Additive manufacturing (AM) of medical devices and scaffolds for tissue engineering based on 3D and 4D printing

CHAPTER 5

Additive manufacturing (AM) of medical devices and scaffolds for tissue engineering based on 3D and 4D printing Sudip Kumar Sinha Department of Metallurgical Engineering, NIT Raipur, Raipur, India

1 Introduction In the last decade, the biomedical field has witnessed massive and sustained growth in various facets of human tissue regeneration. This process is primarily concerned with cell growth and the reconstruction of organs, and therefore tissue regeneration is an area of immense interest for scientists and academicians globally. Based on last fifty years research in the field hard tissue implants, it can be understood that organ transplantation, its substitution, and fixation are the practical alternatives for patients with injured or damaged organs. Long waiting lists for organ transplantation are a common occurrence across the world. Through June 2017, the US Department of Health and Human Services had a list of around 120,000 patients who were in dire need of various lifesaving organ transplants while only about 5200 donors offered such organs [1]. It is clearly seen that there are an increasing number of patients waiting for transplanted organs while the number of transplants that actually took place over the last few decades was somewhat constant across various age groups. Therefore, there is an extreme need to discover unconventional ways to meet the needs for this scarcity of implantable organs [1]. Three-dimensional (3D) printing is a type of additive manufacturing-based technology for the exact 3D construction of engineering components. It is extensively used in biomedical engineering. Although there are various categories of additive manufacturing techniques available, the terms 3D and 4D printing are often interchangeably used in this context for simplicity. The technology began from a liquid-based stereolithography method during the latter part of the 1980s [2]. 3D printing (popularly referred to by other names such as additive manufacturing or rapid prototyping) is primarily a layered deposition and manufacturing technique where materials are overlapped as one layer over another. During the last four decades, several 3D printing technologies have been developed based on this processing approach. The American Society for Testing and Materials 3D and 4D Printing of Polymer Nanocomposite Materials https://doi.org/10.1016/B978-0-12-816805-9.00005-3

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(ISO/ASTM 52900:2015) has designated additive manufacturing techniques in 50 different types, which can be further categorized by seven distinct groups: (i) jetting by binders, (ii) jetting based on various materials, (iii) extrusion of materials, (iv) vat photopolymerization, (v) powder bed fusion, (vi) energy deposition, and (vii) sheet lamination. However, in a different approach, the gross classification of 3D printing techniques is done on the basis of the raw material (or ink) deposition procedures. So far, the extrusion-based techniques are the most widely used 3D printing techniques. These include fused filament fabrication (fused deposition modeling (FDM)) and direct ink writing or direct writing (DIW). In these conventional methods, a 3D object is fabricated by deposition on a line-by-line basis and then a layer-by-layer sequence. The only distinction is that FDM melts a solid string or filament emitted through a nozzle that is heated according to the specification while the DIW method pours a viscous or semislurry ink-based solution that might be treated later. Extrusion-based techniques are advantageous because they are capable of printing a vast range of printable materials. Therefore, this technology aids in the fast and efficient fabrication of any type of customized or complex engineering component by the precise accumulation of feed materials using a solid modeling approach with the help of a digital 3D file such as a computer aided design (CAD)-based illustration or CT (computed tomography) scan imaging. The concept of tissue engineering (TE) was first posited by Langer and Vacanti in 1993 and was published in their landmark manuscript in Science [3]. They demonstrated the primary attributes and applications of 3D biodegradable scaffolds. 3D scaffolds should possess a very porous morphology with well-interconnected networks of open pores, and must have a uniform and ample pore size and their distribution for cell proliferation and penetration [4] of cells. Following this breakthrough invention, a number of manufacturing techniques were used for the creation of these porous scaffolds. However, the traditional techniques suffer major limitations as they don’t have the ability to support or have enough potential to master the scaffold architecture, pore size, and its network, it eventually forms a 3D scaffold that is incompatible and less than ideal from the ideal one. This was the motivation behind the use of 3D-printing techniques to produce tailor-made scaffolds with unique structures consisting of well-controlled and uniformly distributed pores of the desired size and shape [5–7]. The scope of application of 3D printing has enabled scientists and engineers to perform desired modifications with ease without any additional requirement of equipment or manufacturing units. The technique also has the unique advantage for manufacturers to construct devices replicating a patient’s specific anatomy (customized devices) or equipment with a very intricate internal assembly. These potential abilities have prompted enormous curiosity in 3D printing, not only in the field of medical devices but also in other applications ranging from food items to commodities humans use in their daily lives to automotive components. On the other hand, 4D printing is a much more recently investigated field that was developed by considering 3D printing as the base. It has promising applicability in

Tissue engineering based on 3D and 4D printing

advanced biomedical research. A group of researchers at the Massachusetts Institute of Technology [8] first proposed the technique in 2013. The development of a modern 4D printing device is largely based on existing 3D printing technology, a rapid growth in smart materials, and a strong foundation of mathematical modeling and subsequent design. Although similar to 3D printing, 4D printing actually adds an extra dimension of revolution in the course of time, where the as-fabricated products interact with environmental stimuli such as temperature, humidity, and light. This results in a change of the active material structure as per the requirements over time [9–12]. In spite of great improvements in the rapidly emerging and advanced 4D printing technology, its precise application in bioengineering is largely dependent on the creation of multifaceted stimuli-responsive 3D printable materials, highly sophisticated 3D printing technologies with the ability to deposit multiple active materials, and feasible computational design approaches to envisage the 3D printed entity’s performance under external stimuli. The sustained growth and development of 3D and 4D printing as an emerging technology in numerous advanced biomedical applications such as traditional biomaterials, tissues and organ (re)generation, chronic diseases, drug delivery applications, implants for hard tissue replacements, biomedical instruments, and prosthetics have motivated scientists and researchers to seek the next level of invention. The flow sheet diagram shown in Fig. 1 explains the 3D/4D printing technology in correlation with various biomaterials used and the conventional synthesis of scaffolds for tissue engineering.

Fig. 1 Schematic sketch showing relationships of additive manufacturing processes in connection to various materials used for scaffold preparation in biomedical engineering.

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2 Scaffolds for tissue engineering The concept of tissue engineering was first outlined by Langer and Vacanti, as already mentioned, in 1993. It was recognized as an interdisciplinary area that is closely associated with various streams such as biological science, materials science, engineering, and medicine [13,14]. Tissue engineering is directly related to the evolution of restorable biological substitutes, preserve, or develop tissue’s function or an entire human organ that has been severely damaged [15]. The above-mentioned requirement can be met with the use of stem cells. Typically, stem cells are distinctive types of cells that don’t transform or modify to become any cells other than the host one. They exhibit remarkable prospects to expand into various other cell types in the human body during the early stage of life and growth. Keeping this in mind, the major technique to produce these structures is to be able to securely distribute these stem cells, and construct a physically and mechanically robust and steady structure so that these stem cells can grow and be sustained. Scaffolds in human anatomy can be structurally created by the effective use of tissue engineering in combination with regenerative medicine. The scaffolds thus fabricated can provide support or substitute for organs and organ systems that have been injured due to various reasons. The fundamental approach of tissue engineering could be described as follows: In a first step, the benefactor or donor cells and growth elements are seeded on a 3D scaffold that gives initial support, along with providing the desired framework for adhering the cells, multiply and transform. The entire process is maintained and cultured in an in vitro atmosphere in a bioreactor to encourage the growth of a fresh, healthy tissue matrix. In the last stage, the biomimetic structure is transplanted into the patient. Scaffolds generally show biodegradable characteristics because after the end of the specific task, newly grown tissues can obstruct their function [16,17]. In order to develop porous scaffolds for tissue engineering, various type of biomaterials are used, provided a production technology is readily available that can work with the existing biomaterial properties. Typically, polymeric biomaterials are a popular choice to create scaffolds because these materials have the ability to supply the structural support required for the attachment of cells and finally lead to tissue development. Among the polymer materials used as scaffolds, synthetic and natural polymers have been accepted as potential alternatives, owing to their vastness, variety of properties, and biocompatibility. The first biodegradable scaffold materials for clinical applications belong to natural polymers. This arises because of their superior ability to interact with different types of cells, and the deficiency of an immune reaction. In later years, synthetic polymers came into existence because they are cost effective and they permit improved functionality over their counterpart. However, they are susceptible to toxicity or the absence of an immune response. The synthetic polymers are primarily comprised of organic materials

Tissue engineering based on 3D and 4D printing

based on poly(L-lactic acid) (PLLA), poly(glycolic acid) (PGA), poly(caprolactone) (PCL), or poly(lactic-co-glycolic) acid (PLGA) for the formation of 3D structure-based scaffold materials [18,19]. Synthetic polymers are often mixed with natural polymers to eliminate unwanted effects in correlation with cell attachment, hydrophilicity, and biodegradability. Additionally, the functionalization of the scaffold surfaces with the help of particular ligands such as protein molecules often aids in boosting cellular responses. 3D scaffolds developed from synthetic and natural biomaterials yielding nanofiberlike features, hydrogels, and sintered microparticles have been extensively investigated [20,21] over the last decade. The primary role of these highly porous 3D scaffolds is to create an environment in the vicinity of the implant to restore the damaged or missing tissue. Among the various controlling factors in a 3D scaffold, porosity plays a major role for second-generation tissue engineering, which in turn indicates the need for cell infiltration and growth along with vascularization into the 3D pore network within the as-fabricated scaffold [22]. As already mentioned, porosity plays a decisive role because the cellular networks rely on interconnected pathways when there is no engineered blood supply. The overall effect arises owing to the diffusion of O
and various nutrients and unwanted products away from the porous scaffold. All these factors play a decisive role in nutrient supply, cell migration, and proliferation to the scaffold. They also enhance the unfilled surface area required for cell-scaffold bridging and interface with nearby tissues mimicking the native extracellular matrix (ECM) environment in the structure. It is imperative to understand that when pore size is reduced, the available scaffold surface area increases. The rise in exposed scaffold surface thus increases the propensity of scaffold ligands to develop strong bonds with cells and to interrelate with. Conversely, with too small a pore size, the migration of cells within the scaffold structure becomes challenging. In addition, the surfaceto-volume ratio of a scaffold determined by pore size distribution and its networking must not be too large because it deteriorates its mechanical strength [23]. In view of the above observations, the microarchitecture of the scaffolds should precisely be designed with factors favorable to cell viability and fostering tissue ingrowth. This tradeoff between pore size distribution and scaffold properties is one of the essential models of tissue engineering and therefore needs to be taken care of during the development of novel biomaterials. An ideal 3D scaffold material should be comprised of a biocompatible, biodegradable material and its mechanical properties must match as close as possible those of the host tissue within the implant. The scaffold is not anticipated to be considered a permanent implant. In reality, they should enable the host cells to attract an extracellular matrix (ECM) and over time, it should be replaced with that of the scaffold structure. From a clinical viewpoint, it is also expected that the scaffold structure is to be simply deployed into diverse shapes and sizes to permit in situ treatment of specific defects in patient organs and tissues.

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The following list offers the desired scaffold material and structural properties: Cell and tissue compatibility and biodegradability: • The byproducts should display no toxicity once they decompose from the implanted scaffold material. • The synergistic effect of natural and synthetic scaffolds can be logically used in order to control the degradation and thereby improve the biocompatibility in scaffold tissue engineering. • The scaffold material should assist the inbuilt host cells to generate their own extracellular matrix. Bioactivity • Next-generation scaffold biomaterials should actively interact with and connect to the host tissue to stimulate in vivo mechanisms of tissue regeneration. In this way, it could enable itself to self-healing mechanism thus leading to substitute of the scaffold via regenerating tissues. • They should be osteoconductive and osteoinductive in nature. • Biochemical signals and growth regulators arising from the excitement biological tissues in correlation with cell-adhesive ligands for cell attachment and differentiation. As an example, the synthesis of hydrogels by covalent bonding or ionic cross-linking can help in protein entrapment and thus discharge them by swelling of the hydrogels.

2.1 Scaffold architecture Scaffolds containing interconnected pores create a large surface area for inbuilt vascularization, promote the formation and growth of new tissues to facilitate cell migration, and subsequently host tissue unification after implantation. The scaffold biomaterials should be modified to tailor the pore size and distribution to target tissues and cells without significantly weakening its mechanical properties and/or affecting the mechanical stability of the scaffold. Lastly, the scaffold materials must be able to degrade with time, keeping the same pace of the new extracellular matrix by forming newly formed tissue, once the implantation is done.

2.2 Mechanical properties Scaffolds with comparable compressive, elastic, and fatigue strength provide strength and stability to the host tissue. The organs and tissues mechanobiologically mimic the scaffold and therefore must maintain their structural integrity under in vivo conditions.

3 Biomaterials for tissue engineering and scaffold fabrication A vast range of materials has been explored for use in the prospective fabrication of scaffolds in tissue engineering. Typically, they can be classified into three subgroups: natural

Tissue engineering based on 3D and 4D printing

polymers, synthetic polymers, bioceramics, and composites. In the polymeric materials group, the materials are realistically fabricated by integrating with individual functional groups within its molecular backbone to determine its chemical, physical, and biological properties. Each of these various classes of biomaterials offers desired benefits and drawbacks. Therefore, it is logical to utilize composite scaffolds consisting of diverse phases, an idea that is becoming progressively popular in this field of application.

3.1 Natural polymers Research on polymer biomaterials has been a subject of interest both in academia and industry for a span of 60 years or so. Biological materials are used as scaffold biomaterials for the synthesis of natural polymers. Typical examples of natural polymer-based biomaterials include collagen, a range of proteoglycans, alginate-based substructures, and chitosan. All these substances have been tried and used in the fabrication of tissue-engineered scaffolds. Collagen is the most abundant and naturally occurring polymeric biomaterial. Some other commonly listed natural polymers investigated for this purpose include polysaccharides (chitin, chitosan, hyaluronic acid, etc.), silk fibroin, fibrin, alginate, gelatin, fibronectin, etc. They have been found to be extremely effective in tissue engineering, as they have the capacity to be restructured in in vivo conditions. Natural polymeric biomaterials have enormous potential to form scaffolds that retain the extracellular matrix composition of the host tissue. Although natural polymers demonstrate outstanding bioactivity and biodegradability for their abundant use in soft tissue engineering, their poor mechanical strength inhibits them for load-bearing applications. In addition, natural polymers also have insufficient usage where absolute support to injured or healing tissues is a requisite, owing to their fast degradation rate. Keeping these factors in mind, cross-linking of these materials is compulsory in the fabrication of long-standing tissue supports. Cross-linking is the process of attaching one polymer chain to another. Regardless of the inherent structural advantage arising from cross-linking strategies, several cross-linking elements can modify tissues by means of several methods when scaffolds are less expected to be populated by native cells. Lastly, biomaterials based on natural polymers also have the tendency to create xenogenic difficulties because the majority of these result from animal components.

3.2 Synthetic polymers Synthetic polymers, as their name suggests, are produced in the laboratory or in industrial units in the course of a chain of chemical reactions from low molecular weight organic compounds. Numerous synthetic polymers have been tried to produce scaffolds and they are grossly separated into two groups: biodegradable and nonbiodegradeable.

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Biodegradable polymers include polystyrene, poly-L-lactic acid (PLLA), polyglycolic acid (PGA), polylactide and its copolymer poly-lactic-co-glycolic acid (PLGA), polyphosphazene, polyanhydride, poly(propylene fumarate), polycaprolactone, polyurethane etc. Among the widely used nonbiodegradeable polymers are polyvinyl alcohol (PVA), polyhydroxyethymethacrylate (pHEMA), poly (N-isopropylacrylamide) (PNIPA), polymethyl methacrylate (PMMA), etc. The main benefit of this form of polymeric scaffold substitute arises from its controlled and superior physicochemical properties (e.g., porosity, degradability, and mechanical strength), lack of immunogenicity, and ease and abundance of processing owing to various artificial synthesis techniques. On the other hand, there is a possibility of denial of the artificial polymeric scaffolds from target tissues due to their inferior bioactivity. It has been found that acidic byproducts are readily formed while using synthetic polymer scaffolds during the entire degradation process, which subsequently leads to lowering the local pH. This eventually leads to its diminishing strength and results in cell-tissue necrosis.

3.3 Bioceramics Bioceramics are a special type of inorganic and nonmetallic ceramic and could be of both natural or synthetic origin. They are primarily aimed at the restoration and regeneration of affected parts arising from injury or trauma in the musculoskeletal structure as well as for periodontal irregularities. Even though soft tissue regeneration is not vastly promoted by bioceramics, there has been extensive use of these ceramic materials/scaffolds for loadbearing orthopedic applications (hip acetabular cups coatings), bone grafts/cements, and in dentistry [24]. Bioceramics are usually characterized by their high mechanical stiffness (Young’s modulus), corrosion resistance, and a hard and wear-resistant surface. Moreover, from a bone tissue point of view, they exhibit excellent osteoconductivity and biocompatibility, which arises due to the chemical and structural resemblance to the mineral phase of bone or osseous tissue. It is worth mentioning here that the mutual interactions between osteogenic cells and bioceramics are vital for bone regeneration because ceramics are identified as augmenting osteoblast differentiation and proliferation [25,26]. In spite of these numerous advantages, they suffer from severe drawbacks, including inadequate fracture toughness (leading to brittleness), low elastic properties, and exceptionally high stiffness [27]. All these actuate the limited clinical applications of these materials for tissue engineering (TE). Calcium phosphate bioceramics offer a special interest in tissue engineering practices owing to their close likeness with bone and teeth, which arises from the chemical similarity with these hard mammalian tissues. CaP-based bioceramics exhibit excellent biological performance, including osteoconductivity and bioresorbability. The material properties thus assist in integration into existing tissues by a similar process found in bone remodeling. In addition, calcium phosphate-based

Tissue engineering based on 3D and 4D printing

materials are cost effective from a fabrication point of view, and their medical grade certification is simple to achieve. Besides, the success of CaP-based bioceramics is also achieved in some extent to skin and muscle tissue replacement. However, these special categories of bioceramics suffer from poor mechanical properties such as strength, fracture toughness, and fatigue resistance, required for load-bearing applications in the biomedical field. Among the calcium phosphate-based bioceramics, the major focus has been applied to hydroxyapatite (HA), α- and β-TCP, and biphasic CaPs in the biomedical field concerning hard tissue, owing to their structural similarity with implants and bone defects [28]. HA [hexagonal, stoichiometric Ca/P ratio of 1.67 Ca10(PO4)6(OH)2] is crystalline in nature and is a very stable compound with negligible solubility among CaPs when tried in a solution below pH 4.2 [28]. Contrary to that, β-TCP is a high-temperature entity of calcium phosphate compounds that is derived by thermal breakdown at temperatures at least above 800°C. The biodegradable nature of β-TCP has often been tried in bone substitute applications in the form of granules, blocks, or in CaP-based bone cements [29]. Researchers have found that the biological absorption capability of both these CaP-based ceramics substantially differs, even if these species are analogous in terms of chemical composition. Hydroxyapatite shows slow bioresorption dynamics and hence, mostly integrates itself into the newly formed bone tissue once the implantation is done. On the contrary, β-TCP is entirely reabsorbed [30] in the bone tissue.

3.4 Metal-based scaffold materials In situations where human bones are damaged or need to be substituted, porous metallic scaffolds have been a popular material of choice, owing to their outstanding physical properties in addition to their ability to support tissue ingrowth. Out of the numerous alloys available, titanium (Ti) and tantalum (Ta)-based metals and alloys have been proven to be the best candidate material for this purpose. Medical-grade Ti alloys perform much better in bone tissue ingrowth capability when compared to stainless steel due to the 50% greater strength/weight ratio of the former. Its elastic modulus (105 GPa) is similar to that of human cortical bone (7–21 GPa). The metal/alloys also exhibit excellent corrosion resistance [31]. In addition to cast Ti alloys, porous titanium (Ti) scaffolds have also been studied as bone replacement materials. These components of biomaterials are not biodegradable and do not assimilate with biomolecules. Among the Ti-based alloys, Ti-6Al-4V in specific is broadly used in a variety of orthopedic applications because it enjoys excellent biocompatibility and improved mechanical properties over conventional stainless steel, Co-based alloys, and even pure titanium. Recently a 15-year-old boy diagnosed with cancer received a perfect implant manufactured by the German 3D printing technology giant, EOS Technology. The entire

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Fig. 2 Titanium implant fabricated by 3D printing technology for partial replacement in hip joints [32]. (Copyright 2019. Reproduced with permission from Electro optical systems (EOS) GmbH. https://www.eos. info/press/case_study/additive_manufactured_hip_implant.)

process of hip replacement, starting from the initial CAD design to the optimum implant, required only six weeks [32]. Fig. 2 shows the 3D-printed prototype structure fabricated by EOS Technology. Although several advantages have been recorded in this development, the element vanadium in the Ti-6Al-4V based alloy in isolated form still possesses cytotoxic effects, which has motivated researchers to develop novel β-Ti alloys with nontoxic elements such as Nb, Zr, and Ta [33]. On the other hand, porous metallic scaffolds are considered to be the desired material for implants in conjunction with hard tissue engineering in load-bearing applications. Their superior fatigue resistance [34] in addition to favorable compressive strength are found to be extremely effective for load-bearing applications such as the femur, vertebra, skull, and hip and knee joint replacements. These porous configurations also show similar and consistent mechanical properties to that of human bone. In addition, they assist in improving osteoblast adhesion, proliferation, and differentiation [35]. In a different architectural approach, Xue et al. fabricated Ti scaffolds with porosity ranging from 17 to 58 vol.% and with average pore sizes of 800 μm [35]. On a different note, the surface modification of Ti and its alloys has been revealed to enhance osteoconductivity, as demonstrated by Das et al. [36]. Titanium dioxide (TiO2) nanotubes have been grown onto porous Ti scaffolds via a chemical method called anodization. The overall advantage is to increase the apatite formation tendency of these scaffolds in simulated body fluid (SBF). In spite of these above-mentioned advantages, metallic scaffolds suffer from the following limitations: (i) poor biological response on the material surface or bioactivity is

Tissue engineering based on 3D and 4D printing

by far the major weakness of metallic scaffolds, (ii) biomolecules cannot be chemically integrated within the metallic scaffolds, (iii) in general, the biodegradability of metallic scaffolds is extremely negligible or not observed at all, (iv) the possibility of slow discharge of toxic metal ions/particles naturally or through corrosion or wear is another severe concern, and (v) it is not easy to hold/sustain the architecture of a porous metallic scaffold. Apart from the permanent bio-implanted metals, biodegradable metals for implant applications have shown potential for fracture fixation where entire tissue regeneration is likely. Presently, allows based on iron (Fe), magnesium (Mg), and zinc (Zn) are found to be the ones best suitable as biodegradable metals, especially for orthopedic and cardiovascular applications [37], because they exhibit excellent in vivo biocompatibility, a lesser biodegradation profile, and adequate mechanical properties to mimic bone for the period of regeneration. In principle, bioresorbable metals/alloys offer mechanical properties as compared to bioresorbable polymers such as polylactide (PLA), polyglycolide (PGA), or the polylactic-glycolic acid (PLGA) copolymer, owing to their brittleness and strength comparable to that of implants [37,38].

3.5 Biocomposites The above-mentioned single-phase scaffolds/biomaterials suffer from various practical difficulties that limit their applications as advanced biomaterials where the host tissue should be partially or fully replaced in a damaged or diseased organ. Therefore, significant research is being dedicated toward biocomposite scaffolds that consist of numerous fillers or reinforced materials in the matrix phase in the nano- and microscale as an alternative and viable solution. For example, in a particular case, fibrous-like biomimetic scaffold structures are of growing interest, both in the field of advanced soft-tissue engineering and also for state-of-the-art bone tissue engineering applications. In each of the asprepared biocomposite scaffolds, they are formed by at least one part that is not a naturally occurring component of our body. However, they suffer from the usual biomedical problems associated with biocompatibility, biodegradability, or both. Several materials have been studied for the fabrication of 3D biocomposite scaffolds for tissue engineering. The primary characteristics for the composite scaffolds include slow degradation and favorable cell biocompatibility as well as noncytotoxicity, nonantigenic, nonimmunogenic, and nonmutagenic actions. It is observed that about 30% of the presently used biomaterials are of the composite category [39]. Among the various combinations, polymers, when mixed with ceramics, result in the most popular variety of composite scaffold. Polymeric scaffolds show insufficient mechanical properties, for example, tensile strength, elastic modulus, and toughness. Inferior bioactivity is another crucial issue often found in polymer scaffolds. These attributes can be removed by controlled inclusion of ceramic material fillers, such as hydroxyapatite (HA) or tricalcium phosphate

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(α- or β-TCP). The synergistic association of softer polymeric material with stronger and biomimetic ceramic accelerates the regeneration of tissues [40]. In a list of naturally occurring polymers commonly applicable in tissue engineering dealing with hard tissues such as bone, collagen is one of the most extensively used scaffold materials. Because the human bone matrix is 90%–95% composed of elastic collagen fibers, it is a natural alternative for fabrication in a composite bone tissue scaffold. A collagen-HA-based composite scaffold has been fabricated by Villa et al. via a simple coprecipitation and freeze-casting method [41]. The composite scaffold exhibits a large extent of permeability appropriate for cell infiltration, attachment, and osteogenesis arising from its 99% interconnective pore network. Interestingly, the study found that while the in vivo testing is done for three weeks for the scaffolds implanted into a mouse calvarial defect, the defects are found to be near complete filling when compared to pure HA. On the contrary, pure HA scaffolds have not shown any positive outcomes and eventually decayed six years after being placed inside four ailing patients suffering from long bone defects [42]. Apart from collagen, some other commonly used natural polymers such as chitosan [43], chitin [44], alginate [45], and silk [46] have been successfully combined with HA for replacements in bone tissue engineering. Nevertheless, the difficulty that arises from using these natural polymeric materials is their poor mechanical properties [47]. As a result, researchers have made numerous attempts to avoid the potential weaknesses of natural polymers concerning mechanical strength by combining HA with biodegradable synthetic polymers, for instance poly (lactic acid) [48], poly (ε-caprolactone) (PCL) [49], poly (lacticco-glycolic acid) [47], and poly (D,L-lactide) [50]. Zhang et al. [51] showed the osteoconductive characteristics and cell maturation behavior of nano-HA/PCL spiral scaffolds that were synthesized with different HA/PCL weight percentages with the introduction of a controlled amount of porosity. They found that among the diverse nano-HA/PCL spiral scaffolds, the optimal HA/PCL compositional ratio is on the order of 1∶4 (weighted average) for bone tissue regeneration. Hermenean et al. [52] prepared novel 3D chitosan (CHT) scaffolds reinforced with graphene oxide (GO) to investigate osteogenic differentiation for regenerating bone tissue in critical-sized mouse calvarial defects. These composite materials show promise as implants by revealing more of the as-grown bone in the chitosan/GO-substituted defects in comparison to chitosan alone. In combination with GO, an increase in alkaline phosphatase activity equally in in vitro and in vivo trials has been observed by the addition of chitosan in these composite scaffolds.

4 Direct 3D-printing processes The 3D printing technique normally uses materials in various forms [53] under atmospheric conditions, such as fluids with the ability to solidify, flexible filaments, layered or laminated thin sheets, and small powder particles.

Tissue engineering based on 3D and 4D printing

A particular method utilizes a specific form of material to prepare the scaffold. However, if a particular material is prepared in a desired form for 3D printing, it does not ensure that the substance is 3D printable because in order to qualify for printing in the upright direction, it is also crucial to impart the bonding strength in the interlayer of the scaffold material. Consequently, the prime aspects while designing an object for 3D scaffold fabrication is based on the existing types of the material at the initial stages. Moreover, with the aim of adding to the collection of 3D printable bioscaffolds, novel and productive methods should be invented in the future to convert the present class of biomaterials into an appropriate form of feed material to be 3D printable. For instance, gelatin gel solidifies once its temperature is reduced, but working in this reduced temperature-based atmosphere does not favor the effective growth of cells. This paves the way for developing newer methods and mechanisms that involve the ease of solidification of gelatin, for instance enzymatic cross-linking [54], or a novel hybrid method for the growth of hydrogels and cells at much lower temperatures [55]. Rapid prototyping technology has evolved during the last 20–25 years, and consequently a range of diverse 3D printing and additive manufacturing techniques has been tried for medical applications. The most frequently used 3D/4D printing techniques include extrusion printing/bioprinting, stereolithography (SLA), powder deposition printing (FDM), laser-assisted printing (SLS or SLM), and direct ink writing or inkjet bioprinting (DIW, etc.). Fig. 3 shows the schematic representation of the operating principles of three major types of 3D printing techniques used in biomedical engineering. In addition, a comprehensive review stating the basic features with corresponding advantages/limitations and applications is provided in Table 1.

4.1 Stereolithography (SLA) The stereolithography (SLA) 3D printing technique is based on using photosensitive polymers (also called photopolymers) as the feedstock, thereby polymerizing the feedstock solution into a specified pattern. A UV light bulb or laser light is accurately projected onto this photopolymer and a controlled reaction is achieved by digital micromirrors. In this technique, a layer-by-layer process is adopted. Each successive layer must be entirely cured or dried before the second layer adjacent to it can be deposited to obtain the desired 3D printed object. Among the various additive manufacturing processes, stereolithography is one of the oldest techniques applied in bone tissue engineering. This sophisticated method allows very high precision of creating scaffolds for the fabrication of complex structures ranging from the micro- to the nanometer dimension. Other techniques such as extrusion-based processes deal with micrometer scales of higher orders of magnitude [68]. This technique offers the unique advantage of producing sophisticated scaffold architectures with internal complexities and exceptionally high resolution (1.2 microns) [68]. In addition, the high quality of printing, the promptness, and the cell proliferation and differentiation are

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Fig. 3 Schematic representation of three major types of 3D/4D printing technologies: (A) stereolithography (SLA), (B) fused deposition modeling (FDM), and (C) selective laser sintering (SLS).

added advantages that have made the SLA technique a well-accepted process to fabricate bone tissue engineered structures. When a 3D object is constructed by the SLA-based 3D printing technology, additional measures must be adopted to enhance the mechanical properties of the final replaceable structure as per the specific host organ requirement. Therefore, polishing and eliminating the undesirable scaffold structures that remain as an attachment must be performed precisely. The method is carefully designed by fabricating a multilayered 3D heterogeneous architecture with graded mechanical and biocompatible features by controlling the premeasured solution that is accessible to UV or other forms of light from one layer to another. It is thus possible to construct a 3D scaffold extending in both the perpendicular and horizontal directions [69,70].

Table 1 A brief review of common 3D printing techniques for biomedical applications Technique

Advantage

Disadvantage

References

Inkjet printing

• Vast range of biomaterials can be printed • Structural complexities do not require

• Potentially toxic • Low mechanical strength compared

[56]

any additional support

• High concentration of cells included in • • • • • Direct ink writing (DIW)

• • • •

the scaffold Significantly low cost Superior printing speed Creation of composition gradient is easy Multiple solution compositions can be coprinted Bioactive composites can be simultaneously printed Low viscosity material can be printed Use of hydrogels is easily realized Simplicity Multiple inks can be utilized

• • • • • • • • •

Bioplotting

• Prints viable cells • Soft tissue applications

• •

Fused deposition modeling (FDM)

• Lower cytotoxicity content than direct

•

3D printing • Relatively inexpensive

• • • •

[57]

[56]

[58]

Continued

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•

to SLS Time-taking process High setup cost Limited material choice As-printed cells may be affected by piezoelectric printers Continuous printing process is not feasible Vertical structures exhibit inferior activity Cell density is lower Not very suitable for complex operations Combination of thickening and thinning agents in bioink is too crucial Desired microstructure is difficult to be realized Nozzle size limitation Requires support structure for complex designs Limited material processing (often requires thermoplastics) Nonbiodegradable material used Additional support structure should be provided for shape complexity Postprocessing required Little resolution

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Technique

Advantage

Disadvantage

References

Selective laser sintering

• Scaffolds with enhanced mechanical

• It must resist shrinkage and heat

[59,60]

• • • Stereo lithography

• • • • • • • • •

Laser-assisted bioprinting (LAB)

property produced Complex structure can be incorporated with the help of a powder bed Fine and high resolution Biomaterial deposition in solid or liquid phase Very high resolution Fast fabrication Smooth surface finish Complicated internal architectures can be printed easily Can easily print components that are released in an outward direction Shear forces are excluded on print stock Printing time independent of complexity High accuracy Nozzle-free technique

• High-precision printing at ambient conditions • Single cell patterns within scaffold • Diverse bioactive materials are used. • Several different solutions printed at a time

• • • • • • • • • • • • • • •

effects. Very high temp required (up to 1400° C) Expensive and time consuming Thermal damage can happen during processing Photopolymers are generally used Expensive External assistance is required for overhang and objects with complicated design Involves photocross-linkable polymer Poor mechanical strength Limited resolution Mechanical properties that are horizontally graded are inconvenient to produce UV blue light is toxic to cells Deficiency of printing multicells Damages cell during photo curing Costly Limited height of scaffolds

[61]

[56]

3D and 4D printing of polymer nanocomposite materials

Table 1 A brief review of common 3D printing techniques for biomedical applications—cont’d

Powder fusion printing (PFP)

Extrusion printing

• A vast range of materials (metals, polymers, etc.) • Excellent mechanical strength • Complex geometries can be printed • Properties vary in vertical directions

• Precise control over printing conditions • Wide variety of materials • Can print physical and compositional gradients

• Can directly print cells and bioactive factors Vat photopolymerization

• • • • •

can be used Bioactive materials can be incorporated High resolution Cells can be incorporated Raw material base is of solid polymers High resolution

• • •

powder microstructure Horizontal property gradients are inconvenient to produce Multiple fusion steps can generate cracks Build time, material usage, and other factors could increase to provide support to overhanging components Cannot print single cell Only applicable for viscous liquids Slow speed

• • • • •

Limited materials UV source necessary Near UV blue light’s toxicity Damage to cells during photocuring Lack of multicomponent cells

• • •

• High cost • Thermal damage due to high tem-

[60]

[62,63]

[64–67]

[64–67]

perature during deposition Sheet lamination

Indirect 3D printing

• Layered laminate structure is formed • HA (hydroxyapatite), zirconia, osteoblast-like cell, human osteoprogenitor cell, and human umbilical vein endothelial cell can be formed • Shows promises for prototyping/ preproduction • Versatile materials

• Only layered laminates are formed • Requires further processing

[64–67]

• Dedicated waxes are prime requisites

[56]

for biocompatibility • Poor accuracies/resolution • Mold should be provided for casting • Fabrication time is more

Tissue engineering based on 3D and 4D printing

Directed energy deposition

• Ability to print high cell densities • Raw material base is a fluid material • Huge range of photocurable polymers

• Microfractures/voids arise from

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Earlier, this process of fabricating a layered and compositionally variable scaffold structure for a specific requirement was an uphill task for researchers and engineers. This was because the pouring solution basin must be unfilled first and then restocked or substituted with another solution in each successive layer. This dilemma has been successfully overcome by separate groups of researchers [69,71,72]. These groups of scientists have discovered a novel design approach in the SLA instrument by providing an automatic replacing unit that potentially permits layered and heterogeneous biomaterials for scaffold fabrication. However, several weaknesses have been uncovered following the application of this technique. One of the major drawbacks arises from the application of resins that are found to exhibit carcinogenic effects and could therefore be life threatening. Another common problem is that the resins are photosensitive in nature, therefore leads long term stability of the as-fabricated scaffold structure. The extremely slow production rate (1–3 cmh
1) of the 3D printed object is yet another challenging issue in SLA techniques [73]. The commonly used ultraviolet source of light for the polymerization process creates another threat because reports show that this light source damages our DNA cells and could be a potential reason for skin cancer [74,75]. To resolve this matter, visible light has been replaced as an alternative source of light in SLA-based bioprinting systems found substantial acceptance in this field. Recently a group of researchers at the University of British Columbia has successfully demonstrated a custom-made bioprinting unit comprised of a beam projector with combinations of PEGDA, GelMA, and erosin Y made photoinitiator as the injected bioink [76]. Poly(D,L-lactide)-based resin or a poly(D,L-lactide-co-3caprolactone)-based resin has been used as the raw material to fabricate porous scaffolds. The scaffold object’s mechanical properties and stability can be optimized by controlling the pore architecture and polymer compositions. A photocross-linkable PCL-based resin has been successfully fabricated by Elomaa et al. with the help of high gel-containing networks [77]. Similarly, porous scaffolds have been developed by using the resin, inhibitor, and dye. This type of as-fabricated scaffold conforms well to the targeted object to be used and demonstrates the acceptability of the resin for construction of tissue engineering scaffolds. In recent years, a considerable number of novel and promising biodegradable resins have been increasingly used. In this aspect, various materials have been developed as promising alternatives for application in the SLA process, and the list includes poly (caprolactone) [77], poly(D,L-lactide) (PDLLA) [78,79], and poly (propylene fumarate)-diethyl fumarate (PPF-DEF) [80]. In another attempt, Winder et al. [81] developed a self-modulated cranial titanium (Ti) prosthesis directly synthesized by SLA-based resin, thus making the process amenable for use on an industrial scale. In general, the SLA technique is used for tissue engineering applications related to blood vessels, cartilage, muscle-neuron coculture, etc. From futuristic point of view it is essential to develop various novel biocompatible and biodegradable photocurable polymers for successful implementation of this

Tissue engineering based on 3D and 4D printing

technique. At the same time, in order to include the list of polymeric materials, it is crucial to design and develop visible light-based STA systems in the near future.

4.2 Microextrusion-based 3D bioprinting Extrusion-based techniques have been extensively tried as an alternative method for scaffold fabrication, owing to their simplicity, diversity, and predictability. This extrusion-based method makes use of mechanical (piston or screw) or pneumatic forces to supply bioink by means of a nozzle or needles connected to cartridges loaded with ink. The subsequent micropatterning is followed by a computer-generated design. Here, filaments of the organic biomaterial are dropped layer by layer from a μ-extrusion top in 2D, and the stage or the μ-extrusion head shifts down the z axis. In the μ-extrusion printer, several cartridges can be attached to the instrument for printing of heterogeneous structures. Here, cells are mixed together with bioink. Bioink is the material that is used to envelop cells to allow an environment to extracellular matrix (ECM) and protect cells from the disturbances that are experienced by a cell in the course of printing. Inkjet printing is broadly used for 3D printing of cell-laden assembles because it can afford excellent cell viability in contrast to μ-extrusion printing. Nevertheless, the printing of viscous or sticky bioinks is reasonably difficult in most polymeric ensembles. This motivates researchers to utilize μ-extrusion printing to print glutinous bioinks. The viscosity observed in popularly used bioinks printable via μ-extrusion printing falls in the range of 30–6 107 mPas [82]. In order to meet the complexity requirements of various internal organs of our anatomy with substantially high resolution, bioinks should display a shear thinning ability to perform microextrusion via a needle. Also, rapid gelling characteristics should also be experienced to permit a layered deposition and preserve the printed profile [83,84]. On the other hand, the improvement in gelling property and the viscoelastic propensity of the bioink-based materials might contribute to the shear stress increment. This would then heavily affect the cells in the typical μ-extrusion-based printing process. All these factors contribute to the premature death of the targeted cells and damage the μ-extrusion-based cell printing practice [82]. Keeping these factors in mind, it is imperative to understand that a favorable range of bioink viscoelasticity is by far a crucial factor to estimate while fabricating through the μ-extrusion-based printing process. In this aspect, Markstedt and coworkers investigated the rheological functions and printing capability of nanocellulose-alginate-based composite bioink in conjunction with human chondrocytes. They found that the survival rate of these cells was 73% [85]. In another attempt, Kesti et al. studied a combination of poly(N-isopropylacrylamide), hyaluronan, and methacrylated hyaluronan as the prospective bioink material that is entirely derived from its rheological properties, swelling activities, printability, and prior biocompatibility results [86]. Another group of researchers [87] demonstrated the use of polypeptide-DNA hydrogel-based multilayer 3D μ-extrusion-based printing for studying its healing properties and superior mechanical strength as a new-age bioink substitute material.

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The μ-extrusion-based bioprinting tools have been successfully used to create heterogeneous scaffolds required for osteochondral regeneration. Recently, Esfahani et al. [88] applied this method of printing to produce glass-ceramic scaffolds that had adequate strength and toughness for curing of bone defects under a static load. The hexagonal interconnected porous architecture allows the formulation of a widespread surface near or within the vicinity of the contacts surrounded by printed layers and a superior load transfer potential while evaluating with other usual patterns (rectangular, zigzag, or arched). In addition, the state-of-the-art design established superior fatigue resistance, failure reliability, and flexural strength under in vivo conditions. This method of scaffold preparation for hard tissues might create opportunities for handling bone defects in orthopedics in relation to load-bearing applications as well as dental and maxillofacial replacements. In μ-extrusion-based bioprinting, the bioinks fall in both the Newtonian and non-Newtonian categories and have been vastly modified for enhanced printable viscosity. In the case of non-Newtonian fluids, the printable viscosity of the bioinks can be optimized by the strain rate variation during the entire process of printing, which also relies on its concentration and molecular weight. Besides the shape of the orifice, its size and the active temperature (in case of temperature-sensitive bioinks) also affect the strain rate during the entire printing process. Bioinks are typically controlled by shear thinning effects and the viscosity reduces with the enhancement in strain rate and acts to safeguard cells in addition to improving its resolution. The process of shear thinning restricts the entanglement of chains consisting of organic molecules arising from the sliding of chains over one another. This further helps in the smooth extrusion of viscous bioink through the opening of the nozzle. Based on this phenomenon, researchers worldwide have tried to improve the shear thinning properties of viscous bioinks used in μ-extrusion-based bioprinting [89]. While evaluating with inkjet bioprinting, μ-extrusion-based bioprinting provides denser cell growth but at the loss of speed and resolution [90]. In addition, a broad range of biomaterials can be printed with the help of this relatively cost-effective bioprinting instrument, including tissue spheroids, tissue strands, cell pellets, decellularized matrix parts, and cell-laden hydrogels. In a simple but practical alternative, many researchers have tailored the available commercial 3D printing units to print scaffold objects or have fabricated their own printing equipment internally to ease costs [91–93].

4.3 Inkjet based 3D-printing Inkjet printing is an extremely popular bioprinting technique used these days. 3D printing via inkjet is sometimes considered an alternative process for sterolithography [94]. However, the 3D bioprinting via inkjet method has found restricted applications when matched to extrusion-based processes. The prime reason behind this is the printing head moves

Tissue engineering based on 3D and 4D printing

to supply a continuous flow of the printing solution, which confines its relevance in bioprinting. Bioprinting is a type of additive manufacturing process. Moreover, it is an advanced technique to biofabricate 3D structures for functional tissues by arranging cellular components in a 3D space so as to mimic the human tissue function. This process can efficiently fabricate 3D structures as biomaterial scaffolds for tissue regeneration by embedding cells over the 3D-designed scaffold that relies on the computer-aided design model and the process having advantages as it uses bioink by encapsulating cells to prepare the 3D structure [95]. 3D inkjet printing is known for its ability to construct intricate scaffolds in various clinical sizes for biological applications. Furthermore, the developed products are mobile and easy to carry. Synthetic polymers for poly(e-caprolactone)PCL, polyglycolic acid, poly(lactic acid), etc., are mostly used as the biological ink [96]. Walczak et al. first designed and fabricated a microfluidic channel chip for capillary gel electrophoresis. The time taken to construct the biochip was around 3 h, which is the beauty of an inkjet printing process as compared to a conventional process [97]. Krivec et al. constructed a 3D printed prototype package with the adoption of a photopolymer. Additionally, they incorporated silver nanoparticle ink for radio-frequency identification to make better wireless radar signaling [98]. Kyobula et al. used a 3D thermal inkjet printing technique that is solvent-free to make a drug release carrier through natural derivatives. The architecture of the manufactured drug was designed in a honeycomb structure with a controlled cell size. The intention of the writers was to develop drug release tablets through a 3D inkjet printing technique by using beeswax in the shape of the honeycomb structure, which could then be clinically used in a solid dosage form [99]. Zhang et al. applied an inkjet (ij) 3D printing method to develop hydroxyapatite-based scaffolds along with controlled porosity in the micro/macro level. Moreover, mechanical properties in the form of the compressive strength of the 3D-printed bioceramic scaffolds with different diameters of HA nanopowders have been shown [100]. Phillippi et al. developed pattern growth factors by using inkjet printing. In their study, the authors used adult stem cells to modify the cell fate by stem cell differentiation. The biological approach has been applied to osteogenic BMP-2 by immobilizing growth factor. In regenerative medicine, it has the ultimate potential application [101]. Gao et al. developed a bioprinted scaffold and showed excellent biocompatibility. The bioink was composed of poly (ethylene glycol) and PBS, followed by mixing with human mesenchymal stem cells (hMSCs). The 3D inkjet printed polymer scaffold resulted in good cell viability with superior mechanical properties. Moreover, the construction time was reduced because of less process complexity. Hewes et al. developed microvessels with bioink comprising fibrinogen and alginate to fabricate freestanding microvessels through an inkjet 3D printer. The alginate-based solution has been used to bioprint those microvessels. With the use of the alginate template method, it was possible to load cells within the fibrin scaffold via the 3D inkjet printing method [102]. According to Cui et al. [103], there might be a chance of damaging of cells in inkjet

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3D printing process. The authors also have studied 3D printed Chinese hamster ovary (CHO) cells and evaluated cell viability. In addition to that, the authors have checked a possible number of the damaged cell membrane and was found this thermal inkjet printing method is an effective and reliable way to print mammalian cells [103]. Owczarczak et al. [104] used the application of a thermal inkjet printer to incorporate g-actin monomers into 3T3 fibroblast cells. The technique used by them does not harm cells and injected cells together with molecules. Moreover, the method exhibits superior cell viability. Inzana et al. optimized the mechanical property as well as the biocompatibility with the aid of an inkjet 3D-printing approach. The fabrication of collagen-calcium phosphates was optimized by blending a phosphoric acid binder solution with collagen. This enhanced the cell compatibility for better bone regeneration [105]. Xu et al. fabricated a 3D cellular structure in combination with fibrin and NT2 cells with a layer-by-layer approach to form a neural cell sheet, which has a wide function in tissue engineering [106].

4.4 Fused deposition modeling (FDM) FDM is one of the most widely practiced filament extrusion-based additive manufacturing processes to develop low cost complex 3D objects for bioengineering applications. This technique is also alternatively regarded as the fused filament fabrication (FFF) process and was first available for commercial used in 1991 [107]. In this technique, a thermoplastic filament or a small bead polymer is extruded from a small nozzle in a movable platform at a measured rate. An extra heater is used in the nozzle to first melt or soften the filament. A fan is provided as an extra attachment at the nozzle end for controlling the solidification and subsequent deposition rate of the as-fabricated solid 3D object in a layer-by-layer manner. The process parameters that play crucial roles in the successful fabrication of the 3D construct with predetermined pore size, porous network morphology, and internal linkages include raster diameter, gap width, the angle of the raster, etc. [108]. All these factors make FDM a simple process for fabricating 3D complex structures as compared to other traditional techniques with a much more complex approach, such as lithography and micromachining. However, FDM suffers from some critical issues such as a longer operation time and lower resolution as compared to other commonly used 3D printing techniques. Moreover, the application of a higher deposition temperature for thermoplastic materials, in a range varying from 120°C to 300°C, is not comfortable for inserting cells or drugs within the filament while preparing the 3D scaffolds. In addition, this method of 3D printing for scaffold fabrication allows only a limited range of thermoplastic polymers such as polylactic acid (PLA), polycapro-lactone (PCL), poly (lactic-co-glycolic acid) (PLGA), acrylonitrile butadiene styrene (ABS), and polylactic acid (PLA) for commercial applications. Some of these polymeric substances are frequently used in the FDM process in combination with biomaterials to produce tissue-engineered scaffolds with a low melting point.

Tissue engineering based on 3D and 4D printing

Among these, PCL is the best material of choice that has been extensively used in dental applications and wound repair because it offers high printability and biocompatibility with replaceable tissues. For the fabrication of scaffolds in bone regenerationbased applications, FDM has been found to be very effective, owing to its favorable mechanical strength and physicochemical properties that closely match human bone structure [109–111]. The mechanical and biomedical performances of natural or synthetic biopolymers are modified by doping biocompatible reinforced units such as HA and β-tricalcium phosphate (β-TCP) [112] or by simple coating [113,114]. Korpela et al. developed PCL/bioactive glass (BG)-based composite scaffolds by this method. They demonstrate superior properties equally in compressive modulus and the required biocompatibility [115]. Oladapo et al. [116] developed a new hybrid scaffold structure with a poly lactic acid (PLA) matrix supplemented with carbohydrate particles (cHA) in different proportions (PLA:cHA ¼ 10/0, 95/5, 90/10, and 80/20) for replacement in bone tissue. The hybrid composite mixed with an 80:20 proportion offers better results than the others. The pore size variation and porosity distribution network of the 3D-printed complex architectural scaffolds created by computer-aided design and subsequent layer-by-layer printing and solidification are shown in Fig. 4.

Fig. 4 Schematic illustration of bone formation by the PLA/cHA-based hybrid scaffold by the FDM technique, (A) optical images of a two-layer scaffold produced by 3D printing, (B) 3D structure showing tetragonal symmetry in successive layers, (C) radial architecture of the 3D-printed structure with different layers filled in concentric patterns, (D) SEM image of 3D-printed scaffolds with consistent 1 mm fiber spacing [116]. (Copyright 2019. Reproduced with permission from Elsevier. https://doi.org/10.1016/j.compositesb.2018.09.065.)

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4.5 Selective laser sintering (SLS) SLS is a widely used laser-assisted additive manufacturing (LAAM) technique developed by Carl Deckard in the 1980s at the University of Texas and subsequently patented in 1989. Selective laser sintering (SLS) employs a high-power CO2 laser to melt thin layers of polymer or ceramic powders to construct structures and replicas for parts from 3D CAD models, 3D digitizing system-acquired data, CT scans, and MRI scan records. The anticipated layers of deposited material are chemically bonded mutually throughout the SLS process to produce a specific shape, as depicted by the CAD-based 3D model. In SLS, the powder deposition system usually contains a set of rollers or a scraper that permits the instrument to deposit consecutive powders layer by layer with a thickness range of 20–150 μm, forming solid 3D objects. The entire fabrication operation in the SLS process is most often performed under an inert atmosphere (e.g., Ar, N2) to produce a contamination-free 3D object and also to avoid the unwanted oxidation of feed granular powder particles during the process. SLS often suffers from the deposition of layered objects with low-grade surface contours and substantial fluctuations in various dimensions (X, Y, or Z) of the as-fabricated parts. Therefore, as a secondary alternative, costly postprocessing surface treatments such as machining, heat treatment, polishing, etc. [117] are required. Like any other AM process, here also the final phase aggregate and porous microstructures of the composite scaffolds can be determined by controlling the essential processing parameters, including the laser power, the temperature of the platform, the laser scan speed, etc. [118]. Additionally, the size and shape of the powder granules are identified as deciding factors in the SLS process [119]. These parameters have shown their efficiency on the densification behavior and flowability of powder granules. The flowness of the powder particles is an important aspect because these powders must be dispersed evenly at elevated temperatures and with a thickness of 100 μm. All these factors lead to the fact that the powder particles used in SLS should possess a specific granulometry and a superior sphericity. In SLS, to fabricate the 3D structure with interlayer compositional differences, the successive base composed of the powder particles must be cleaned and a new biomaterial substance should be added. This deposition method is slow and time consuming while also contaminating the new surface layer. Heterogeneous mixing of the powder particles creates associated problems with the as-fabricated scaffolds. Some researchers [120] have revealed an alternative approach by presetting the powder layers and physically adding each deposit one at a time. Their study is successful for the synthesis of PCL- and PCL/Hap-based combinations of scaffolds with varying layers of stoichiometric HAp in order to resemble the graded configuration of the osteochondral unit. The SLS process has been successfully used to fabricate biomaterial scaffolds in various forms, including biopolymers, bioceramics, biocomposites, and even metallic components, for potential tissue regeneration applications. It is already established that biocompatibility, bioactivity, and biodegradability are essential properties for the

Tissue engineering based on 3D and 4D printing

development of tissue engineered scaffold constructs. However, the available materials that include the above properties are restricted enough for this purpose. Nevertheless, research efforts have been made for TE 3D scaffolds for biomedical applications. Some examples that have been successfully used in the SLS process to create complex structures in bone and other hard tissue engineered scaffolds include poly(ε) caprolactone (PCL) [121,122], poly-L-lactide (PLLA) [123], poly-D-lactide (PDLA) [124], polyhydroxybutyrate (PHB) [125], and poly(3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV) [123]. On the other hand, nonbiodegradable polymers such as polyetheretherketone (PEEK) [126] and ultrahigh molecular weight polyethylene (UHMWPE) [127] have been effectively transformed in sintered 3D scaffolds; these have been investigated recently. Similarly, in order to obtain the advantages of both biodegradable polymers and bioactive ceramics, for example hydroxyapatite (HA) and β-tricalcium phosphate (β-TCP), in blended TE composites, HA/PCL, HA/poly(L-lactide-co-glycolide) (PLGA), and β-TCP/PLGA [128,129] have been fabricated by the SLS technique. Researchers have implanted nano-HA/PCL and PCL-based composite scaffolds developed by the SLS process into rabbit femur defects. Their study shows excellent biocompatibility and a positive indication of curing of bone defects [122]. In a recent attempt, researchers developed an advanced version of the SLS technique with superior features, known as surface selective laser sintering (SSLS), to resist biopolymer degradation and assist the absorption of cellular entities and bioactive agents [130,131].

5 Indirect 3D-printing processes As compared to the direct additive manufacturing techniques, 3D printing of sacrificial molds and indirect 3D printing are also versatile techniques that are now becoming well accepted for scaffold fabrication in tissue engineering. These approaches of 3D scaffold architecture design and fabrication have shown great promise for structures with uniform interconnectivity, desirable pore size, and multifaceted internal/external construction using a similar (compared to direct methods) and sophisticated CAD/CAM model. In this method of scaffold fabrication, the flexibility of scaffold fabrication is enhanced toward a broad range of biomaterials ranging from polymers to ceramics. In principle, this indirect 3D fabrication technique utilizes a sacrificial mold. The mold is a predesigned dummy structure that possesses the negative framework of the target scaffold architecture. The as-prepared mold is consequently used for solidifying the scaffold of the final biomaterial. This indirect 3D-printing method is applicable for the fabrication of both polymeric and ceramic scaffolds of various biomaterials. Ceramic scaffolds includes alumina [132], HA, etc., whereas polymeric scaffolds are fabricated by poly(ε-caprolactone) (PCL), poly(lactic-co-glycolic) acid (PLGA), poly(L-lactide), collagen, and chitosanalginates [133–136]. While fabricating a polymeric scaffold, indirect 3D printing

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techniques largely utilize a solvent-based design with temporary molds. The material to be injected in the mold cavity is prepared in the form of an organic solvent. This leads to an increase in scaffold preparation time as it requires more time for mixing the proposed material in the solvent and further to solidify it within the mold by solute-solvent reactions or thermal evaporation in ambience. From this standpoint, it must be mentioned that the nature of the organic solvent and the solution composition might influence the inner construction of the scaffold, for example, the size and morphology of the interconnected pore network. Consequently, these types of 3D structural/architectural anomalies arising from scaffold fabrication can result in unwanted mechanical properties and can lead to an undesirable effect on the mechanochemical performance of the tissue engineered configuration once the implantation is done in the patient. An indirect method of printing shows a unique advantage by creating a tailored scaffold that accurately replaces the host organ that is affected. In a study [137], scientists printed a customized scaffold that exactly matched with a human mandibular condyle. By using the indirect 3D printing technique, they produced a biomimetic scaffold with an appropriate combination of PCL and chitosan. The PCL was solution-treated in chloroform and subsequently a gelatin mold prepared from 3D printing was used to pour the solution. Upon drying, the 3D-printed gelatin mold was eliminated by inserting the scaffold in dH2O at 50°C for 6 h. Once the PCL scaffold was fabricated, a coating of thin apatite layer was applied to this scaffold in order to sustain the cell growth of the bone marrow stromal cells (BMSCs). In this way, the as-fabricated PCL apatite-coated scaffolds exhibited much better proliferation characteristics of BMSCs when compared with uncoated composite scaffolds. Natural polymers such as collagen or alginate have also been used for scaffold fabrication in the indirect 3D printing method. The prime advantage of these naturally occurring superior scaffolds arises from their potential ability to develop the hMSCs for a specific time (up to 4 weeks) in vitro with substantially lower or no cytotoxicity on the other grown cells [138].

6 4D printing for biomedical applications The fast development of AM technology and advances in 3D bioprinting to build complex biocompatible structures have motivated scientists and engineers all over the world to develop similar but more effective techniques in the fourth dimension. The revolution of 4D printing technology has been highly inspired by the development of 3D-printable “smart” materials. The scope of 4D printed objects was first successfully demonstrated by the MITs Self-Assembly Lab in 2013. Skylar Tibbits, the codirector of that lab, is also regarded as the pioneer of 4D printing technology [139]. In his 2013 TED Talk (TED stands for technology, entertainment, and design), Tibbits explained

Tissue engineering based on 3D and 4D printing

the phenomenon of shape transformation within the structural design of the material by self-motivation when exposed to a programmed stimulating factor, which includes osmotic pressure, heat effect, current, ultraviolet light, or other energy sources [139–142]. Earlier, it was found that one major weakness of 3D bioprinting is that it is primarily based on the initial condition of the object that is to be 3D converted, supposing that it is stationary and inert. In contrast to that, 4D bioprinting has materialized, where a “time” factor is incorporated with 3D bioprinting as the fourth or final dimension. However, in this classification, time does not indicates the actual time to finish the job, but rather it determines that the 3D-printed biocompatible scaffold or living cellular assembly can be made to “self-transform” over time after being printed. Fig. 5 shows the statistical documentation of publication on the topic of 4D printing or 3D printing + biomedical + FDM/SLS/Inkjet from the Scopus database from 2012 to January 2019.

Fig. 5 Statistical record (from 2012 to January 2019) of publications on the topic of 4D printing or 3D printing + biomedical + FDM/SLS/Inkjet etc. from the Scopus database: (A) publications, (B) publications in different countries/regions, and (C) list of total publications of various 3D printing techniques.

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Improving the functionality and/or performance-driven capabilities of the final tissue product [143,144] has always been a bottleneck in tissue engineering research. With the implementation of 3D printing of specially designed shape memory scaffolds, an enormous opportunity has been explored in regenerative medicine by the favorable integration of 3D printing and the shape memory effect that arises from the time-dependency factor. It is worth mentioning, therefore, that mechanical stimuli (e.g. strain pattern, perfusion) have a pronounced affect on restructuring specific cells [145,146]. These cells often respond to biological stimuli and eventually control cellular activities and the efficiencies of the regenerated tissue [147,148]. However, the popularly used biochemical devices in this application are complicated and costly while their range of application is insufficient for this purpose [149]. Also, materials that exhibit selfevolving structures that execute geometric folding, curling, expansion, and some other scheduled shape changes when given a mechanical stimulus, including shape memory polymers (SMPs), are being considered for this purpose [150]. A finite number of investigations have been executed by using shape memory polymers to show the consequence of a one-time mechanical stimulus on cells in contact with 2D sheets [150,151]. In the recent past, the world has seen immense growth in 4D printing. It is worth mentioning here that the requirement for the development of various soft active 3D printable biomaterials (SAMs), sophisticated 3D printing technologies that achieve fast fabrication of multiple active materials, and advanced computational methodologies to transform rational material models to envisage the 3D printed object for its futuristic expansion in designing novel biomedical devices and bio-inspired architectures, is very much essential. As already mentioned, 4D printing adds time to 3D printing and accordingly offers some additional benefits as compared to 3D. It is capable of developing intelligent instruments that assist in transforming the profile and functions of the smart as-printed structures of the bio-inspired materials. Second, the method helps to print thin-walled construction or lattice structures, and thus has the prospect to reduce the overall time of printing and materials. A recent study confirms that self-folding can speed up the rapid prototyping of 3D objects, possibly leading to a reduction of 60%–87% of printing time and expenditure of materials [152,153]. Apart from these notable advantages, the shape-changing phenomenon triggered by various stimuli can be wisely exploited to engage less space and transportation. For instance, shape memory polymers can be converted into a horizontal surface by 4D printing for simple posttreatment operations and moving as well for storage purposes. As far as the prospective materials are concerned, 4D printing technology uses a diverse range of materials aimed at developing devices/ objects such as smart devices, metamaterials, and origami for a variety of functional applications in the aerospace, biomedical, and biomedicine fields, among others [154,155]. These range from the micro- to the macroscale dimension. It has been found that while matching with smart metals and ceramics, smart polymeric materials can readily react upon exposure to diverse stimuli with substantial deformability [156]. Therefore, until now,

Tissue engineering based on 3D and 4D printing

4D printing has primarily focused on smart active polymer-based bio-inspired materials. The following two categories of materials are extensively studied as SAMs for 4D printing: shape memory polymers (SMPS) and hydrogels. In the following section, these two groups of materials, applied in 4D printing technology, will be discussed in detail.

6.1 Soft active shape memory polymers Shape memory polymers (SMPs) are emerging smart materials that have seen major advances in the past decades. This special class of intelligent materials possesses the distinctive capability to recover into a 3D printed stable “memorized” structure once exposed to an appropriate stimulus (mostly of thermal origin, although the chemical or electromagnetic type also affects). The shape memory phenomenon can be found in our everyday life, too. For example, heat shrinkable tubes, state-of-the-art medical parts, and self-deployable components in spaceships are some of the potential areas where the technology can be adopted. The shape memory effect in these special class polymers, as found, is not intrinsic, and the overall effect arises in combination with the polymeric material’s special shape memory property as well as with mathematical models to program the deformation. Initially, a thermoresponsive shape memory polymer is allowed to transform into a temporary shape and subsequently put beyond its transition temperature (Ttrans). Once the temperature is attained, the shape memory effects are produced and, consequently, a stable memorized structure is recovered [157]. In biomedical engineering, these materials exhibit enormous possibilities concerning a variety of applications, for example negligibly invasive implants (e.g., stents), self-knotting sutures, orthodontics, and drug delivery devices [158–160]. In comparison to metal/alloy-based or ceramic-based shape memory materials, polymer-based shape memory materials have numerous benefits that exceed the two above-mentioned types of materials. These include good shape recoverability (capable of 400% stain recovery), low density, processing flexibility, control of properties (e.g., Tg, stiffness, biodegradability, and functionally grading capacity), deformation in a controlled and programmed manner, and cost effectiveness. The above-mentioned features of a typical SMP are mentioned as follows: SMPs have the following advantages over conventional materials:  Possess excellent recovery strain.  Lower density.  Cost-effective construction. However, these materials suffer from the following drawbacks as well:  Modulus of elasticity is low in comparison to other classes of materials (metals, ceramics).  Inferior strength.

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SMPs show widespread significance in the field of biomaterials and bioinspiration [161,162]. Polyurethane-based shape memory polymers achieve outstanding biocompatibility. Therefore, this can be exploited in various clinical components when placed in the human body [163,164]. Wache et al. [165] recently completed a viability study of SMPSs on sustainability on a polymer vascular stent in the drug delivery system. Significant improvements in restenosis and thrombosis have been noticed while using the SMP stent in drug delivery systems. Likewise, a polycaprolactone dimethacrylate-based shape memory tracheal stent has been fabricated by the use of an UV-LED digital light processing (DLP) printer [166]. The UV-LED light source used in this study has a wavelength of 405 nm. Stereolithography uses a laser source to fabricate the layered objects. A number of photocurable methacrylate-based copolymer networks were synthesized and printed by means of high-resolution (μm range) projection microstereolithography (PμSL) [167]. One special category of SMPs belongs to the biodegradable-based polymers, which might be fabricated from a few monomers and show a lot of promise in this field [168,169]. A typical example is the poly(ε-caprolactone) (PCL)-based biodegradable polymer, which has prospects in medical applications [170].

6.2 Hydrogel-based 4D printing Hydrogel-based 4D printing technique is primarily based on the integration of crosslinked networks of long polymer chains that bulge as a bigger volume after water or other solvent disperses into them. Upon immersion of these 3D-printed configurations in a solvent, the hydrogel structure swells at first, and thus with the generation of a strain mismatch among the two different substances, an overall shape change takes place [171–173]. Although hydrogels are easier to use, they are too soft to be transformed into rigid configurations of the tissue engineered matrix. In addition to this, the quick degradation and the overall integrity of the printed construct possesses other drawbacks of these materials. Typically, these materials are utilized in 4D printing because of their superior toughness and volume changeability when suffering environmental alterations. Some typical examples of hydrogels based on natural polymers are dextran, chitosan, and collagen while those based on synthetic polymers are poly (vinyl alcohol), poly(ethylene glycol) diacrylate (PEGDA; hydrophilic), poly(propylene glycol) dimethacrylate (PPGDMA), etc. The first applications of hydrogel-based materials in 4D printing were successfully demonstrated by Stratasys multimaterial polyjet 3D printing [174]. In this very first attempt, the researchers have practically shown that a straight strand changed into the script in “MIT” pattern in the fluidic atmosphere. The precise control of spatial distribution and the temporal arrangements of this self-assembly were recognized by the accurate printing of the hydrogel/elastomer-based smart hinges. Later, Bakarich and coworkers fabricated a hydrogel-based smart temperature-receptive valve. In order to acquire a smart valve with actuating temperature ranging from 20°C to 60°C, alginate/poly

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(N-isopropylacrylamide) (PNIPAAm) inks were synthesized. These special smart/intelligent valves clogged to decrease the flow rate while changing the length between 41% and 49% at 60°C. This occurs because the transition temperature of the smart valve falls in the range of 32–35°C [175]. Ikegami et al. [176] seeded human mesenchymal stem cells (hMSCs) onto a 3D bioprinted tissue assembly with a pattern framework. This enhances the transformation time of the tissue assembly from two days to two weeks, caused by matrix deposition and environmental restructuring from hMSCs. The resultant effect of the matrix deposition phenomenon from hMSCs is the formation of a highly robust grid-like pattern that could be handled easily. Overall, hydrogels loaded with bioactive multifarious materials, for example, drugs and antibodies, have shown promise as an area of promising research while they also show a great opportunity for artificial implants [177,178]. On the other hand, self-healing hydrogels can inherently and routinely cure damaged tissue and reestablish a normal state of organ functioning. These are in conjunction with other similar 4D printing technologies, and thus significantly inspire the study of self-healing hydrogels in tissue and organ regeneration applications.

7 Factors affecting 4D printing Smart or intelligent materials for biomedical applications are based on their unconventional activity to mimic the specific tissue or organ to be cured. These materials have the ability to change their biological and mechanochemical receptiveness to environmentally and naturally occurring stimuli, for instance, temperature, stress, humidity, ionic strength, pH, and electric or magnetic fields. Essentially, they are intended to be varied in a predetermined mode, and accordingly, stimulate the cells and their attachment, proliferation, and differentiation. In the following section, we will discuss the effect of two externally stimulating factors that decide the kind of smart materials to be used in the 4D-printed structure and their properties.

7.1 Effect of temperature The most frequent external stimulus to activate the 4D-printed constructs is temperature [178,179]. Thermoresponsive materials, including shape memory thermoplastics and shape memory thermosets, can fold, contract, or swell while changing their temperature. Several other categories of polymers demonstrate phase-transition temperatures, such as glass transition temperature (Tg) in thermosets or melting temperature (Tm) in thermoplastics similar to physiological temperature. This kind of behavior is found in poly (N-isopropy-lacrylamide) (PNIPAAm), which is a commonly used candidate material for drug delivery applications as well as tissue regeneration [180]. A bilayered object printed by PNIPAAm and a water-insoluble polymer poly(e-caprolactone) (PCL) exhibit the self-folding or self-unfolding characteristics once exposed to slightly elevated

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temperature; therefore, the phenomenon of yeast cell encapsulation and release is found in it [24]. In another attempt, a thermoresponsive combination involving poly(acrylic acid)-PNIPAAm and Al2O3-based ceramic powder can experience a sol-gel transition by the application of a thermal stimulus [181]. In a different study, Hendrikson et al. [182] prepared a polyurethane shape memory structure (Tg 32°C) to construct active scaffolds by the fused deposition method. A custom-made stretcher was used to print the scaffolds at 65°C, chilled down to 4°C, and released to retain a transition shape. Then, the injected cells were grown at 30°C to help in cell adhesion and proliferation. In the final step, the temperature was enhanced to 37°C and thus, a stable shape was achieved. In the entire process of dimensional recovery, the as-received cells were transformed in an elongated and aligned configuration. In general, the process involves the integration of a thermosensitive polymer with targeted cells, nutrients, or growth factors, and subsequently the whole combination is infused into the body. While being injected, the polymer experiences a phase shift owing to its temperature increment and transforms into a gel-like substance, releasing elements to form a 3D structure [183].

7.2 Effect of water or solvent Another frequent stimulus that drives the deformation (or swelling) in shape memory effects in bioprinted shape memory polymeric scaffolds is water (or solvents). Watersoluble or hydrophilic polymers typically display the SME when they come into contact with humidity. The final shape is obtained after drying. The glass transition temperature of hydrophilic SMPs may slash down to below room temperature once the water molecules are diffused into [184] it. A lower glass transition temperature (Tg) that is almost close to room temperature permits the printed objects to reveal the SME in this temperature range. In comparison to the thermally induced shape memory effect, it has been found that water-induced SME displays a slower shape recovery of active 3D printed objects. In spite of these drawbacks, including a slow reaction time, the humidity response of the SMPs for water-triggered actuation has been effectively used for various 3D printed architectures, especially in aquatic environments or at ambient temperature. For instance, Raviv et al. [185] have shown three classic examples of deformationlinear stretching, folding, and ring stretching-where water would react and develop various configurations with the help of multiple parts of the materials. In the linear stretching operation, the stiffer substance acts as a scaffold while the lively (hydrophilic) deposit creates the force for stroke. This scheme was accomplished by gathering a chain of firm disks while the center material stretches. Tibbits and other research group [186,187] have fabricated a sequence of primitives (or hinges) with a firm base of plastic material and a hydrophilic rubber (hydrogel) by the 4D printing technique that swelled while in contact with water with the use of a Connex Objet500 printer. Apart from the conventional

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fabrication of layered structures, biomimetic 4D printing of multifaceted, morphologically diverse and naturally inspired structures, for instance leaves and flowers that convert in contact to water, was executed with nanoinks [188]. During the process of extrusion of the ink out of the nozzle exit, a portion of the cellulose constituent was, as experienced in woodfiber extrusion [189], self-ordered in the hydrogel by shear stresses. This eventually resulted in swelling in a specific dimension of the extruded structure along the longitudinal direction. Zhu et al. [190] have shown a thermally effected shape memory effect for watersensitive smart active polymers based on cellulose nanowhiskers (CNWs) and thermoplastic polyurethane. Here, the reversible formation and breakage of the CNW percolation network has contributed to the shape memory effect when exposed to water.

8 Conclusions As a promising manufacturing technology of the 21st century, 3D printing, is an additive manufacturing-based process for exact 3D construction, revolutionize the world with its ability to make automated complex structures. It is a very exciting technique to develop functional devices such as laboratory chips and cell-laden scaffolds for tissue growth with respect to biomedical applications. Various efforts at developing novel biocompatible materials for bioprinting demonstrating fast cross-linking properties are fundamental requirements toward practical implementation of 3D printing technology in tissue engineering. The function and realistic potential of the 3D printing technique has not been fully utilized, owing to its lower speed and subsequent time factor for the manufacturing of 3D objects. On the other hand, 4D printing, based on the 3D printing fabrication scheme with the added time factor, has shown great interest and potential to the scientific community for practical purposes. In the future, this technique has the possibility to replace many existing methods and materials involved in developing smart, customized implantable medical devices that offer competent output for surgeons. Because it is quite possible to fabricate various medical models, the surgeon can create smart anatomies according to patient requirements any time, which has not been feasible until now. Multiresponsive biocompatible, topographical, and 4D dynamic shape-changing tissue scaffold configurations activated by diverse stimuli demand greater efforts and proficiency. Above all, the structures have to perform based on a particular application for tissue and organ regeneration matters, for example biodegradability and/or biocompatibility properties. In a nutshell, the 4D bioprinting technologies open up an excellent opportunity of an unexplored world of research with a wide range of devices of multifaceted printing, not only in tissue and organ regeneration but also in other fields in connection to scienctific understanding and technological development.

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Acknowledgment The author is grateful to Shashank Poddar, Ujjwal Chitnis, and Anish Ash for helping in the literature survey and preparing preliminary information for completing this chapter.

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