grafting and encapsulating adhesion-promoting molecules into the surface have made great effort to improve the interaction between the material surface and the cell, which will be capable of promoting cell adhesion. In recent years, a great number of breakthroughs have been accomplished in the field. In addition to covalently graft various adhesion-promoting molecules on the surface, it is reported that the adhesion performance would be significantly promoted by precise control of the sequence of the adhesion-promoting molecules. The sequence modulation represents a new strategy of surface modification, which can greatly enhance cell adhesion without the introduction of more powerful cell-promoting molecules or increasing their density. Besides, the intelligent change of adhesion-promoting and anti-adhesion surface can be accomplished according to the stimuli of external signals, which has been successfully applied in cell sheet tissue engineering and shows promising potential to be put into clinical application. Consequently, the review comprehensively summarizes the influences of material surface properties on cell adhesion by physical and chemical surface modification as well as the regulation of stimuli-responsive surface, the designs of adhesion-promoting surface, the technologies of adhesion-promoting surface preparation, and discusses their prospects.

Contents

1 Introduction

2 Cell adhesion regulated by surface physical properties

2.1 Surface roughness

2.2 Surface topography

2.3 Surface modulus and viscoelasticity

3 Cell adhesion regulated by surface chemical properties

3.1 Hydrophilicity and surface charge

3.2 Surface chemical modification by adhesion-promoting molecules and their sequence modulation

3.3 Surface encapsulation of adhesion-promoting molecules

4 Cell adhesion regulated by stimuli-responsive surface

4.1 Cell sheet tissue engineering

4.2 Supramolecular hydrogel with dynamic surface modulus

5 Conclusion and outlook

Graphical abstract

Keywords

cell adhesion / biomedical materials / surface modification / sequence regulation / stimuli-responsive surface

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Xuanshu Zhong, Zongjian Liu, Xue Geng, Lin Ye, Zengguo Feng, Jianing Xi. Regulating Cell Adhesion by Material Surface Properties. Progress in Chemistry, 2022, 34(5): 1153-1165 DOI:10.7536/PC210510

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1 Introduction

Biomedical materials are a class of special functional materials used to contact and interact with living systems, and can diagnose, treat, replace, repair or induce regeneration of their cells, tissues and organs, covering three major material fields of organic, inorganic and metal[1,2]. The interaction between materials and cells is the key factor to determine their efficacy and application prospects. When the material contacts with cells, the first thing that occurs is the adhesion of cells on the surface of the material, which has a very important impact on the regulation of the interaction between the material and cells. The number, adhesion force, state and morphology of adherent cells can affect and stimulate a series of molecular signaling pathways, promote the further adhesion of other cells, and then regulate their subsequent behaviors including proliferation and differentiation, and ultimately have an important impact on whether biomedical materials can achieve therapeutic and repair functions[3]. Therefore, the regulation of cell adhesion process is of great significance for biomedical materials[4].
Because cell adhesion occurs on the surface of materials, the surface properties have a very important impact on the process of cell adhesion. In organisms, cells are connected and adhered to other cells and intercellular matrix through transmembrane adhesion molecules including cadherins, integrins, selectins, etc[5][6][7][8,9]. Similarly, cell adhesion on the surface of materials also requires the recognition of cell adhesion molecules on the surface of materials. As shown in Fig. 1, cell adhesion can be divided into nonspecific and specific adhesion. When tissue fluid, blood or culture medium contacts the surface of the material, the infiltration of water molecules will occur first, followed by the adsorption of proteins, and finally the non-specific adhesion of cells on the surface of the material will occur through the interaction between cell adhesion molecules and proteins adsorbed on the surface. Surface physical properties, including roughness, morphology and modulus, have important effects on cell nonspecific adhesion. By introducing bioactive molecules that can be recognized by cells on the surface of materials through chemical modification, specific interactions between the surface and cells can be established, and specific adhesion of cells on the surface of materials can be achieved. Several reviews have reported the regulation of cell adhesion behavior by physical and chemical factors, respectively[2,10,11]. In recent years, there have been many new research directions and progress in the work of cell adhesion. For example, researchers are increasingly inclined to promote cell adhesion through a combination of chemical and physical factors. In addition to modifying various bioactive molecules with cell adhesion promoting properties on the surface, further regulating the physical properties of these molecules on the surface of the material, such as sequence and mobility, can further improve the cell adhesion without increasing the density of surface modification. In addition, smart material surfaces with stimuli-responsive properties have attracted more and more attention. It will produce intelligent response behavior according to the change of external signals including light, pH value, field intensity, temperature and other signals, so that the surface of the material can realize intelligent conversion between promoting cell adhesion and preventing cell adhesion. In cell sheet engineering, after cell layer culture in vitro, it is necessary to change the surface of the material from pro-adhesive to anti-adhesive, which is conducive to the removal of the cell layer and transplantation in vivo. In summary, in addition to reviewing the regulation of cell adhesion behavior by physical and chemical properties of material surfaces, this paper focuses on the latest research progress in the regulation of cell adhesion behavior by combining physical and chemical properties in recent years, as well as the use of stimuli-responsive surfaces to regulate cell adhesion.

2 Regulation of Cell Adhesion by Physical Properties of Material Surface

2.1 Surface roughness

Surface roughness has a great influence on cell adhesion. For osteoblasts (Osetoblast) and osteoblast-like cells, in general, material surfaces with higher roughness promote cell adhesion. Titanium alloy is one of the most common biomedical metal materials, which is widely used in the field of artificial bone and artificial joint because of its good mechanical properties and biocompatibility. Generally speaking, the rougher the surface of titanium alloy, the more it can promote the proliferation and adhesion of osteoblasts[12⇓~14]. However, surface roughness is a very complex concept, and there are many parameters to characterize surface roughness from different perspectives. Although the profile arithmetic mean deviation Ra (the arithmetic mean of the absolute value of the profile offset within the sampling length) is the most commonly used parameter to characterize surface roughness, sometimes material surfaces with the same Ra value have different cell adhesion behavior. This is often because, although the Ra values are the same, the other roughness parameters of these surfaces are not the same[15⇓⇓⇓~19].
Micro-nanofiber scaffolds fabricated by electrospinning have been widely used in tissue engineering because of their Excelluar Matrix (ECM) -like structure[20]. Stachewicz et al. First prepared electrospun scaffolds composed of smooth micron polycaprolactone (PCL) fibers, and then prepared electrospun scaffolds composed of porous micron PCL fibers by humidity control. The existence of porous structure on the fibers increased the roughness of the scaffold surface[21]. The results of in vitro cell culture confirmed that the porous PCL fiber scaffold could promote the adhesion of human osteoblast-like cells (MG-63) more than the smooth PCL fiber scaffold.
At present, interventional stent placement has become the mainstream treatment for all kinds of vascular stenosis and occlusion. Stent placement can cause some damage to the vascular wall, resulting in the loss of the endothelial layer. Vascular restenosis caused by excessive proliferation of smooth muscle cells may occur if the injury site is not rapidly re-endothelialized. Recently, it has been found that the surface roughness of magnesium metal stents has a good selective effect on the adhesion of endothelial cells and smooth muscle cells. As shown in fig. 2, when the surface roughness parameter Sa (the arithmetic mean height of the surface, which refers to the arithmetic mean or geometric mean of the distances between the points in the profile surface and the center plane) increases from nanometer to micron,The scaffold surface could significantly promote the adhesion and proliferation of endothelial cells and greatly inhibit the adhesion of smooth muscle cells, and the discrimination effect was the most obvious when the Sa value was 1. 0-2. 0 μm[22]. If the stent with endothelial cell selectivity is placed in vivo, it is expected to greatly accelerate the re-endothelialization process through the selection of adhesive cell types by surface roughness, thus greatly reducing vascular restenosis after stent implantation[2,17,22⇓⇓⇓⇓⇓⇓⇓⇓⇓⇓ ~33].

2.2 Surface morphology

Because the concept of surface roughness is too complex, it is difficult to guide the preparation of biomaterials and the control of surface morphology.Therefore, many researchers have turned to construct regular surface morphologies such as ridge and groove structures, micropillar arrays, nanosphere arrays, nanowire arrays and nanotube arrays on the surface of materials (Fig. 3), and studied the effects of morphological parameters on cell adhesion, proliferation and differentiation[34]. It should be pointed out that these topographic parameters can only affect cell behavior if they are close to the cell size, that is, in the range of submicron or 1 ~ 10 μm. In general, submicron-sized morphologies are more favorable for osteoblast adhesion, proliferation, and spreading[35,36]. Taking the adhesion of osteoblasts on the surface of titanium scaffolds as an example, the regulation mechanism of submicron surface topography on adhesion can be roughly divided into direct mechanism (cell-surface response) and indirect mechanism (protein-surface response). The direct mechanism is mainly considered from the kinetic point of view. Submicron topographic parameters can better match the cell size, induce osteoblasts to stretch filopodia and platopodia, promote the formation of filamentous actin and focal adhesion, and facilitate the adhesion and diffusion of osteoblasts. The indirect mechanism is mainly considered from the perspective of molecular recognition. The submicron structure of titanium surface can promote osteoblasts to express integrin subunits (such as α5β1, αvβ3) and secrete proteins containing RGD sequence (Arg-Gly-Asp). The high wettability of titanium surface due to its submicron structure allows RGD protein to adsorb on it and promote the formation of focal adhesion by specifically binding to integrin subunits on the surface of osteoblasts, which is conducive to the adhesion and proliferation of osteoblasts. At present, many literatures have reviewed in detail that micron-scale morphology has a very important impact on cell differentiation, especially the osteogenic differentiation of stem cells[24,35,37]. Therefore, the control of surface topography is one of the factors that need to be considered in the design of biomedical materials.
Porous material surfaces with appropriate dimensions are generally conducive to cell adhesion. On the one hand, the porous structure means an increase in the surface roughness of the material; On the other hand, appropriate pore size and surface layer "pattern" can tightly embed cells in the material, which is beneficial to cell adhesion and homing. It should be pointed out that when the pore size is equivalent to the cell size, the regulatory behavior of porous structure on cell adhesion is more obvious. PEEK (polyetheretherketone) is increasingly used in orthopedics and plastic surgery because of its easy processing and manufacturing, excellent wear resistance and mechanical properties similar to bone[38]. Wang et al. Used a certain proportion of nitric acid and sulfuric acid to form a multi-level porous structure on the surface of PEEK, and found that this porous structure was conducive to improving the cell affinity of PEEK[39].
The random overlapping of fibers in electrospun scaffolds results in a large number of porous structures, usually with thinner fibers forming small pores and thicker fibers forming large pores. The pore size has little effect on the adhesion of cells on the surface of the material, but when the scaffold is used in vascular tissue engineering, the adherent cells need to migrate from the surface to the inside of the scaffold. Small pore scaffolds block the migration of cells into the scaffold and cause macrophages to be in a pro-inflammatory M1 phenotype, resulting in a strong inflammatory response to the scaffold. Zhao et al. prepared a macroporous scaffold with a fiber diameter of 5 ~ 6 μm and a pore size of 30 μm by controlling various parameters in the electrospinning process, and proved that the macroporous scaffold could not only regulate macrophages (RAW264.7) to express M2 type of healing in vitro, but also had a very good effect on promoting angiogenesis in vivo[40].

2.3 Material surface modulus/viscoelasticity

Inside an organism, various tissues have completely different moduli/viscoelasticity. Brain tissue, for example, has a low modulus and appears to flow like a liquid even on long time scales, while bone tissue is rigid and has a high modulus. Obviously, different cells will have a range of moduli suitable for self-adhesion. Therefore, when preparing biomaterials, we should try our best to control the modulus of materials within the range of suitable modulus for cell adhesion[41⇓~43]. Zhang Tian prepared polyacrylamide gels with moduli of (11.15 ± 0.1) kPa, (32.51 ± 1.91) kPa and (80.64 ± 2.11) kPa, respectively, to simulate the stiffness requirements of normal (about 30 kPa) and diseased vessels[44]. Subsequently, HUVEC (human umbilical vein endothelial cells) cultured on substrates of different moduli were co-cultured with THP-1 (vascular endothelial cells adhering to human acute monocytic leukemia cells), and it was found that THP-1 cells adhered to the substrate at a near-physiological hardness (about 30 kPa). Because THP-1 cells are associated with vascular diseases such as atherosclerosis, this result not only reveals the regulation of vascular surface modulus on cell adhesion, but also shows that vascular surface modulus is an important factor affecting the occurrence of cardiovascular diseases.
In fact, modulus not only affects cell adhesion, but also stimulates related biomechanical signaling pathways through mechanical stimulation of cells, which has a very important impact on cell proliferation, differentiation, apoptosis, and even cell-related disease progression and cell response to drugs. Recently, Nature published a review that reviewed the complex mechanical behavior of tissues and extracellular matrices, discussed the effects of extracellular matrix modulus/viscoelasticity on cells, and described the potential of using material modulus/viscoelasticity to regulate various cell behaviors, including adhesion, in regenerative medicine[45].

3 Regulation of Cell Adhesion by Surface Chemical Properties of Materials

3.1 Hydrophilicity and surface charge

The hydrophilicity and hydrophobicity of the material surface have a regulatory effect on cell adhesion. Too strong hydrophilic or hydrophobic surface is not conducive to cell adhesion. In the process of cell nonspecific adhesion, there are two steps before adhesion: water molecule infiltration and protein adsorption. The hydrophilicity and hydrophobicity of the surface will directly determine the degree of water molecule infiltration. For strong hydrophobic surfaces, water molecules are difficult to infiltrate, which makes it difficult for proteins in aqueous solution to contact and adsorb on the surface of the material, thus not conducive to cell adhesion. However, the hydrophobic hydration between proteins and materials is the key to adhesion, so most proteins prefer to adhere to hydrophobic surfaces, so strong hydrophilic surfaces are also not conducive to protein adhesion[46]. To sum up, the material surface that is more conducive to cell adhesion is usually obtained by appropriate hydrophilic modification of the hydrophobic surface[47,48]. Hydrophilic groups such as amino, carboxyl and hydroxyl groups can be introduced onto the surface of materials by surface plasma modification, electron sputtering, surface etching and other methods to improve the hydrophilicity of hydrophobic surfaces. Zhao Qiang et al. Improved the material to a hydrophilic surface by heparinizing the PCL surface, which not only maintained a certain growth rate of vascular smooth muscle cells, but also showed a contractile phenotype[49]. Hydrophilic molecules can also be introduced into the surface through covalent bonds by click reaction, EDC/NHS bonding, etc. As shown in Fig. 4, PCL is a typical hydrophobic material. Our research group adopted the strategy of "corrosion first and then grafting" to graft hydrophilic heparin molecules on the surface of PCL fiber by EDC/NHS bonding. Water contact angle experiments confirmed that heparin surface modification improved the hydrophilicity of PCL electrospun scaffolds, and the results of cell proliferation rate further confirmed that the improvement of hydrophilicity could promote the adhesion of endothelial cells on PCL electrospun scaffolds[50]. In addition, heparin was directly grafted onto the two terminal groups of hydroxyl-terminated PCL by EDC/NHS bonding, which could also improve the hydrophilicity of PCL surface[51⇓~53].
Because cell membranes are usually electronegative, it is clear that positively charged material surfaces are conducive to cell approach and adhesion, while negatively charged surfaces are not conducive to cell adhesion due to electrostatic repulsion. Therefore, it is generally believed that chitosan, polylysine and other materials with positive electricity will be conducive to cell adhesion. However, cell adhesion is the result of multiple factors, and it is one-sided and inaccurate to judge whether a single factor on the surface of a material is conducive to cell adhesion. For example, although negative surface charges are not conducive to cell adhesion, we mentioned in the previous paragraph that negatively charged heparin grafted onto the surface of PCL electrospun scaffolds can promote cell adhesion. This is because the introduction of heparin not only improves the hydrophilicity of the material surface, but also the polysaccharide structure of heparin is conducive to cell adhesion. Under the combined effect, the negatively charged material surface modified by heparin still shows the effect of promoting cell adhesion. More importantly, the surface of the negatively charged material can also adsorb positively charged growth factors, including basic fibroblast growth factor (bFGF), vascular endothelial growth factor (VEGF), hepatocyte growth factor (HGF), and so on, and then control the sustained release of these growth factors in vivo through electrostatic interaction. After the adsorption of growth factors, the performance of promoting cell adhesion on the surface of the material will be further greatly improved. After the surface of heparin-modified PCL electrospun scaffolds was further adsorbed with bFGF, the adhesion of the material surface to endothelial cells was higher than that of the surface of heparin-modified PCL electrospun scaffolds without bFGF, and also much higher than that of PCL electrowoven scaffolds. Therefore, although the negatively charged material surface itself is not conducive to cell adhesion, the use of negatively charged material surface to adsorb growth factors is a common strategy to promote cell adhesion through surface modification in organic polymer biomaterials[54].

3.2 Modification of Adhesion Promoting Molecule and Its Sequence Regulation

Whether in specific or non-specific adhesion, it is necessary to form specific interactions between various transmembrane adhesion proteins of cells and the surface of materials or proteins previously adhered to the surface of materials. Naturally, people think that if we find adhesion molecules that can specifically interact with transmembrane adhesion proteins and modify them on the surface of materials, cell adhesion on the surface of materials can be greatly enhanced. Because cells are mainly tightly connected with the extracellular matrix and other cells in vivo, a large number of bioactive/adhesion molecules, including polypeptides, growth factors, nucleic acid ligands, adhesion proteins, glycoproteins and sericin, have been found in the extracellular matrix.It was found that they could effectively promote the adhesion of cells on the surface of the material after they were modified on the material[55][10,56⇓⇓⇓⇓⇓⇓⇓ ~64]. Research in this area is very popular, and a large number of literatures have been published in recent years. Various bioactive molecules that can be modified, their cell specificity, and their biological mechanisms of promoting cell adhesion have been reported in detail[65].
In order to further enhance the adhesion promoting properties of materials, it is a common idea to find bioactive molecules with stronger adhesion promoting properties and cell specificity. However, with the deepening of research work, new bioactive molecules will become more and more difficult to find, their structures will become more and more complex, and it will be more and more difficult to synthesize them by chemical methods and modify them on the surface of materials[66⇓~68]. Increasing the density of bioactive molecules modified on the surface of materials is another common idea, but the increase of modification density is also limited by existing chemical technologies[69]. The above limitations have prompted people to actively look for new ideas to improve the adhesion promotion ability of material surfaces. Interestingly, some researchers have pointed out that the order of bioactive molecules on the surface of materials plays a key role in cell adhesion. Polypeptide ligands, for example, recruit and capture cells mainly by forming specific interactions with the transmembrane molecule integrin. The binding of a single integrin to a ligand is not sufficient for cell adhesion on a material surface. Cell adhesion on the surface of materials can only be achieved by multiple interactions between integrin clusters formed by multiple integrins and multiple ligands. This requires a certain density and order of ligands on the surface of the material, so that the integrin cluster can easily form multiple interactions with multiple ligands. As shown in Fig. 5, Spatz et al., University of Heidelberg, Germany, confirmed that the distance between ligands that can make osteoblasts adhere should be less than 70 nm. At the same time, due to the random arrangement of integrins in integrin clusters, the surface of materials with random structure is more conducive to osteoblast adhesion than the surface with regular structure[70,71]. When the ligand on the surface of the material is fixed, cells need to constantly adjust the arrangement of integrins in the integrin cluster to adapt to the ligand order and form multiple interactions. When the ligand is disordered, the integrin can establish multiple interactions between the integrin cluster and the material surface through less movement. Recently, Choi and Bian of the Chinese University of Hong Kong have fixed RGD ligands on gold nanorods with different aspect ratios through Au-S bonds to obtain different ligand sequences. They found that the RGD ligand on gold nanorods with large aspect ratio can more easily form multiple interactions with integrin clusters, thus better promoting the adhesion of human bone marrow mesenchymal stem cells. More importantly, gold nanorods with high aspect ratio can also activate biomechanical conduction signals, thereby promoting cell adhesion and better extension and growth[72]. These results clearly indicate that regulating the order of ligands on the surface of materials contributes to the formation of multiple interactions between ligands and integrin clusters, and promotes cell adhesion on materials.
As shown in Figure 6a, there is a kind of polyrotaxane formed by cyclic molecules including cyclodextrin, cucurbituril, calixarene, etc. And linear molecules through mechanical interlocking, in which the cyclic molecules and linear axial chain molecules are not connected by chemical bonds, so they can freely slide and rotate along the axial chain[73][74⇓⇓⇓⇓⇓~80]. When various functional motifs are introduced into the cyclic molecule, these motifs will also move with the movement of the cyclic molecule on the main chain. Nobuhiko Yui of Hokuriku Advanced University of Science and Technology, Japan, and our research group jointly synthesized a triblock supramolecular copolymer with a polyrotaxane chain as the middle segment, as shown in Figure 6B[81]. In the triblock copolymer, a hydrophilic polyrotaxane segment is arranged in the middle, and hydrophobic polybutyl methacrylate segments are arranged on both sides. When it is coated on a hydrophobic substrate and contacted with an aqueous solution, the hydrophobic segments on both sides are "anchored" to the substrate, while the hydrophilic segments in the middle extend into the aqueous phase to form a "polyrotaxane loop". Quartz crystal microbalance (QCM) technique was used to confirm the formation of dynamic polyrotaxane rings with motional properties when the triblock copolymer was in contact with aqueous solution[81,82]. When the RGD ligand was introduced into the side chain of the CD in the copolymer, it was found that the polyrotaxane ring could significantly increase the binding constant between the ligand and the receptor (Fig. 6C)[83,84]. In this work, the receptor is in solution and has a high freedom of movement, so the binding constant is mainly affected by the mobility of the ligand. The increase of the binding constant is related to the freedom of movement of the ligand brought by the "fluidity" of the hydrophilic surface layer on the one hand, and the mobility of the ligand endowed by the polyrotaxane ring on the other hand. Furthermore, QCM was used to study the interaction between the ligand on the polyrotaxane ring and endothelial cells, and it was found that the ligand on the polyrotaxane ring could achieve more rapid adhesion of endothelial cells[84]. Apparently, this is because that ligand on the polyrotaxane re can rapidly adjust its distribution through its movement during adhesion, thereby establish multiple interactions with the integrin more rapidly. This result reveals that the mobility of ligand molecules on the surface of materials is also an important factor in regulating cell adhesion.
However, when Yui et al. Coated the polyrotaxane ring on the surface of the cell culture plate to culture cells, they found that although cell adhesion could be promoted on the first day, the number of cells on the polyrotaxane was less than that on the control material after further culture[84]. They believe that the surface layer of polyrotaxane ring is too hydrophilic and has a certain "fluidity", which makes its surface modulus extremely low. As mentioned earlier, cell adhesion is the result of a combination of factors. Therefore, after the polyrotaxane ring initially promotes cell adhesion through the movement of ligands, the low modulus hinders further sustained cell adhesion. Although the effect of polyrotaxane ring on cell adhesion is not obvious, it has a significant regulatory effect on cell proliferation and differentiation. Yui et al. Further reported that the polyrotaxane ring can not only promote the proliferation of fibroblasts, preosteoblasts and preadipocytes, but also promote the osteogenic differentiation of preosteoblasts[85]. Combined with the results of DNA microarray, they believe that the movement of cyclodextrin on polyrotaxane can produce certain mechanical stimulation to cells, stimulate biomechanical signals and promote better proliferation and differentiation of these cells. Furthermore, they used the type and degree of substitution of cyclodextrins to regulate the motility of polyrotaxane rings, and found that polyrotaxane rings with lower motility could better promote the proliferation of HUVEC and the expression of M2 configuration in liver macrophages (Kupffer cells), which is conducive to tissue regeneration[86][87]. The polyrotaxane ring with moderate exercise performance can promote myocyte differentiation, while the polyrotaxane ring with high exercise performance can promote osteoblast differentiation[88][82].
Adhesion-promoting molecule modification is the mainstream strategy to modify the surface of materials to promote cell adhesion, which is essentially based on the specific interaction between cell adhesion proteins and the surface of materials during cell adhesion. Compared with the traditional method of simple hydrophilic-hydrophobic or charge modification of material surface, this method involves the understanding and preliminary utilization of cell adhesion mechanism. Furthermore, whether it is to regulate the sequence of ligands on the surface of materials or the mobility of ligand molecules to establish multiple interactions between ligands and integrin clusters more quickly and improve cell adhesion performance, it involves a deeper understanding of cell adhesion mechanism and a higher level of application. Therefore, on the basis of a deeper understanding and revealing the mechanism of cell adhesion, the design and synthesis of new material surfaces that meet the requirements of cell adhesion is a very important research idea and direction for regulating material surfaces to promote cell adhesion in the future.

3.3 Surface loading of adhesion molecules

Coating the surface of materials with multifunctional coatings to promote cell adhesion is also an effective strategy to improve the surface adhesion performance. Coatings can often simultaneously control a variety of properties of the material surface, including hydrophilicity, hydrophobicity, charge performance, and so on. More importantly, the coating can effectively improve the adhesion promoting performance of the surface by loading adhesion promoting bioactive molecules[89]. Since Lee et al. Found that dopamine structure is an effective component in mussel adhesion protein, the construction of functional adhesion promoting coating (Fig. 7) using dopamine structure has become a hot research topic in material surface modification[90][91⇓~93].
Dopamine self-polymerization can successfully construct functional coatings on the surfaces of metal, inorganic and organic polymer materials, indicating that this strategy is universal for biomedical materials. Immobilization of antimicrobial peptide KR-12 coating on the surface of polyetheretherketone implant by polydopamine (PDA) bonding well promoted osteoblast adhesion and bone tissue integration. The polysaccharide modified dopamine was grafted and fixed on the cobalt-chromium stent, and the efficiency of rapid re-endothelialization of blood vessels was significantly improved. Because the extracellular matrix contains a large number of polysaccharide structures, the polysaccharide-loaded dopamine coating can regulate the bioresponsiveness of integrin and myosin in cells and promote cell adhesion[94].
Hydroxyapatite is the main inorganic component of minerals in human and animal bones, which is characterized by high mechanical strength, chemical bonding with body tissues at the interface, and certain solubility in the body.It can release ions harmless to the body, participate in metabolism in vivo, stimulate hyperosteogeny, and promote the repair of defective tissues, so it is widely used in osteogenic differentiation of cells and induced regeneration of bone tissues[95⇓⇓⇓~99]. Ti metal is widely used in bone tissue medicine, but bare Ti metal is difficult to be directly grafted on bone tissue in the dental field, and hydroxyapatite coating can improve the efficiency of osseointegration. As shown in Fig. 8, Jalali et al. First coated a titanium dioxide film (TiO2) on a Ti substrate by sol-gel dip-coating method, and then uniformly coated a mesoporous hydroxyapatite film (MHF) containing phosphoric acid, calcium nitrate tetrahydrate and nonionic surfactant[95]. The MHF has the advantages of smoother surface, uniform layer thickness, clear surface pores and the like, is beneficial to the adhesion and proliferation of the mesenchymal stem cells, induces the osteogenic differentiation of the mesenchymal stem cells, and has important application in bone regenerative medicine[100,101].
As mentioned above, electrospun scaffolds have been more and more widely used in various tissue engineering, so it is also a hot research topic to modify electrospun scaffolds by various methods to improve their cell adhesion promoting properties. Electrospun scaffolds are obtained by random overlapping of micro-nano fibers, so the modification of fibers is the main method to achieve the modification of electrospun scaffolds[102,103]. On the one hand, grafting various bioactive molecules on the fiber surface by various chemical methods can improve the adhesion promotion performance of the scaffold, for example, as mentioned in Section 3.1, heparin was grafted on the surface of polycaprolactone fiber by the method of "corrosion first and then grafting" to improve the cell adhesion performance. On the other hand, the core-shell structure fiber shown in fig. 9 is prepared by using the coaxial electrospinning technology, which can conveniently realize the loading of the fiber on the adhesion promoting active molecules. In the core-shell structure, both the core and the shell can load adhesion molecules and promote cell adhesion. For example, our research group prepared a core-shell fiber with bovine serum albumin and VAPG (val-gal-pro-gly) as the core and polycaprolactone as the shell, which can achieve the slow release of bovine serum albumin and VAPG, thus promoting the adhesion of cells on the scaffold[104]. Subsequently, our research group also used this technology to prepare core-shell fibers with polycaprolactone as the core and collagen as the shell. Collagen is the main component of extracellular matrix, which can promote cell adhesion, so the collagen shell can well promote the adhesion of smooth muscle cells on the electrospun scaffold[105]. Recently, in order to enhance the mechanical properties of electrospun vascular scaffolds, our group proposed a new idea of compounding hydrogel with porous electrospun scaffolds, that is, the hydrogel precursor was injected into the pores of porous electrospun scaffolds and gelled in situ[106]. Because the hydrogel is a very good drug controlled release platform, if the corresponding adhesion molecule is added to the hydrogel, it will be slowly released from the hydrogel and promote the adhesion of cells on the surface of the vascular stent. Therefore, the system of hydrogel composite electrospun scaffolds can also be used to improve the cell adhesion promoting performance of electrospun artificial vascular scaffolds.

4 Modulation of cell adhesion by stimuli-responsive material surfaces

4.1 Cell sheet tissue engineering technique

In recent years, cell sheet tissue engineering is a new type of tissue engineering technology. Cells, degradable scaffolds and cell regulators are the three elements of traditional tissue engineering. However, the interaction regulation between cells and materials, especially the matching of cell proliferation rate and material degradation rate, has not been well solved. In addition, how to better simulate the in vivo environment in vitro, regulate the interaction between cells and scaffolds, and make the cells in scaffolds have the same density and function as in vivo is also a huge challenge. In order to bypass the above complex material control requirements, some researchers have proposed the concept of cell sheet tissue engineering for tissues with two-dimensional layered structure, such as cornea and myocardium[107]. As shown in fig. 10, cell sheet tissue engineering means that after cells are cultured in two dimensions on the surface of a material to form a cell layer with an extracellular matrix, the cell layer is peeled off from the surface of the material and taken out, and then directly transplanted into an organism. Compared with the traditional tissue engineering route, cell sheet tissue engineering does not require the use of degradable three-dimensional scaffolds, but only requires two-dimensional culture of cells, which is relatively simple and easier to achieve clinical transformation as soon as possible. The culture of cells on the surface of materials and the peeling of cells after the formation of cell layer are the key technologies of cell sheet tissue engineering[108]. However, this technology puts forward opposite requirements for the surface properties of the material. In the cell culture stage, a pro-adhesion surface is needed to enable cells to adhere and further grow and proliferate to form a cell layer, and then an anti-adhesion surface is needed to enable the cell layer to be completely peeled off from the surface. Obviously, it is necessary to construct a "stimuli-responsive" material surface, that is, to regulate the surface properties of materials from "pro-adhesion" to "anti-adhesion" by changing an external condition. The most typical and earliest application is the temperature-sensitive "stimuli-responsive" surface. As shown in fig. 10, poly (N-isopropylacrylamide) (PNIPPAm) is a phase change material with strong temperature dependence. It has good solubility in water below 32 ℃, but it will quickly precipitate into a solid and is insoluble in water above 32 ℃. The cell culture plate prepared by using the temperature sensitivity of PNIPPAm can control the formation and shedding of the cell layer through the recognition of hydrophilicity and hydrophobicity of cells. At the cell culture temperature of 37 ℃, the surface of the material showed hydrophilic and hydrophobic properties which were beneficial to cell adhesion due to the precipitation of PNIPPAm, while at the temperature below 32 ℃, PNIPPAm changed into a water-soluble material, which made the surface of the material change into a hydrophilic anti-adhesion surface, which was beneficial to the peeling of the cell layer from the material surface. The surface of PNIPPAm modified materials can achieve complete peeling of the cell layer, while the traditional enzymatic hydrolysis and other means can lead to the dissociation of cell and cell, cell and extracellular matrix in the cell layer, resulting in the complete destruction of the cell layer. With the continuous development and progress of technology, a variety of stimuli-responsive surfaces have been developed and applied to cell sheet tissue engineering, such as stimuli-responsive material surfaces regulated by electricity, light and ROS (reactive oxygen species)[109].

4.2 Dynamic modulus hydrogel based on supramolecular self-assembly.

As mentioned above, the modulus of the material surface can regulate the cell adhesion behavior. Therefore, it is another feasible idea to use the surface modulus of stimuli-responsive materials to influence the cell adhesion behavior. In the late 1990s, Pelham and Wang synthesized polyacrylamide hydrogels with different elastic modulus and found that the modulus had a very important effect on cell adhesion, diffusion and migration[110]. Since this study, many groups have used polyacrylamide gels as a material platform to study the effect of modulus on cells. In recent years, Harada et al. Introduced cyclodextrin and the corresponding guest molecules into the side chain of polyacrylamide respectively, and then crosslinked polyacrylamide by supramolecular self-assembly between cyclodextrin and guest molecules to obtain a series of stimuli-responsive hydrogels[111⇓~113]. The degree of supramolecular self-assembly can be regulated by the stimulation of external conditions, which can affect the macroscopic properties of hydrogels, including modulus. For example, they introduced β-CD and adamantane into the side chain of polyacrylamide, and used the supramolecular self-assembly of β-CD and adamantane to prepare hydrogels with "reversible" cross-linking points (Figure 11)[114]. Then the free β-CD was added to the system, and the free β-CD would compete with the β-CD on the side chain of polyacrylamide and self-assemble with part of adamantane, resulting in the reduction of crosslinking points in the hydrogel and the reduction of the modulus of the hydrogel. On the contrary, washing out the free β-CD with water will lead to the increase of crosslinking points and the modulus of the hydrogel. The results of cell experiments showed that when the modulus of the hydrogel was adjusted to 4 ~ 11 kPa by controlling the amount of free β-CD, the surface of the material showed a good adhesion promoting effect on mouse myoblast cells. Too low or too high modulus is not conducive to the adhesion of the cell. The controllability of hydrogel modulus provides a suitable material platform to study the adhesion behavior of cells at different moduli. In addition, the dynamic change of hydrogel modulus can be regulated by adding or eluting free β-CD at different times, which makes it possible to monitor and study the adhesion behavior of cells during the dynamic change of modulus[115].

5 Conclusion and prospect

With the continuous application of various advanced processing methods, people are increasingly able to precisely control the physical properties of metal and inorganic non-metallic biomedical materials, including roughness, surface morphology, porous structure, surface modulus and so on, and then regulate the cell adhesion behavior on the surface of materials. The cell adhesion promoting performance of organic polymer biomedical materials has been successfully improved by continuously finding and synthesizing adhesion promoting molecules with stronger adhesion promoting performance and cell specificity and modifying the adhesion promoting molecules on the surface of the materials by chemical methods. In addition, the combination of stimulus-responsive surface and cell sheet tissue engineering is expected to be applied to the clinical treatment of two-dimensional structures such as cornea and myocardium.
Stem cells, including embryonic stem cells, adipose-derived stem cells and mesenchymal stem cells, are considered to be important seed cells in tissue engineering because of their ability to be induced to differentiate into a variety of cells, but they are difficult to adhere to the surface of biomaterials[116]. Recently, some researchers have begun to pay attention to and solve this problem, such as enhancing the wettability and swelling of hydrogels by polydopamine coating, which is beneficial to increasing cell proliferation rate, focal adhesion gene expression and cytoskeletal protein adhesion, and regulating the morphology and proliferation of adipose-derived stem cells[117][118].
The process of cell adhesion is determined by many factors. Therefore, the design of the surface of the new generation of strong adhesion promoting materials in the future should not only start from one point of view, but also consider comprehensively from various aspects, and improve the adhesion promoting performance of the materials through the synergistic effect of various adhesion promoting effects. In recent years, there have been some positive trends in the regulation of cell adhesion through the combined effects of multiple factors. For example, because polydopamine coatings can be successfully deposited on the surfaces of metals and inorganic non-metallic materials, people can easily regulate the surface chemical properties of these two types of biomaterials in addition to using traditional methods to regulate the surface physical properties. In addition to the functional modification of organic polymer materials by chemical methods, the molecular structure control of adhesion promotion belongs to the category of surface physical property regulation. Therefore, the comprehensive consideration of multiple factors is an important direction for the future design and preparation of new material surfaces with stronger adhesion promotion properties.
In addition, cell adhesion processes occur at any time in organisms, and these cell adhesion processes successfully regulate various physiological functions. Obviously, based on a clearer understanding of the cell adhesion mechanism, a biomimetic perspective is another important development direction for the future design and preparation of new surface materials with stronger adhesion promoting properties.

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Funding

National Natural Science Foundation of China(2017YFC1104101)

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