Porosity is an important material feature commonly employed in implants and tissue scaffolds. The presence of material voids permits the infiltration of cells, mechanical compliance, and outward diffusion of pharmaceutical agents. Various studies have confirmed that porosity indeed promotes favorable tissue responses, including minimal fibrous encapsulation during the foreign body reaction (FBR). However, increased biofilm formation and calcification is also described to arise due to biomaterial porosity. Additionally, the relevance of host responses like the FBR, infection, calcification, and thrombosis are dependent on tissue location and specific tissue microenvironment. In this review, we discuss the features of porous materials, and the implications of porosity in the context of medical devices. Common methods to create porous materials are also discussed, as well as the parameters which have been used to tune pore features. Responses towards porous biomaterials are also reviewed, including the various stages of the FBR, hemocompatibility, biofilm formation, and calcification. Finally, these host responses are considered in tissue specific locations including the subcutis, bone, cardiovascular system, brain, eye, and female reproductive tract. We highlight the effects of porosity across the various tissues of the body and emphasize the need to consider the tissue context when engineering biomaterials.
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Porosity is an important biomaterial feature which enables biomaterial functionality and the capacity for cell integration, which is commonly characterized as a response with greater biocompatibility. In this review, we discuss methods used for fabricating porous materials, the effect porosity and specific pore size has on various host responses, and the responses which occur in tissue-specific microenvironments.
In this review, we discuss the medical applications for porous materials and various methods for fabricating porous materials for research. We further summarize the biological responses to biomaterial implants including the FBR, hemocompatibility, biofilm formation, and calcification, as well as the specific effect porosity has on these pathologies. Tissue-specific responses to porous biomaterials are additionally discussed, highlighting the beneficial and negative applications of porosity in foreign bodies. Here, we emphasize the need for greater synchrony between the knowledge of biomaterial engineering and local tissue physiology. Information presented in this review should also motivate the need to characterize materials in relevant tissue compartments, as well as the existing need for developing improved tissue models. Further research of the biological mechanisms behind specific host responses could also inform the development of such models, as well as the design of materials which can exploit these responses. Considering the dedication of this special issue, we highlight the pioneering and ongoing work by members of Professor Buddy Ratner&#;s Biomaterials Group.
Although the FBR is the main consideration for material biocompatibility, the resulting FBR can be beneficial in some cases. For devices requiring osseointegration, calcification and an encapsulating FBR can be beneficial for generating integrated and mechanically stable tissue. [ 1 , 20 , 21 ] Generalizing the FBR has also caused clinical issues, especially for porous devices in the female reproductive tract, where dampened tissue-integrative responses commonly lead to rejection of implants like surgical mesh. [ 22 ] Porous material can also be at an increased risk of biofilm formation &#; an infection of bacterial communities on the material surface &#; especially where tissue integration is low, and microflora tolerance is promoted. [ 23 ] Although the FBR, calcification, and biofilms are all important concerns for cardiovascular implants, hemocompatibility must be considered for any material that is designed to be blood-contacting. Blood is protein rich and composed of leukocytes and platelets which can initiate inflammation or thrombosis in response to biomaterial surfaces, potentially resulting in fatal thromboembolization. [ 24 &#; 26 ] Local tissue microenvironment and implant function are therefore crucial considerations when engineering material features such as porosity.
Indeed porosity, and specific pore sizes, have shown to be a determining factor of biocompatibility with reduced fibrous capsule formation and increased angiogenesis when subcutaneously implanted. [ 10 , 15 , 16 ] Depending on the intended function of the implant and the local tissue microenvironment, biological responses other than tissue integration or the FBR may be more significant in determining implant acceptability. Further, the host response has been observed to vary between different tissue microenvironments. [ 17 &#; 19 ] Therefore greater consideration of tissue specific responses towards material porosity and specific pore size is needed to better inform truly biocompatible outcomes.
Biomaterial selection dictates the functionality of medical devices &#; such as drug delivery systems, grafts, sensors, electrodes, and protheses. One of the most important determining factors of device success is implant safety and biocompatibility, which can be regulated by biomaterial composition and architecture. [ 1 ] Tissue mimicry can inform functional and biocompatible material designs. [ 2 ] One of the most notable features of tissue matrices is porosity, which is defined as having an architecture with interconnecting open spaces. [ 3 ] Pores allow for a multitude of material functions that include: cell integration with and into the material; [ 3 , 4 ] inward dispersion of oxygen, nutrients, analytes, and outward diffusion of pharmaceutical agents; [ 5 , 6 ] angiogenesis; [ 5 , 7 , 8 ] and pro-healing responses from immune cells [ 9 &#; 11 ] that can affect the foreign body reaction (FBR). [ 6 , 12 &#; 14 ] The FBR describes the innate response towards foreign materials, which results in immune cell infiltration at the material surface and the development of a fibrous capsule, or avascular scar tissue growth which encapsulates the implant. [ 1 , 12 , 13 ] Such a response must be controlled, as a fibrous capsule can cause the patient pain, deform materials, alter the functional mechanics of the implant, limit drug release, and impede signals from electrodes and sensors. [ 1 ] Therefore, biocompatibility and the FBR not only have implications for patient wellbeing, but also material functionality.
Factors of porosity such as the increased surface area, decreased path of diffusion within the matrix, and greater inward or outward flux may increase drug release rates, and specifically drug release mediated by polymer scaffold erosion. [ 29 , 122 ] Experimentally, materials with larger surface-to-volume ratios &#; like materials with high porosity &#; have shown faster, diffusion mediated release. [ 29 , 122 ] However, this correlation is material dependent, as polymers with acidic byproducts such as poly(lactic acid) (PLA) and poly(lactic-co-glycolic acid) (PLGA) show increased drug release for materials with lower porosity, lower permeability, and thicker material walls. Accumulation of carboxylic groups within pores accelerates the hydrolysis of these degradable polymers. [ 139 &#; 141 ] Higher porosity can also imply a decreased deliverable dose or a reduction of the polymer which controls agent release. [ 121 ] Therefore, the release kinetics of drug delivery systems are heavily influenced by both material composition, material pore morphology, and the interaction of these factors.
This form of the Higuchi equation accounts for a change in the effective diffusion constant for porous mediums, which increases with higher porosity and lower tortuosity, thereby also increasing drug release. [ 136 , 138 ] The diffusion coefficient also changes dynamically in response to matrix swelling. [ 136 ] In the case of porous mediums, swelling alters pore size and tortuosity overtime. [ 138 ] The presence of porosity in a drug delivery system described by the Higuchi equation also modifies the drug solubility factor (C S ), which accounts for the partition of drug concentrated inside and outside of the pores. [ 135 , 136 ]
Drug delivery is one important application of biomaterials and is impacted in various ways by porosity. The release of drug into biological solution requires both drug dissolution and diffusion. Within a polymer matrix, drug release can occur by the erosion of the encapsulating polymer or by diffusion through a polymer matrix - both of which may be impacted by porosity. [ 135 ] The Higuchi Equation separately considers non-porous and porous drug release systems ( Equation 3 ), is widely used, and simply estimates the flux of total drug release (Q) over time (t) per unit area (A) considering drug dissolution and diffusion. [ 135 , 136 ] Further, this equation is modeled one-dimensionally for a rate-limiting ointment film releasing drug into skin under sink conditions, where D is the diffusion coefficient of drug into the biological solution, ε is the porosity of the material, τ is the tortuosity &#; or the size and branching of the interconnected pores, C 0 is the initial concentration of drug, and C S is the drug solubility within the matrix. [ 137 ]
While attachment, drug release, material erosion, and mechanical properties are impacted relatively by porosity, the specific selection of the material will also have a critical impact on these attributes. Further, specific material or polymer composition will provide features like erodibility or swellability that can alter pore geometry in situ over time. [ 129 , 130 ] Specific pore shape, strut curvature, and alignment has also shown determining effects on cellular responses. [ 33 ] To delineate the contribution of porosity alone, research commonly assesses chemically identical materials with multiple pore sizes (see for methods to alter pore size per method), [ 10 , 15 , 80 , 131 ] or screens various materials and assesses the correlation of pore size effects in addition to the composition. [ 16 , 132 &#; 134 ] The effect by pores as opposed to wall structure has also been studied by compressing porous materials after identical fabrication methods. [ 80 ]
Porosity also impacts material surface-to-volume ratios &#; as greater percent porosity increases material surface area, and larger pores with constant percent porosity decreases material surface area. [ 122 ] Wall or strut size is therefore also a determinant of the surface-to-volume ratio. Materials with larger surface-to-volume ratios can enable greater access for cell attachment, [ 115 , 121 ] or increase drug release rates (described in section 2.3 ). [ 122 ] The presence of pores and walls also implies the presence of texture or roughness at the material surface, which additionally impacts cell attachment, protein absorption, inflammatory responses, platelet activation, and bacterial adhesion. [ 25 , 37 , 125 &#; 128 ] Increased porosity also decreases scaffold support which can inversely affect the mechanical integrity of the material. [ 36 ]
When considering material porosity, context is important for evaluating porosity effects on performance outcomes. For implants and tissue engineering scaffolds, voids are typically intended to permit cell migration, the influx of oxygen and nutrients to sustain these cultures, and the outflux of metabolic wastes. [ 5 , 120 ] For drug delivery, these voids are intended to be saturated with biological fluids, thereby enabling drug release by modes of outward drug diffusion or hydrolysis mediated material erosion. [ 29 , 121 , 122 ] Thus, permeability is an important effect of porosity and influenced by pore size and their interconnections, bulk material dimensions, surface characteristics of the material such as hydrophilicity, and pressure differences. [ 3 , 27 ] Capillary action is one mechanism of permeability, which mediates uptake of fluids within pores due to pore size and surface tension of the solution with the material walls. [ 123 , 124 ] While decreased pore or throat size may limit permeation of agents such as cells via size exclusion, [ 5 , 41 ] capillary action dictates an increased fluid permeation distance with smaller pore diameters. [ 124 ] Therefore, the impact of pore size on biological fluid permeation depends on the target permeating agent and the surface interactions of the fluid and material walls.
Meshes are commonly used in biocompatibility studies of porous features as these materials are truly porous, thin, tunable, and are clinically relevant to tissue support and hernia repair surgeries. [ 79 ] Electrospinning yields sheets of materials with nano- or micro-scale fibrous architectures, and is therefore commonly used for drug delivery research, tissue engineering, and study of the host response ( ). [ 18 , 80 , 81 ] Specifically, some of these applications include wound dressings, [ 82 &#; 85 ] vascular grafts, [ 86 &#; 91 ] and intravaginal dosage forms for antiretroviral drugs. [ 92 &#; 99 ] Additionally, this method is amendable to virtually any natural or synthetic polymer and can efficiently encapsulate physicochemically diverse pharmaceutical agents or growth factors. [ 81 , 100 , 101 ] Conventional spinning methods start from a solution composed of desired polymers, agents, and a volatile solvent, which is loaded into a positively charged syringe. This charged solution is then extruded across from a negatively charged collector, so that fine fiber protrusions form from the tip of the needle, and deposit onto the collector as a sheet of solid, solvent-free material. [ 81 , 101 ] Various factors including polymer concentrations, flow rate, voltage, solvent, and modified collection targets with perforations or rotation can be altered to change the dimensions and alignment of the electrospun fibers, and therefore pore size and porosity. [ 80 , 102 &#; 104 ] However, the resulting porosity from electrospun fibers is typically lower than other scaffold fabrication methods ( ). Much like 3D printing, the electrospinning process can be adapted with subtractive methods like porogen leaching and freeze casting to improve porosity. [ 32 , 105 , 106 ]
Recent advances in 3D printing technology have made this an increasingly popular method for creating porous scaffolds ( ). [ 31 ] 3D printing broadly includes additive manufacturing techniques capable of precision patterning materials. [ 30 ] Thus, pores can be formed through the spacing of deposited or cured materials. For a comprehensive description of these 3D printing methods, the reader is directed to the review by Ngo, et al.. [ 30 ] Various materials are amenable for 3D printing techniques, including polymers, ceramics, metals, and even living cells. [ 30 , 31 ] The ease of making larger pore features with this method makes printing bone scaffolds relevant, [ 36 , 75 ] but bioprinting has made a variety of tissues and vasculature structures possible, [ 76 ] and 3D printing methods can be even amenable for drug delivery. [ 77 ] Some printing methods like stereolithography have a high feature resolution of less than 10 μm, and greater porosity can also be achieved by combining subtractive methods like salt leaching. [ 78 ]
Decellularizing tissues is another strategy commonly investigated to create biocompatible and porous materials, used for notable tissue scaffold applications like bioartificial hearts. [ 68 ] However, the resulting scaffolds often lack mechanical integrity, or contain pores only a few nanometers in diameter which are too small for host cells to repopulate. [ 69 &#; 72 ] To mitigate this, decellularized tissues have been additionally processed with subtractive methods to increase porosity, but this further sacrifices mechanics. [ 72 ] Additive methods combined with these scaffolds can reinforce tissues, but at the disadvantage of decreased porosity. [ 70 ] Emerging research is therefore more focused on utilizing extracellular matrix (ECM) components from decellularized tissues as a medium with additive methods like electrospinning, 3D printing, or molding for both porous and robust materials. [ 71 &#; 74 ]
Comparably, gas foaming methods use supercritical carbon dioxide as a porogen in polymer solutions, which expands to create pores when returned to a gas state ( ). [ 5 , 62 ] One of the main benefits of this technique is that organic solvents are not used, which can have adverse biological and environmental effects. [ 62 &#; 65 ] Due to low solubility of CO 2 in hydrophobic polymers, the use of surfactants and water emulsions are commonly used, although these factors alter pore size. [ 62 , 63 ] The resulting high porosity from both gas foaming and freeze casting make both of these methods widely used for tissue engineered scaffolds in general, and can generate materials with pores sufficiently large for bone ingrowth. [ 36 , 66 , 67 ]
Freeze-drying or freeze casting is another subtractive method used commonly for making porous materials ( ). The freeze-casting process starts from a crosslinked or colloidal suspension commonly composed of an aqueous polymer solution. This solution is added to a mold, frozen, then sublimated to remove regions of solidified ice crystals. [ 56 , 57 ] Various parameters in the freezing process can modify pore size and various polymers can be utilized. However, pore size can be difficult to tune with freeze-drying, the processing time can be long, energy intensive, and the resulting materials often have poor mechanical integrity. [ 5 , 36 , 56 , 58 ] Alterations on the freeze-drying process include directional freezing to control pore morphology, [ 5 , 56 , 59 ] freeze-extraction and freeze-gelation to eliminate the time and energy needed for the freeze-drying process, [ 60 ] as well as added freeze-thaw cycles to increase scaffold toughness. [ 5 , 61 ]
Perhaps the most frequently employed progeny leaching method for researching the effect of porosity on biocompatibility is sphere tinplating ( ). [ 4 , 9 , 10 , 38 &#; 43 ] Devices including hemodialysis grafts [ 44 ] and supraciliary drainage implants for glaucoma treatment [ 45 , 46 ] are fabricated using sphere templating, illustrating the translatability of this method. Sphere templating is a subtractive method which creates pores by first forming an array of sintered sphere microparticles, then filling the spaces of the array with polymer precursors, solidifying the polymer, and finally removing the beads so that just the polymer scaffold remains. [ 47 ] Pore size and porosity is therefore controlled by the geometry of the beads, which are commonly composed of polymers like poly(methyl methacrylate) (PMMA). [ 7 , 9 , 10 , 40 , 48 &#; 50 ] Similarly, salt particles are commonly used as a porogen in solvent cast and hydrogel materials ( ). [ 5 , 51 &#; 53 ] Although a wide range of pore sizes have been achieved, [ 5 , 51 &#; 53 ] salt crystals yield irregular pore geometries. [ 40 ] Various porous materials have been created with sphere templating, commonly UV polymerizable polymers like poly(hydroxyethyl methyacrylate) (poly(HEMA)), as well as silica, carbon, and even metals. [ 4 , 9 , 10 , 38 , 39 , 41 &#; 43 , 47 , 54 ] Selective UV polymerization can also be used as a strategy for patterning voids into materials using photolithography, which creates channels in materials using a 2D mask pattern with macro- or nanoscale pore diameter precision, and can be combined to create complex structures in combination with other fabrication methods. [ 54 , 55 ]
Materials with macroscale pores can be fabricated using a variety of techniques. These methods can be generalized as either subtractive or additive: either creating pores through the removal of sacrificial materials called porogens, or manufacturing walls/struts to surround open spaces, respectively. [ 27 , 30 , 36 , 37 ] Porosity can be an inherent feature of crystalline structures, but the resulting pores typically have molecular dimensions. [ 27 ] Here, we will focus on some of the most utilized methods that create tissue-scale, macro porous biomaterials. A summary of porous fabrication methods and resulting pore characteristics is listed in and can be seen in .
Measurements of porosity, or void fraction, are reported as either a percent or ratio (ε) of the pore volume (V P ) to the total bulk volume of the material (V T ). Materials are considered to have low porosity when ε < 30%. [ 35 ] Pore volume is typically calculated using the density of the material used (ρ m ) and the mass of the sample (m S ). [ 3 , 27 ] Another simple way to calculate porosity uses Archimedes principle, which states that the volume of a fully emersed body is equal to the volume of water displaced. Pore volume can be calculated by the difference of wet and dry material weight, while total volume can be calculated by the difference of wet and fully submerged material. However, this method is not accurate for hydrophobic materials due to poor water penetration. [ 3 ] Equation 2 shows these calculations of percent porosity.
Materials with porous features are mainly characterized by pore size or by the percent or ratio of porosity in the medium. For fibrous porous materials, strut or fiber diameter is also an important feature to characterize. Pore and strut size is most commonly estimated using imaging. [ 3 ] Methods such as optical microscopy, scanning electron microscopy (SEM), and transmission electron microscopy (TEM) can capture sections or regions of the material, then image analysis software can measure the pore or strut size features within the region of interest. [ 3 ] While these methods are simple, intuitive for assessing the material morphology, and avoid preparation methods that may modify native microstructures, imaging methods fail to capture the complete 3D structure of the material. [ 3 ] Other methods such as mercury and flow porosimetry can also quantify pore size as measured by the differential gas pressure of mercury or the measured pressure of a wetting fluid (P). However, these methods cannot measure closed pores and may have variable accuracy depending on surface and wetting fluid interactions. [ 3 , 27 ] Using the Washburn equation ( Equation 1 ) or similarly the Young-Laplace formula, pore diameter (D) is calculated as a function of the surface tension of mercury or the wetting liquid (γ), and the contact angle of mercury or the wetting liquid and the sample (θ). [ 3 ]
Porosity is defined as the void of material, and is further characterized by the interconnections, or throats, between these pores, and walls or struts of the medium which forms the 3D structure ( ). [ 3 ] Features of pores, throats, or struts are characterized by their size, shape, organization, density, and homogeneity. Traditionally, pore scale is delineated as macroporous when greater than 50 nm, mesoporous between 2 nm and 50 nm, and microporous when smaller than 2 nm. [ 27 &#; 29 ] Nanoporosity is also used to describe materials within the nano-scale range, with pore diameters between 1 and 100 nm. [ 28 ] Spherical, tubular, and random pore structures are commonly observed in porous materials, but novel fabrication methods are continuing to develop complex, high-resolution geometries, [ 30 &#; 33 ] and even materials with transforming topologies. [ 34 ] The geometry of these pores and interconnections are further defined as either closed, open, or blind-ended. [ 3 ]
In the literature, it is wideley recognized that materials with higher porosity and pore size can initiate greater calcification ( ). [ 157 &#; 159 , 161 ] Comparing pHEMA hydrogels with varing porosities in buffer solutions, Lou, et al. found small deposits of calcium around implants with 58% porosity, large clumps of mineralization on implants with 96% porosity, and increased calcification around material defects. [ 159 ] In micro- and macroporous gels, Šprincl et al. found larger pores allowed for deeper calcification of subcuteanous implants ( ). [ 157 ] Also studied subcutaneously, Golomb, et al. found that greater calcium deposition towards porous films was an effect of the increased water capsity. [ 158 ] However, vascular graft implants with decreased porosity have been observed to initiate degenerative effects that initiated greater calcification. Wesolowski, et al. therefore concluded that materials with porosities greater than mL of water per minute per square centimeter of material at a pressure of 120 mmHg could prevent calcification and other responses. [ 132 ] Although it is true that porosity in general leads to greater calcium deposition, it is important to remember that inflammatory responses also initiate mechanisms of mineralization. Therefore, in complex biological systems, it is important to consider the local tissue response, and to prioritize material designs that reduce inflammatory responses.
Biomaterial calcification or mineralization can commonly arise on devices contacting bone, blood, and urine and biofilms. [ 161 ] For orthopaedic implants, calicification results in osseointegration and therefore implant success. [ 192 ] For devices such as heart valves, calification can change the mechanical properties of the material, causing implant failure. [ 161 ] In general, mineralization occurs through electrostatic interactions between calcium and phosphate ions and an anionic material surface. For tissues or naturally derived scaffolds, this can occur at amino acids with anionic functional groups. [ 193 ] Osteoblasts mediate calcification in bone tissues. [ 192 ] In soft tissues, cytokines like TNF-α released by macrophages can cause osteogenic differenciation of local progenitor cells which enables mineral deposition. [ 161 , 193 ] Bacteria within a biofilm can release enzymes that lead to an increase in pH and promote hydroyapeptite crystal formation. [ 161 ] Calficication is not always cell mediated and can also form from free minerals in serum or urine. [ 158 , 161 ]
In an investigation of biofilm formation towards various dense and porous implants with 100 - 200 μm pores, Merritt, et al. found that porous materials were more susceptible to biofilm formation if the material became contaminated before or at the time of implantation. However, the porous materials were more resistant than dense materials if the infection occurred after 28 days, when the tissue had integrated into the material ( ). [ 156 ] Sclafani, et al. assessed biofilm formation towards high-density polyethylene (PHDPE, Medpor) materials with pores 100 - 250 μm and ePTFE with internodal pore sizes ranging from 10 - 30 μm. For both materials, implants became infected when inoculated with S. aureus immediately following surgery. When inoculated 14 days after implantation, PHDPE materials with larger pores were found to be more resistant towards infection due to faster tissue integration. Other studies have shown tissue integration with ePTFE materials, so infection may not arise with these implants at later timepoints. Specific material composition may also confound the effect by pore size alone. [ 134 ] While porosity does increase the risk of biofilm formation, porous materials still provide benefits for tissue ingrowth which can decrease the risk of infection ( ). Therefore, surface treatments can be an effective strategy to control pathogen growth while still gaining the other advantages of biocompatibility that porous materials provide.
The correlation between porosity and infection is not as simple when studied in complex biological systems. The &#;race for the surface&#; describes the competition between host tissue and bacterial communities for space on the implant surface. [ 79 , 191 ] Tissue integration is therefore one of the most effective mitigation strategies for biofilm formation ( ). [ 79 , 160 , 191 ] Although porous mediums with large available surface areas can promote the adherence of biofilms, these porous materials also encourage tissue ingrowth, as described in section 3.1.5 .
Overall, implants with greater porosity have been implicated with a higher risk of biofilm formation ( ). This is especially true for small-scale porous materials that may be size exclusionary towards leukocytes but not bacteria ( ). [ 160 ] On titanium surfaces, Braem, et al. found greater bacterial adhesion to materials with pore sizes up to 150 μm or porosities greater than 15% because of greater surface roughness ( ). [ 155 ] Antibiotic prophylaxis can reduce the risk of infection, as well as using materials surface modified with nonadherent properties, anti-microbial agents, or specific topographies. [ 160 , 190 , 191 ] Feng, et al. also studied the effect of pore size on biofilm formation towards alumina materials with 0, 15, 25, 50, and 100 nm pore diameters in vitro and via computational modeling. Materials with either 15 or 25 nm pore sizes were found to reduce bacterial attachment through repulsive forces from the densely packed vertical pore sidewall within the anodic surfaces. [ 154 ] This effect of porosity on bacterial attachment is therefore also dependent on pore geometry and chemical composition of the material.
Biofilms are formed through a cycle with four stages. First, motile bacterial cells adhere to the material surface. [ 23 , 160 ] Bacterial cells can adhere reversibly, with non-specific forces or irreversibly with specific interactions with lectin or adhesin. [ 160 ] Next, adhered cells form a colony and secrete EPS. These bacterial colonies continue to grow and form into a mature biofilm, where cells can remain dormant until favorable conditions for infection arise. Dormancy, along with the colony structure, make biofilms characteristically resistant to antibiotics. [ 23 , 160 , 190 ] Finally, the biofilm disperses as aggregates or sessile bacterium to escape regions of accumulated bacterial waste and to start new regions of infection. [ 23 ]
Biomaterials can also initiate adverse responses by creating a new niche for foreign pathogens within the body. Indeed, the bacterial colonization of medical devices, known as biofilm formation, is one of the most frequent complications of clinical biomaterial use. [ 160 , 189 ] While the host immune system might normally clear invading bacteria, the local fibrous capsule generated in response to the biomaterial creates an immune depressed environment, and bacterial colonies within a biofilm are protected by extracellular polymeric substances (EPS). [ 23 , 160 ] Thus, biomaterial surfaces can enable the persistence of bacteria such as Staphylococcus aureus within the body that lead to device failure and chronic disease. [ 160 ]
The fluid dynamics of blood is another consideration for hemocompatibility. Porosity mediated water permeation is beneficial for transferring nutrients to cells within the tissue. However, greater permeation can increase platelet activation and lipid infiltration. [ 150 ] Additionally, materials with porosity higher than 50 mL of water min &#;1 cm 2 of material at a pressure of 120 mmHg can cause hemorrhage when treated with anti-coagulants. [ 132 ] Porosity and pore size importantly contribute to the complex interactions which occur between biomaterials and blood, and these generalized conclusions for thrombogenesis and platelet activation are summarized in . In general, smooth surfaces are typically preferred for improved hemocompatibility, but porous materials are often a necessary deign feature for blood contacting materials to permit suturing or for local tissue integration. [ 152 ] Section 4.3 includes a discussion of both hemocompatibility, endothelium passivation, and local tissue healing responses towards porous materials in the cardiovascular system.
Interestingly, the materials which are the current standard for hemocompatibility &#; expanded polytetrafluoroethylene (ePTFE, Gore-Tex) and poly(ethylene terephthalate (PETE, Dacron) &#; have micron scale pores (~5 μm). [ 152 , 153 , 183 ] Porous materials permit tissue ingrowth and can affect mechanical compliance, both which can contribute to more favorable outcomes for hemocompatibility. [ 87 , 150 ] Clot formation is still observed on the surface of these materials, but this response is considered acceptable as long as the clot does not embolize. [ 152 , 183 ] The complex response of porous materials in the cardiovascular system are further discussed in Section 4.3 .
Mechanical activation of platelets can also arise with both high and especially low fluid shear forces. [ 184 ] Indeed, simulated and experimental results have shown that 75 μm crevices in a surface can induce shear mediated thrombosis ( ). [ 151 ] Milleret, et al. found electrospun scaffolds with 5 μm material struts and an average pore diameter approximately 35 μm yielded greater platelet adhesion and activation than non-porous controls ( ). Materials with approximately 10 μm pores and struts less than 1 μm behaved comparably to non-porous materials ( ). Thrombosis was also independent of polymer used or hydrophilicity of the material, so the effect is attributed to the changes in topography. [ 133 ] Zhao, et al also found that materials with pores 2-5 μm induced statistically similar low platelet adhesion and high whole-blood compatibility as compared to smooth surfaces, while 35-45 μm porous materials yielded statistically greater platelet adhesion and lower whole blood compatibility ( ). [ 153 ] Overall, surface roughness associated with greater porosity can be considered more thrombogenic, but several studies support opposite findings. [ 186 &#; 188 ] Interestingly, these in vitro studies that find reduced platelet activation towards materials with greater porosity, and specifically materials with approximate 30 μm pore sizes, use static testing conditions and sodium citrate instead of heparin as an anti-coagulant to reduce fibrin formation. [ 187 , 188 ] Although the use of any anti-coagulant will alter thrombogenesis, sodium citrate is a calcium chelator, which is especially problematic for blood-biomaterial interactions because the removal of calcium will reduce platelet-surface interactions. [ 152 ] Responses may also be attributed to differences in strut geometry arising from different fabrication methods. The varied platelet response to porosity illustrates the complexity of blood-material interactions, as well as the need to standardize hemocompatibility assessment methods.
In addition to the cellular components, blood is composed of important plasma proteins like albumin, fibrinogen, and immunoglobulins. [ 24 ] As described in the first phase of the FBR ( section 3.1.1 ), blood serum proteins spontaneously absorb onto biomaterial surfaces after implantation due to the inherent blood contact which occurs with surgical procedures. [ 13 ] This phase is also the first step of the intrinsic clotting cascade involving a series of enzymatic response initiated by surface contact with a foreign material. [ 13 , 25 ] As stated in section 3.1.1 , the larger surface area provided by greater material porosity can lead to greater protein adhesion on the material surface. [ 163 ] However, clotting responses are also protein specific, as albumin is generally considered passivating and fibrinogen is thrombogenic. [ 25 ] Protein conformation, and the change of protein composition on the surface dictated by the Vroman effect, also impacts hemocompatibility over time. [ 24 , 152 ] The second mechanism which coagulation can also occur by is the extrinsic pathway, which is activated by the release of tissue factor (TF) from injured vascular endothelial cells (ECs). ECs line blood vessels and are likely to be damaged with an incision or implant placement. [ 184 ] The activation of TF may also arise from inflammatory mechanisms from leukocytes at the material surface or by complement activation. [ 24 , 181 , 185 ]
Medical devices that contact blood &#; such as dialyzers, drug delivery systems like nanoparticles, and cardiovascular devices including vascular grafts and stents &#; have unique considerations for biocompatibility. [ 181 ] Blood is composed of cells including erythrocytes, leukocytes, and platelets. [ 24 ] Materials that contact blood are therefore susceptible to inducing cell lysis, immune recognition (similarly described in sections 3.1.2 and 3.1.3 ), and especially thrombosis. No standards currently exist for assessing material hemocompatibility, nor is there a consensus on which materials are hemocompatible. Assessments of blood biocompatibility typically focus on thrombogenesis prevention, [ 152 , 182 ] but these assessments lack predictive value on biocompatibility. Excessive thrombosis can result in lethal thromboembolism or device occlusion, which can halt device function and downstream blood flow. [ 152 ] On the other hand, blood clotting on biomaterial surfaces can be beneficial to prevent hemorrhage and to act as a matrix for cell attachment .[ 152 ] In fact, adhered but non-activated platelets can be a natural passivating surface against thrombogenicity, so blood contacting materials like woven vascular grafts are commonly pre-clotted prior to implantation. [ 152 , 183 ] Thrombosis is also dependent on more than just surface interactions, but also specific blood chemistry and the mechanics of blood flow. [ 152 , 184 ]
Studies of larger pore-size materials have indicated that the benefit of increased pore size on pro-healing and angiogenic responses has an upper limit. Sussman, et al. found that porosity overall reduced fibrosis. However, materials with 34 μm pores were permeated by cells with minimal collagenous growth and a greater density of blood vessels, while 160 μm pores contained a greater fraction of fibrotic tissue ( ). [ 10 ] Similar pHEMA scaffolds have found that implants and pore features remain intact out to 28 days. [ 4 ] In a study by Bezuidenhout, et al., materials with pore sizes ranging from 63 to 180 μm showed no statistical difference in vascularization. [ 131 ] Overall, porosity and pore sizes approximately 30-40 μm in diameter significantly impact the acceptability of local tissue responses when studied in rodent models ( ). Interestingly, this pore size is near twice the size of relevant cells &#; with rodent macrophages and fibroblasts measuring approximately 13 and 18 μm, respectively. [ 179 , 180 ] A summary of the effects from pore size on fibrous capsule formation and angiogenesis is illustrated in .
A reduction in fibrous capsule size in response to porous materials was first observed by Karp, et al., where a difference in FBR was demarcated between commercial Millipore filters with pores 0.025-0.1 μm and 0.22-8.0 μm. In the materials with larger pores, a less developed fibrous capsule was observed with cells found inside the pores, more non-adherent macrophages between the implant and capsule, and many FBGCs. [ 15 ] Cells have been found to penetrate materials with pores as small as 0.8 μm, but an increase in pore size up to 9 μm further promoted vascularization. [ 16 ]
In contrast to fibrous capsule formation, the desired outcome for most implanted biomedical materials is to functionally integrate into the surrounding tissue environment. Beneficial wound healing responses promote the regeneration of local tissue cell types, and the ingrowth of new vasculature, known as angiogenesis, to support the transport of oxygen and nutrients into developing tissue. [ 1 , 5 , 13 , 178 ] Angiogenesis is especially important in biomaterial scaffolds for tissue engineering, where cells must functionally inoculate the material structure. Without supporting capillaries, oxygen and nutrients can penetrate a distance of approximately 150-200 μm. [ 5 ]
In the final stage of the FBR, collagen rich scar tissue growth, known as fibrous capsule, commonly develops around biomaterials. In response to factors such as TGF-β released from cells like M2 macrophages, fibroblasts migrate to the surface of the biomaterial and generate ECM proteins like collagen which forms the fibrous capsule. [ 1 , 13 ] Macrophages may also secrete other factors such as matrix metalloproteinases (MMPs), which degrade ECM proteins as needed for tissue remodeling at the implant site. [ 13 , 168 ] Biomaterials become encapsulated by this largely acellular tissue, which can act as a barrier in applications like drug delivery, sensors, or electrodes, and additionally can alter local tissue mechanics and therefore tissue function. [ 1 ] Further, fibroblasts may differentiate into myofibroblasts, which can contract the tissue surrounding the biomaterial, potentially causing patient pain and device damage. [ 1 , 13 ]
The discrepancy in the FBGC response towards porosity across these different studies could be confounded by additional factors which may influence the FBR. First, pore size is not reported for PLA meshes with high FBGC adhesion, and highly porous PU materials with low FBGC adhesion are not compared to a non-porous control. [ 131 , 149 ] Therefore, a window of high FBGC adhesion may exists for porous materials with small void diameters, as is observed for macrophage polarization. Second, FBGCs effect degradable PLA and bio-stable PU differently. While FBGCs can induce PU cracking, the scaffold is not degraded like PLA, which will change in structure as the material is eroded. Further, the resulting release of lactic acid from degraded PLA is known to promote inflammation. [ 131 , 149 ] However, reported FBGC responses from other in vitro studies of PLLA contradict what is observed in vivo for PLA electrospun materials, although differences in chirality and resulting changes in polymer crystallinity may confound this comparison. [ 148 , 149 ] Finally, the difference in implant site may also contribute to FBGC responses. Indeed, greater inflammatory responses have been observed towards materials implanted intramuscularly, as compared to materials implanted subcutaneously. [ 17 ] Thus, the effects of porosity on FBGCs, and the effects of FBGCs on the FBR, are not simply positive or negative for wound healing. Future assessments to clarify the role of porosity on FBGCs and the resulting FBR should consider additional factors like material and local tissue environment.
In comparison to macrophage polarization, the relationship between pore size and FBGC formation is less consistent in the literature. Saino, et al. showed greater FBGC formation in vitro on PLLA non-porous films as opposed to electrospun nanofibrous and microfibrous materials. [ 148 ] Similarly, a study of highly porous (~85%) sphere templated polyurethane (PU) materials by Bezuidenhout, et al. found significantly reduced FBGC formation on materials with larger pore diameters (150-180 μm) when implanted subcutaneously in rats. [ 131 ] However, Lucke, et al. showed greater FBGC formation on porous electrospun PLA mesh surfaces than non-porous membranes, which were implanted intramuscularly in rats. [ 149 ] At early timepoints, the macrophage populations within and surrounding porous materials are consistent with the observations made by Sussman, et al. describing high M2 populations at the surface and high M1 populations within pores. However, after 56 days, the FBGC response was seen to be greater towards porous mesh materials, as compared to smooth, membrane implant controls. [ 10 , 149 ]
The fourth phase of the FBR is foreign body giant cell (FBGC) formation. Multi-nucleated FBGCs form by the fusion of macrophages at the biomaterial surface. [ 1 , 13 ] This is a result of frustrated phagocytosis, where adherent macrophages attempt and fail to phagocytose the large implant, and is thought to be a mechanism to avoid apoptosis. [ 1 , 13 , 174 ] Cytokines IL-4 and IL-13 are implicated in FBGC formation and are considered to be derived from M2 macrophages, but distinctively express both pro-inflammatory and anti-inflammatory cytokines ( ). [ 13 , 168 ] Further, pro-healing growth factors, such as VEGF, are secreted by FBGCs, [ 175 , 176 ] as well as reactive oxygen species (ROS) and enzymes which contribute to material degradation and device failure. [ 13 ] The function of FBGCs in either a pro-healing or destructive FBR is not clearly defined, although these cell types likely contribute to both pathways depending on material features and time. [ 149 , 175 , 176 ]
Bartneck, et al. also studied the effect of electrospun porosity on macrophage polarization, and inversely found 20 μm porous materials induced more M1 polarization and 100 μm materials induced greater M2 polarization as measured by 27E10+ and CD165+ surface markers, respectively. However, the cytokine profiles suggest the 20 μm materials initiate greater pro-angiogenic signaling, while greater pro-inflammatory expression is measured in response to 100 μm materials. [ 171 ] These results are still consistent with the consensus of work, as signaling cues have the largest downstream effect on the FBR, and illustrates the over simplification of the polarization model for macrophages. Moving to even larger pore size, Yin, et al. found increased M2 polarization and VEGF expression in response to materials with 360-μm pores as compared to 160-μm pores. [ 147 ] Combined with existing research, this suggests that a window of unfavorable immune responses may exist for pore sizes around 100 &#; 160 μm. In studies of meshes with pore diameters ranging between 460 μm and up to μm, more favorable responses have been seen with larger pores. [ 172 , 173 ] Pores larger than μm are said to prevent contact between filament associated inflammatory infiltrate, and therefore prevents the bridging of scar tissue. [ 173 ] However, the capability of having large scale pores is dependent on the three-dimensional thickness of the scaffold and the required mechanical properties. Overall, existing research suggests that despite difference in polymers, fabrication methods, and architecture, pore sizes approximately 30-40 μm in diameter yield greater pro-healing immune cell responses at the material surface, and have been further associated with greater vascularization, a reduced fibrous capsule, and greater tissue integration. [ 4 , 7 , 9 , 10 , 14 , 38 &#; 40 , 80 , 146 ]
Garg, et al. used electrospun polydioxanone materials with randomly conformed pore diameters averaging approximately 2, 22 and 30 μm, which increase with strut diameter. Macrophages seeded onto materials with larger pore and strut diameters resulted in greater M2 macrophage attachment ( ). Additionally, compressed materials with equivalent strut size, but reduced pore diameter, had reduced M2 attachment, indicating that the pore size is indeed the main contributor to macrophage polarization in these materials. [ 80 ] Using 3D printed polycaprolactone (PCL) box-shaped pores, Tylek, et al. illustrated that M2 differentiation is accompanied by macrophage elongation. Further, elongation is promoted with smaller, 40 μm pores (illustrated in ). [ 33 ]
The macrophage response to porous materials has been extensively studied. Sussman, et al. characterized macrophage phenotypes in response to sphere templated poly(hydroxyethyl methyacrylate) (pHEMA) materials with either 34 μm or 160 μm pores as well as non-porous pHEMA, as illustrated in . Inflammatory infiltrate was significant at the surface of non-porous implants but was minimal at the surface of either 34 or 160 μm porous materials. Within the fibrous capsule at the material interface, non-porous materials had M1 dominated responses, whereas porous materials had M2 dominated responses. Within the 34 μm pores, macrophages were more significantly M1 phenotypes, but were associated with greater vascularization and lower intrapore fibrosis. This distinction was not observed significantly in 160 μm-porous materials. [ 10 ] Sphere-templated pHEMA pores 100 μm in diameter have also been found to induce greater gene expression for inflammatory Th1 cells in comparison to 40 μm pores. [ 9 ]
These subtype delineations are known to be an oversimplification of macrophage behavior in vivo. [ 1 , 168 ] Additionally, absence or over-abundance of either pro-inflammatory and anti-inflammatory macrophages can lead to unfavorable FBR. [ 168 ] Beyond macrophages, cells of the adaptive immune system also play a role in the FBR. [ 9 , 169 , 170 ] Regulatory T-cells (Tregs) have been implicated in pro-healing responses. Th1 cells increase inflammation and inhibit collagen deposition, while Th2 cells induce fibrosis and inhibit inflammation. [ 9 , 169 ] Further, T-cells may act on macrophages and influence their polarization through the release of cytokines, or by antigen presentation. [ 169 ]
Chronic inflammation defines the third phase of the FBR, and persists for two to three weeks with notable local responses from macrophages. [ 13 ] Macrophages have been characterized as the primary arbiter of the FBR, and are activated by cytokine signaling and by adhesion to the provisional matrix. Activated macrophages are described to exists in two main polarization states: the pro-inflammatory M1 phenotype and the anti-inflammatory M2 phenotype (illustrated in ). [ 1 , 12 , 13 ] M1 macrophages kill pathogens, activate T-cells, degrade ECM proteins, are implicated in T helper 1 cell (Th1) responses to intracellular pathogens, and are associated with a FBR that has greater acute and chronic inflammation. M2 macrophages are further divided into M2a, M2b, M2c, and M2d phenotypes. All M2 subtypes are implicated in Th2 responses to extracellular pathogens, ECM synthesis, angiogenesis, and control over acute and chronic inflammation. [ 1 , 12 , 168 ] Although, M2b macrophages may also promote inflammation and M2d macrophages are implicated in wound healing as opposed to fibrosis. [ 168 ] includes known inducers of these phenotypes, and factors which are secreted from these cell types.
Following protein adhesion, the FBR enters the second phase known as acute inflammation. [ 1 , 13 , 162 ] Neutrophils are the first immune cells to infiltrate the material surface and act as the primary cell population during acute inflammation, along with mast cells, for a period of hours or less than a week. [ 13 , 162 ] Tissue damage from implantation, recognition of the provisionary matrix, recognition of the foreign body and/or bacterial infection may initiate neutrophil migration. [ 1 , 13 ] Studies have rarely focused on the biomaterial impact on acute inflammation, therefore not much is known about the effect of porosity or topography on this phase. [ 13 ] One study by van Tienen, et al. distinguished the one-week acute inflammatory response towards materials with porosities of 73% or 86% &#; both with pores 150-300 μm in diameter but varied by the throat diameter. A similar magnitude of neutrophil infiltration was observed towards both materials, despite differences in tissue ingrowth at later timepoints. [ 167 ] These studies focused on quantifying cell accumulation, but additional information on the cytokine release profile by neutrophils may better inform the factors that recruit, activate, and direct macrophage phenotypes. [ 13 ]
Beyond pore size, the protein adhesion is also dependent on the geometry of the material walls or struts. Woo, et al. demonstrated that porous poly(L-lactic acid) (PLLA) scaffolds with nanofibrous walled pores (diameter = 0.05-0.5 μm) absorbed 4.2 times the protein compared with materials with solid walled pores (diameter = 250-420 μm). Additionally, these nanofibrous materials specifically absorbed greater quantities of proteins implicated in cell attachment including albumin, fibronectin, vitronectin, and laminin. [ 166 ] Stochastic roughness, which may be created from materials with random pore morphology, has shown to increase absorbance of fibronectin, an ECM protein that promotes macrophage attachment. [ 162 , 163 ] Although informative to understand early mechanisms of the FBR, studies of protein adhesion alone are conducted in vitro. These studies therefore exclude competitive protein binding dynamics that will have determining effects on immune cell signaling and the final host response.
The magnitude of protein absorption is highly dependent on the type of material used, but is also increased by surface roughness. [ 13 ] Increased absorption is likely a factor of a greater surface area to volume ratio. [ 163 ] Greater porosity can therefore induce greater protein adhesion as well as the specificity of proteins that adhere to the surface. Jansson, et al., showed that titanium surfaces with pores between 0.2-0.3 μm absorbed two to eleven times more albumin and IgG than smooth titanium surfaces. While IgG may activate immune cells, albumin reduces platelet and neutrophil activation. [ 164 ] However, this effect is altered at the scale of protein-size, as was shown in a study by Richert, et al.. Nanoporous titanium surfaces with a mean pore diameter of 0.011 μm promoted fibrinogen, lysozyme, and human growth/differentiation factor-5 (GDF-5, osseogenic cell promoter) binding, and decreased absorbance of bovine serum albumin, IgG, fibronectin, and collagen as compared to non-porous surfaces. [ 165 ] The dimensions of IgG and fibronectin surpass this pore diameter, which likely explains this difference in absorption. [ 165 ]
The FBR can be characterized by five phases. [ 1 , 12 , 13 ] First, local tissue and blood proteins will adhere to the material surface within an order of seconds following implantation. In this first phase, a provisional matrix is formed around the implant, including clotting proteins like fibrinogen, fibronectin, and vitronectin, as well as opsonins such as proteins belonging to the classical complement system pathway and immunoglobulins (IgG). [ 12 , 13 ] This provisional matrix develops rapidly into a thrombus composed mainly of fibrin. [ 13 ] The composition of the provisional matrix changes dynamically over time following the Vroman Effect, as proteins absorb and desorb on the surface. [ 1 , 13 , 162 ] Bioactive agents in the provisional matrix &#; including cytokines, growth factors, and chemoattractants &#; can promote the activation, migration, proliferation, and polarization of immune cells and fibroblasts, thereby impacting the subsequent stages of the FBR. [ 162 ]
A modern definition of biocompatibility put forth by Crawford et al., is &#;the ability of a material to locally trigger and guide the proteins and cells of the host toward a non-fibrotic, vascularized reconstruction and functional tissue integration.&#; [ 1 ] The effect of material porosity on biocompatibility has been extensively studied, repeatedly showing that porosity improves tissue healing responses and decreases scar tissue growth. [ 4 , 10 , 13 , 33 , 38 , 39 , 80 , 142 &#; 144 ] Considering our tissue-specific focus here, we have broadened our definition of biocompatibility to consider the other factors which are critical to the success of the medical device. Beyond the FBR, this includes hemocompatibility, biofilm formation, and calcification &#; all of which have nuanced impacts by porosity. and includes a summary of porosity effects on the various aspects of biocompatibility discussed in this section.
As reiterated throughout this review, porosity can yield a more favorable host response, and many different methods exist to engineer biomaterials with precise porous architecture. Yet, medical devices are not always composed of porous materials. This is because various tissues have unique functions and necessitate different aspects of biocompatibility, which impacts the design requirments for materials attempting to achieve local homeostasis. Further, tissues have unique microenvironments with different local immune cell populations that contribute to known variations in tissue specific host repsonses.[17&#;19,194] In this section, we review the impact of porous devices within specific tissue syestesm such as the subcutis, skeletal system, cardiovasuclar system, eye and nervous system, as well as the female reproductive tract. summarizes these tissue-specific host reponses and functional impacts towards porous materials and shows some example surfaces relevent to the various tissues discussed. For details on tissue-specific biomaterial factors beyond porosity that are important for ECM-mimetic scaffolds, the reader is directed to an excellent review by Tonti, et al..[195]
The subcutis exists as an adipocyte rich layer of tissue just beneath the skin surface, and primarily serves the body for thermal regulation, energy storage, and for protection from injuries.[219] Due to the presence of blood capillaries,[220] lymphatic plexus,[219] and the minimally invasiveness of accessing this compartment, the subcutis is widely used as a model to study the biomaterial FBR.[6,8&#;10,15,18,126,221] As detailed above (Section 3.1), porous materials with pore sizes approximately 40 μm in diameter show low fibrotic encapsulation, higher vascularization, and pro-healing responses from local immune cells, such as M2 polarized macrophages ( ).[7,9,10,80,126,222]
The vascularization and peripheral location of the subcutis makes this tissue compartment ideal for long-acting drug delivery implants for applications such as contraception, treatment of schizophrenia, management of opioid addiction, and HIV prevention.[220,223&#;227] Subcutaneously implanted porous devices have also shown greater fibrotic encapsulation than devices implanted in the intraperitoneal space or the epididymal fat pad in mice.[194] Although porous materials demonstrate improved biocompatibility over dense, non-textured surfaces,[15] some subcutaneous drug delivery systems are intentionally designed with low porosity to reduce drug release rates, as described in Section 2.3.[129,228] Sustained, long-acting drug release is critical for implantable drug delivery systems to improve patient adherence compared to a frequent daily dosing schedule, and fibrous capsule formation can impede drug release.[1,8,220,229] Implant re-insertion is also not practical in a timeframe of approximately less than six months.[224] Biocompatibility is also essential for enabling patient compliance towards these drug delivery systems, as patients will discontinue use of elective devices that cause discomfort.
Contraceptive implants &#; such as Nexplanon and its precursor Implanon, or Jadelle and its precursor Norplant &#; are the most widely used intradermal systems.[224,225] Nexplanon is a single cylindrical implant (2×40 mm) made of an ethylene vinyl acetate (EVA) polymer core loaded with the contraceptive agent etonogestrel and a drug-free EVA rate-controlling membrane, which has an effective duration of three years.[224,225,230,231] Jadelle is composed of 2 rods (2.5×43 mm each) that release the contraceptive levonorgestrel for up to five years within a polymer core and silicone rate-controlling membrane.[224,225] Information concerning the porosity and surface topography of these implants is limited. However, similar hot-melt fabricated, drug-free EVA membranes exhibit microtextured but non-porous features.[232&#;234] Additionally, silicone implants have been described as smooth and solid, as opposed to porous.[235] Therefore, surface features of these implants lack microporous features which are associated with improved healing responses. Although adverse reactions have been reported,[236] such contraceptive implants have largely proven to be safe and tolerable.[237] In a study of tolerability of both levonorgestrel and etonogestrel subdermal implants, López del Cerro, et al. found one non-tolerable case of a FBR, out of 221 implants, which resulted in implant expulsion.[237] In another assessment of local side effects of Norplant by Alvarez, et al., 108 (35.6%) patients reported local hyperpigmentation of the skin, and 68 (22.4%) patients reported skin depression at the site of the implant due to a loss of subcutaneous tissue,[236] both which may be attributed to a FBR.[238]
A study of HIV preventative implants by Barrett, et al. indeed showed that dense implants created by hot-melt extrusion could achieve long-acting release of the investigational antiretroviral drug MK- beyond 6 months. Interestingly, the implants became more porous over time, developing from the implant surface eroding into the core, with pores being created and increasing to ~2 μm wide with random geometry after full drug release. This possible porogen effect of drug particles within materials traditionally considered non-porous has been noted in other subcutaneous implant studies as well ( ).[239] The mechanism of drug release was therefore said to be mediated by solution permeation through voids evacuated by solubilized drug. The study does not include an assessment of the local tissue response.[129] In fact, another HIV implant study by Su, et al. describes overall lack of FBR assessment across other HIV implant studies, and reports an unacceptably adverse inflammatory reaction towards their tenofovir alafenamide fumarate loaded polyurethane membrane implant, despite lower inflammation observed against placebo implants.[240] Pharmaceuticals can contribute to adverse host responses, and therefore specific material and drug combinations must be assessed for their biocompatibility. Although porosity is not the sole arbiter of biocompatibility, material microarchitecture has shown to have a robust effect on resulting tissue outcomes and has been extensively studied in the subcutis. Therefore, porosity &#; and the balance of its effects on pharmacokinetics and biocompatibility &#; should be considered in the development of safe, long-acting devices.
Bone interfacing implants such as prosthetic hips, knees, plates, pins or nails for fixation, bone cement, and oral implants represent some of the most common devices and biomaterials clinically implemented.[21,160] Such implants serve biomechanical functions, and therefore a mechanically stable, osseointegrative host response is needed for long-term biocompatibility.[192] The structure of bone tissue is either cortical (compact) or cancellous (spongy or trabecular, ).[192,241] Compact bone comprises the hard outer surface of bone and is porous, yet dense with 3-12% porosity and pores 100-200 μm in diameter.[192,242] Osteons, or haversian systems, are the microscale structural unit of compact bone and are approximately 150-250 μm in diameter.[241,243] Cancellous tissue inside bone constitutes most of the bone tissue and has 50-90% porosity with irregularly patterned trabeculae structures and bone marrow within the voids, which can measure up to 1 mm in diameter ( ).[241,242,244]
Bone tissue is continuously remodeling and is composed of blood vessels, cells, interstitial fluid, collagen fibers, and the mineral hydroxyapatite.[192,241,243] Bone building cells start as osteogenic stem cells, and differentiate into osteoblasts which synthesize the ECM components of bone, and range in diameter between 20-50 μm.[241,243] Osteoblasts mature into tissue maintaining osteocytes, and are embedded in bone lacuna measuring 15-20 μm in diameter, which have small radiating channels known as canaliculi filled with extracellular fluid and osteocyte processes for exchange of nutrients and wastes.[241,245] Conversely, osteoclasts enzymatically degrade bone for resorption as a means of continuing bone maintenance, and are derived from the fusion of many monocytes which results in a cell diameter ranging between 10-300 μm.[241,246] Thus, biomaterials that mimic the architecture of native bone and promote bone cell proliferation and integration could be ideal for biocompatible orthopedic devices ( ).
The primary pathway of bone regeneration and material osseointegration is similar to the traditional FBR (Section 3.1) but differs by cell specific responses. First, blood clotting occurs at the site of bone loss and the implant. In this first phase known as osteoconduction, the hematoma and specific activated platelets drive host reactions through the release of growth factors like TGF-β, IL-6, VEGF, fibroblast growth factor (FGF), and insulin growth factor (IGF) that recruit osteocytes and influence osteoblast action.[66,192,241,247] The callus then forms from collagen and cartilage to bridge fractured bone tissues and scaffolds. Through hormone signaling, the callus mineralizes requiring mainly calcium and phosphorous, but also magnesium, fluoride, and manganese.[241,243] Vitamins A, C, D, K, and B12 also contribute to osteoblast activity, bone protein synthesis, or calcium uptake.[241] The mineralized callus is then ideally remodeled into mature bone tissue.[192,241,247] Although bone tissue is mineralized, bone is also highly vascularized.[241] Therefore vascularization is essential for healthy scaffold-bone tissue integration.[248]
Due to the common use of orthopedic devices and the natively porous structure of bone, extensive research has been conducted with porous materials for osseointegration.[5,58,66,192,198&#;202,249&#;252] Unlike the subcutis, which has shown optimal host responses with approximately 40 μm pores ( ),[10,33,80] the critical pore size for bone tissue is suggested to be larger, yet a consensus on the specific size remains debated in the literature ( ). The necessity of porosity for osteogenesis has been well described in a review by Karageorgiou and Kaplan.[75] Various studies and reviews have defined 100 μm as the minimum required pore size for bone integration,[5,58,192,198&#;201] while other studies specify pores greater than 300 μm are optimal for bone growth.[196,197,203] However, other studies have stated that smaller pore sizes may be sufficient[5,66,202,249&#;251] Specifically, Hulbert, et al. found that pores greater than 100 μm promoted the greatest mineralized bone growth, and pores greater than 150 μm facilitated osteon formation within calcium aluminate implants placed midshaft of dog femurs.[199] However, Itälä, et al. claimed new bone ingrowth is independent of pore size within a range of 50-125 μm, which was assessed in titanium implants placed in non-load bearing regions of rabbit femurs. The authors also point out their observed response may be dependent on weight bearing conditions.[202]
Despite the known benefit of open pore structures, surgeons have been said to prefer the handling and stability of solid biomaterials.[242] Indeed, the addition of porosity can come at the sacrifice of mechanical stability, which is vital for orthopedic implants which often function for providing loadbearing support or mechanical fixation.[58,199] Clemow, et al. assessed porous coated titanium implants with pore sizes ranging between 175-350 μm and found percent bone ingrowth to be proportional to shear strength of the implant at the interface within the femoral medullary canal of dogs. Bone ingrowth was therefore also inversely proportional to the square root of pore diameter. This suggests that larger pore size is detrimental to bone ingrowth if the specific material choice has inadequate structural support, although all pore sizes assessed here were greater than 100 μm.[253] To mitigate the detrimental effects of porosity on material mechanics, modern approaches have focused on developing bone scaffolds with biomimetic architecture. One strategy to better mimic bone structure used polymer scaffolds with a porosity gradient, which captures both the mechanical integrity of dense cortical bone, and the higher porosity of cancellous bone, which enables tissue ingrowth.[58] Finite element modeling has also been used to elucidate how the specific design of pore microgeometry can effect parameters like mechanics, permeability, and surface area.[254] Future research into bone scaffold design therefore must consider material composition and microgeometry for stable osseointegration.
As is true for all implants, biofilm formation is also a concern for orthopedic implants.[20,160] Biofilm formation is an especially important consideration for dental scaffolds, considering the exposure to bacteria from the oral mucosa.[20,255] Osseointegration is important for dental implants as artificial teeth are often implanted into the jaw bone and anchored into place by healthy bone ingrowth.[247,255] Inflammation caused by dental implant biofilm, known as peri-implantitis, is known to contribute to bone tissue loss.[255] For these reasons, emerging research into bone integrative scaffolds show increased focus on biofilm prevention,[255] with strategies such as combined antibacterial nanoparticles[256] or probiotic biofilms, improving osseointegration.[257] In addition to material functionalization methods, it should be noted that implant osseointegration also reduces the available surface for biofilm infection, known as the &#;race for the surface&#; described in section 3.3.[79,191,255] Initial control over biofilm formation is needed to mitigate potential bone loss, which would better promote osseointegration, and in turn inhibit future infections. Future research is needed to further investigate effective prevention methods for biofilms, and to study the competing effect of large pore size on both biofilm formation and osseointegration.
The cardiovascular system consists of the heart, the connecting blood vessels, and blood.[241] The heart and vessels are each composed of three tissue layers.[258] Blood contacts the heart endocardium or vessel tunica intima, which are continuous with each other, and are both composed of the endothelium and the basement membrane which these endothelial cells attach. Cardiac or smooth muscle cells compose the inner layers, known as the myocardium and tunica media in the heart and vessels, respectively. The external layers include the epicardium and the tunica adventitia which both consist of fibro-elastic connective tissue.[241,258] Biomaterial interactions between these tissue layers as well as blood (for blood tissue composition, see section 3.2) make cardiovascular devices a unique challenge for biocompatibility. These complications include: thrombosis and device occlusion by clot (section 3.2); sepsis from biofilm formation (section 3.3); material calcification which deteriorates mechanical function (section 3.4); as well as fibrosis and stenosis caused by excessive tissue growth known as neointima hyperplasia.[150] Unlike bone tissue, where responses like calcification promote material biocompatibility, all these pathologies in the cardiovascular system can yield fatal responses. Porosity can increase tissue integration and decrease inflammatory responses ( ), but porosity can also promote platelet activation ( ) and calcification ( ). Here, we focus on biomaterials that are porous and intended for cardiovascular tissue integration.
The FBR occurs in cardiovascular tissue and has been studied especially in the heart. However, the resulting response also been observed to vary somewhat from what occurs in the subcutis. Luttikhuizen, et al. characterized a greater pro-inflammatory response towards supra-epicardially implanted collagen, as compared to subcutaneous implants. Further, greater MMP activity, especially MMP-9, was found within super-epicardial tissue, which resulted in greater degradation of these implants.[19] Despite this, the pore-size dependent FBR is comparable between cardiovascular and subcutaneous tissues. In a study by Madden, et al., sphere-templated, poly(2-hydroxyethyl methacrylate-co-methacrylic acid) (pHEMA-co-MAA) scaffolds with pore diameters 30-40 μm resulted in lower fibrous encapsulation and greater neovascularization, as opposed to non-porous, 20, 60, or 80 μm porous materials.[50] This response is similar to what was observed subcutaneously towards 34 μm pore diameter pHEMA implants by Sussman, et al. (see Sections 3.1.3 and 3.1.5).[10] Porous materials with controlled pore size therefore show positive implications when in contact with the cardiovascular tissue microenvironment, although hemocompatibility is not specifically addressed in these studies.
Appropriate cell integration within scaffolds also has implications for hemocompatibility. At the interface of blood and tissue, ECs act as an inert barrier against protein adhesion, platelet activation and consequently thrombosis.[181,205] ECs also change morphology in response to shear forces by elongating and aligning in the direction of blood flow. Cell alignment inhibits inflammatory responses in regions under high laminar shear by altering the interactions of flow forces on the ECM.[259] Creating biomaterial surfaces that can support endothelial cell growth is therefore an effective strategy for promoting hemocompatibility. As described in section 3.1.5, porosity contributes to microvascular EC migration and proliferation during angiogenesis. Studies of macrovascular endothelium growth in vitro generally show that lower porosity and smaller pore size materials facilitate greater EC attachment.[187,260&#;262] ECs are 10 to 40 μm in diameter and their growth is adherent-dependent. Proliferation is therefore limited on materials with pores that exceed cell size, specifically greater than 30 μm in diameter.[262,263] It should be noted that these monoculture models fail to capture the proliferative effect that local cell types like fibroblasts and the intimal ECM has on ECs. Pre-seeding porous scaffolds with fibroblasts has shown improved endothelialization and increased expression of growth factors like FGF, which is otherwise supplemented in EC growth media.[261,262,264&#;267] Therefore the effect of porosity on endothelialization should be considered using in vivo models, where the combined effect from the improved capacity for tissue infiltration can also be assessed.
While porosity is implicated in greater tissue integrative responses, neointima hyperplasia is known as an unfavorable response in the cardiovascular system.[206] In cardiovascular devices like stents, which open or support blood vessels in cases of coronary artery disease or aneurysm, stenosis by neointima hyperplasia is one of the main challenges for biocompatibility.[268,269] Stenosis can also occur at graft anastomosis.[270] Stents are often composed of porous metal meshes, and lower stent porosity is implicated with greater neointimal growth and less flexibility for device placement.[269] Wesolowski, et al. described the effect of porosity for vascular grafts, finding that increased porosity rather than material composition, had the greatest impact in reducing graft stenosis, which lead to tissue calcification in pigs, dogs, and even human.[132] Thrombosis may also arise from stenosis to altered hemodynamics and flow stasis.[271] Therefore, tissue ingrowth should be controlled not only to prevent occlusion, but also secondary responses such as thrombosis and calcification. The mechanism for neointimal hyperplasia is thought to arise in part by inflammatory mechanisms including pro-inflammatory cytokine signaling from macrophages.[270] Further, a dysfunctional or disrupted endothelium may enable neointimal growth.[268,270] Drug eluting stents used clinically effectively curb stenosis, and rely on agents which inhibit smooth muscle cell growth.[268] However, further improvements on cardiovascular host responses, including reduced stenosis, will require an environment which supports healthy EC growth.[268]
The effect of porosity on thrombus formation, endothelial cell growth, and neointimal hyperplasia in vivo has been extensively studied with ePTFE vascular constructs, as ePTFE grafts greater than 6 mm in diameter have shown great clinical outcomes.[205,207,272] For ePTFE, pore size is commonly reported as the internodal distance, as the microarchitecture is uniquely composed of nodes of material interconnected by thin polymer fibrils ( ).[142,205&#;207] Internodal distance and pore size is therefore also the same as fibril length.[206] Synthetic, small diameter vascular grafts with inner diameters less than 4 mm historically perform poorly in humans, with thrombus generation on the luminal graft surface being the most common causes of failure.[152,272] Biocompatible, small-diameter vascular grafts therefore remain an unmet clinical need, and a complex challenge for controlling host responses.
The unique surface topology of ePTFE has shown important impacts on cellular responses. However, the response to pore size in ePTFE grafts was also dependent on specific study methods and graft preparation. Campbell, et al. screened untreated ePTFE materials with internodal pore diameters ranging between 9 and 65 μm, and found increased pore size decreased the rate of patency of small diameter grafts tested in carotid or femoral arteries of dogs. Specifically, pores less than 22 μm resulted in greater rates of grafts with healthy tissue ingrowth, angiogenesis, and a thin neointima as opposed to grafts greater 34 μm.[206] Boyd, et al. found pre-seeding small diameter ePTFE grafts with ECs indeed increased the thrombus-free surface area as opposed to non-cellular ePTFE implanted in dog carotid arteries. Further, pre-clotted scaffolds with 40 μm internodal pore diameters had the greatest patency and largest thrombus-free surface with or without endothelial cell pre-treatments compared to grafts with either 28 or 52 μm pores.[205] Pore diameters are quantified prior to pre-clotting, so the resulting porosity of the treated scaffold is not reported. Porous, small diameter ePTFE grafts were also implanted in circulation and studied in baboons by Golden, et al., which is the most relevant model for human hematology.[152,204,207] Here, implants were pre-clotted, and animals were treated with anticoagulants. Grafts with internodal pore sizes 60 and 90 μm enabled full integration of the endothelium across the graft length, while grafts with 10 or 30 μm pores failed to gain full luminal endothelium coverage. Interestingly, pore size also effected the mechanism of cellular ingrowth, as 10 and 30 μm porous grafts allowed for cell migration from the ends of the vessel whereas 60 and 90 μm porous grafts enabled transmural tissue ingrowth. However, at 3 months, 90 μm porous grafts also caused degradation of the intimal layer of the vessel and platelet accumulation, potentially caused by the lower mechanical integrity of the more porous material or by shearing effects.[207] While these studies highlight the various ways pore-size specificity of ePTFE can enable greater cardiovascular biocompatibility, the determined optimal pore sizes across these studies lack congruency.
In ePTFE and other vascular graft materials, porosity has been identified as a positive material feature which provides needed elasticity and potential tissue ingrowth to sustain long-term tissue stability. However, if porosity or pore size increases above some threshold, the increased tissue ingrowth leads to low patency, increased thrombosis, and poor biocompatibility. In a review of the effect of porosity on various vascular protheses, White proposes materials with 45 μm pore sizes may be optimal for maintaining viable tissue growth, avoiding fibrosis, and providing mechanical compliancy.[204] In a recent clinical trial, Drews, et al. studied autologous bone marrow-derived cells seeded onto large diameter vascular grafts made of fibrous poly(glycolic acid) and poly(caprolactone-co-lactide) (PCLA) with a comparable average pore diameter of 40 μm. Although grafts experienced high incidence of asymptomatic stenosis in humans within 6 months, simulations and ovine studies found this stenosis self-resolves and could be a part of the natural vessel healing process.[208] While this study further supports the use of 40 μm pores for supporting biocompatibility, this work also illustrates the need to better understand the pathology of biomaterial neovascularization in humans.
Other strategies have focused on designing grafts with materials other than ePTFE. One notable porous graft intended for hemodialysis is the sphere-templated silicone STARgraft developed by Healionics (Seattle, WA), which is currently being investigated in a clinical trial (NCT).[44] Grafts with differing regions of porosity have also been recently investigated to mimic the layers of a native vessel. Wang, et al. developed a three-layered porous polyurethane graft using salt-leaching to mimic blood vessel anatomy.[273] Matsuzaki, et al. studied two layered grafts composed of a heparin PCLA co-polymer sponge surrounded by an electrospun PCL layer with pore sizes ranging from 4 to 15 μm.[89] In this study, larger pores did indeed allow for greater cell infiltration than 4 μm porous materials, but materials with pores greater than 4 μm also dilated under arterial blood pressure.[89] In both studies, poor mechanical matching resulted in the overall failure of both materials when studied in ovine models.[89,273] These studies illustrate that porosity is not the exclusive mediator of cardiovascular biocompatibility. Rather, mechanics of the graft &#; and the specific materials and fabrication methods that dictate these properties &#; are important design requirements. In another two-layered graft, de Valence, et al. assessed the effect of low and high porosity in either the luminal or external layers. Electrospun grafts with an approximate 11.3 μm pore size (81% porosity) in the luminal layer and 2 μm pore size (62% porosity) in the external layer yielded the most favorable responses placed in the abdominal aortas of rats, with cellular ingrowth in the adventitia and without blood leakage ( ).[91] In addition to polymer selection and fabrication method, control over regional microgeometry may enable positive tissue integrative effects of porosity in cardiovascular materials, while mitigating the host responses including hemocompatibility and excessive tissue ingrowth.
Tissues of the brain and eye &#; specifically the anterior chamber and cornea &#; are immunologically unique from other tissues and are described to exist under a state of immune privilege.[274,275] The cells which compose these tissues have a limited capacity for regeneration. Therefore under this immune privileged state, inflammatory responses are dampened to prevent cell death.[274] Although brain and eye tissues are similar in this way, anatomical differences between the two tissues regulate this tolerogenic state through different mechanisms. For both tissues, this privileged state is maintained in part by the blood-brain barrier (BBB) or the blood-ocular barrier, which restricts the migration of inflammatory cells into these tissues.[276] The BBB also limits the passage of large proteins like cytokines and antibodies.[277]
Cells present within the eye and brain also dictate host responses in these unique tissue microenvironments. Like the subcutis, the eye cornea and sclera are rich in fibroblasts and collagen fibers. The cornea, however, is also lined with nonkeratinized stratified squamous epithelium on the outer surface and simple squamous epithelium at the interface of the anterior chamber.[241] Cell populations within the brain are especially unique. The primary functional cell type of the brain is the neuron, which responds to stimuli and transmits electrical signals across other neurons in the circuit.[241] Glial cells support neurons and include astrocytes and microglia, which are considered the main immune effector cells in the brain and are also the most relevant cell-types in the host response.[278] Similar to peripheral macrophages, microglia are phagocytes and secrete cytokines that will influence astrocyte responses.[241,277] Astrocytes are the most common glial cell sub-type and are known to have functions such as tissue support and maintenance of the BBB. In a process known as gliosis, astrocytes can also encapsulate foreign bodies or regions of tissue damage.[241,278,279] Gliosis is analogous to the traditional FBR, where microglia &#; along with macrophages &#; first adhere to the foreign material and signal for the activation of encapsulating astrocytes, a response which is comparable to fibroblasts in the fibrous capsule.[279] Indeed, biomaterial gliosis results in a thin surrounding layer of collagen and astrocyte feet known as the glial scar.[277,279] From an evolutionary standpoint, given the reduced capacity to fight foreign materials with inflammatory responses, &#;walling off&#; the material becomes the host&#;s best defense for protecting the surrounding tissue.
Immune cells like Tregs also influence the tolerogenic state of the eye, as these cells mount tolerogenic responses towards foreign antigens, as opposed to immunogenic.[275] In the anterior chamber of the eye, Tregs are also associated with the production and surface association of the anti-inflammatory cytokine TGF-β, both which enable immune suppression.[275,280] Tregs exist in small quantities in healthy brain tissues, but Treg accumulation in brain tissues is also associated with protective responses against gliosis.[281] In the eye, anti-inflammatory soluble factors like cytokines are a necessary component of the immune tolerant state.[274,282] Cytokines like TGF-β, macrophage migration inhibitory factor (MIF), and IL-10 are found within the aqueous humor of the eye, and are known to suppress innate and adaptive immune responses like the complement system and natural killer (NK) cell cytolysis.[274,275,282&#;284] There still exists a need to better elucidate the role of these soluble, immunosuppressive factors in brain tissue.[274]
Although the eye and brain are described to be immune tolerant, biomaterial implants still undergo encapsulation either by fibrosis or glial scarring.[46,277&#;279,285,286] The reduced capacity to fight pathogens may also imply an increased risk of implant associated biofilm formation. Bacterial growth on intraocular lenses during cataract surgery can occur especially by the adhesion of bacteria from the conjunctival flora, thereby causing endophthalmitis.[213,214] Biofilms in the brain have been described to be a concern for neural probe signal impedance,[286] but otherwise there exists little research on this topic. Similarly, calcification towards biomaterials in the eye have been studied,[287] and although neural tissue calcification occurs under pathological states,[288] biomaterial calcification in the brain has not been described.
Various ophthalmic devices exist and are routinely used or implanted in patients, including but not limited to intraocular lenses (IOLs), contact lenses, glaucoma drainage devices, and orbital prostheses. Porous materials composing these devices have shown to provide greater mechanical flexibility, fluid flow, oxygen permeation, and cellular integration ( ).[46,48,212] Ophthalmic devices like IOLs do require a material structure with optical transparency. However, macroporous features are observed in some hydrogel contact lenses, and is thought to improve comfort by improving gas and water permeability to the cornea.[212] One promising application of porous materials in the eye is as a supracilliary drainage system to reduce intraocular pressure in uncontrolled glaucoma ( ). The iSTAR Medical (Wavre, Belgium) MINIject implant is made of flexible silicon with 27 μm sphere-templated pores. The flexible and porous structure of the MINIject has shown to conform to the shape of the eye, allow natural outflow rates of fluid through the pores, and reduces fibrosis. Glaucoma surgeries have a high risk of fibrosis, which causes device failure, and therefore must be mitigated.[45,46,48,289] In a recent clinical trial, the MINIject implant was found to effectively reduce intraocular pressure and reduce the need for medications across the two year study, without any serious adverse events.[46]
As described in Section 3, porosity can increase the risk of biofilm formation and calcification, which is also true for ophthalmic implants. Antibiotic loaded scleral bandages with 38 μm diameter pores have been proposed to mitigate both infection and fibrosis.[49] In an in vitro assessment of various orbital protheses, Toribio, et al. observed greater bacterial attachment on high density porous polyethylene implants with 100-500 μm pores, as compared to than non-porous silicone. However, the effects by material chemistry were not separately assessed from porosity, and despite any increased risk of infection, porous orbital implants are most commonly used.[214] Vijayasekaran, et al. also found samples of pHEMA sponges with pores on the order of 10 μm in diameter could calcify after 12-weeks implanted in a rabbit cornea model. Although this response was not compared to smooth surfaces or pores with different sizes, this calcification was thought to arise by the penetration of physiologic fluids in the eye and growing tissue into the scaffold.[287] While porosity can be beneficial for ocular biocompatibility, devices in the eye can serve a wide variety of functions. The selection of smooth or porous topography therefore must consider the specific region, application, and risk.
For brain biocompatibility, gliosis has been presented as the main challenge.[277] Neural probes are the device with the most research and interest for brain biocompatibility, as the glial scar impedes the capacity to receive or transmit signals to neurons.[277&#;279,285,286] Long term access to the neural circuit is needed for brain computer interface applications, which hold the potential for restoring motor function in cases like paralysis or limb loss.[279] One factor which contributes to the magnitude of glial scarring is the mechanical mismatch between the probe material and brain tissue.[143,278,279,290] Porosity is one strategy which can improve material compliance.[48] Topography has also proven to impact gliosis, as nanoscale features have shown greater selectivity of neuronal coverage, rather than astrocytes, in vitro.[210,291] The increased surface area of textured and porous implants has even shown to reduce the electrical impedance of microelectrodes.[209,210] Bioactive, cell-seeded electrodes have also been promoted as a strategy for improving tissue responses.[292] Porous scaffolds can additionally permit such bioactivity, and has the potential to reduce fibrosis, as has been observed across other tissue microenvironments.[10,48,50] Indeed, lower glial encapsulation has been observed in early assessments of materials with pores 40 μm in diameter as compared to 100 μm or non-porous implants.[211] Porous materials therefore show great potential for improved host responses within brain tissue. More research is still needed to study such devices, as well as the mechanisms behind host responses in the brain.
Like the eye and brain, both the male and female reproductive systems are immune privileged. For both these reproductive systems, this privileged state is necessary to prevent immune recognition of haploid germ cells as non-self.[274,293,294] The female reproductive tract (FRT) is especially interesting regarding the FBR, as tolerance towards paternal haploid cells, and embryos is required for fertility, and balanced tolerance towards the commensal vaginal microbiota is essential for general wellbeing and protection against disease.[294&#;296] Tregs are implicated as the primary mediator of this privileged state, especially during early pregnancy.[294,295] Ovarian steroids like estrogen and progesterone are known to regulate immune cell responses either directly or indirectly through cytokine signaling, and immune cell populations change temporally during the menstrual cycle.[294] Other factors such as MMPs are also known to change in response to the menstrual cycle and pregnancy, and contribute to the immune state of the FRT.[297,298]
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The FRT is also interesting in the study of the FBR, as it is a common compartment for device placement &#; including tampons, diaphragms, and vaginal rings for drug delivery in the vaginal cavity, and contraceptive intrauterine devices (IUDs) in the uterus. Immune tolerance is observed towards medical implants like IUDs, which do not undergo fibrotic encapsulation, yet do induce inflammation as one mechanism of contraception.[299&#;302] However, intrauterine fibrosis can arise, and is especially studied in cases of fallopian tubal occlusion. The occurrence of tubal occlusion has been observed to occur under acute stimuli such as material and mechanical stress from permanent contraceptive devices such as Essure,[303] large and repeated administration of chemical sclerosing agents,[304,305] or by chronic chlamydia tachomatis infection.[304,306] Considering that tissue integration is uncommon in the FRT and even detrimental to tissue function, as well as the bacteria tolerant nature of the FRT, biofilm formation is perhaps the most significant risk against material biocompatibility. Device biofilm formation can promote pathogen proliferation in the FRT, and potentially inhibit drug release for devices like contraceptive vaginal rings.[23] Thus, porous materials can cause negative health effects when resident in the FRT, and non-porous materials are commonly favorable for preventing the integration of bacteria into devices, especially IUDs and contraceptive rings ( ).[234]
Tampon use is a known risk factor for menstrual toxic-shock syndrome (mTSS).[215,216] Common use of extra-absorbent tampons starting in created a new niche for the common bacteria Staphylococcus aureus to interact with the human host environment, thereby contributing to the - epidemic of mTSS.[215,216] However, mTSS can arise without device use,[216] and reports of pessary use to treat uterine prolapse as early as the 19th century also describe possible cases of mTSS.[215] As a result of the porous structure which enables device absorbency, it is thought that a tampon with absorbed menses increases oxygen content within the typically anaerobic vaginal environment. This can allow for S. aureus growth, which releases TSS toxin 1 (TSST-1).[216,307] TSST-1 binds to vaginal epithelial cells, which induces chemokine signaling to macrophages and CD4+ T-cells. Consequently, these immune cells release a cytokine storm which manifests as mTSS &#; causing fever, hypotension, gastrointestinal effects, and/or alterations in consciousness.[216] However, incidence of mTSS with tampon use is rare. Efforts to reduce mTSS cases have included a standardization for labeling device absorbency, as well as recommendations for reduced wear time and use of the lowest needed absorbency rating.[216]
The IUD Dalkon shield is another interesting case of the possible detrimental effects of porosity in the FRT. The Dalkon shield was clinically available from until and varied from other IUDs in its shield-like shape of the device body. Most importantly, the Dalkon Shield also uniquely had a porous, multifilament removal string, or tail, which was needed as the larger body of the device required greater force to remove ( ).[217,218,308,309] Unlike other IUDs, which feature a monofilament removal string, the tail was found to wick bacterial species from the vaginal cavity into the uterus by capillary force.[217,218] Further, the small pore size between filaments excluded the passage of polymorphonuclear leukocytes through the IUD tail, so bacteria within the string remained protected against host immune responses.[218] In the United States, 11 deaths from generalized sepsis and 209 cases of septic spontaneous abortion were reported for women using the Dalkon shield, as well as a fivefold increase in pelvic inflammatory disorder cases as compared to other IUDs of the time.[217,309] Despite setbacks from this device, modern IUDs are safe and are the most effective contraceptive method available, with less than a 1% failure rate.[309]
Regardless of the history concerning porosity related health issues in the FRT, research on porous devices have been studied to address various FRT tissue disorders. Asherman syndrome is one uterine tissue disorder which causes abnormal endometrial tissue growth which causes the lining of the uterus to adhere. One strategy to mitigate this tissue growth is the placement of a biomaterial barrier within the uterus, such as an IUD.[310&#;312] To improve local coverage, a compressible porous scaffold was developed by Cai, et al. and studied in an intrauterine adhesion rat model. Scaffolds enabled the delivery of basic fibroblast growth factor (bFGF), which together led to statistically equivalent endometrium and gland growth as undamaged uterine tissue, and greater neovascularization. However, the porous scaffold alone initiated low endothelium and gland growth, and higher fibrosis &#; all comparable to the intrauterine adhesion group.[311] The added growth factor therefore appears to be the main contributor to healthy endometrial regeneration, rather than the porous scaffold.
Surgical mesh is another example of porous material commonly used to provide tissue reinforcement. In the case of pelvic organ prolapse, where pelvic organs herniate into the vagina, reinforcement with transvaginal mesh was previously a solution. However, due to mesh erosion and material exposure from chronic inflammation and poor tissue integration, the use of transvaginal mesh &#; but not intra-abdominal mesh &#; is now banned countries like the United States, United Kingdom, and Australia.[22,313] Mukherjee, et al. proposes the use of ECM-mimicking electrospun meshes seeded with endometrial mesenchymal stem cells (eMSCs) to improve vaginal tissue integration. Although electrospun meshes with eMSCs promisingly increase gene expression associated with angiogenesis, cell adhesion, and ECM regulation, these implant studies were conducted subcutaneously due to limitations of mouse FRT size.[313] Therefore, these results are not truly indicative of the FRT response. Accurate models of the human FRT are limited, but the specific microenvironment is necessary to capture this nonstandard host response which has been frequently misunderstood in the past.
As shown in Table 2 , polyester (mainly polyethylene terephthalate&#;PET&#;but also polybutylene terephthalate&#;PBT&#;and polyethylene naphthalate&#;PEN) is the predominant fiber used in manufacturing of automotive textiles. It has a share of 42%, whereas polyamide 6.6 (PA 6.6) has a share of 26%. These fibers are preferred because of their good physical properties and high mechanical performances, being dyeable as well as inexpensive [ 1 ].This chapter deals with the use of polyester in automotive applications by discussing the basic properties and performance aspects of the fiber. It presents the recent studies about the production of textile-based automobile components made of polyester and recycled polyester fibers. In addition, recyclability of automotive components made of polyester fiber as well as the usage of polyester in the production of natural fiber composites enabling reuse of waste materials is indicated.
In the automobiles, fibers are used in different forms of textile materials such as circular knitted, warp knitted, woven and nonwoven structures. Fibers are also used as a component in multi-layer composite structures. Textile products used in automobiles are expected to fulfill different performance requirements regarding the application area. These requirements are generally fulfilled by using man-made fibers as summarized in Table 2 [ 3 ].
The amount of fiber usage in a standard passenger car is sizable, reaching around 25 kg; even the safety and comfort requirement can increase this amount [ 1 ]. In the automotive industry, fibers are used in the manufacturing of textile products, which are given and described in Table 1 [ 2 ].
Moreover, due to the increasing environmental awareness to protect resources, reuse and recycle of the products have gained importance in the manufacturing of automobile components. Polyester fiber can be used in the recycled form in car upholstery, especially in the automobile carpets and seat covers [ 11 , 12 ]. Polyester nonwovens can be recycled to produce new materials [ 13 ]. Moreover, polyester resin can be used as a matrix material for especially the natural fiber composites, enabling the reuse of waste materials used in the pre-assembled parts of the automobiles such as headliner, bootliners and parcel shelves [ 14 , 15 ]. Thus, polyester can be considered as a select fiber for recycling and sustainability.
In addition to mechanical performance properties, PET has good sound insulation property, which is important for carpets and pre-assembled interior components such as headliners, bootliners, parcel shelves and door panels because of the requirements for increased driving and traveling comfort. PET fiber has a potential for use to increase sound transmission loss within a wide range of frequency (i.e. 100&#;500 Hz) similar to that achieved by using fiberglass [ 8 ]. Besides, nonwovens produced with a high percentage of hollow PET fibers (e.g. 45&#;50%) in the blends with PET fiber has demonstrated higher sound absorption rate when compared to the samples produced from 100% PET fibers [ 9 , 10 ].
Wool fiber provides a high level of thermal comfort owing to its high moisture absorption ability. Together with its high level of resiliency, these features make wool fiber attractive and appropriate especially for seat cover fabrics. However, due to its high cost, PET has been replaced with that fiber and has become the predominant fiber used in seat cover fabrics. Thus, wool fiber is generally used only in high-end cars [ 5 , 6 ]. Despite the inappropriateness of polyester fiber for use in applications where thermal comfort has priority, its low moisture absorption property can be an advantage when dimensional stability is required under changing environmental conditions (e.g. when it is used in seat belts). Although it is not shown in Table 3 , cellulosic fibers can also be used especially in seat covers due to their good esthetic and thermal comfort properties. However, very good abrasion resistance of PET provides durability and makes that fiber appropriate for use in seat cover fabrics [ 5 ].
PA has greater toughness, excellent tensile strain recovery and excellent adhesiveness when compared to PET [ 1 ]. However, PET has higher modulus, higher heat stability, higher resistance to color change, higher durability to sunlight exposure and it is less expensive than PA.
When compared to PP, PET and PA fibers have better dyeability characteristics, temperature resistance and dimensional stability. For example, 3 M&#; Thinsulate&#; acoustic insulation material is composed of 35% of PET staple fibers and 65% PP fibers. While the material has low weight, thanks to the polypropylene fiber, the unquestionable role of PET fiber is strengthening of the material for enhancing its usability in automobiles [ 7 ].
Although PP seems to be a good choice for several automobile components with its low cost and lightweight, this fiber has several disadvantages in terms of use, such as low melting point and moderate abrasion resistance. Besides, difficulty in PP&#;s dyeability overshadows the advantage of low cost [ 5 ].
Advantages and disadvantages of textile fibers commonly used in automotive applications are summarized in Table 3 [ 1 , 5 , 6 ]. As stated before, polyester and polyamide are the leading fibers in automotive applications; however, the very low cost of polypropylene makes it attractive; thus, it is also included in the table. Besides, as a natural fiber, wool is also preferred even in low amounts in automotive applications.
AdvertisementCarpets used in the automobiles are basically required to be durable against soil and abrasion with high color fastness and they are expected to insulate the noise as well as having a pleasing appearance [16]. In addition, environmental friendliness is a recent trend found in automobile carpets. In this regard, lightweight carpets for decreasing the total weight of the car and carpets from recycled materials have been developed for decreasing the environmental pollution [11].
Carpets used in automobiles basically can be categorized into two types: needle punched nonwoven carpets and tufted carpets. Fifty-five percent of the interior carpets are composed of nonwoven fabrics, whereas 45% are tufted carpets, depending on the car-manufacturing region or country [16, 17]. In both types of carpets, PET is mainly used in the facings [1]. On the other hand, needle punch facings are made from PET fibers in Europe and from PP fibers in the USA. In PET velour constructions, different fiber linear densities can be blended together to provide stability of the piles and esthetic appearance [17]. Carpets produced by using tufting method have backing surfaces for supporting the piles on the facing surface. In the backing surfaces, thermally bonded spunlaid PET and/or PP nonwovens are used [16]. For example, Colback® backings are thermally bonded spunlaid nonwovens, made from bi-component filaments with a PET core and either a PP or a PA surface. The product offers processing stability, high tear strength and uniform elongation as well as excellent thermal and dimensional stability [18]. For many years, automotive carpets have been produced by molding to the shape to match the dimensions where the carpet is used [19]. PET and/or PET blended nonwovens are appropriate to meet this requirement, since they can easily be molded into shape at relatively low temperatures and have outstanding dimensional stability [16].
Recently, the most important trend in automobile carpets is using recycled PET in order to reduce the environmental pollution and promote sustainability and efficient use of resources. The analysis of sound transmission loss of needle punch nonwoven automotive carpets made up of recycled PET revealed that comparable quality levels and specifications could be achieved using recycled PET instead of pure PET in terms of mechanical properties [11].
Made from recycled PET, Freudenberg Performance Materials from Germany produces automobile carpets with an environmentally friendly production process. In addition to using recycled fiber, they have eliminated the use of chemical binders and also obtained lighter carpets when compared to conventional automobile carpets, which promotes further environmental protection [12].
The required characteristics of the seat covers used in automobiles can be durability, soil resistance, UV resistance and appearance retention [17].
Automobile seats are produced in a three-layer structure. Seat cover is at the uppermost layer, a foam layer is in the middle and a scrim backing at the bottom, and these three different components are connected with adhesive layers. The seat cover at the top layer is usually made from woven, warp knitted or circular knitted fabrics, in which PET is used as a dominant fiber. Since late s, PET has been used in 95% of seat covers in automobiles due to its high strength and modulus, high performance against abrasion, UV radiation and heat, outstanding anti-aging performance, good shape retention and dimensional stability and low cost [20]. PET is generally used in pure form; however, sometimes it is blended with wool. Other characteristics that make PET appropriate for seating fabric are high tear resistance, easy care property and wrinkle resistance. On the other hand, very low moisture absorption capability of PET fiber (around 0.4%) is a disadvantage in terms of thermal comfort, especially in hot weather [5, 21]. In the usage of 100% PET constructions, fabrics are also produced from waste materials blended with thermoplastic to meet price points [16].
Banex is a fabric that uses a special type of PET yarn used as a seat cover that achieves the cushioning effect of springs with the machine-made warp knitting and finishing [22].
The foam layer beneath the top layer is generally made from polyurethane (PU) and acts as an adhesive agent [23]. However, the materials composed of PET and other polymeric fibers have been recently used in order to replace the foam material [17]. Nonwoven PET fabrics made from recycled fibers and the knitted structures can be used in the cover laminate instead of PU foam. Major knitted structures preferred over PU foam are the spacer fabrics, knit, multiknit and struto. Within these knit structures, multiknit comprises two stitch layers with piles in between, whereas knit consists of a stitch layer with a pile on the top [24].
The third layer at the bottom is the backing layer, which is generally produced from either PET or PA [5, 23, 25]. In the backing layer, bolster fabrics and reinforcement nonwovens such as needle punch, hydroentangled and spunbond nonwoven fabrics are preferred [17].
Pre-assembled parts are the fixtures in the car other than upholstery and carpeting, which are produced by molding in shape and covered with a fabric. Among them, there are headliners, bootliners, door panels and parcel shelves. In the production of these components, knitted, woven and nonwoven fabrics are used generally as the facing layer. On the other hand, polyester can also be used as a thermoset resin with vinyl ester in the production of natural fiber composites in automotive applications such as door panel, seat backs, headliners and dashboards. Whereas the natural fiber composites produced using vinyl esters are tougher, the orthophthalic polyester provides rigid products with low heat resistance and isophthalic polyester provides moisture resistance [14, 15].
Headliners are the parts which are tightly fitted into the interior roof from the rear window to the front of the car. They are usually produced by molding and are given the shapes to house sunroofs, lamps and coat hooks [16].
The characteristics, which are expected from headliners, include esthetic properties, sound absorption, thermal insulation and cushioning. Moreover, they are required to be produced from lightweight materials. The headliners are also required to be soil resistant [17].
Headliners are produced to have at least three layers, which are esthetic-facing fabric, a foam backing and the core [17]. As the facing fabric for covering, knit tricot fabrics and woven materials are used in majority. Nonetheless, nonwoven linings can also be preferred in the headliners as the facing fabric. Whereas knitted and woven linings provide better appearance, nonwoven materials are cheaper, easy to process and they show more resistance to abrasion. The other components of headliners are selected from porous structures such as PET foam sheets and fiber-reinforced porous polymer sheets, since the sound absorbing capability and heat insulation is needed [24].
On the other hand, the construction of the materials used in the production of headliners may change in different regions. For instance, regarding the facing fabrics, it is stated that half of the facing fabrics are made up of warp-knitted tricot fabrics, which are followed with needle punched and stitch bond fabrics in Europe [17]. Contrary to this, PA and PET warp-knit tricot fabric or woven materials are selected as facing fabrics to cover the headliners [16]. Nonetheless, having the ability to be easily assembled into the roof, headliners covered with dope dyed 100% PET needle punch fabrics seem to replace conventional warp-knitted fabrics. The other reason for this selection is the cost advantage, acceptable abrasion resistance and good thermal molding characteristics of the needle punch fabrics. Moreover, the weight of the needle punch fabric can be decreased by decreasing the fiber linear density. In parallel with this weight reduction, the softness of the fabric increases as well [16]. The other fabric type which drew attention is the hydroentangled fabrics because of better durability and softness. Being made up of spunlaid and hydroentangled splittable PA/PET bicomponent fibers, Evolon® fabric has excellent strength and softness. Produced using Apex technology and used in the production of headliners, Miratec fabric can copy the fabric patterns of 3D textiles and they have high strength in both horizontal and vertical directions [17, 26].
There are also some studies in which the headliner was produced solely from polyester fiber. The headliner material was formed from two layers of polyester fibers which include the binder and non-binder fibers. Whereas the layer including 20&#;30% binder fiber provided the loftier part and better sound absorption properties, the layer including 40&#;60% binder fibers provided rigidity to the headliner material [27].
In fact, polyester fiber is suggested to be a convenient material for the recycling process because of its thermoplastic characteristics. The headliners, which are produced from PET fibers, can be ground, melted and spun into new headliners [13]. Besides, needle punch PET facing fabrics were developed with PET core as an alternative to foam type materials. Although the foam type materials control the stiffness, increase the sound absorption and provide the cushioning effect, their recycling processes are complicated. By bonding directly the facing fabric to the core, the use of foam type materials can be eliminated [17].
The boot of the car can be considered as the extension of the cabin. The coverings such as rubber matting and a carpet joined to hardboard base is preferred as the bootliners [16].
The required properties for the bootliners can be wear resistance, abrasion resistance, durability, stiffness, lightweight and ease of cleaning. Moreover, the bootliner should not have excessive recovery since this may result in a tendency to shrink over time [16] and this property gains importance within the useful life of the product.
PP and PET are the mostly used fibers in the boot linings. Whereas PP is partially stable and because of this reason less preferred, PET is more dimensionally stable since it can be molded at high temperatures.
Because of economic concerns, nonwovens are mostly used in boot and luggage compartments. Again, in the nonwoven structures, PET is the preferred fiber for bootliners. Bootliners are usually produced from needle punch fabrics for which the facing fabric is made from a staple fiber PET or needle punch fabrics in either flat or velour construction [17].
On the other hand, the sound insulation properties of the bootliners can be improved with the underlay fabric produced from fiber batts composed of recycled fibers [17]. The fibers obtained from clothing wastes can be the blends of PET, PA, PP and acrylic or natural fibers such as wool, coconut, sisal jute and hemp [16].
The other application of polyester in the bootliners is the integration of it into the natural fiber composite materials. While natural fiber composites are preferred in automobiles since they cause a reduction in weight, energy production and cost at 10, 80 and 5%, respectively, polyester is used as a matrix material for these types of composites. Even different materials other than textile materials like sunflower can be used as the core material of these composites [28].
The door panels are the third type of preformed structures within the car interiors. The lower section of the door panel is produced as an extension of floor covering and the upper part is upholstery fabric or vinyl [16].
Having higher modulus, good heat stability, excellent resistance to color change and high durability for sunlight degradation and being less expensive, PET is used for making door panels [17].
For the panel trim in the door including the inserts or bolster, both the underlying reinforcement fabrics and lower facing fabrics can be constructed from nonwovens. The facing fabric can be selected from flat or random velour needle punch or hydroentangled materials [17].
Whereas needled fabrics produced from PP are used in the USA, the fabrics produced from fiber spun dyed PET fiber are preferred in Europe and Japan. Since interior fabrics are subjected to UV exposure, spun dyed fibers are preferred. But, in fact, spun dyed PET fiber has lower abrasion resistance when compared with PP fiber [17].
The second component of the door trim is usually made of hydroentangled fabrics. They are joined with the facing fabric or foamed PU and usually composed of 100% PET, 75% PET, 25% viscose or 50% PET/50% PP [17].
In fact, polyester is used in the door panel constructions which have parts differing from the ones explained earlier. In a study, the vehicle door panel was patented, which is mainly made up of rigid plastic panel, a paper-backing material attached with rigid plastic panel, a polyester fiber pad providing a cushioning surface, a cloth membrane and nylon adhesive membrane placed on the cloth and finally a vinyl membrane supported by cloth fabric [29].
The parcel shelf is the part of the car that encloses the area between the rear seats and boot. The requirements of this part are light fastness, lightweight and sound absorption to some extent.
Usually a needled fabric made up of PA, PET or PP fibers is used in the parcel shelf [16]. Often matt is included in the parcel shelf structure to increase the sound absorption.
Tires are the interface between the car and the road. The components of tires can be classified as the tread, belt package, ply, inner liner, apex, bead bundle, sidewalls and chafer [30]. Other than the rubber and steel components, textile fibers such as rayon, PA, PET or Kevlar are used in the ply component of the tire coated with rubber.
The tire ply cords are classified into three parts, which are named as bias, belted bias and radial ply cords according to their configuration. In diagonal (bias) tires, ply cords are laid at angles less than 90° to tread centerline. Belted bias tires have the belt added in the tread region. The radial tires have body ply cords, which are laid radially from bead to bead at 90° to the centerline of the tread [16]. The ply and tire ply cords transmit the braking and steering forces and withstand burst loads [30]. The requirements of the ply cords are thus tenacity, flexibility, shrinkage at high temperature, heat resistance, wear and abrasion resistance [16, 31]. Being used on the rougher roads, and requiring lower wearing resistance, PA fiber is used in bias tires as ply cord because of its excellent toughness. On the other hand, PET fiber is used in radial tires as ply cord because of having higher modulus and reducing the flat spotting [1]. PET is strong and stiff, and it provides excellent dimensional stability [16]. Moreover, it has high tenacity, good heat resistance, good wet resistance and low water absorption [32]. When compared to PA tire ply cords, the thermal shrinkage and flat spotting characteristics are superior. However, it lacks bonding with rubber when compared to PA [24]. The PET fiber type used in the tire cord is a multifilament fiber with high modulus and low shrinkage [1]. In radial tires, rayon is also used because of its superiority to PA in high-speed impact [33]. PET is also good at high-speed impacts but loses modulus and strength faster than viscose [33]. When these three fibers, PET, viscose and PA, are compared to each other, PA has the highest tenacity, whereas viscose has the least shrinkage at 160°C. On the other hand, heat generation is very low in rayon and it depends on driving conditions for the PET fiber [16].
In comparing PA 6, PA 6.6 and PET, Naskar et al. [34] applied cyclic compression and tension onto the cord-reinforced rubber composite specimens at different strain levels and time intervals on Goodrich compression and tension fatigue tester and found out that PET tire ply cords had excellent dimensional stability but poor fatigue resistance.
PET fibers are also developed to have distinguishing properties for being used as tire ply cords. A high modulus yarn was prepared by spinning polyethylene naphthalate (PEN) or other semi-crystalline PET polymers to an optimum crystallinity state. The resulting yarn had high tenacity, dimensional stability of less than 5% and shrinkage lower than 4% [35].
Finally, PET fibers were proposed to be used in the other parts of the automobile tires rather than tire ply cords. A limited twisted PET yarn having low polymerization degree was used in the belt breaker [32].
An average car involves many different types of filters to prevent air, fuel, oil and water from contaminants such as carburetor air filter, cabin interior filter, crank base breather filter, ABS wheel/brake filter, power steering filters, engine oil filters, fuel tank filter, transmission filters, wiper washer screen filters, air conditioning recirculation filter and diesel/soot filters [16].
The basic air filters in the automobiles can be classified as the engine air filter and the cabin air filter. The major purpose of air filters is cleaning the air and preventing the impurities within the air, which is used by the engine during combustion. Thus, the air filters indirectly protect the components of the engine from wear. Cabin air filter prevents the airborne pollutants and allergens within the cabin and they improve the quality of cabin air.
The air intake filters are usually produced from wet-laid, resin-impregnated cellulose papers. The other media used in the air filters are PU foam, nonwovens from synthetic, natural fibers or both in hybrid systems. In fact, the usage of nonwovens in air filtering is advantageous because they are more durable and have higher bursting strength. Besides, it is possible to control the parameters such as thickness, porosity and fiber diameter [17]. Moreover, the nonwoven filter media can be constructed to have specific characteristics such as being flame retardant, antibacterial property and so on.
PET fiber can be used in different forms as a filtration media. Mainly made up of PET fiber, QualiFlo® is developed as a gradient filtration media which is produced to have a trilaminate structure. It has exceptional dirt/dust-holding capacity consisting of continuous filament PET web with filaments having a trilobal cross section in one outer section and fully bonded air-permeable high loft batt with a randomly dispersed blend of crimped PET fibers in the other outer section. Moreover, requiring no additional binder, StarWeb, which is made up of spunbond PET, was proposed to be used in filtration with Qualiflo®. StarWeb is constructed as a trilaminate filter medium containing PET fibers as in the case of Qualiflo®. Within this filter, a top layer was produced from trilobal PET spunlaid fabric, the middle layer was produced from 100% PET homopolymer and copolymer PET filaments and the bottom layer was produced from high-loft PET crimped fiber fabrics with isotropic fiber arrangement [36].
Although there are filtration media produced only from PET fiber as is discussed in previous examples, PET is also preferred to be used in nonwoven and hybrid nonwoven structures in the air filter media.
A high-capacity hybrid, multilayer automotive air filter was developed and patented [37]. This air filter was designed to have a fluid filter media containing porous natural fibers and a porous synthetic fiber media containing absorbent spunbond PET. In the natural fiber filter media, two cotton mesh layers with different densities were used in a way that the first cotton mesh, which was placed closer to receiving end of the air stream, has lower density than the second cotton mesh layer, which was placed closer to the filtered effluent air stream. In parallel with this, synthetic fiber filter media, the first spunbond PET fiber filter placed closer to the second cotton mesh has lower density than the second spunbond PET mesh placed closer to the filtered effluent air stream creating a gradient density.
Placed between the fuel reservoir and the engine, the fuel filters are also used to protect the engine of the cars from dirt, water and the other contaminants.
The fuel filter media is usually made from PA. The reason for this is that, although the thermoplastic materials such are fluoropolymers, PET and ethylene-vinyl alcohol copolymers are resistant to fuels, the impact toughness especially at low temperatures is not found as satisfactory as PA [38]. Nonetheless, PET fiber can be used in different forms as a component of fuel filter media.
In fact, in a patent for a multilayer plastic fuel filter having at least three layers, the inner and outer layers are suggested to be produced from plastic material, which is made conductive using additives (PA6, PA6.6, PA11 and PA12). Embedded between them, PET was listed as one polymer which can be used with the other plastics which is not made conductive (fluoropolymer other than PVDH, PET or impact-modified PET) [38].
In a patent for developing depth media in-tank fuel filter with extruded mesh shell [39], it is stated that an in-tank fuel filter is developed to have two panels of filtration media. Each panel is comprised of an outer layer of extruded mesh, a pair of spunbond filtration media and an inner layer of meltblown filtration media in between. PET is used in both the spunbond and the meltblown filtration media.
Moreover, a diesel fuel filter is patented with a smoke suppressant for which the smoke suppressant is adsorbed onto the strip of nonwoven PET fiber, which is placed equidistantly around the round container [40].
Seat belts and airbags are the two main items used as safety equipment in automobiles. A seat belt is used to fix the passengers on their seats and decrease the import shock by absorbing [1]. High tensile strength and stability under static and dynamic loadings are needed in seat belts. The narrow fabric used in seat belts has mostly the weave structures such as plain weave, twill weave and sateen weave used in single-layer or double-layer structures. Seat belts are manufactured in a needle loom where the weft is inserted through the warp sheen and a selvedge is formed. On the other hand, filament yarns made of PA or PET are used to produce seat belt webbing [24].
Although PA was used as a major fiber in seatbelts for years, PET has been replaced with PA due to some advantages [40]. Having higher-impact energy-absorbing capability, less discoloration against sunlight and better dimensional stability under changing temperature and humidity conditions, PET is preferred to PA [1, 41].
When the static and dynamic loads have been applied on the seatbelts made from multifilament PET and PA yarns and compared, it has been observed that PET is superior due to lower extensibility that prevents stretching of the belts under loading during impact situations and higher stiffness [24, 41, 42].
As one other important safety equipment in automobiles, airbags are required to have extremely low gas permeability by means of a combination of high-density weaving and a thin coating treatment to resist high temperature, to have high extensibility and to be durable for storage in a compacted state for many years [43].
Demands for airbag yarns have been increasing recently, as a result of global rise in safety requirements in automobiles. Airbags are usually made up from PA 6.6 filament yarns. Nonetheless, there are some attempts to find alternative fibers. There is a noticeable trend toward PET and PA 4.6 filament yarns for the airbag fabrics and sewing threads [44, 45].
In Table 4, characteristics of the fibers used in airbag yarns are given [44].
CharacteristicPA 6.6PA 6PETDensity (kg/m3)Specific heat capacity (kJ/kg°K)1.671.671.3Melting point (°C)Softening point (°C)Heat to melt (kJ/kg)Characteristics of the fibers used in airbag yarns.
North America&#;s airbag market in comprised around 45,000 tons of PA 6.6-based fabrics and an additional tons of PET-based alternatives. On the other hand, in Europe, the use of PA 6.6-based fabrics has stayed around 33,000&#;35,000 tons per year while the use of PET fabrics has increased. It is forecasted that around 12,000&#;13,000 tons of PET will be employed in European airbags by [46].
Textile materials are used in engine compartment items such as driving belts (V-belt, synchronous belts and serpentine belt), hoses (brake and clutch hoses) and lines (power steering lines and bonnet lines).
In automobiles, the mechanical parts of the engines are driven by belts. The belts used in engine compartment of automobiles are required to be resistant to fatigue, abrasion, heat, chemicals as well as have high tensile strength and good dimensional stability [16].
A typical V-belt cross-sectional scheme is shown in Figure 1.
In order to enhance the fatigue properties of the belts that proposed mechanical and thermal loads, cords are used for reinforcement. The major reinforcing element used in the belts is PET cord, which is composed of twisted filament yams [16]. PET fiber is applied both in the cord of V-belt and in the fabric cover of its upper part, whereas p-Aramid fiber is applied to the cords of V-belts, V-ribbed belts and metal-combined belts [1].
The bonnet line and the fabric linings with it in the engine compartment require both thermal and sound insulation functions and it usually is constructed from metal or fiber-reinforced plastic composite. PET spunbond is usually used as a nonwoven facing to cover stiffening components such as glass fiber foam or resin bonded nonwoven fabric [17].
The brake and clutch hoses are required to prevent absorption of the lubricant fluid and to resist the fluid. As an example of those kinds of hoses supplied by Fenox, the PET yarn is used as the reinforcement component that prevents fluid absorption and extends the service life of the fluid and also it improves the ability of the hose to withstand pressures as a result of increased rigidity [48].
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