Biointerface Coatings With Structural and Biochemical Properties Modifications of Biomaterials

Designing a successful biological interface for a biomaterial requires the application of concepts from multiple fields, including materials science, surface science, nanoscience, physics, and biochemistry. With current biomaterials, such as polymers, ceramics, and natural materials, the original inherent properties of these materials no longer suffice for the needs of advanced and stringent biotechnological and medical applications that require controlled precision of physical and biochemical properties. Thus, the development of surface modifications to provide coatings that alter the interface properties of the original materials is pursued with much success to render desired and sophisticated interfacial properties. With acknowledging the available processing materials and fabrication techniques of the biointerface coatings, perspectives from advanced achievements in both physical and biochemical properties of the coating interfaces, and anisotropic presentations of these properties such as gradient of specific properties, the control of interface properties in confined micro/nanodomains and/or in introduced topographical geometries to mimic relevant biological activities is discussed in depth. The mentioned interface properties and the intended interactions toward biological microenvironments in vitro and in vivo are reviewed. Future prospective biointerface coatings with combinatorial and complementary properties from the discussed disciplines are suggested by the authors’ opinions. Last, challenges of current works demonstrating practical and important applications are discussed.

developed. These concepts have fueled modern schemes for the design of prospective biointerface science and the development of new surfaces for novel biomaterials and for biotechnological applications. Biointerface represents a wide range of applications, with biointerface coatings capable of detecting cells, viruses, and small molecules of biochemical significance, with DNA and protein microarrays forming the bio-MEMS platform. Biosensors measure the changes experienced by a biometric layer as it interacts with a target molecule. A similar concept is used in biorobot, which drives the properties of cells or tissues to bind functional biomaterials appropriately in response to external stimuli, enabling complex biological activities. On the other hand, drug carriers use surface modification in their release characteristics to achieve controlled release and targeted drug delivery, which is an interfacial phenomenon. The variety of biointerface devices also has numerous applications in the diagnosis of medical conditions.
In this review, we highlight relevant studies and advances in research on the development of interface coatings for biomaterials but may not comprehensively cover the literature in the entire field. Topics ranging from the fundamental interactions of the coating materials selected with the underneath substrates to the existing classifications of coating methods are discussed, with prospective discussions concerning the physical and biochemical properties of the coatings. Last, practical applications and current technological challenges are presented, and potential future directions are suggested according to the opinions of the authors. We envision a prospective biointerface coating is developed with combinatorial and complementary properties from the discussed disciplines, as shown in Figure 1.

Biointerface Coating Materials
The combination of coatings and substrates is a major aspect of interface engineering, with the main objective being to improve the surface properties of the substrate. There are additional factors to consider with regard to biointerfaces for in vivo applications, such as biocompatibility and biodegradability of the materials in vivo. In particular, there are also several criteria that need to be fulfilled so that the biodegradable materials would not lead to future clinical problems, such as i) all by-products that will be present due to the degradation process should be nontoxic, and ii) gradual degradation should be obtained instead of burst degradation to prevent inflammatory responses due to the change in pH. [19] This section reviews the functions and applications of common types of biocoating materials, as shown in Table 1.
As metallic biomaterial implants are susceptible to corrosion in the human body, [20,21] surface modification, such as biocoatings on the surface, is performed to extend the service life of the implant. [20] By maintaining mechanical properties and improving the problem of biointerface incompatibility, bioceramic materials are a common type of surface functionalized coating for biomedical applications. [22] For instance, hydroxyapatite (HAp) coatings have been used extensively in orthopedics over the last few decades. The coatings are used on metal implants in clinical applications to accelerate and enhance fixation for bone growth; however, there are some problems with the long-term stability of HAp coatings. [23,24] A biomimetic coating can be formed following a specific procedure and facilitates bone formation around the implant. [25] HAp coatings can also be integrated with other materials, such as tricalcium phosphate, which has shown better promotion of cell adhesion to the implant surface. [26] HAp modification or substitution of HAp coatings are interesting strategies to improve the performance of implants. Calcium phosphate (CaP) coatings are also commonly used at the tissue-implant interface to improve the reaction between the material and the organism, promoting inward bone growth and enhancing biocompatibility. CaP coating has been demonstrated to enhance cell adhesion, proliferation, and differentiation for bone regeneration. [27,28] Plasma-assisted technology is a www.advmatinterfaces.de frequently used method of preparing CaP coatings, which have dense, homogeneous, and highly adhesive properties that prevent toxic ion leakage and stimulate osteogenesis. [29] CaP coatings have been used on a variety of metallic and nonmetallic substrates, such as dental implants and joint replacements, and enhanced the osseointegration of devices with CaP coatings. [30] In addition, the deposition of protective ceramic coatings on a substrate surface, such as TiO 2 , a material with very low electrical conductivity, provides excellent corrosion resistance. TiO 2 nanocoatings have been prepared using sol-gel and dip coating technology, resulting in a homogeneous and dense surface with high hydrophobicity of surface water contact angle (WCA) 150° and excellent corrosion resistance in solution. [31] In addition, PCL/TiO 2 nanocomposite coatings have been applied to modify the surface of titanium implants to improve antibacterial properties, which can prevent implant-associated infections and promote cell adhesion to orthopedic devices. [32] Although bioceramics have been shown to enhance bone regeneration in vivo, [33] its use as in vivo coating materials are still limited and not widely studied. [34] Meanwhile, it is known that organic macromolecules are the most basic building blocks of organisms and that organic macro molecules are the basis of life. The ubiquitous existence of polymer compounds in the biological world determines their special status in the medical field. Among various materials, the molecular structure, chemical composition, and physicochemical properties of polymer materials are the closest to those of biological tissues; thus, they are most likely to be used as biomaterials. For example, polyethylene glycol (PEG) has great advantages in medical products and biomaterials and has been favored by researchers due to its unique properties. First, gradient PEG coatings can help immobilize growth factors and achieve specific click reactions on biointerfaces. [35] PEG-modified surfaces also show good anti-fouling and hydrophilic properties and are thus becoming popular for use in biosensors [36] and biomedical applications. Among biomedical implants, PEG coatings are commonly seen in dental implants due to their additional antibacterial properties, enhancing the functionalities and biocompatibility of the original implants. [37,38] PEG-based polymers can also be incorporated into the surface of contact lenses to create a wetted surface on the lens material for improved comfort, which is a major clinical advance in the use of PEG. [39] Due to its excellent mechanical properties and stable physicochemical properties, [40,41] polyurethane (PU) has great potential in biomedical applications, and it is a significant biomaterial that has long attracted the attention of researchers.
As the elasticity and wear resistance of PU materials allow it to withstand repetitive and long-term motion, it is mainly suitable for applications such as artificial hearts and heart valves. PU coatings also induce good cell adhesion and are well accepted by the host system, expanding their applications to insulators and catheters. [42] On the other hand, polyvinyl alcohol (PVA) has good film-forming properties and possesses great elasticity and chemical resistance; due to these unique properties, PVA is widely used in various industries, especially in cardiac tissue engineering. PVA coatings have the ability to optimize the characteristics of cardiovascular tissue replacements, as revealed by both in vitro and in vivo experiments. [43] Apart from cardiovascular devices, the mechanical properties of PVA are the most critical parameters to consider in orthopedic devices to repair or replace components of the musculoskeletal system that must withstand different degrees of compressive stress. [44] Poly-pxylylene (parylene) is also often used as a medical coating and has several advantages, such as uniformity, corrosion resistance, chemical inertness, and biocompatibility with human tissue and blood, making it widely used in the field of biomedical applications. Parylene coatings can be utilized on the surface of 3D substrates and can even define patterns to produce specific regions. Currently, parylene coatings are commonly used on human implants, such as coronary stents, [45] pacemakers, [46] and neural probes, [47] because they can effectively avoid infection in the body and are very suitable for devices that require long-term implantation. Furthermore, an innovative intraocular lens device made of parylene materials has been investigated. [48] The synergic functions of parylene provided with precise surface chemistry may pave the way for the next generation of biomedical optical products. Even though some polymers have been used as coating materials in biomedical devices, such as PEG to resist bacterial adhesion, [49] and PU and parylene to increase biocompatibility of biomaterials, [50,51] most of the studies regarding the use of polymers as coatings are still studies conducted in vitro, which limits its application for clinical studies. [49,52,53] Furthermore, as mimicking the natural environment in vivo is also of great importance, there is a search for natural ways to perform surface modification. Extracellular matrix (ECM) coatings have been of interest due to their biocompatibility and ability to provide biomaterials an environment similar to that in vivo. The ECM contains various growth factors that may enhance cell adhesion, [54] osteogenesis, [55] angiogenesis, [56] and endothelialization, [57] among other processes. One of the most common ways to obtain ECM coatings is via decellularization, and culture dishes coated with tissue-specific ECM have been shown to enhance adhesion and proliferation and the ability to maintain cell phenotypes. [58] Decellularized ECM coatings on titanium surfaces have also resulted in increased proliferation and osteogenesis, [55] while coatings on hydrogels lead to increased vascularization, with higher expression of hepatocyte markers. [56] Although surfaces can be coated with intact ECM, sometimes only one or several components of the ECM are chosen for biomaterial surface modification. A commonly used component is collagen, which is the main structural element of the ECM that is capable of modulating various functions, such as cell adhesion, endothelialization, [59] osteoconduction, [60] and angiogenesis. [61] Collagen has various applications in bone tissue engineering. For instance, collagen type-I coated on titanium alloys has been shown to enhance osteoblast adhesion and spreading. [62] Collagen can also be coated as a composite coating with other materials, such as cerium-doped hydroxyapatite, which results in the rapid generation of bonelike apatite. [63] Apart from bone tissue engineering, collagen also has applications in wound healing. Specifically, collagen coated on ostholamide electrospun nanofibers has been shown to promote wound healing and is a possible suitable biomaterial for wound dressings. [64] Another common ECM component used as a coating is hyaluronic acid (HA). Hyaluronic acid, also known as hyaluronan, is a group of polysaccharides often found in epithelial, connective, and nerve tissues. [65] Hyaluronic acid is capable of regulating inflammatory processes and angiogenesis and is hemocompatible, which has led to the use of hyaluronic acid for wound healing and as a coating for bloodcontacting medical devices. [65,66] For instance, hyaluronic acid is coated together with polydopamine via Michael conjugation to implant substrates, which results in better hemocompatibility to support adhesion and proliferation of endothelial cells. [67,68] Moreover, hyaluronic acid is often applied as a coating layer for targeted drug delivery, such as electrostatic coating of liposomes to achieve targeted delivery of paclitaxel for cancer therapy. [69] In terms of in vivo application, ECM, collagen, and hyaluronic acid have all resulted in reduced inflammatory responses and increased biocompatibility when applied as biomaterials coating in vivo. [70][71][72]

Coating Fabrication Techniques
Coating materials and coating techniques are used in many fields, and this chapter will focus on the different coating fabrication techniques. Surface coatings have many roles, such as substrate protection, [100] surface modification, [101] and www.advmatinterfaces.de functionalization. [102] By using a variety of coating methods, the advantages of various techniques can be combined to overcome existing problems. The performance of the various coating fabrication techniques has been combined and compared with other techniques by using spin coating as a standard. Spin coating can be carried out at ambient temperature and vacuum condition is not necessarily required to produce large quantities of coating on substrates, making it relatively simple to operate and obtain. Processing time is typically short, usually within seconds or minutes, with the exception of a few high viscous solvents. The cost depends on the type of solvent and the size of the substrate, and is lower compared to other methods, usually a few thousand dollars, and other coating techniques are compared as shown in Table 2.
The traditional coating techniques include spraying and dip coating. Although these techniques are well developed and easily performed, there are some disadvantages to each method. For example, thermal spray coatings often exhibit porosity with poor adhesion between the coating and the substrate, causing gases or liquids to flow through the interface. [103,104] Dip coating is a convenient and facile method; however, owing to the simplicity of the equipment, the quality of the coating is inconsistent. Uniformity is a major problem for dip coating; and thus, this method is not suitable for industrial applications. [105] To prevent uniformity problems, spin coating utilizes centrifugal force to spread the gel or solvent flat on the surface of the substrate, resulting in a uniform film. [106] However, for 3D substrates, spin coating provides poor coverage or accumulation of coating material. The other common coating technique is brushing, where polymers are often used to bind with the substrate physically or chemically. For instance, molecular brushes can be grafted onto solid surfaces via a combination of atom transfer radical polymerization and click reactions. [107] Grafting of polymer brushes onto surfaces to make them functional can be distinguished as "grafting to" or "grafting from" depending on the type of covalent interactions and is a surface modification involving a chemical reaction. [108] In addition, many techniques are accompanied by the process of solvent evaporation, which leaves a thin coating, and the selection of the solvent has a significant impact. This method is widely applied for drug delivery, and emulsion solvent evaporation is often used to prepare nanoparticles. [109] Some coating techniques require additional energy involved in the fabrication process, such as electric and photo energies. In the use of electricity, electrical chemistry and electrospinning are the two most commonly used in coating syntheses. Neither method has high equipment requirements nor do they have specific demands for temperature and pressure. Electrical chemistry, usually electrodeposition, features colloidal particles suspended in a liquid medium that will migrate to and deposit on electrodes under an external electric field. However, the coating resolution obtained via electrodeposition cannot be precisely controlled. The other technique, electrospinning, is able to make up for this shortcoming. In electrospinning, an electric force is applied to draw thin threads from a polymer solution that is particularly suitable for producing fibers of large and complex molecules. Elongation and thinning of the fibers caused by back and forth oscillation of the liquid  [111,115] results in a very constant fiber diameter that can be controlled in nanometers or micrometers. In addition to electrical use, photoenergy is widely applied in coating fabrication. Herein, photolithography and photochemistry are mentioned as examples. Some polymer coatings, such as polydopamine, have been fabricated via photochemical reactions in the presence of UV light. [110] Usually, materials that are sensitive to UV light will be chosen to prepare the coating of biomaterials in this way.
On the other hand, photolithography is a coating process that is more precise but requires more sophisticated equipment and has a higher cost. Photolithography uses exposure and development to delineate a geometric pattern on a photoresist layer and then transfer the pattern on the photomask to the substrate through the etching process. Thus, more precise control over coating resolution can be obtained with photolithography. Other coating techniques involving energy requirements under milder processing conditions are soft lithography and 3D printing. Unlike photolithography, which requires high-end equipment and technology, soft lithography utilizes a flexible imprinting mold made of soft polymer material; after coating a self-assembling monomer on the mold, it is slightly pressed on the substrate like a stamp, and the nanopattern is formed. This method can be performed at low cost, can produce large quantities, and is especially suitable for very large planes, but it still has the disadvantages of pattern errors and misalignment. Currently, 3D printing is receiving much attention in scientific research. Materials used in 3D bioprinting include living cells and biological materials, commonly referred to as "bioinks." It is a convenient and increasingly mature technology; however, building 3D structures in arbitrary patterns and shapes has still been challenging. After the development of vacuum pumps in the 17th century, various processes that require vacuum conditions became possible, and the vapor deposition method was one of them. Vapor deposition is mainly categorized into chemical vapor deposition (CVD) and physical vapor deposition (PVD), with many subdivisions in each category. CVD is a thin film deposition method in which a thin film is formed through chemical reactions of vapor-phase precursors. CVD is capable of producing a highly uniform coating layer, even for complex geometries. [111] However, it usually requires a very high temperature, which limits the types of substrates and coating materials that can be used. [112] On the other hand, PVD differs from CVD in that it uses physical methods, such as evaporation and sputtering, to produce thin films on substrates. PVD benefits from the lower temperature required in comparison to that required for CVD but is unable to produce the uniform coatings produced by CVD. [111,112] Another method for producing thin films is to use plasma, such as in plasma-enhanced chemical vapor deposition (PECVD) and plasma immersion ion implantation and deposition (PIIID). PECVD is superior to both conventional CVD and PVD in terms of lower temperature and uniform deposition; nevertheless, unintended ion implantation might occur, and undesired species might be present in the reactor. [113] On the other hand, PIIID is a process in which substrates are immersed in plasma, followed by the application of negative voltage to remove electrons from the surface, where the positive ions then bombard the surface. [114,115] PIIID has various applications in biomaterial engineering due to its ability to improve biological responses, such as bioactivity and hemocompatibility. [116]

Structural Properties of Biomaterial Interfaces
The interaction between cells and their surrounding environment, which has been studied since as early as 1911, [136] is an essential part of cell development and function. [137] As in biomaterials, the main interaction between cells and biomaterials occurs at the interface of biomaterials, and the biointerface of a biomaterial is of great importance to improve the interaction between cells and biomaterials. [138] In particular, surface chemistry, mechanical properties, and structural properties are influential factors that contribute to cell activities. [139,140] Although surface chemistry plays a major role in altering cellular behavior on a substrate, studies have shown that, even with a similar surface chemistry, the structural properties of an interface, including surface roughness and surface topography, also affect cellular behavior, such as adhesion, proliferation, and differentiation, [141] which indicates that structural properties are an integral part of cell functioning. This is because biological tissues in the body also present morphological structures, such as pores and fibers; [142] thus, the morphology of the biomaterial interface affects cellular binding to the surface, which in turn affects cell shape and cell processes. [1,143] In addition, structural properties might influence other physical properties, such as surface wettability and surface charge. [144,145] As a result, the influence of the structural properties of a biomaterial interface to best mimic the in vivo environment has been widely studied, and methods to modify the interface of biomaterials have become of great importance. In this section, how structural properties might affect cell adhesion and cell differentiation as well as current methods to achieve the desired cell fate through varying structural properties are reviewed. In addition, with respect to the rise of 3D cell culture models, the discussion regarding the influence of interfacial structural properties was extended to the interface of 3D culture biomaterials to show how the structural properties also affect cells cultured in these systems. Moreover, the fabrication of biomaterials is inseparable from their possible application in vivo. Therefore, recent developments in the structural properties of biomaterial interfaces that have been studied in vivo are also included in this segment. The structural properties discussed in this chapter are summarized in Figure 2.

Cell Adhesion
It is widely accepted that surface topography plays a critical role in cell adhesion. [155,156] Based on different cells, different types of topographies may be needed to enhance the attachment of cells to the surface. In fact, different cells seem to have their preferred surface roughness, for example, bone cells attach better to rough surfaces, [146] while epithelial cells attach better to smooth surfaces. [157] Therefore, various methods and materials have been used to modify surfaces to obtain the desired surface roughness. For instance, Wang et al. employed passivation treatment of titanium alloy to alter the surface properties www.advmatinterfaces.de and observed fibroblast attachment on the surface quantitatively and qualitatively by measuring detachment force and assessing cell morphology. They reported that a rougher surface resulted in a lower initial cell adhesion force. In addition, because the results contradicted the results of their previous reports, it was concluded that surface roughness plays a vital role in the determination of cell adhesion force. [158] Similarly, it was reported that higher roughness surfaces with a chitosan/ hyaluronan coating fabricated through layer-by-layer coating hampered tumor cell adhesion, where the surface roughness could be varied by varying the charge density of hyaluronan molecules. [159] On the other hand, using the cold spray method, Lee et al. coated hydroxyapatite on the surface of PEEK to produce a rougher surface with better biocompatibility. Their results showed that with the increase in interfacial area due to increased surface roughness, human bone marrow stromal cell (hBMSC) adhesion and viability were increased. [160] In addition, when modified TiO 2 was dip coated on glass slides, leukemia cell adherence was increased on surfaces with a high surface roughness and layer homogeneity with minimum surface skewness and kurtosis. [161] In another study, 100 µm-square textured patterns on hydroxyapatite and 45S5 bioglass-coated zirconia surfaces were created with laser sintering, and MC3T3-E1 cell adhesion on the surface was observed. In comparison to smooth surfaces, the cell viability was increased by 90% on surfaces with the square patterns. [162] In addition to surface roughness, porous structure is another key aspect in topography that affects cell attachment to the surface. Porous structures are capable of providing larger surface areas to let cells migrate into the inside of the porous structures, which leads to better cell adhesion. [163] For instance, Liu et al. scattered chitin nanocrystals in chitosan solution to produce a composite scaffold with pore sizes ranging between 100 and 200 µm and a porosity of 80%. They observed that the addition of the nanocrystals led to an increase in osteoblast adhesion and proliferation. [164] In another study, Zhou et al. fabricated ellipse nanopores onto poly(L-lactic acid) microfibers and explored the response of human vascular smooth muscle cells. Increased protein adsorption was observed in comparison to surfaces without nanopores, and cell adhesion was escalated on the surface with nanopores. [165] Although porous structures are beneficial in enhancing cell adhesion, an optimal pore size range to obtain enhanced cell adhesion seems to exist as studies have shown that pores that are too small limit cell migration, [166] while those that are too large result in decreased surface area. [167] Ganguly et al. studied the response of rat cortical astrocytes on nonporous surfaces and on anodic aluminum surfaces with pore diameters of 20 and 90 nm, where they observed that surfaces with an average pore size of 20 nm provided the best cell adhesion among the three groups, with more focal adhesions observed on the samples. [168] Ti6Al4V scaffolds with pore sizes of 500, 600, and 700 µm have also been fabricated, and those with smaller pore sizes (500 and 600 µm) provided better cell adhesion. [147]

Cell Differentiation
Structural properties affect not only cell adhesion but also cellular differentiation. After adhering to the surface and receiving molecular signals, cells may start to undergo differentiation; [169] thus, a difference in surface topography also affects the differentiation of cells. For instance, Vilardell et al. produced a highly rough pure titanium coating with cold spray coating technology and compared it with a sand-blasted Ti6Al4V alloy substrate. Alkaline phosphatase activity results showed better differentiation on the highly rough titanium coating. [170] Song et al. fabricated structures with different geometries and dimensions to study their effects on the neuronal differentiation of human induced pluripotent stem cells. Their studies showed that both aspects had considerable effects on neuronal differentiation, with the 560 nm nanogratings promoting neuronal cell generation and increasing TBR1 expression. [171] Interestingly, different surface topographies provide cells with different cues that allow cells to undergo different paths; thus, determining the fate of the cells seeded or cultured on the surface of the biomaterials. For example, in a study by Chen et al., it was shown that human mesenchymal stem cells (hMSCs) showed enhanced expression of different genes when cultured on electrospun fiber surfaces with similar porosity and fiber size but different surface roughnesses. On surfaces with a higher roughness average of 71.0 ± 11.0 nm, an increase in osteogenic-related genes, such as osteopontin (OPN), bone morphogenetic protein 2 (BMP-2), and run-related transcription factor Reproduced with permission. [146] Copyright 2007, Elsevier. In the bottom figure, the proliferation of cells on surfaces varied with pore size. Reproduced with permission. [147] Copyright 2020, Elsevier. c) Cell differentiation on surfaces with different structural properties. In the top figure, different surface roughness values obtained by varying the humidity conditions during the electrospinning process led to differential adhesion of skeletal cells, which in turn affected the osteogenesis of the cells. Reproduced with permission. [148] Copyright 2017, Elsevier. Meanwhile, the lower figure shows that different pore sizes may determine the differentiation of cells: cartilage formation was favored in scaffolds with small pore sizes, while endochondral ossification was preferred in scaffolds with larger pore sizes. Reproduced with permission. [149] Copyright 2018, Elsevier. d) Several studies on the culture and differentiation of cells in 3D cultures. On the top, halloysite nanotubes were coated on the surface of 3D-printed polylactic acid (PLA) scaffolds to align human mesenchymal stem cells (hMSCs). Reproduced with permission. [150] Copyright 2019, Elsevier. Beneath is a comparison of hMSC chondrogenic differentiation on surfaces with nanofibrous topographies and on smooth surfaces obtained through direct 3D printing. Reproduced with permission. [151] Copyright 2009, IOP Publishing. e) In vivo applications of surface-modified implants. In the uppermost picture, hybrid silica-chitosan coatings were fabricated on titanium implants to improve osseointegration ability. Reproduced with permission. [152] Copyright 2019, Elsevier. In the middle, an octacalcium phosphate coating containing BMP-2 increased the biocompatibility and osteoconductivity of coralline hydroxyapatite. Reproduced with permission. [153] Copyright 2019, Elsevier. At the bottom, Notch3 knockout (N3KO) mice have anatomical deficits in the spinal cord compared to wildtype (WT) mice, indicating a cellular environment that promotes differentiation. Reproduced with permission. [154] Copyright 2014, John Wiley and Sons. www.advmatinterfaces.de 2 (RUNX2) was observed; meanwhile, on surfaces with a lower roughness of 14.3 ± 2.5 nm, higher chondrogenic gene expression was observed. [148] On the other hand, studies regarding the effect of TiO 2 nanotube pore diameter on hMSCs and human osteoblasts (hOBs) demonstrated that a pore diameter of 20 nm favored the induction of osteogenic differentiation, while a pore diameter of 50 nm favored hOB osteoblastic maturation. [172] Similar to cell adhesion, porous structures also affect cellular differentiation by providing a larger surface area for protein adsorption and ion exchange. [173] The difference in proteins adsorbed on the surface then influence the fate of the cells, resulting in different cell differentiation paths. [174] For example, cell osteogenesis was shown to be preferred on Ti6Al4V scaffolds fabricated with a pore size of 500 µm in comparison to those with larger pores, [147] while cartilage generation, as indicated by type II collagen and aggrecan expression and production, was favored on surfaces with pore sizes between 150 and 250 µm in comparison to fabricated surfaces with pore sizes from 250 to 500 µm. [175] In another study, the influence of pore size on chondrogenesis and endochondral ossification of rabbit and hBMSCs was studied, with the results indicating that chondrogenic differentiation and cartilage formation were favored on surfaces with small pores, while endochondral ossification was limited on surfaces with small pore sizes. [149]

3D Culture
In recent years, cell culture has slowly transitioned from 2D to 3D culture because 3D culture provides an environment more similar to that in vivo and allows cells to grow inside the 3D cultures; thus, resulting in better cell culture in comparison to 2D culture systems. [176,177] 3D culture can be divided into several categories, including 1) suspension on nonadherence plates, 2) gel-like substances, and 3) scaffolds. [178] However, most 3D cultures have undefined surface characteristics and are not well studied in comparison to 2D culture systems, which has led to the fabrication of highly structured surfaces to study the effects of surface characteristics on cellular behavior. [179] In this section, we review the effects of the surface structures of 3D scaffolds on the behavior of seeded cells. Current fabrication methods for scaffolds with highly structured surfaces include direct fabrication with 3D printing [180,181] and coating a layer of a substance with defined structural properties on fabricated 3D scaffolds. [182] In direct fabrication via 3D printing, scaffolds with desired structural properties, such as pore size and shape, on a scaffold interface can be directly fabricated to study the effect of specific properties on cell behavior without having to use a mold during the fabrication process. [151] For instance, Prasopthum et al. showed that scaffolds with a nanofibrous topography led to better protein adsorption, which eventually resulted in better cell adhesion and differentiation. [151] His group also fabricated scaffolds with micro/nanoporous surfaces with pore sizes ranging from 0.2 to 2.4 µm, and chondrogenesis and osteogenesis of hMSCs were shown to be increased when cultured on the scaffolds even without the addition of differentiation factors. [180] In addition, Ran et al. utilized selective laser melting to fabricate scaffolds with pore dimensions of 400, 600, and 800 µm to study osteogenesis. Their results indicated that smaller pores were preferred for cell adhesion, larger pores were beneficial for cell proliferation, and scaffolds with a pore size of 500 µm were best for bone ingrowth. [181] The influence of pore geometries on mesenchymal stem cell differentiation was investigated by Ferlin et al., and their results showed that cylindrical pores increased osteogenic marker expression at an early stage, but at a later stage, osteogenic protein expression was instead enhanced by culture on surfaces with cubic pores. For adipogenesis and chondrogenesis of the cells, cubic pores resulted in higher relevant gene expression than cylindrical pores. This was likely due to the increased elastic modulus, pore size, and porosity of surfaces with the cubic pores. [183] Another method to obtain the desired surface topography is performing surface modification on fabricated scaffolds. For example, Liu et al. encapsulated black phosphorus in graphene oxide and then coated the composites on poly(propylene fumarate) 3D scaffolds, where the addition of graphene oxide provided more surface area that resulted in an increase in earlystage cell attachment. [184] A homogeneous nanoapatite coating on alginate/gelatin scaffolds was achieved via in situ mineralization, which resulted in a higher Young's modulus in comparison to that without the coating. The coating also resulted in an increase in surface roughness, which enhanced the proliferation and osteogenesis of rat bone marrow stem cells. [182] In another example, halloysite nanotubes (HNTs) were incorporated on the surface of a patterned polylactic acid (PLA) 3D scaffold. The PLA patterns displayed cell orientation induction ability, while coated HNTs improved the surface roughness and hydrophilicity of the scaffold and enhanced the adhesion and proliferation of hMSCs. [150]

In Vivo Activities and Implant Biomaterials
Biomaterials have profound effects on the host immune response, with the material eliciting a host response when implanted in living tissue. Long-term survival and biomolecular responses play an important role in improving the integration of implants in vivo to avoid chronic inflammation and foreign body reactions. [185] This section explores the immunological outcomes triggered by different coatings on implants and growth factors and provides insights into host responses to implant materials.
Xue et al. prepared wollastonite coatings via plasma spraying for implantation into the muscle, cortical bone, and marrow in dogs, and the implants integrated well with the bone without producing fibrous tissue. The wollastonite coating was found to stimulate surface bone formation better than a titanium coating, demonstrating good biocompatibility and osseointegration properties. [186] Titanium and titanium alloys, despite providing good biocompatibility and mechanical properties, still present a significant infection risk around implants. [187] Bacteria can easily attach to the surface of biomaterials and form biofilms; thus, precautions need to be taken to mitigate the associated infection problems, leading to the development of surface coatings that produce an antibacterial response. [188,189] For example, chitosan exhibits bactericidal properties. It was www.advmatinterfaces.de reported by Palla-Rubio et al. that silica-chitosan coatings adhered effectively to metal substrates, with the addition of tetraethyl orthosilicate regulating silica release, promoting bone regeneration, and increasing the rate of cell proliferation on the surface of the materials. [152] Kazemzadeh-Narbat et al. showed that microporous calcium phosphate coatings formed on the implant surface via electrolytic deposition as drug carriers could sterilize gram-positive (Staphylococcus aureus) and gram-negative (Pseudomonas aeruginosa) bacteria within 30 min, effectively preventing infection around the implants. [190] Plasma-sprayed HAp coatings have attracted widespread interest for clinical applications, but there are many controversies surrounding the performance of the implants, particularly with regard to long-term stability. [191] Xue et al. prepared HAp coatings with different crystallinity levels. The high crystallinity HAp coating remained intact for 3 months after implantation into bone, whereas the low crystallinity HAp coating separated from the surface, causing the coating to disintegrate. [192] In addition, HAp particles embedded in the joint surface of polyethylene can cause bone abrasion, which could lead to osteolysis. [193] Decellularized ECM has a wide range of applications in tissue engineering, such as surgical patches and soft tissue repair. In clinical models, it has been used at various sites, including the heart, [194] spinal cord, [154,195] liver, [196] meniscus, [197] and brain. [198] The combination of an ECM hydrogel and densely seeded myocardial cells showed the potential to repair the myocardium in vivo. [199] Gothard et al. constructed a blended hydrogel containing selected growth factors, including vascular endothelial growth factor (VEGF) and bone morphogenetic protein 2 (BMP-2); bone marrow stromal cells; and alginate for structural support to increase bone formation, and the hydrogel exhibited highly osteoinductive properties and contributed to the longevity of osteointegrated implants. [200] The ECM coating had a significant impact on cell growth, differentiation, and function [201] and effectively improved the biocompatibility of vascular grafts through surface modification. [202] Understanding the interaction between the ECM and cells in the in vivo environment assisted in the modification of the coating, and coatings of ECM substances have potential for drug development and cell therapy development. [201] Encapsulating growth factors into bioactive coatings is an innovative approach in bone tissue engineering [203] and enhances osteoinduction, which is an enabling condition for bone healing and is essential for the fixation of implants. Lin et al. found that an octacalcium phosphate coating incorporating BMP-2 improved the biocompatibility of coralline hydroxyapatite, showing that the BMP-2-incorporated coating enhanced osteoinductive efficiency and reduced inflammation. [153] Angiogenesis is a crucial process in dental osseointegration, and VEGF is considered to be an angiogenic factor. Guang et al. found that VEGF facilitates osteoblast proliferation and protein secretion on a titanium surface, with VEGF-coated implants promoting OCN-and CD31-positive cells around the bone cavity in vivo. [204] We explored the immunological processes occurring at the interface with different coatings. Biomaterials facilitate the integration of implants into the body and reduce tissue damage by improving the physiochemical surface properties.

Biochemical Properties of Biomaterial Interfaces
On the interfaces of biomaterials, biochemical properties have always received attention, and many studies have been conducted to evaluate the functionalities of biomaterials. [205,206] The current primary interest in biochemical properties significantly influences the future due to the versatilities of different biomaterial surfaces and the modification methods to enhance the probability of achieving the desired biochemical properties in biomaterial interfaces. Biochemistry is important and necessary to either aid in the expression of different interface functionalities or to give a biomaterial the ability to perform reactions with specific targets and biomolecules, [207,208] mimicking the complicated environment in a living system. In the following section, some classic and critical biochemical functionalities are discussed in detail, such as anti-corrosion, anti-fouling, immobilization, and therapeutic properties, followed by a discussion of multifunctional biomaterial interfaces, which exhibit more than one biochemical property simultaneously. [209][210][211] To adapt to environmental changes in nature, biomaterials combining host-guest chemistry are a novel solution to fabricate multifunctional biointerfaces with multistimulus responsiveness. [212] Moreover, gradient modification that allows the creation of a specific distribution on the surface to present different biochemical properties is also discussed in this section. In addition, cell cultures on biomaterials that promote cell adhesion, differentiation, and proliferation are included in this review topic because cell performance is highly related to the surface of biomaterials. [213] The concepts are described in Figure 3. Further examples of in vivo implants in different portions of the human body are strongly correlated with biomaterial interfaces, which have a great impact on biological activity.

Biochemical Functionality
It is very common that most biomaterials are exposed to many biomolecules and various cells; [217] therefore, the biochemical properties become very important and crucial to the interface between material surfaces and target molecules, [218,219] directly determining how the interface will perform and what functionalities it will have. In this section, single biochemical functionality is emphasized in detail, including anti-corrosion, anti-fouling, immobilization, and therapeutic properties, which are mainly described for materials and applications that possess these four properties to affect the biomaterial interfaces.
Implant material surfaces usually have basic properties, such as biocompatibility, [220] mechanical properties, [221] and the ability to customize implant shapes. [222] However, these materials still suffer from some significant difficulties, and the most common is surface corrosion after implantation. Due to the long-term contact with tissue and under rising temperature and saline conditions, corrosion will inevitably occur and produce harmful substances that affect the human body, especially alloy materials, which can release harmful metal elements. [223,224] Corrosion resistance is an essential property of metal implants as it determines the interaction between the metal surface and tissues. Thus, appropriate anti-corrosion coatings for the biomaterial surface are believed to be a promising strategy for www.advmatinterfaces.de  [214] Copyright 2016, Wiley-VCH. c) A smart biointerface in response to external stimuli for bacteria-releasing ability. Reproduced with permission. [215] Copyright 2017, American Chemical Society. The performance of cell cultures on surfaces with specific biochemical properties, including d) osteogenesis enhancement induced by RGD and BMP-2. Reproduced with permission. [216] Copyright 2016, Elsevier. e) Cell proliferation synergistically induced by black phosphorus and graphene oxide nanosheets. Reproduced with permission. [184] Copyright 2019, American Chemical Society.

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implant surface protection and continue to be the subject of many studies. For instance, Cieślik et al. demonstrated silaneparylene C coatings on SS 316L implant material, whose interface was between the metal implant and human tissues. The double-polymer coatings not only resulted in successful corrosion protection but the optimum thickness exhibited excellent wear resistance properties. [225] Moreover, Golda-Cepa et al. developed an implant coating using oxygen plasma-modified parylene C and drug-loaded, biodegradable poly(dl-lactide-coglycolide) (PLGA) to prevent implant corrosion. For practical purposes, polymer coatings should be easy to deposit, firmly adhere to the implant surface, and be chemically inert in the physiological environment. Parylene C meets these conditions and has been used in many medical applications. [226] Biofouling is a difficult problem to overcome in medical, marine, and underwater work equipment. [227] Due to longterm immersion in the sea, many microorganisms or algae in the ocean adhere to the bottom of equipment, and biological adhesion increases the attached load and resistance and causes functional problems. [228] In addition, the safety of many invasive medical materials and the prevention of infection have gradually been considered. When medical materials are placed into the human body, nonspecific biomolecule adhesion on the material surface often causes severe problems, such as an immune response and bacterial infection, resulting in the functional loss of medical materials and even endangering the patient's life. [229] Cordeiro et al. fabricated a biofunctional coating for Ti15Zr alloy via plasma electrolytic oxidation (PEO) to enhance surface characteristics. PEO appeared to improve the adsorption of albumin while decreasing bacterial adhesion. [230] Polyetheretherketone (PEEK) was coated with high purity magnesium by Yu et al. via the vapor deposition method to advance bioactivity, such as antibacterial properties. Studies have shown that the degradation of magnesium coatings can effectively kill S. aureus with an antibacterial rate of 99%. Magnesium coatings can be used to coat bioinert biomaterials to provide specific bioactivity. [231] Buxadera-Palomero et al. presented PEG coatings on titanium using three different methods: plasma polymerization, electrodeposition, and silanization. All methods were demonstrated to be great options to produce a PEG coating on titanium that has anti-fouling properties against bovine serum albumin (BSA) and significantly reduces bacterial adhesion of Streptococcus sanguinis and Ligilactobacillus salivarius. [37] Immobilization refers to the use of physical or chemical methods to prevent biomolecules from being lost from an inert carrier. Compared with natural free biomolecules, the advantages of the immobilization process can be expressed as increased stability, easy separation from the reaction system, manageable control, and repeated use. [232] Immobilized biomolecules are convenient for transportation and storage and are conducive to automated production. However, their activity may be reduced, and even narrowing the applications are the current problems that may be faced. [233] Nevertheless, immobilization is an application technology developed in the past 10 years and has attractive application prospects in industrial production, chemical analysis, and medicine. [234] Taking polydopamine as an example, polydopamine coating is a very promising approach to immobilizing biomolecules on almost all types of substrates. Zhou et al. designed a way to speed up the self-polymerization of a polydopamine coating on a QCM chip surface via higher temperature and vigorous stirring. [235] In addition, a tannic acid-3-amino-propyltriethoxysilane (TA−APTES) coating was investigated by Zhou et al. for enzyme immobilization. The nanostructured particle provided a larger specific surface area for enzymes, while quinone groups were in charge of covalent binding, resulting in much better enzyme immobilization performance. [236] Suo et al. used ionic liquidmodified cellulose-coated magnetic nanoparticles as supports for enzyme immobilization. With the induction of ionic liquids, the water contact angle indicated an increase in hydrophobicity, enhancing the affinity between lipase and the substrate immobilized on the prepared nanoparticles. [237] Therapeutic refers to the treatment and care of a patient to prevent or combat disease or to relieve pain and injury. Broadly speaking, therapeutic means the comprehensive service and care of the patient, the prevention of illness, and the management of specific problems. Targeted approaches used to treat particular symptoms include the use of medications to relieve pain or treat infection, surgical removal of diseased tissue, and replacement of poorly or nonfunctioning organs with appropriate implants. [238] Golda-Cepa et al. used a PLGA/Parylene C coating for implant surfaces with additional therapeutic functions, which controlled the drug release kinetics. In addition, when designing drug delivery systems, it is essential to always keep in mind the appropriate therapeutic window for each drug used. When the level of the released drug falls below its critical value, the therapeutic function is lost, and overdose can lead to serious and untreatable side effects. [226] Kim et al. used a cellsheet-graphene hybrid to offer a new strategy for implantable biomedical devices. The hybrid consisted of a sheet of C2C12 myoblasts on a buckled mesh graphene electrode, featuring high-quality electrophysiological sensing, electrical or optical stimulation therapy, and regenerative therapy with hybrid cells. These cell sheets offered practical therapeutic functions to the implanted tissues and simultaneously formed a highquality biological interface in the absence of a host immune response. [214] Moreover, Zhu et al. reported PCL conduits that were functionally modified with molecules containing RGD and YIGSR, and these conduits were able to promote axonal regeneration and functional recovery and to enhance vascularization in regenerated nerve tissue. This made the RGD/YIGSR-PCL conduit a potential candidate for treating critical neurological defects. [239]

Multifunctional Biointerface
Multifunctionality means that materials have more than one functionality on the same biointerface. In biological systems, the surface often faces many stimuli and must show corresponding responses to different targets to be able to cope with the complex environment in the organism. Therefore, many researchers are interested in the performance of biointerfaces. [240] In order to selectively conjugate to corresponding biomolecules, bioorthogonal chemical conjugations, such as copper-free click reactions, Staudinger ligations, and other specific bioorthogonal reactions, appear to be an ideal way to achieve reactions that can be carried out in living cells or www.advmatinterfaces.de tissues without interfering with the biochemical reactions of the organism itself. [241] Furthermore, surface science, including host-guest chemistry and gradient distribution, is also discussed, which involves modification of surfaces to express multiple functionalities, mimicking living systems more completely. With the appropriate combination and complement of the above properties, the multifunctionality of biointerfaces can be achieved more precisely and comprehensively.
The necessity for precise and adaptable conjugation methods has increased as a result of recent breakthroughs in biotechnology, [242] regenerative medicine, [243] and medical implants. [244] Assume that precise immobilization of two or maybe more biomolecules in specified ratios is desired. In that case, the individual reactions must not only be orthogonal with respect to ongoing biological events but also with respect to one another. This is necessary for chemical reactions with high specificity toward the molecule of interest. [245] The reactions created for application with multifunction in bioorthogonal chemistry are covered in this section of this article. The bioorthogonal reactions include a number of examples involving Staudinger ligation, as well as copper-free click reaction, condensation of aldehydes and ketones, and others involving thiol, NHS ester, and maleimide. The conventional Staudinger reduction of azides with triphenylphosphine is modified by the Staudinger ligation. The peptidic natural substance yaku'amide B (1), which had antiproliferative properties against several cancer cells, might now be wholly synthesized using a solid-phase method, according to the study by Itoh et al, [246] wherein traceless Staudinger ligation allowed enamide to develop between newly created phosphinophenol ester and sterically impounded alkenyl azides. Hu et al. demonstrated that photoirradiation in the visible spectral region could activate traceless Staudinger ligation, which was helpful for bioconjugation applications. Through the conjugation of amino acid/oligopeptide building blocks by the distinctive peptide linkage obtained by traceless Staudinger ligation, several visible-light-induced oligopeptide syntheses were synthesized. [247] In addition, as a result of not requiring the use of dangerous copper ions as a catalyst and having the benefit of speed, copper-free click chemistry makes a great alternative reaction for the Staudinger ligation. Using a modular approach based on the strain-promoted azide-alkyne cycloaddition click reaction, deoxycholic acid-or octanoic acid-modified N-azido propionyl-N, O-sulfate chitosan was synthesized in the study by adjusting the hydrophobic groups. [248] It was expected that this modular approach could easily and effectively deliver many chitosan derivatives for developing multifunctional drug carriers. Moreover, Deng et al. presented a synthetic method for creating reactive coatings for 1,3-dipolar cycloadditions that were copper-free and showed how a bioorthogonal reaction strategy based on two sequential click reactions could be achieved. The double-click chemistry suggested for multifunctional coatings would likely find applications in the controlled immobilization of several ligands. [245] Moreover, due to the small size and compatibility in living systems, the carbonyl groups in aldehydes and ketones are appealing for biomolecule tagging and these electrophiles will easily condense with α-nucleophiles, such as hydrazides and aminooxy groups. The 6-amino-acid consensus sequence recognized by the formylglycine-generating enzyme was utilized in the approach for the site-specific introduction of aldehyde groups into recombinant proteins by Carrico et al., which could be employed for a variety of protein labeling applications. [249] Other examples of critical bioorthogonal reactions were also mentioned as follows. Chen et al. demonstrated protein labeling or site-specific protein immobilization on biochips could also both be accomplished successfully using the photochemical thiol-ene reaction. [250] Ma and the group also developed a method through the carefully regulated irradiation of particular areas of the gel, a photoinitiated thiol-yne reaction was used to systematically integrate complementary moieties for the two bioorthogonal pairs in a spatiotemporal way. [251] In addition, research showed the production of protein-polymer hybrid hydrogel that could serve as a platform for immobilizing functional proteins. This material was be used in the thiolmaleimide procedure, which enabled the easy one-pot synthesis of a functionalized hydrogel by cross-linking the hybrid network and conjugating proteins to the gel backbone. [252] Many processes in living systems start from interfacial molecule recognition, primarily [119][120][121] influenced by various external stimuli at biological interfaces. To mimic the complex molecular systems observed in nature that accommodate environmental changes, a combination of host-guest chemistry and surface science is fundamental to fabricating future generations of multifunctional biointerfaces. Yang et al. demonstrated biointerfaces and biosurfaces with great responses to external stimuli, such as UV light, pH, redox chemistry, and competitive guests, and further presented feasible strategies for the future design of stimuli-responsive multifunctional biointerfaces. By expertly integrating the two critical technologies, host-guest chemistry and surface science will be the key developments in interface chemistry. [212] In their work, Zhou et al. developed switchable bioadhesion platforms, including protein adsorption and release and bacteria and cell attachment and detachment. The multistimulus-responsive biointerfaces responded to changes in temperature, pH, or sugar content of the medium, either individually or all simultaneously. [240] Herein, the different stimuli for host-guest reactions, properly designed and matched functions, and a great multifunctional biointerface able to adapt to the natural complex environment are discussed. Wei et al. developed an intelligent supramolecular surface capable of reversibly switching between bactericidal activity and bacterial release in response to ultraviolet-visible light. Surface-immobilized trans-azo groups could specifically bind CD-QAS to achieve a robust bactericidal surface that killed up to 90% of attached bacteria. Under UV light irradiation, the azo group was converted to the cis form, resulting in the release of dead bacteria from the surface via dissociation of the azo/ CD-QAS complex. [215] Moreover, Yu et al. prepared supramolecular hydrogel microcapsules using a host-guest assembly approach. With the help of UV irradiation, the supramolecular hydrogel network converted from a noncovalent to a covalent structure, offering a mechanism to control permeability on the surface of microcapsules. [253] PBA groups form complexes with diol-containing biomolecules under neutral or weakly basic conditions. To obtain these complexes, lowering the pH value or adding other diol-containing biomolecules that have a higher affinity for PBA groups helped in preparing a pH-responsive biointerface. [240] Similarly, Wan et al. fabricated a pH-responsive reactivated biointerface for the reversible adsorption of www.advmatinterfaces.de cytochrome c (cyt c). Using the structured change of proteinresistant PEG in the stretched state at pH 7.0 to immobilize cyt c and PEG in the relaxed state at pH 4.0 to release cyt c, a switchable interface was realized. [254] Zhou et al. reported multistimulus-responsive biointerfaces, in which PNIPAAm polymer underwent a phase transition when the temperature rose or fell across its lower critical solution temperature. This thermoresponsive PNIPAAm has been used to produce biointerfaces for various applications, including biosensors, anti-fouling surfaces, and biochromatography. [240] Zhang et al. successfully synthesized a novel dianthracene calix [4] arene surface exhibiting a solid ability to bind with tryptophan. In response to a tryptophan solution, this surface showed outstanding reversible wettability and fluorescence and could be broadly applied in tryptophan identification and other applications, including nanodevices and biochips. [255] Functionally gradient materials are a new type of complex material in which two or more materials are composited with continuous gradient changes in composition and structure from one direction (1D, 2D, or 3D) to another direction. [256] From the perspective of material structure, functionally gradient materials are different from homogeneous materials and composite materials. Two or more materials with different properties are selected; by continuously changing the composition and structure of the material, the interface disappears and causes the properties of the material to slowly change, forming functionally gradient materials. [257] Such gradient-manipulated surfaces can produce more functionalities at biomaterial interfaces, leading to more viable applications. Guan et al. demonstrated that a uniform polymer biointerface generated continuous and multifunctional gradients with a controlled click reaction (copper-free alkyne and azide click reactions) and other orthogonal conjugates (thermoactivated thiol−yne reaction). In this study, two gradients of PEG and RGD with a countercurrent distribution were achieved by these two reactions and enabled precise immobilization and distribution of the biomolecules FGF-2 and BMP-2. [35] Furthermore, a surface with a continuous anti-fouling gradient was created by Deval, P., C.-H. Lin, and W.-B. Tsai. pSBAE gradients were prepared on glass slides over 7 min by aspirating the reaction solution in the presence of the oxidizing agent ammonium persulfate. The specific functional chemical gradient was correlated with surface wettability, determining the gradient of subsequent cell adhesion and protein adsorption. [258]

Cell Culture and Biochemical Activities
Cell culture is a modern biological technology in which biological cells are cultured in a controlled state to allow their growth and performance, providing an excellent model system to study the normal physiology and biochemistry of cells. The main advantage of cell culture is the consistency and reproducibility of results that can be obtained by using artificial cells before in vivo experiments. [259] In addition, cell culture performance often closely depends on the properties of the biochemical interface. In this section, we further discuss the performance of cells at the interface of different biomaterials in terms of attachment, differentiation, and survival.
Cell adhesion is important and necessary for cellular regulation and communication and is essential for developing and maintaining tissues. Cell adhesion refers to the ability of a cell to adhere to another cell or to the extracellular matrix. [260] The interaction between a cell and its extracellular matrix significantly influences the cell's behavior and function. The fundamental function of cell adhesion has generated great interest in developing and studying methods to examine cell adhesion properties. Herein, we mainly discuss how the biochemical performance of a biomaterial surface affects or improves cell adhesion. RGD peptide is a cell adhesion sequence that mimics cell adhesion proteins and binds to integrins. RGD peptides can be used in synthetic scaffolds in tissue engineering to enhance cell adhesion, mimicking in vivo conditions. Bilem et al. used RGD to functionalize biomaterial surfaces, which promoted adhesion of bone marrow stem cells through integrin receptors. [216] Another group developed a method to functionalize PAA with RGD peptides via a conjugation reaction. Through this RGD-modified interface, a surface that used to inhibit cell adhesion became prone to cell adhesion. [261] Kim, B.-S., S.-S. Yang, and C.S. Kim coated bone morphogenetic protein-2loaded nanoparticles (BMP-2/NPs) with an ε-polycaprolactone polymer emulsion, which enhanced the cell attachment performance. [262] Furthermore, Rengaraj et al. fabricated two bioactive film coatings to form miniature tumor tissues and mimic the early stages of metastatic cancer. BMP-2, BMP-4, and fibronectin were induced to investigate the behaviors of two (immortalized and patient-derived) human pancreatic cell lines. It was shown that pancreatic cancer cell adhesion and BMP signaling were affected differently by biochemical signals in different cell types. [263] Stem cells are cells with self-renewal and differentiation capabilities. [264] Mesenchymal stem cells have the ability to differentiate into specific tissues and can repair damage in tissues or organs. Most mesenchymal stem cells in adults are in the stationary phase of the cell cycle. Only when the tissue is stimulated by damage or remodeling signals will they enter the cell cycle, further promoting tissue regeneration and repair. Mesenchymal stem cells have an irreplaceable position in cell therapy, regenerative medicine, tissue engineering, and clinical therapy. Using ECM-derived peptide-grafted surfaces with RGD and BMP-2 ligand crosstalk, Bilem et al. were able to stimulate stem cells and prompt their specific differentiation. [216] Raftery et al. designed a scaffold with a collagen and chitosan ratio of 75:25, and significant increases in calcium production and sulfated glycosaminoglycan (sGAG) production were observed when osteogenic and chondrogenic differentiation of mesenchymal stem cells were induced on the scaffold surface. [265] Osteogenesis on 3D-printed scaffolds was also reported by Liu et al. Black phosphorus was initially encapsulated in negatively charged graphene oxide nanosheets and then adsorbed onto positively charged poly(propylene fumarate) scaffolds. The encapsulated black phosphorus nanosheets were slowly oxidized to produce sustained phosphate release, which is an important osteoblast differentiation promoter to stimulate cells to form new bone. [184] Cell proliferation refers to the process by which a cell grows and divides into two daughter cells, which doubles the number of cells; thus, it is a fast mechanism for tissue growth. In cell www.advmatinterfaces.de research and analysis, whether to evaluate the compatibility of the material with cells or the promotion of cell growth or to analyze the effect of a growth factor on cell growth, cell proliferation assays can be used to detect changes in the number of living cells or dividing cells. Liu et al. demonstrated that 3D-printed scaffolds, in which black phosphorus was encapsulated in graphene oxide nanosheets, could release phosphate groups to enhance the proliferation of preosteoblasts as the scaffolds degraded. [184] In addition, Lilge and Schönherr synthesized a well-defined polymer brush-functionalized surface via surface-initiated atom transfer radical polymerization, which enabled good cell proliferation. [261]

Bio-Nano Interface
Biofilm formation is one cause of implant infections, which result from the binding and subsequent proliferation of bacteria on the surface. In the case of implants in biological fluids, for example, proteins are nonspecifically adsorbed onto the surface and form a regulatory layer that allows microorganisms to easily adhere to the surface and promote bacterial settlement. [266] Understanding the mechanisms of bacterial activity on surfaces is essential for effective infection prevention. Based on the process of bacterial colonization of a substrate, different antimicrobial coating strategies have been explored. The key to antimicrobial surfaces is the inhibition of bacterial uptake, which is based on the principle of reducing the nonspecific uptake of proteins. In addition, a nanostructured surface topology is effective in controlling bacterial adhesion, with superhydrophobic and hydrophilic nanosurfaces contributing to low bacterial adhesion, accompanied by variations in surface tension. [267,268] Hydrogels are porous, 3D structures that have been widely used in biomedical applications due to their similar properties to biological tissues but that have the disadvantage of poor mechanical properties. To improve the stability and adhesion of a hydrogel coating, Zhao et al. introduced a chemical method to fix the hydrogel layer to the substrate and construct an antimicrobial coating that included surface grafting polymerization and chemical crosslinking of the coating layer by layer. [269] In addition, hydrogels provide an anti-fouling effect through functionalized modifications. For instance, Tang et al. synthesized a new anti-fouling hydrogel based on thiolated chitosan and maleic acid-grafted dextran through the Michael addition reaction. The hydrogel showed good cytocompatibility with NIH3T3 cells and demonstrated that VEGF and PCNA accelerate wound repair in diabetes and inhibit bacterial adhesion for wound dressing, which is beneficial for the treatment of diabetic ulcers. [270] Hydrogel coatings are promising antimicrobial materials with the ability to accommodate a wide range of materials.
In recent years, significant advances have been made in the use of nanoparticles (NPs) in medical applications as a delivery system for therapeutics, particularly in the interaction between NPs and biological interfaces. When NPs enter the physiological environment, a protein corona surrounds them to create an interface with the cell, and the surface is controlled by a variety of physical forces: van der Waals forces, hydrogen bonding, and electrostatic interactions. [271,272] The biological properties of NPs are affected by the protein corona, and the size, surface charge, and solubility of NPs determine the composition of the protein corona. Depending on the binding force between the protein and the nanoparticle, protein coronas are divided into soft coronas and hard coronas. [273] Monopoli et al. demonstrated protein adsorption on different NP compositions, sulfonated polystyrene, and silica NPs, using physicochemical concepts to correlate studies of protein corona composition at different plasma concentrations with structural data from in situ and excess plasma-free complexes. [274] In this section, the binding of different proteins to nanoparticles following surface modification is exemplified in Figure 4. PEG is a common ingredient in drug delivery and is widely used for surface modification of NPs. Grenier et al. showed that anti-PEG immunoglobulins modified the settling of protein coronas through activation of the complement cascade and that immunoglobulin M binding to the PEGylated surface of NPs changed the properties of the proteins. [275] Studies have also investigated the effect of various serum proteins on surfaces with hydroxyl regulation using graphene/ gold as an example, evaluated the relationship between surface properties and protein coronas, and verified in mice that apolipoprotein E (ApoE) preadsorption enhances the ability of nanomaterials to survive prolonged blood circulation. [276] Fleischer et al. used polystyrene NPs functionalized with amine or carboxylate groups to create a cationic or anionic surface, while NPs exhibited different binding behaviors with cells after the adsorption of BSA. The enhancement of the cell binding capabilities of BSA-NP complexes formed by cationic substances and the inhibited behavior of BSA-NP complexes formed by anionic substances were attributed to differences in the structure of the adsorbed proteins. [277,278] As another example, apolipoprotein A-I (APOA1) can be bound to nanoparticles through surface modification. Cormode et al. reported that nanoparticles synthesized from high-density lipoprotein, combined with modified lipids and APOA1, could be used to model the in vivo physiology of high-density lipoprotein. [279] Many factors influence the corona of NP biomolecules, and there is increasing research into the role of NP-protein interactions in generating biological responses in vivo [280] and the role of particle surface chemistry in cell binding. [281,282] Protein corona studies have been developed as a field of bio-nano research and extended to environmental matrices to establish the molecular mechanisms of engineered nanomaterials and protein transformations. [283] A perspective on the bio-nano interface will enable more accurate prediction of particle behaviors in vivo, contributing to improved safety and clinical translation.

Applications and Devices
Developments in micro and nanoprocessing technologies have enabled the production of increasingly small mechanical sensors that are capable of detecting the interactions between biological molecules, and this section describes the application of various devices in different fields of biointerface engineering. Bio-MEMS devices have numerous potential applications in disease diagnostics, including medical research, drug delivery, www.advmatinterfaces.de and biosensors. DNA and protein microarrays provide a powerful Bio-MEMS platform for rapid detection, drug discovery, and selection, especially with integrated microfluidic and sensitive detection technologies. [284] Biosensors are analytical devices that combine a biosensitive element with physicochemical sensors to selectively detect the presence of specific molecules in a given external environment. [285] Taking advantage of the ability of biomolecules to specifically bind complementary biomolecules, the most typical interaction resembles antibody-antigen binding. Electroactive polymers have attracted significant interest for a variety of biomimetic applications, [286] including bionic robots and tactile and dynamic sensors. With the emergence of wearable electronics, a higher level of tactility and skin friendliness is expected for more secure human-computer interactions. To date, drug delivery systems have been well developed and used in various therapies. [287] In vivo drug delivery is an interfacial phenomenon, and modulation of biointerface interactions allows for the in vivo delivery of drugs via a carrier system, depending on the physical and chemical properties of the carrier molecules. In the past, tissue engineering tended toward biological or engineering techniques to produce tissue structures to facilitate bone regeneration, [288] whereas modern approaches are converting to an integrated strategy of bioengineering interfaces, often involving biomaterials to provide a biocompatible bionic platform for tissue regeneration as an essential ingredient to support or induce bone regeneration.

Bio-MEMS
Bio-MEMS is an abbreviation for biomedical microelectromechanical systems. It refers to biomedical components made with MEMS process technology according to the needs of the biomedical market. Bio-MEMS devices have the advantages of being light, thin, and short, reducing tedious examination procedures and sample inspections and providing high-efficiency inspection capabilities. Therefore, they have a broader application range and possess multiarray potential to complete multiple inspections and analysis tasks on a single device, and the goal of the so-called Lab-on-a-Chip is achieved. In this section, the interaction of bio-MEMS surfaces and the molecules between the interface are emphasized in detail. Hanein et al. patterned a tetraglyme (pp4G) coating onto silicon substrates via standard photolithography and a plasma polymerization process. Not only did the coating have excellent anti-fouling properties but it also had good adhesion to various substrates. This method allowed precise control of polymer shape, size, and arrangement, providing a reliable process for protein sheets and cell culture and resulting in excellent coating quality, good biocompatibility, and ease of integration into MEMS fabrication. Therefore, the pp4G patterned coating was a considerable advantage in the surface chemical engineering of bio-MEMS to better connect devices to the biological environment. [289] Another bio-MEMS device was designed by Wilson et al. and consisted of cantilevers and  [278] Copyright 2012, American Chemical Society. c) Combination of IgMs with the PEGylated surface of nanoparticles. Reproduced with permission. [275] Copyright 2018, Elsevier. d) Interaction of surface functional groups with ApoE. Reproduced with permission. [276] Copyright 2019, Springer Nature. e) Binding of nanoparticles with ApoA1. Reproduced with permission. [279] Copyright 2008, American Chemical Society. Selective functionalization of nanoparticles enables biomedical applications.

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AFM detection systems for the analysis of cultured myotubes. While selectively stimulating the contraction of myotubes on the cantilever, physiological responses were collected with the AFM apparatus to improve the integration of cellular components and a bio-MEMS device in the future. [290] Moreover, Dhayal et al. fabricated a microfluidic device coated with a plasma-polymerized acrylic acid-arrayed thin film on silicon substrates formed via an optical lithography process to change the surface properties from nonpolymeric to polymeric, and these bio-MEMS devices could be used in different biomedical applications. [291] As shown in Figure 5, by immobilizing and patterning bioluminescent bacteria in a microfluidic chip, Yoo et al. developed a bio-MEMS device to detect specific toxicities. As the luminescence intensity of the bioluminescent bacteria DK1 varied with the presence of chemicals at the interface, this bacterium is often used as a toxicity indicator. [292]

Biosensor
Biosensors measure the physicochemical changes experienced by the biometric layer attached to the solid sensor as it interacts with the sample containing the target molecule. For immunosensor applications, the affinity biosensor interface consists of an antibody, preferably covalently linked to the sensor surface.
With the prevalence of novel coronaviruses, electrochemical immunosensors for virus detection are attracting attention. The primary biomarker for the identification of SARS-CoV-2 is the spike glycoprotein, and a surface made up of recombinant antibodies and a nanoparticle molecularly imprinted polymer was designed to detect this glycoprotein. [293] As a proof of concept, Martins et al. constructed an immunosensor using 3D conductive filamentary carbon black and polylactic acid that provided a natural carboxylic group to anchor biomolecules for detection of the Hantavirus Araucaria nucleoprotein, which could be applied in human serum for virus detection. [294] Antigen testing was performed with a 3D-printed COVID-19 immunosensor by J. Munoz and M. Pumera, who covalently anchored recombinant proteins onto a nanocomposite electrode surface. The electrical readout depended on the monitoring of changes at the electrode/electrolyte interface, following interaction with monoclonal COVID-19 antibodies via competitive assays and contributed to reducing the spread of emerging infectious diseases, [295] as shown in Figure 6. There have been more applications for 3D-printed biosensors in the field of bioanalytical platforms. Sensitive detection tools have been designed for the identification of cancer-related protein biomarkers, combined with fluorescent, electrochemical, and chemiluminescent sensors. [296] Kadimisetty et al. reported that a supercapacitor-powered electrochemical luminescent protein immunoarray allowed short-term detection of cancer Figure 5. A bio-MEMS-based cell-chip that was able to detect specific toxicities was fabricated using bioluminescent bacteria in a microfluidic chip, and the increase in the bioluminescence intensity with the injection of hydrogen peroxide confirmed the capacity of the chip for toxicity monitoring. Reproduced with permission. [292] Copyright 2007, Elsevier.

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biomarker proteins in serum, and this was the first 3D-printed immunosensor for detecting three proteins. [297] In electrochemistry, direct electron transfer enzyme biosensors were constructed for hydrogen peroxide detection via 3D-printed graphene/polylactic electrodes. [298] Multiplexed immunosensors are also available for agricultural applications, with simultaneous detection of the herbicides atrazine and acetochlor. [299] A novel micro 3D-printed electrode biochip enabled simultaneous measurement of multiple biomarkers, including cholesterol and choline. [300] Microfabricated microfluidic devices or microchips used as sensors for immunoassays provide a useful platform for screening drugs or biomolecules. For instance, Yakovleva et al. utilized different modified silica surfaces to immobilize antibodies on silicon microchips for microfluidic enzyme immunoassays with chemiluminescence detection of horseradish peroxidase as the enzyme marker, [301] and Yang et al. demonstrated a microfluidic immunoassay based on lipid bilayers containing dinitrophenyl (DNP)-conjugated lipids that bound to bivalent anti-DNP antibodies and enabled the flow of the fluorescently labeled antibodies into a linear array channel. [302] Another example is combining fluorescence polarization with microfluidic channel technology on polydimethylsiloxane for fast and accurate detection of concanavalin A bound to lectindextran and glycoprotein-acetylcholinesterase at low concentrations. [303] The affinity biosensor interface was the connection between the surface and the transducer to selectively detect the presence of specific compounds; and was therefore, critical to the general performance of the biosensor. We summarized the developments in 3D-printed electrochemical immunosensors and biosensing in microfluidic channels.

Drug Delivery
Drug delivery systems have been used to solve the limitations of conventional drug administration systems. [304] In recent years, the use of carriers in drug delivery systems has increased significantly due to their ability to extend circulation time and residence time and their potential for site targeting and drug protection. [305] However, drug carriers still have limitations regarding their burst release profile. [306] Currently, surface modification of drug carrier systems is often employed to obtain controlled release of drugs and to achieve targeted drug delivery. [307,308] Controlled release of drugs is of great importance in drug delivery because it prevents drugs from being removed from circulation before reaching the target [309] and prevents burst release, which may lead to high initial drug concentrations in the host. [310] In tackling this problem, Ding et al. coated polydopamine on drug-loaded nanoparticles, making the nanoparticles thermally responsive to protect the drug from burst release until adequate heat was generated for therapy. [311] Chang et al. utilized a polydopamine coating on mesoporous silica nanoparticles to produce pH-dependent drug delivery systems. The coating resulted in controlled desipramine release based on the pH of the system, where a higher pH led to a higher release rate. [312] Furthermore, Ghasemi-Mobarakeh et al. coated clotrimazole-loaded substrates with poly(2-hydroxyethyl methacrylate) (p-HEMA) and poly(methacrylic acid) (p-MAA) using the initiated CVD method, which prevented sudden drug release despite the porous structure of the substrate. [313] To produce an effective drug delivery system, targeted delivery is a necessary feature, especially in cancer treatments. Targeted delivery not only ensures that the drugs are delivered to the necessary treatment site but also reduces undesirable side effects that may come with the treatments. [314] Thus, researchers have been looking for a way to achieve targeted drug delivery. [315] Bekaroglu et al. investigated the influence of polymers of different ionic charges on cancer cell targeting. Their results showed that the cationic biopolymer hydroxyl ethylene cellulose notably decreased cancer cell viability but had a nonsignificant effect on healthy cells. [315] In another study, galactosamine was coated on polydopamine-modified nanoparticles  [295] Copyright 2021, Elsevier.

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via ligand-mediated endocytosis to target liver cancer cells, with in vivo results showing that the surface-modified nanoparticles were particularly targeted to liver cancer lesions and were best at reducing tumor cell size. [316] Cao et al. utilized a more biological method to target cells by coating liposomes with macrophage membranes to confer macrophage functions on the carrier. With the macrophage membrane-coated liposomes, in vivo cellular uptake by metastatic 4T1 breast cancer cells was significantly enhanced. [317] Similarly, to achieve self-recognition targeting of cancer cells, drug carrier nanoparticles coated with cell membranes from the same cell line led to significantly increased selectivity to the target cells despite the nanoparticles being in an environment surrounded by various other tumor cells, as shown in Figure 7. [318]

Biorobot
Actuation is critical for a material surface to interact with the surrounding biomolecules and fulfill its designed function. Actuation relies on the properties of cells or tissues to appropriately combine them with functional biomaterials, which can respond to various external stimuli and achieve complex biological activities. Novel materials with good biocompatibility and sensitive actuation capability are attracting much attention for use in emerging biomedical applications and biologically actuated devices. In this section, several specific bioactuators with different material choices in response to different stimuli are introduced. Park et al. developed a novel biohybrid system, a ray that incorporated tissue engineering with the ability to follow light directions and exhibit swimming movements, as shown in Figure 8. To understand the actual batfish fin deflection performance, patterned rat cardiomyocytes were enclosed in microfabricated gold scaffolds. By adjusting the light stimulus, the engineered cardiomyocytes could respond to and allow the bioactuator to move through the obstacle course. [319] Another study reported combining single motor cells with magnetic microtubes and thin films of magnetic materials to develop a micro biorobot that could be guided to a defined location when applying an external magnetic field. This micro biorobot was able to move forward without any fuel using the cell's own flagella. [320] In addition, Tanaka et al. demonstrated that a PDMS micropillar array adhered to neonatal rat cardiomyocytes after surface modification. Through attachment to these pulsatile cardiomyocytes, the array exhibited a typical spontaneous pulsatile phenotype, forming a biological actuator that did not require triggering. [321] Moreover, a muscular bioactuator constructed by culturing muscular cells on the surface of polydimethylsiloxane thin films with micropatterned extracellular matrix proteins was presented by Feinberg et al. These muscle film structures acted as bioactuators by changing the tissue structure, film shape, and electrical pacing when released from heat-sensitive polymer substrates. [322]

Tissue Engineering
Tissue engineering is a field that aims to replace or repair damaged cell tissues or organs by utilizing cells, growth factors, and scaffolds. [323] One of the main reasons for the rise of tissue engineering is the lack of transplant donors when a tissue or organ is so seriously damaged that a transplant is needed, which leads to the idea of using the self-healing potential of the human body to heal patients. [324] To give you an idea, a bioartificial heart may be a conceptual substitute for transplantation due to the long wait for a heart donor. A recent method to mimic the heart involves decellularizing the heart via coronary perfusion, followed by recellularization with cardiac or endothelial cells.  [318] Copyright 2016, American Chemical Society.

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The main advantage of this method is the ability to conserve the natural extracellular matrix and architecture of the heart. A study by Ott et al. showed the pump functions of the heart were stimulated when the heart was provided a physiological load and electrical stimulation. [325] Some examples of tissue engineering research are shown in Figure 9.
Recently, two major methods in tissue engineering development are stem cell-based and scaffold-based methods. [326] In scaffold-based methods, the basic aim is to obtain a scaffold that is capable of mimicking the natural environment of the body to allow cells to adhere and to act as a support to carry cells to the desired site. [327,328] Nonetheless, current developments in tissue engineering have led to the use of tissue engineering scaffolds beyond their basic aim by loading scaffolds with various biomolecules that result in functionalization of the scaffolds. [329] A commonly used method to achieve the functionalization of tissue engineering scaffolds is surface modification of scaffolds, which will be discussed in this section.
In general, there are various methods to functionalize surfaces to immobilize and deliver biomolecules. One way is by modifying the chemistry of the surface to immobilize biomolecules. Lee et al. coated catechin hydrate on scaffolds through self-assembly via cation-π interactions. The catechin coating bestowed antioxidative and calcium-binding abilities on the scaffold, which in turn increased hADSC adhesion, proliferation, and osteogenesis. [330] In another study, a poly(ethyl acrylate) coating applied via the plasma polymerization method caused fibronectin to assemble into nanonetworks to sequester BMP-2, thereby increasing the osteogenesis of hMSCs. [331] Kim et al. coated decellularized ECM from human lung fibroblasts onto PLGA/PLA-based scaffolds, followed by coating with heparin to immobilize BMP-2. hMSCs were then cultured on the scaffold, and the osteogenesis of the cells was measured. Their results indicated that osteogenesis-related markers, such as alkaline phosphatase (ALP) activity and mineralization, were higher in cells cultured on the scaffold than in those cultured on the bare scaffold and fibronectin-coated scaffold. [332] Another surface functionalization method is encapsulation of biomolecules in the coating layer. For instance, Kim et al. loaded BMP-2 in nanoparticles, which were then incorporated Figure 8. The design of a tissue-engineered ray. When the frequency of light pulses was adjusted, the artificial ray changed direction and passed obstacles by creating an asymmetric wave motion between the left and right fins. Reproduced with permission. [319] Copyright 2016, American Association for the Advancement of Science.
www.advmatinterfaces.de Figure 9. Examples of tissue engineering scaffolds: a) schematic illustration depicting the fabrication of a multifunctional scaffold obtained through vapor sublimation and deposition that allows for the incorporation of biomolecules in the scaffold. The fabricated scaffold can also be compartmentalized into various geometries, with oil red-O and fluorescein-5-isothiocyanate used to show the compartmentalization. Reproduced with permission. [333] Copyright 2021, Springer Nature. b) Bioartificial organs are regarded as a way to solve the problem of the lack of transplant donors. In the figure, the mouse heart was decellularized and then recellularized with cardiac and endothelial cells in the hope of application as a bioartificial heart in the future. Reproduced with permission. [325] Copyright 2008, Springer Nature.

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on the surface of hydroxyapatite scaffolds to enhance cellular behaviors on the scaffolds. Their in vitro experiments showed that the osteogenic properties of hMSCs were enhanced on coated scaffolds, and in vivo results showed an increase in bone ingrowth in comparison to uncoated scaffolds. [262] Similarly, polylactic acid scaffolds were coated with BMP-2-loaded polyelectrolyte film to induce cell differentiation and bone regeneration. BMP-2 induced markedly increased bone formation on the scaffold. [177] In another example, black phosphorus was loaded in graphene oxide, which was then coated on the surface of poly(propylene fumarate). Phosphate is slowly released following the oxidation of the black phosphate nanosheets, which results in enhanced osteogenesis. [184] Wu et al. utilized vapor sublimation and deposition processes to build a 3D scaffold with distinct compartments incorporating multiple biomolecules and living cells. With the addition of biomolecules and living cells, the scaffolds induced enhanced proliferation, osteogenesis, and neurogenesis. The compartmentalization of biomolecules and cells also allows cell coculture to be performed in different compartments in the same scaffold construct. [333] In addition, altering the physical structures of scaffold surfaces is a common way to enhance differentiation in tissue engineering. Govindan et al. utilized alginate, chitosan, and gelatin to obtain a scaffold with enhanced mechanical properties for bone tissue engineering, and the fabricated scaffold induced a sevenfold increase in comparison to the pure scaffold without the coating. [334] On the other hand, to obtain scaffolds with an enhanced Young's modulus and good protein adsorption, Luo et al. coated hydroxyapatite on alginate/gelatin scaffolds, which resulted in a twofold increase in the Young's modulus in comparison to that of the pure scaffold. [182] In a study by Wu et al., poly(3,4-ethylenedioxythiophene) polystyrene sulfonate (PEDOT:PSS) was incorporated through vapor sublimation and deposition processes to produce a scaffold with electrical conductivity potential. They reported an increase in Nestin and Tuj-1 expression, which indicated neurogenic differentiation enhancement. [335] In another study, conductive polypyrrole/ poly(D,L-lactic acid) scaffolds were fabricated via emulsion polymerization and the dip coating method, and PC12 cells were seeded and then stimulated with 100 mV for 2 h. The cells seeded on the scaffold showed increased neurite-bearing cells and average neurite length. [336]

Future Outlook
Biomaterials are usually based on the most basic properties of biocompatibility and the ability to exist stably for a long time in the living body. More important is the particular function of the biointerfaces, which in turn affects the interaction with the surrounding body tissues so that the materials can achieve the ideal therapeutic effect as an implant. In the selection of biomaterials, the decision will be made according to the needs of the expected target. The three common types of materials are bioceramics, polymers, and natural biomaterials. To date, bioceramics seem to be an excellent choice as an interfacial coating between implants and bones due to their superior osteoconductivity and chemical stability. By enhancing the adhesion of the interface, bioceramic coatings are widely used in bone tissue engineering because they maintain a good interaction with the bone after implantation. However, a major challenge in clinical application still remains, particularly the problem of infection around bone implants. Pathogenic infections occur during surgery and result in secondary surgery. Hence, the development of antibacterial coatings is an important issue, which can be enhanced by surface modification or the incorporation of other substances, preserving the mechanical properties of the implant and increasing its biocompatibility. Bioceramic coatings are an attractive option in the field of biomedicine and offer great promise. Polymer materials are the focus of modern medical materials because they have high mechanical strength, are easy to process, and exhibit physicochemical stability, which cannot be easily achieved with metal or ceramic materials. Polymers can also form porous structures that allow for tighter bonding between materials and tissues. However, challenges, such as poor degradability or the problem of blood coagulation, will be encountered after long-term implantation. If the host response can be minimized, it is still the primary focus in the development process of biomaterials. The prospects of ideal medical polymers with biomedical functions are extensive. In addition, due to their biocompatibility and natural ability to mimic tissue-specific in vivo microenvironments, natural biomaterials have found various applications as coating biomaterials for tissue and regenerative engineering scaffolds, as well as in cell and drug delivery systems. However, many of these biomaterials lack mechanical strength, especially in vivo, due to their rapid degradation. Thus, studies to improve the mechanical properties of these biomaterials would be highly beneficial for further in vivo applications. Moreover, studies to identify novel natural biomaterials that are superior to the currently used biomaterials are also under way, with the hope that these novel biomaterials can also find applications in vivo. In addition, biomaterials must be matched with the appropriate fabrication technique to obtain the best coating quality, and the equipment requirements and processing conditions of each method are different. The coating resolution that can be achieved is highly related to the complexity of fabrication, which usually needs to be sacrificed for convenience. For example, spray coating, dip coating, and soft lithography are methods that are easy to perform, cost little, and are suitable for fast and mass production; in contrast, when high-precision and specific coating thicknesses are needed, photolithography and CVD are better choices.
Surface properties significantly influence biomaterials and greatly determine the performance of the material's functionality. In terms of structural properties, the surface topography of material may provide different cues for cells that may lead to different cell fates. In addition, porous structures on a biomaterial interface may provide an environment that is more similar to that in vivo. Thus, modification of the structural properties of biomaterial interfaces may help improve the desired cell adhesion and differentiation fate. Cell culture systems have also transitioned from 2D to 3D systems as 3D cultures provide cultured cells with a microenvironment that better mimics their natural extracellular matrices in vivo. In terms of biochemical properties, to better adapt biomaterials to complex biological microenvironments, special functional groups on the surface of www.advmatinterfaces.de materials are often used to provide conjugation for specific biomolecules to achieve particular functionalities, including anticorrosion, antifouling, immobilization, and therapeutic functions, and even multiple functionalities on the same material surface. For example, in vivo studies on many clinical applications of dental implants have been reported. Physical modification can change the surface morphology and roughness of the implant to induce cell adhesion and stimulate osteogenesis, while surface chemistry can be modified through adsorption or chemical reactions to change the surface structure, such as the preparation of nanosurface structures that promote collagen synthesis in gingival fibroblasts and improve the adhesion of surrounding tissues. Moreover, from the perspective of the nanointerface, the interaction between proteins and surfaces can be improved through various coatings on nanoparticles. Accurate prediction of particle behavior in the body can reduce nonspecific protein adsorption and bacteria adhesion on the surface to prevent infection. In the future, more optimized coatings will be developed to achieve significant progress in clinical applications. With the current focus on biomaterials and the continuous growth of the field, many biomedical applications have been developed. The demonstrated examples include bio-MEMS devices, biosensors, drug delivery agents, biorobots, and tissue engineering approaches. These devices are well and comprehensively combined with either the structural or biochemical properties of biomaterials and have great potential and promise for the development of biocoatings in the field of biomedical engineering in the future.
With more sophisticated interfacial properties being pursued due to the more precise and stringent requirements for prospective biomaterials, the development of a unique and universal interface coating technology that produces a specific property may seem outdated, but a combination approach to obtain advantageous properties is feasible using current coating methods, considering i) precision presentation of biofunctional molecules on a relevant scale, ii) reduced background nonspecific binding of undesired molecules, iii) multifunctionality, iv) a dynamic response of interface properties to external stimuli, v) an optimized compatibility with a specific microenvironment or fabrication process, vi) a controlled anisotropic property (e.g., a physical and/or chemical gradient, hierarchical composition, or compartmentalization), and vii) biodegradability. To date, flawless control and combination of these interface properties for coatings is still difficult to achieve but should pave the way for modern biointerface science and biomaterials science.