Scaffolds mimicking the native structure of tissues

Scaffolds mimicking the native structure of tissues

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R. Gater, W. Njoroge, H.A. Owida, Y. Yang Institute of Science & Technology in Medicine, University of Keele, Stoke-on-Trent, United Kingdom 

3.1   Introduction During embryonic development, primary cell layers known as “germ layers” work to give rise to all cell types required by the body. In humans, three germ layers develop known as the endoderm, mesoderm, and ectoderm. Once these cells aggregate, begin to communicate, and function together, tissue structures are formed [1]. The endoderm gives rise to the epithelial linings of digestive tract and lung tissues, as well as structures such as the stomach, liver, and pancreas. Alongside this the mesoderm aids in production of the cardiac, skeletal, and smooth muscle, as well as red blood cells and tissue surrounding the kidneys. Finally, the ectoderm is known to give rise to tissue with an epidermis such as hair, nails, and tooth enamel. Additionally, parts of the ectoderm develop further into the neural crest and neural tube, which aid to produce tissues such as the nervous system and brain. Because of this crucial role, some literature considers the neural crest as a fourth germ layer, despite being derived from the ectoderm. For these tissues to provide the specialized support mechanisms required for cellular function, native tissue structures contain an intricate microenvironment known as the extracellular matrix (ECM). Tissue ECM differs considerably between different cell and tissue types, playing key roles in the generation of diverse tissues and organs. Research into the generation of functional tissues through tissue engineering approaches necessitates a thorough understanding of the diversity of native ECM, in terms of its chemical and physical properties. This understanding gives rise to the design of artificial ECM in structures known as scaffolds, which can lead to the “development” journey of tissue formation in vitro, while still following in vivo paths.

3.2  Characterization of native tissues 3.2.1  Common chemical components in ECM ECM is a complex mixture of molecules secreted by cells to provide their surrounding structural and biochemical support. ECM characteristics can differ depending on the requirements of a particular tissue. Tissue-specific expression and synthesis of structural proteins and glycoprotein components result in unique functional and biological characteristics at distinct locations. Handbook of Tissue Engineering Scaffolds: Volume One. https://doi.org/10.1016/B978-0-08-102563-5.00003-4 Copyright © 2019 Elsevier Ltd. All rights reserved.

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The most abundant fibrous protein within ECM is collagen, which functions to provide the “scaffolding” of tissues for the attachment of proteoglycans and other ECM components. The name “collagen” is used as a generic term for proteins forming a characteristic triple helix of three polypeptide chains. All members of the collagen family form these supramolecular structures in the extracellular matrix, although their size, function, and tissue distribution vary considerably. There are 28 different types of collagen identified in vertebrates, with fibrillar collagen being the most abundant accounting for approximately 90% of all collagen in the body. Fibrillar collagen includes types I, II, III, V, XI, XXIV, and XXVII, with type I being the most common. Type I collagen is found in various tissues such as skin, connective tissue, vasculature, and the organic part of bone. Meanwhile, type II is found abundantly in tissues such as cartilage and the vitreous humor of the eye, accounting for up to 90% of collagen in these structures [2]. Collagen type III is widely distributed in collagen I containing tissues with the exception of bone and is an important component of reticular fibers in the interstitial tissue of the lungs, liver, dermis, spleen, and vessels. This homotrimeric molecule often contributes to mixed fibrils with type I collagen and is most prominent in highly compliant connective tissues [3–5]. Furthermore, type IV is the main collagen component of basement membrane, being a network-forming collagen that underlies epithelial and endothelial cells and functions as a barrier between tissue compartments [3,6]. Across all tissues, glycosaminoglycans (GAGs) are an important ECM component. GAG molecules are long unbranched polysaccharides with a repeated disaccharide unit. They attach to proteoglycans, adhesive glycoproteins, and fibrous proteins (e.g., collagen) to provide structural support [3,7] and play a major role in maintaining the equilibrium of healthy tissue by protecting and preserving ECM proteins and cytokines. The major GAG types with physiological significance in tissue can be categorized into four groups: chondroitin/dermatan sulfate, heparin/heparan sulfate, keratan sulfate, and hyaluronan.

3.2.2  Specific characteristics in ECM The characteristics of a tissue are ultimately determined by the type of cells, GAGs, proteoglycans, and proteins present. For example, connective tissue may typically have collagen fibrous protein present, surrounded by the proteoglycans decorin and biglycan, with chondroitin/dermatan sulfate GAG chains attached. Different tissues have differing properties and functional adaptations in ECM. There are diverse variations in ECM, for instance, soft or stiff ECM, stratifying (epithelial) ECM, layered/zonal topographic ECM, and vascularized ECM embedding multicellular tissues. These native scaffolds offer the principles in smart design and guidance for tissue engineers to produce artificial ECM (scaffolds) for tissue regeneration in vitro.

3.2.3  Mechanical properties - hard versus soft tissues The properties of native hard tissue (e.g., bone, teeth) are particularly unique in comparison with other tissues. Interestingly, water is a key component of bone tissue accounting for approximately 25% of its mass [8]. Approximately half of the dry weight of bone tissue is made up of mineral salts, mainly calcium and phosphate.

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The remaining 25%–30% of bone tissue mass is then comprised of fibrous proteins, primarily collagen. Research aiming to characterize the GAG and proteoglycan content of bone tissue has previously determined the presence of chondroitin sulfate and hyaluronic acid GAGs, with the associated proteoglycans (decorin, biglycan, and fibromodulin) and glypican [9]. This mixture of hard and soft material appears to be what gives bone its strength, with the hard material providing rigidity and the soft material providing some room for compression to avoid breakages. The mineral content of bone is believed to be the main influence of its mechanical properties, with increased mineralization making the bone stiffer, yet less capable of compression to withstand shock forces. Cancellous or “spongy” areas of bone tissue are believed to be better at withstanding strain forces, since it is porous and contains fluids such as blood vasculature and bone marrow. On the other hand, cortical or “compact” bone is stiffer and therefore better at withstanding stress forces, in comparison with strain, because of its high mineral content. These biological and mechanical properties are important factors to consider during the development of scaffolds aiming to mimic native hard tissue, such as bone. A common example of native soft tissue is the skin. Skin is a large, complex organ that provides the tissues, it encloses with protection from environmental stresses and serves as a sensory interface between the body and its surroundings. On a basic level, the skin provides a physical barrier that serves the dual function of excluding external stressors, including potential pathogens, and maintaining homeostasis. For the most part, this comprises heat regulation through sweating and vasoconstriction/vasodilation, control of fluid entry and loss by serving as a semiimpermeable barrier, synthesis of vitamin D through ultraviolet radiation in sunlight, sensory perception, and immune cell action [10–13]. Skin is a soft tissue exhibiting key mechanical behaviors. It must be flexible enough not to impair body motion, while being tough enough to resist tearing and piercing, as well as maintaining the ability to return to its original state. Skin can be described as anisotropic, viscoelastic, and nonlinear, resulting in an ability to endure large deformations. Because of the viscoelastic properties, skin undergoes a phenomenon known as preconditioning, where under cyclic loads the stress–strain relationship continuously alters until a steady state is reached [11,12,14,15]. The stratum corneum is the stiffest of the skin layers, therefore the least extendible under applied load. This layer exhibits less viscoelastic and preconditioning behavior compared with other layers but still maintains a nonlinear stress–strain relationship under applied tension. The underlying dermis contributes a large amount to the overall mechanical characteristics of skin; this layer consists of a dense network of collagen and elastin fibers, which allow for high levels of deformation. The underlying hypodermis is the softest of the three layers and evenly transfers loads from the upper skin layer to the underlying tissues. Hence it is important when applying loads perpendicularly to the surface of the skin [15–18].

3.2.4  Tissue with stratified epithelium (skin, lung, cornea, conjunctiva) Some native tissue types within the body possess a layer of cells lining the tissue surface, known as an epithelium. Functions of these cells include barrier protection,

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selective absorption/secretion, transcellular transport, and environmental sensing. Tissue epithelium does not contain blood vessels and therefore receives nourishment from the tissue it lines via a membrane known as the basement membrane. A common example of tissue with an epithelial layer is the lungs. In lung tissue, the alveoli are lined with thin, flattened cells known as the squamous epithelium. This is advantageous for lung gaseous exchange because it creates a short distance between air inside the alveoli and the blood capillaries. The lung trachea and respiratory tract are also lined with cells known as ciliated epithelium. These have tiny hairs known as cilia that function to sweep away any inhaled debris. All GAG types are believed to be present within lung tissue, with hyaluronic acid being the most abundant nonsulfated GAG and heparan sulfate being the most abundant sulfated GAG [19]. The skin is a highly organized, stratified structure consisting of three main layers: the epidermis, dermis, and hypodermis. The superficial epithelium layer, the epidermis, is approximately 75–150 μm in thickness [12,13,20] and consists largely of outward moving cells, keratinocytes, which are formed by division of cells in the basal layer of the epidermis. The second layer is the dermis, which is a dense fibroblastic connective tissue layer of thickness 1–4 mm. It mainly consists of collagen fibers, ground substance, and elastin fibers, and it forms the major mass of the skin. As a result of the collagen and elastin fibrils, the dermis has a mainly mechanical function, allowing for high levels of deformation, as the fibers stretch and reorientate [10,21]. The most numerous of the dermal cells are fibroblasts. They are responsible for the manufacture of all dermal connective tissue elements or their precursors. The third layer is the hypodermis, which contains areolar connective and adipose tissues, connecting the underlying muscles to the skin. The thickness of the layers varies considerably over the surface of the body. The areolar tissue consists of collagen and elastin fibers much like the dermis, and many migrating white blood cells that aim to destroy any pathogens that enter the body. The adipose tissue stores fats and nutrients as a potential energy source and provides cushioning for bony prominences [10,13,17]. Other examples of tissues, which possess an epithelial layer, are eye tissues, such as the cornea and the conjunctiva. Corneal epithelium functions to offer a barrier of protection for the ocular surface, including resistance from excess tear fluid and prevention of bacteria from entering the corneal stroma. Corneal epithelium contains several cell layers, including columnar/basal cells, polyhedral/wing cells, and squamous cells with flattened nuclei. It is reported that the corneal epithelial layers undergo constant mitosis because of the migration of columnar and polyhedral cells, as well as the constant loss of squamous cells as they are washed away by tear film. Underneath the corneal epithelium, the fibrous protein present within corneal tissue itself is primarily collagen. This collagen is highly aligned to enable tissue transparency for vision. Meanwhile, the GAG content of the corneal tissue is reported to be a combination of chondroitin sulfate/dermatan sulfate, heparan sulfate, and keratan sulfate [22]. Conjunctival tissue is a mucous membrane that covers the inside of the eyelids and fore part of the sclera [23]. The area of conjunctiva most commonly associated with disease is the bulbar conjunctiva, making it an interesting area for clinical research. The bulbar conjunctiva sits loosely on top of the sclera, separated by a thin, fibrous tissue layer known as Tenon’s capsule. Similar to other tissues, conjunctiva consists

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of three layers; fibrous, adenoid, and epithelium. Conjunctival epithelium contains several cell types including goblet cells, melanocytes, Langerhan’s cells, and the epithelial cells. Similar to the cornea, conjunctival epithelium functions to offer a barrier of protection for the ocular surface. Here, the goblet cells secrete mucus to trap and prevent bacteria from entering ocular structures such as the cornea. The fibrous protein present within conjunctival ECM itself is primarily collagen, with a sparse distribution of elastin [24]. Minimal previous literature reports the precise GAG content of conjunctival ECM; however, as a connective tissue with the presence of collagen, it can be assumed that chondroitin sulfate/dermatan sulfate GAG chains are present and attached to the proteoglycans decorin and biglycan. One study confirms a content of mainly chondroitin sulfate within conjunctival tissue, with no other GAG molecules reported [25]. The GAG content of scleral tissue (beneath the conjunctival layer) is reported to contain primarily chondroitin sulfate/dermatan sulfate, with a small quantity of hyaluronan [26].

3.2.5  Zonal, layer-specific tissues One example of layer-specific tissue is blood vessels. Blood vessels facilitate gaseous exchange, waste/nutrient transport, and cellular immune defense. Native blood vessel walls, excluding capillaries, are arranged into three concentric layers, which include intima, media, and adventitia. The innermost layer (intima) is comprised of a continuous single layer of endothelial cells that are in contact with the blood. This layer is the central barrier to the escape of plasma and is sealed with tight junctions. Underneath the intima is a dense elastic layer known as the internal elastic lamina, separating it from the subsequent layers. This is then followed by the middle layer (media), which is composed of smooth muscle (spindle-shaped) cells that provide structural support, mechanical strength, and contractility. Smooth muscle cells are arranged helically surrounded by elastin and collagen fibers forming a matrix (subendothelial matrix). The medial layer is separated from the adventitia (outer layer) by a dense elastic membrane, theexternal elastic lamina. The outermost (adventitial) layer consists of a connective tissue sheath without a distinct external border and has a function of anchoring the blood vessel to extracellular matrix (ECM). Thus, this layer has a role in tethering the blood vessel to surrounding tissues [27,28]. Articular cartilage is anisotropic in nature and is also organized into distinct zones. There are at least three architectural zones (superficial, middle, and deep) with striking variations between their structure, chondrocyte phenotype, ECM composition, and mechanical properties [29–31]. In the superficial zone, chondrocytes are elongated as a result of the tightly packed collagen fibers, which have a parallel orientation to the surface to dissipate high tensile strength. Also the concentration of proteoglycans is lower than that in middle and deep zones [30,32]. The middle zone contains rounded chondrocytes with a random collagen fiber orientation, while also having a large amount of collagen II and proteoglycans. In this zone, the collagen fibers are randomly oriented to provide resistance to the multidirectional compressive force [30,33]. The deep zone contains chondrocytes stacked in columns with radial collagen architecture and high proteoglycan concentration. This radial orientation provides cartilage with a high compressive stress resistance [30,32].

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3.2.6  Vascularized tissues In the human body, the vascular system forms an extensive, highly branched, and hierarchically organized network. This majority of tissues and organs are vascularized. The highly organized blood vessel network maintains tissues’ homeostasis through the mass transport of gas, liquids, nutrients, cells, signaling molecules, and waste products. Bone is a composite organ that fulfills several interconnected functions, including locomotion, involvement in phosphate and calcium metabolism, synthesis of endocrine molecules, and hematopoiesis, which may conflict with each other in pathological conditions. Vasculature is indispensable for appropriate bone development, regeneration and remodeling, and is a key interface between the various interconnected functions. It serves as a source of oxygen, nutrients, hormones, neurotransmitters, and growth factors [34,35]. The main blood supply of healthy long bones is derived from the principal nutrient arteries, which penetrate the cortex and perfuse the medullary sinusoids and then exit via multiple small veins [36]. Estimates of the proportion of the cardiac output received directly by the skeleton range from approximately 5.5%–11%. The rich perfusion of bone reflects not only the requirements of the bone cells (i.e., osteoblasts, osteocytes, and osteoclasts), but also the requirements of the marrow (i.e., hematopoietic lineage cells, stromal cells, and adipocytes), as well as endothelial cells. The vascular supply of bone enables the rapid growth and remodeling (including mechanical responsiveness) that would not otherwise be possible in essentially avascular cartilage [36,37]. Bone formation and development occurs through two distinct processes: intramembranous and endochondral ossification. The vascularization is the prerequisite of two different processes of ossification. In intramembranous bone formation, mesenchymal stem cells (MSCs) can be transported through capillaries and differentiate into mature osteoblasts, which then deposit bone matrix and lead to bone formation. During the endochondral ossification, the chondrocytes secrete angiogenic growth factors promoting the invasion of blood vessels, which then transport a number of highly specialized cells and replace the cartilage with bone and bone marrow [38]. In the absence of adequate blood supply, bone has been shown to display reduced growth and repair, experience loss of density, and eventually undergo necrosis. The blood vessels in skin tissue are comprised of a number of cellular and noncellular elements. Endothelial cells (ECs) form the interface between intravascular and extravascular compartments and serve as a selective barrier for the diffusion of cells and macromolecules between compartments. Along with a basement membrane that surrounds the ECs and additional cells such as smooth muscle cells and pericytes, dermal blood vessels are also surrounded by a variety of cells originating from the bone marrow such as mast cells, macrophages, dendritic cells, and mononuclear cells. These bone marrow–derived cells preferentially reside in a perivascular location in the skin. Blood vessels in the skin form a deep plexus in the subcutaneous fat and provide nutrition to the sweat glands and the hair papillae, and a superficial plexus in the papillary dermis [39,40]. The skin is supplied by arteries that originate from their underlying source vessels and are either destined for the skin (direct perforators)

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or from secondary branches of vessels that supply other underlying tissues (indirect perforators). The plexuses that form cutaneous vasculature are formed from either the direct or indirect perforators. The arteries and their associated microvasculature are referred to as angiosomes. The density of angiosomes varies from site to site, and this is primarily determined by whether the skin served is fixed to underlying tissues or is freely moving. Fixed sites tend to have more angiosomes that demonstrate limited branching, whereas large areas of movable skin may be served by a single feeder vessel that branches extensively in the superficial structures [40,41].

3.3  Scaffold designs to mimic the native structure of tissues A current strategy in the tissue engineering field is to induce formation of the specialized features of native tissue, through the design of biomimetic scaffolds. It is hoped that the ability to produce natural tissues through biomimetic scaffolds will aid in the development of cell and tissue replacement therapies, as well as provide a better understanding of health and disease development mechanisms. However, the development of scaffolds to mimic native ECM structurally and chemically is particularly challenging. Attempts have been made to replicate the ECM utilizing a variety of materials both natural and artificial, and these have been used with varying levels of success. Learning from nature, various culture and fabrication techniques have been developed to generate diverse tissues with differences in composition and architecture.

3.3.1  Scaffolds for soft tissue To mimic native soft tissue structures such as skin and conjunctiva, scaffolds using hydrogels and fibrous materials are common. Biocompatible polymers for these applications are considered to be either natural, semisynthetic, or synthetic. Polymers containing natural ECM components provide enhanced biocompatibility to hydrogels, providing mechanical integrity to tissues, as well as enhanced support and regulation for cellular processes. Although synthetic polymers can aim to mimic natural ECM components, a naturally occurring polymer such as collagen, elastin, or fibrinogen is thought to provide cells with a more physiologically relevant substrate on which to reside. A suitable native skin substitute should ideally have the ability to be kept sterile, act as a protective barrier, have a low inflammatory response, and allow for water vapor transmission across the material [42]. There are several scaffold designs attempting to mimic native skin tissue, a common type being porous scaffolds and hydrogels. Most recent examples include a fibrinogen-modified sodium alginate sponge scaffold developed by Soloveiva et al. [43], as well as a silk fibroin and functionalized citrus pectin derived scaffold developed by Turkkan et al. (2018) [44]. High porosity allows for better ECM secretion and nutrient supply for cells, as well as ensuring an even spread of cells across the scaffold without clustering together. However, optimization for the best pore size intended for particular cell types can be challenging.

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Hydrogel-based scaffolds are another common scaffold design in skin tissue engineering. Some of the most recent examples include a functional chitosan-based hydrogel developed for skin wound dressing and drug delivery by Liu et al. [45], plus a poly vinyl alcohol–poly acrylic acid–derived double layered hydrogel for wound dressings by Tavakoli et al. [46]. Hydrogel scaffolds can also allow for good porosity, as well as a high biocompatibility and controlled biodegradation rate. However, because hydrogels are soft structures often with a high water content, reduced mechanical strength can be a limitation. Fibrous scaffold designs can also offer a highly microporous structure, as well as improved mechanical strength for the material. Recent examples include a soy protein/cellulose-derived nanofiber scaffold to mimic skin ECM developed by Ahn et al. [47] and an electrospun poly(lactic-co-glycolic-acid)-derived nanofibrous scaffold by Zhu et al. [48]. However, the surface functionalization of nanofibers can be challenging to develop in this type of scaffold design. Future work in this area may eventually aim to incorporate drugs and biological molecules into fibrous scaffolds, for growth factor or drug release applications to aid in skin wound healing. Model constructs to mimic native conjunctiva are advantageous for the study of wound healing mechanisms in conjunctival tissue after glaucoma surgery [49], as well as for the study of drug treatments to treat diseases of the conjunctiva. Notable studies for polymeric scaffold designs to model conjunctiva include He et al. (2016) [50], who found poly-d-lysine-coated silk to be the most biocompatible polymer substrate for maintaining human conjunctival goblet cells in culture. Being a fibrous tissue, fibrous scaffold designs are also advantageous for mimicking conjunctival tissue to study ocular surface disease. Garcia-Posadas [51] also previously developed a fibrinbased 3D model for the successful culture of human conjunctival fibroblasts. Because of the fibrous protein content of native conjunctival tissue being primarily collagen, collagen hydrogel-based scaffolds to seed conjunctival cells have also been popular in past research [52]. Similar to scaffold design mimicking skin tissue, future work to incorporate biological molecules, such as growth factors, into fibrous and hydrogel scaffolds modeling conjunctiva would be advantageous. This would allow researchers to more closely mimic disease mechanisms of native conjunctival tissue and improve the testing of new drug treatments.

3.3.2  Tissue models with epithelium (coculture + multilayer scaffolds) As outlined previously, an epithelial layer can offer an additional barrier of protection, as well as improve the transcellular transport and environmental sensing of particular tissues. In an attempt to mimic this, multilayer scaffold designs and coculture models are common in the tissue engineering field. Development of lung tissue models can vary depending on the area of lung tissue that researchers would like to mimic. For example, models to mimic the alveoli may include a squamous epithelium, whereas models mimicking the trachea/airway may incorporate a ciliated epithelium instead. There are several multilayer and coculture scaffold designs attempting to mimic areas of native lung tissue. Some recent

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examples include an epithelial–mesenchymal coculture model of human bronchial tissue by Ishikawa et al. [53], as well as a triple layered coculture model of human respiratory tract by Blom et al. [54]. By including multiple cell layers, these models offer an improved opportunity for the study of full-thickness lung tissue. Similar to lung tissue, multilayer and coculture scaffold designs are also popular in research trying to mimic native skin tissue. As outlined previously skin tissue consists of three main cell layers, the epidermis, dermis, and hypodermis, making multilayer model designs particularly appropriate. A previous study of this includes [55] who assembled a multilayer model with hyaluronic acid and poly-l-lysine (for the epidermal layer) onto a porous hyaluronic acid scaffold (for the dermal layer), with human keratinocytes seeded on top. Models such as these are an example of how several scaffold design ideas are frequently combined in current research, such as using coculture techniques in addition to porous and hydrogel scaffold designs. Wilson et al. [56] have developed a multiple-layered corneal model with epithelial cells grown on a stromal layer (Fig. 3.1). The stromal layer was constructed by collagen type I scaffolds. The stromal cell orientation could be controlled by adding orthogonally aligned nanofiber [57]. Epithelial cells could be easily incorporated to the top of the stromal layer. The cross-talk of epithelial–stromal cells regulated stromal cells’ phenotype significantly.

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Combined scaffold design ideas are also frequently seen in the development of recent conjunctival membrane models. Examples include a model developed by Garcia-Posadas et al. [51] consisting of conjunctival fibroblasts seeded within a fibrin-based scaffold with epithelial cells seeded on top. Yao et al. [58] also previously developed a coculture model consisting of a collagen and poly(l-lactic acid co-ε-caprolactone)-derived porous nanofibrous scaffold seeded with conjunctival epithelial cells to stimulate stratification. Furthermore, a recent ocular surface and tear film model was developed by Lu et al. [59] using the coculture of rabbit conjunctival epithelium and lacrimal gland cell spheroids to mimic the aqueous and mucin tear film layers within an optimized air–liquid interface. Although multilayer and coculture methods offer a more realistic mimic of native tissue structures in comparison with monocultures, these are still limited and do not completely mimic a natural ECM structure. To overcome this challenge, continuous research into more accurate scaffold organization, as well as improved mechanical strength, porosity, and biocompatibility for the desired tissue type, is required [60]. Application of surface modifications is one of the approaches to modify scaffolds and cell sheets. For example, Yang et al. [61] recently found the use of collagen and alginate nanofilms improved the mechanical property of cell sheets, which could then be assembled as multilayers. 3D printing technology is another approach, which can be used to produce ECM scaffolds with improved accuracy to native tissue. For example, Hoshiba and Gong [62] recently worked on 3D printed poly(d, l-lactic acid)-derived scaffolds. These could be deposited with ECM constituents using the culture of fibrosarcoma cells on the scaffolds, before subsequently being decellularized for use in other applications. Improved techniques such as these are likely to revolutionize the way in which researchers design scaffolds to mimic native tissue structures in the future.

3.3.3  Scaffolds with zonal, layered structure Considering the prevalence and importance of zonal variations in normal articular cartilage, recent studies have aimed to engineer cartilage with zonal structure or function, or both. Replication of the zonal organization of tissue-engineered cartilage is one of the multiple strategies to generate functional tissue. Yucekul et al. [63] previously demonstrated the potential repair and regeneration of cartilage tissue by using a multilayered biomimetic scaffold. This study proposed a biodegradable, trilayered (poly(glycolic acid) mesh/poly(l-lactic acid)-colorant tidemark layer/collagen I, and ceramic microparticle-coated poly(l-lactic acid)-poly(ε-caprolactone) monolith) osteochondral plug indicated for the repair of cartilage defects. The porous plug allowed for continual transport of bone marrow constituents from the subchondral layer to the cartilage defect site for a more effective repair of the area. Owida et al. [64] developed zonal-specific three-dimensional hybrid scaffolds and a simple method for biomimetic cartilage regeneration, which can induce the generation of zonal-specific cellular morphology and ECM composition within articular cartilage zones. Scaffolds were formed using polylactic acidnanofiber meshes with different alignments: aligned nanofibers for the superficial zone, randomly aligned fibers for the middle zone, and microchannels in hyaluronic acidhydrogel (Fig. 3.2).

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Figure 3.2  Schematic illustration of the assembled three separately cultured zonal scaffolds with corresponding thickness (a). An OCT image of the assembled 3D zonal scaffold showing the aligned nanofiber in superficial zone, random nanofiber in middle zone, and the vertical channels in deep zone (b). Scale bar is 250 μm [64].

McCorry et al. [65] developed a simplified enthesis model, which can be used to mimic the native enthesis morphology and serves as an ideal test platform via test collagen integration with decellularized bone. Injecting collagen into tubing loaded with decellularized bone plugs resulted in a scaffold with three regions: bone, bone–collagen, and collagen. Furthermore, collagen formation was directed in the axial direction using mechanical fixation at the bony ends. Tellado et al. [66] provided new insights into the combinational effects of scaffold internal architecture and growth factors; biphasic silk fibroin scaffolds with integrated anisotropic (tendon/ligament-like) and isotropic (bone/cartilage like) pore alignment, which presented a promising strategy for tendon/ligament-to-bone regeneration. Human primary adipose-derived mesenchymal stem cells were cultured on biphasic silk fibroin scaffolds with integrated anisotropic (tendon/ligament-like) and isotropic (bone/cartilage like) pore alignment. Furthermore, the scaffolds were functionalized with heparin and the ability to deliver transforming growth factor ß2 and growth/differentiation factor 5 was explored.

3.3.4  Scaffolds to promote vascularization In the ever advancing field of tissue engineering, vascularization of tissues is an important aspect that allows the creation of specialized tissues that are able to maintain growth, function, and viability. Vasculature has a vital role to play in creating and maintaining healthy tissues. As such, creating artificial tissues that are able to effectively mimic and integrate with native tissues is a key challenge with tissue engineering. Common approaches to achieve this include the use of cells, growth factors, ECM proteins, and biophysical stimuli [67].

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Bone grafts are crucial for the treatment of a number of conditions that represent a great global burden, including segmental bone defects caused by trauma, tumor excision, or chronic osteomyelitis to name a few. The need for bone constructs stems from the limited availability of donor tissues, which can be categorized as autografts, allografts, and xenografts, with the current gold standard treatment being autografts. However, these are limited in availability, prolong the operation times, and are often associated with donor site morbidity [68,69]. As an alternative, allografts are widely available. However, they place the patients at risk for infections and rejection by the immune system. More importantly, allografts demonstrate decreased integration with the host tissue when compared with autografts, with high failure rates of approximately 25% and reaching up to 60% in patients requiring large grafts. Allografts, as well as synthetic grafts, are associated with complications such as osteonecrosis, as these options lack an endogenous vascular network to facilitate host integration. The current focus looks at developing vascularized bone grafts to counter these limitations [68]. The aim of bone tissue engineering is to provide a bone environment rich in functional vascular networks to achieve efficient osseointegration and accelerate restoration of function after implantation. To attain both structural and vascular integration of the grafts, a large number of biomaterials, cells, and biological cues have been evaluated [70]. An example of an approach that has shown some measure of success in improving graft vascularization is the use of a polymethyl methacrylate (PMMA) cement block, which is placed on the bone defect and is surrounded with the adjacent soft tissues. The host slowly creates a membrane rich in vasculature around the block, which is then removed and replaced with an autograft. Bone repair is then prompted by host remodeling in this vascularized environment [70,71]. Another established approach is the use of synthetic fillers, which provide an osteoconductive scaffold into which bone-depositing cells migrate and deposit new bone tissue [70,72]. Although promising results have been obtained, it is also increasingly evident that new research should be aimed at hierarchical integration of bone and vascular devices to yield fully functional bone tissue engineering constructs with enhanced biological properties that can promote concurrent osteogenic and angiogenic growth and seamlessly assimilate with the host bone matrix and vasculature [70]. The materials selected as the scaffolds for tissue engineering need to be inherently biocompatible, biodegradable, and promoters of cell adhesion. The materials also need to be porous and mechanically stable and exhibit a 3D structure that can be obtained through facile manufacturing processes [73]. Various strategies have been attempted to enhance the establishment of vascular networks within engineered constructs for bone regeneration. These include (1) directing cell behavior through growth factor delivery, (2) using coculturing systems, (3) applying mechanical stimulation, (4) using biomaterials with appropriate properties, and (5) incorporating microfabrication techniques. The development and use of 3D printed scaffolds represents a huge opportunity for the world of tissue engineering. This methodology represents an accurate, in terms of design, and reproducible method for the fabrication of porous scaffolds, which facilitate cell attachment, vascularization, and tissue ingrowth. Development of a vascular network has been shown to be successful with porous, degradable scaffold sleeve designs with a lumen [69,74].

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With numerous established and emerging methods of scaffold fabrication available, the incorporation of vascular cells within the scaffolds provides an ideal opportunity to develop sufficient vasculature within the engineered tissue. The use of endothelial cells in coculture with other cell types can be used to achieve in vitro prevascularization and generate vascularization within the scaffold construct, and it has been reported that there is a mutual cell–cell interaction between ECs and osteoblast-like cells during osteogenesis. Cross-talk between human umbilical vein endothelial cells (HUVECs) and MSCs not only enhanced the osteogenic differentiation of MSCs but also increased the proliferation of differentiated MSCs [75,76]. In addition, cocultures of HUVECs and primary human osteoblasts (pOB) in 3D spheroids seeded into collagen gels demonstrated the formation of microcapillary-like structures resembling capillaries. It was found that during cell–cell contact between spheroids of HUVECs and pOB, an upregulation of VEGFR-2 was observed on the HUVECs, whereas a downregulation of VEGF production and an upregulation of alkaline phosphatase activity were observed in the pOB. This is an example of one of the first studies showing a bidirectional regulation of gene expression of HUVECs in coculture with pOB that required a direct contact with pOB to form capillary-like structures [77]. The choice of endothelial cell type for prevascularization of grafts is crucial. Mature ECs have been traditionally thought to be responsible for postnatal angiogenesis by which capillaries sprout from preexisting structures, until recent identification of peripheral blood-derived endothelial progenitor cells (PB-EPCs) by Asahara et al. [78]. These cells were found to be 10 times more proliferative than HUVECs. The use of human progenitor-derived endothelial cells (PDECs) in a 3D coculture with MSCs showed new osteoid formation as demonstrated by Guerrero et al. [79]. The parallel formation of a vascular network, surrounding the tissue when implanted in vivo, represents a promising approach for achieving vascularization [76,79]. The proangiogenic effect exerted by EPCs appears dependent on their ability to stimulate the growth of endothelial cells and interact with mature endothelial cells to support vascular anastomosis. Nevertheless, the mechanism by which the EPCs exert beneficial effects on endothelial cell growth is likely multifactorial and might include the transdifferentiation of subpopulations of EPCs into mature endothelial cells [68]. Vascularization is a key process in skin tissue engineering, which determines the biological function of artificial skin implants. Consequently, efficient vascularization strategies are a vital prerequisite for the safe clinical application of these implants. Current approaches include (1) modification of structural and physicochemical properties of dermal scaffolds, (2) biological scaffold activation with growth factor–releasing systems or gene vectors, and (3) generation of prevascularized skin substitutes by seeding scaffolds with vessel-forming cells [80]. Attempts are being made to incorporate vasculature into engineered skin although applications in humans are thus far limited. An example of vascular inclusion in tissue-engineered skin is provided by Tremblay et al. [81] where they incorporated keratinocytes, fibroblasts, and endothelial cells (ECs) in a collagen sponge implanted in mice. The implanted skin substitute then spontaneously formed a network of capillary-like structures (CLS) in vitro. After transplantation to nude mice, they demonstrated that CLS containing mouse blood were observed underneath the epidermis in the endothelialized reconstructed

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skin (ERS) in less than 4 days versus 14 days required to have the same observation in non-endothelialized reconstructed skin. This result was associated with inosculation of the CLS network with the host’s capillaries, rather than neovascularization. Another approach is tissue engineering of a vascular network in human dermoepidermal skin substitutes (DESS). Klar et al. [82] utilized adipose stromal vascular fraction (SVF)–derived endothelial cell population to tissue-engineered DESS containing a highly efficient capillary plexus. To develop vascular networks in vitro, they employed optimized 3D fibrin or collagen type I hydrogel systems and upon transplantation onto immune-deficient rats; these preformed vascular networks anastomosed to the recipient’s vasculature within only 4 days. As a consequence, the neoepidermis efficiently established tissue homeostasis; the dermis underwent almost no contraction and showed sustained epidermal coverage in vivo. They conclude that adipose-derived SVF represents a convenient source of endothelial cells and pericytes and when submerged within an appropriate 3D environment, these cells allow the in vitro prevascularization of human autologous dermoepidermal skin grafts.

3.3.5  Scaffolds from decellularized tissues A convenient alternative to manufactured scaffolds is the use of decellularized tissues to obtain ECM. The most obvious advantage of decellularized tissue is the resultant ECM is perfectly designed to host cells in a 3D environment in vitro. This ECM has the natural conformation and chemical composition of specific tissues, generating the specific cues that cells need to grow in a more native environment. The removal of cellular content and antigens from the tissue-derived scaffolds reduces foreign body reaction, inflammation, and potential immune rejection. A disadvantage of this approach is the possible structural alteration of ECM or loss of some important components in ECM because of overexposure to chemicals during the decellularization process [83,84]. The clinical use of decellularized scaffolds has been applied for blood vessels, cardiac valves, and renal bladders. The current applications may be limited to tissue-level and anatomically simple organs; despite this, they ultimately provide the foundation for future complex and functioning organ regeneration [85]. There are a variety of approaches for decellularizing tissues. Effective decellularization methods include chemical, enzymatic, physical, or combinational approaches. The working principle of these methods is that the cell membrane is disrupted, and the cellular contents are released and rinsed away [84]. One method is the use of sodium dodecyl sulfate (SDS) as described by Schaner et al. [86]. They used intact human greater saphenous vein specimens that were decellularized using sodium dodecyl sulfate and assessed their viability for use as scaffolds for vascular tissue engineering. By evaluating burst and suture-holding strength and handling and durability of decellularized vein in vivo, they determined that the vein rendered acellular with SDS has well-preserved extracellular matrix, basement membrane structure, and strength sufficient for vascular grafting [86]. SDS has been used in a number of studies as the preferred chemical agent for decellularization. The method described by Guler et al. [83] involves the use of SDS to decellularize porcine aorta. Dimethyl sulfoxide (DMSO) was used as a penetration enhancer in the

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decellularization process to enhance the penetration of SDS, consequently reducing the exposure time of SDS to treated tissues. They determined that the addition of DMSO resulted in the removal of 64.4% more DNA compared with SDS on its own over a 3 hours period. The inclusion of DMSO also resulted in denser and more organized collagen fibers [83]. Another approach to decellularization is described by Tapias and Ott [87]. Their approach involved the use of perfusion decellularization. This method uses perfusion via the intrinsic vascular network as the most efficient way to deliver decellularizing agents, even to thick tissues, as it greatly decreases the diffusion distance of the decellularizing agent while preserving the three-dimensional macro- and microarchitecture [87]. The work done by Quint et al. [88] demonstrates practical use of decellularized scaffolds. Their approach involved the use of tissue-engineered vessels (TEVs) grown from banked porcine smooth muscle cells that were allogeneic to the intended recipient, using a biomimetic perfusion system. The engineered vessels were then decellularized, leaving behind the mechanically robust extracellular matrix of the graft wall. The acellular grafts were then seeded with either endothelial progenitor cells (EPC) or endothelial cells (EC), which were derived from the intended recipient, on the graft lumen. The TEVs were then implanted as end-to-side grafts in the porcine carotid artery, which is a rigorous testbed because of its tendency for graft occlusion. The EPC- and EC-seeded TEV all remained patent for 30 days in their study, whereas the contralateral control vein grafts were patent in only 3/8 implants [88]. Deviating from animal sources of decellularized matrix, Gershlak [89] utilized decellularized plant tissue as a prevascularized scaffold for tissue engineering applications. Their study exploited the similarities between the vascular structure of plant and animal tissues and using perfusion-based decellularization modified for different plant species, provided different geometries of scaffolding. After decellularization, plant scaffolds remained patent and able to transport microparticles. Plant scaffolds were recellularized with human endothelial cells that colonized the inner surfaces of plant vasculature. Human mesenchymal stem cells and human pluripotent stem cell–derived cardiomyocytes adhered to the outer surfaces of plant scaffolds. Cardiomyocytes demonstrated contractile function and calcium handling capabilities over the course of 21 days [89]. Decellularized scaffolds have vast potential for the regeneration of either small sections of tissue or whole organs and have demonstrated regenerative capabilities in vivo and in vitro, suggesting their value in emerging treatment approaches. There are still limitations with the technique such as the achievement of fully functional organs that bear all the necessary native properties and the limitations of tissue-specific decellularization methods and the resultant variation in efficacy. A combination of multiple approaches, as demonstrated by Guler et al. [83], led to improved outcomes. With an ever increasing understanding of underlying mechanisms and development of new materials and techniques, the generation of engineered tissues and organs that are as good as “the real thing” becomes less of a possibility and more of a certainty. Recellularization of the decellularized scaffolds by corresponded cell types is an other big challenge. The dense ECM network prevents seeded cells infiltrating to

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the related layers, especially the media layer. Recellularization of right phenotype cells in right layers for multiple cellular tissue is even tough challenge. A few nice projects have been undertaken for thin tissue with single cell type. Chani et al. for example,recellularized decellularized rat kidney [90]. The pattern of distribution of the injected cells was organized into vascular structures within the glomerulus, similar to freshly isolated kidney. Dynamically and sequentially seeding with perfusing system is one of the feasible techniques to achieve high seeding efficiency.

3.4  Summary The characteristics of native tissue structures are ultimately determined by the type of GAGs, proteoglycans, and proteins present in the ECM environment, in addition to the arrangement of cell types present. Factors such as mechanical properties, cell layer arrangement, vasculature, and the presence of an epithelial layer can each influence how a tissue adapts to its function. To mimic these native tissue structures in vitro, tissue engineers have developed various scaffold designs with increasing complexity over recent years. These include hydrogel, porous and fibrous scaffolds, coculture and multilayered scaffolds, zonal structure scaffolds, as well as scaffolds to promote vasculature. Continued work in this area is likely to revolutionize the way in which we study and treat disease mechanisms in the future.

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