Progress in the Advancement of Porous Biopolymer Scaffold: Tissue

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Progress in the Advancement of Porous Biopolymer Scaffold: Tissue Engineering Application Rushikesh Ambekar, and Balasubramanian Kandasubramanian Ind. Eng. Chem. Res., Just Accepted Manuscript • DOI: 10.1021/acs.iecr.8b05334 • Publication Date (Web): 21 Mar 2019 Downloaded from http://pubs.acs.org on March 21, 2019

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Industrial & Engineering Chemistry Research

Progress in the Advancement of Porous Biopolymer Scaffold: Tissue Engineering Application Rushikesh S. Ambekara, Balasubramanian Kandasubramaniana* a

Rapid Prototype & Electrospinning Lab, Department of Metallurgical and Materials Engineering, DIAT (DU), Ministry of Defence, Girinagar, Pune-411025, India. *Corresponding Author E-mail: [email protected]

Abstract Tissue engineering is derivative of biomedical engineering which deals with the repair and regeneration of organs or their tissues. Cell embedded porous polymeric scaffolds unveil proficiency in rectifying mechanical trauma of skin, bone erosion and neurodegenerative diseases like spinal cord injury. Archetypically, pristine cell based biomaterial scaffolds are utilized, however, investigations over the periods have validated that incorporation of additives such as silver nanoparticles or bioactive glass augments anti-microbial characteristics in conjunction with cell adherence. Ideal scaffold should exhibit ease of processability, biocompatibility, biodegradability, non-cytotoxicity and excellent mechanical property, these properties can be achieved with biodegradable polymers. Researchers have extensively explored copious advanced as well as conventional fabrication technologies such as electrospinning, 3D and 4D bioprinting, etc. for contouring of porous polymeric scaffolds in tissue engineering functions like skin, bone, liver, cardiac, and neural tissue regeneration, which we have consolidated and discussed in the presented review. Keywords: Tissue Engineering; Biopolymer; Biodegradability; Cell proliferation; Scaffold

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1. Introduction Tissue loss keeps huge burden on the medical and health care department of the country, for example, only America has expended $400 billion/year to tackle the tissue loss cases1. Most of the times the tissue loss due to mechanical trauma, infectious diseases, tumours and fire accidents causes tissue loss in civilian whereas on field soldier due to blast, tactical operations and survelience1,2. It is always not possible to arrange autologous tissue or donor in the posttraumatic accidents for damaged tissues, here polymer scaffold can play a vital role in saving the patient life by restoring the tissue and maintaining the environment suitable for accelerated healing of tissues3. Nowadays, it is possible to transplant the organ or tissue from person to person via biopsy. The present technology of tissue engineering has an advantage over traditional technology such as immunological infection, decreases in susceptibility towards infections etc. due to beneficial characteristics of scaffold-based approach making the tissue engineering a promising solution1,3. Therefore in recent years researcher focused on the materials (as shown in Figure 1) which are highly compatible with tissue as well as human body and also useful for delivery of tissue and promotes the cell adhesion and proliferation increasing the life expectancy with safe treatment. Advanced therapy medicinal products (ATMPs)4–6 come under complex regulations of cell-based medicine since they are responsible for crucial characterization of new pharmaceutical products called guidelines of the cell-based medicinal products (CBMP)7,8. This characterization consists of seeding of different organ cells in the scaffold and then examine the growth of the cells and cell viability these factors determine the quality of the scaffold for safe medicinal practices9.


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Figure 1. Graphical representation of literature data depicting number of publications per year



Tissue engineering is an interdisciplinary branch of science which combines the biologist, chemist and materials scientist to work together for an understanding of polymer-cell interaction, the effect of scaffold on human body and design of the scaffold for higher seeding efficiency2. This interdisciplinary branch consists of proliferation of stem cells in the scaffold in the simulated body fluid medium to confirm the scaffold compatibility in the human body. Initially, the lower amount of tissue cells are tested to study their behaviour on the scaffold2,3. Biopsy cells are collected from a cell bank and cultured on the porous polymer scaffold. The growth factor and cytokine also helps in the enhancement of the cells activity. The basic concept of the tissue engineering as illustrated in Figure 2 where, the whole cycle consists of extraction of the cell from the body, cell biopsy, monolayer cell culture, proliferated cells, scaffold fabrication, cell culture and differentiation, degradation of scaffold and remodeling of extracellular matrix2. In ancient times organ implantation and regeneration was considered as a myth however, later they mentioned the generation of Prometheus in Greek mythology. In 1510-1590, The book Dix livres de la chirurgie of Ambroise Pare` mentioned the reconstruction of the nose, teeth and various body parts10. In 1792-1847 Johann Friedrich Dieffenbach a German surgeon developed the skin which can be grafted to the skin human body, where he confirmed the compatibility of the synthetic skin by implanting in animals. Jolly and C.A. Ljunggren were the first researchers who tested the in vitro cultured cells2,3,10 but the tissue engineering word was first mentioned in the year 1987 and W.T. Green gave their notable contribution in the field of cartilage and bone scaffold via chondrocyte culture method2,3.


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Figure 2. Basic principle procedures for tissue engineering Tissue engineering scaffold can be fabricated from various kinds of materials like metals, ceramics and polymers, out of which metallic scaffold are used as they possess osteoinductivity in the application of bone tissue engineering11–14 however they are non-biodegradable. On the other hand, ceramic scaffold are considered to be rigid, limited biodegradability and poor processability15, whereas polymers scaffold, in contrast, have excellent processability, tunable degradation rate, mechanical properties and wettability making the polymer scaffold potential candidate for skin, bone, liver, cardiac, neural tissue regeneration as well as for bladder regeneration also16. The scaffolds are the essential component of the tissue engineering which is made up of biomaterial having porosity, as the scaffold is a temporary biodegradable framework for organ



and tissue regeneration. The scaffold mimics the extracellular matrix and pores helps them to cell adhesion and proliferation17–20 and simultaneously growth factors are also incorporated in the scaffold as they are essential for cell regeneration21,22. The tissue engineering is the restoration of damaged tissue with the help of cell culture23,24. The scaffold can be of biodegradable and nonbiodegradable materials wherein, we have discussed the biodegradable scaffold in this review. The biopolymers are the best-suited material scaffold as they possess excellent biocompatibility and biodegradability. Most widely used biopolymer for tissue engineering include chitosan25–29, poly(lactic-co-glycolic acid) [PLGA]30–34, poly(caprolactone) [PCL]35–39, poly(lactic

acid)

[PLA]40–44,

poly(urethane)

[PU]45–49,

poly(3-hydroxybutyrate-co-3-

hydroxyvalerate) [PHBV]50 and alginate51. A biological molecule embedded fabrication techniques are rapid prototyping, electrospinning, thermally induced phase separation and solvent casting/particulate leaching52–55. Apart from tissue engineering polymer scaffold are also widely used for wound dressing application56–61 and drug delivery62,63. The reinforced scaffold can definitely be helpful for enhancing properties such as antibacterial, multi-drug resistance and water diffusion to utilize in tissue engineering application. Silver nanoparticles can be easily reinforced in scaffold due to their high surface to volume ratio, ease of diffusion across the membrane that is facilitated by the smaller size of silver nanoparticles and it is also popular for antibacterial activity64 and cytotoxicity65. Carbon nanofibers reinforced scaffold exert enhanced water diffusion and compression property66 and additionally, it also promotes cardiomyocyte growth67 and neural regeneration68 in tissue engineering application. Graphene oxide reinforced scaffold provides biocompatibility, antibacterial effect51, cytotoxicity69 along with accelerated water diffusion70 and mechanical properties e.g. tensile strength71, compressive modulus72 and storage modulus73.


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High porosity of polymeric scaffold helps in the cell adhesion and cell proliferation, on the contrary, polymer scaffold is fragile hydrated at 37˚C (body temperature)74. There are copious methods available to enhance the properties of polymer scaffold such as block copolymer (where the wettability of the polymer scaffold can be manipulated by the optimized hydrophilic and hydrophobic domains)75, Crosslinking density (where the water sorption increases with crosslinking content if crosslinking agent has hydrophilic functional group. Higher crosslinking density stabilizes the pore geometry while processing)76–79. Mechanical properties and dynamic swelling behaviour are manipulated by the encapsulating cells in the interpenetrating polymer networks80. Cell adhesion and proliferation depends on the wettability of the scaffold and it can be enhanced by plasma treatment81–83. Sol-gel method utilized for fabrication of non-biodegradable scaffold possess high water uptake and good mechanical properties84,85. This review discusses the conventional scaffolds fabricated by methodologies like solvent casting/particulate leaching in succession with freeze drying, gas foaming, thermally induced phase separation, electrospinning and advanced technique like 3D printing, 4D bioprinting, electrohydrodynamic jet printing, melt electrospinning writing which provides the shape memory effect and has enhanced cell adhesion and proliferation capability. Further, we have focused the extensive application of the porous scaffold in tissue engineering e.g. Skin, bone, cardiac, liver, neural tissue engineering. Owing to the high surface area of nanomaterials enhances their bioactivity by the interaction with biological molecules such as a cell, growth factors.



2. 3-D Fabrication techniques for polymeric scaffold An attempt is taken to fabricate porous scaffold to achieve one of the important property to accomplish interconnected pore which is later useful in cell adhesion and proliferation. Polymer scaffold has various topological formulations such as film, sponges, meshes, fibers and foams. The selection of scaffold fabrication method is critical owing to method determines the bulk as well as surface properties of the polymer scaffold. The selected method has an impact on the properties such as biodegradability, specific surface area, tensile strength, porosity etc. The methods used for the fabrication of a conventional scaffold are depicted in the Figure 3. 2.1 Conventional fabrication Methodology


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Figure 3. Conventional fabrication methodology for polymer scaffold, a) Solvent casting/ particulate leaching. Reprinted with permission from ref86. Copyright 2003 Elsevier, b) Freeze drying. Reprinted with permission from ref87. Copyright 2016 Elsevier, c) Gas foaming. Reprinted with permission from ref88. Copyright 2006 Elsevier, d) Thermally induced phase separation. Reprinted with permission from ref89. Copyright 2014 Elsevier, e) Electrospinning. Reprinted with permission from ref90. Copyright 2014 ACS Also, apart from biodegradable polymers, scaffolds can be developed by nonbiodegradable biopolymers in polymerisation in solution wherein Gómez Ribelles et al. fabricated polymethyl methacrylate (PMMA) sponge via polymerisation in solution. PMMA/ethanol with different proportion was tried along with benzoin as a photoinitiator and ethylene glycol dimethacrylate as crosslinker. They found that PMMA/ethanol (20:80) was the best ratio for properties desired for scaffold. The increases in concentration of ethanol in polymerisation produces smaller PMMA microsphere with more porosity and narrow diameter



dispersion91. Mar Llorens-Gámez et al. studied porous PMMA prepared via polymerisation in solution method where the authors elucidate that higher cross-linker concentration (5 wt%) provides greater porosity of 78%. Water sorption study shows that increase in porosity helps in higher water sorption due to enhancement in specific surface area76. Gómez Ribelles et al. reported Porous poly(2-hydroxyethyl acrylate) Hydrogels fabricated via polymerisation in solution technique. Microscopic study revealed that the as-fabricated hydrogels had 0.2 µm pore size with interconnected pore structure92. 2.1.1

Solvent casting and particulate leaching In solvent casting and particulate leaching method, first step is solvent casting in which

polymer solution (a mixture of polymer and solvent) is casted in the mould. There are two types of solvent casting method53,93: 1) Pouring polymer solution in the mould and provide sufficient time to evaporate the solvent and then the solid product is detached from the mould. 2) Polymer solution is poured in the mould wherein mould wall acts as a membrane that absorbs the solvent and enhances the solidification of the polymer. The absorbed solvent is extracted from the membrane and it can be reused several times with vacuum drying treatment93,94. Subsequently, in particulate leaching step (as shown in Figure 4) to acquire porous scaffold, where it utilize the leaching agent called porogens e.g. polymer microspheres95, salt, sugar or wax96. These porogens are mixed with polymer matrix and moulded in the desired shape and then to create pores, this moulded product is treated with porogen-soluble solvent and care should be taken that matrix should be stable while treatment. Size of porogen and geometry determines the size of pore and geometry as well as the higher concentration of porogens can be used to get greater pore density but on the contrary interconnected pores cannot be controlled in this method93,97.


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Hou et al. fabricated the poly(d,l-lactide), poly(Ɛ-caprolactone) and polyethylene oxide/polybutylene terephthalate (PEOT/PBT) polymeric scaffold for tissue engineering. Porous scaffold was achieved by the addition of leaching agent salt and then eliminated salts via soaking in the demineralised water for 4-5 days. The porosity of the scaffold was altered with a particle size of leaching agent and polymer to salt ratio86. Ishaug et al. reported the poly(DL-lactic-coglycolic acid) foam for bone tissue regeneration. The as-prepared porous scaffold was prepared by a mixture of polymer and NaCl salt composite in which polymer is a backing material and NaCl acted as a sacrificing material, where the porous scaffold was recovered from the composite mixture by wetting in ethanol for 24 h. The osteoblast cell proliferation study shows that 11.8 x 105 cells/cm2 cell density after 24 h of cell culture out of the 22.1 x 105 cells/cm2 to the 300-500 µm foam resulting in 53 ± 1% cell attachment whereas post-cell culture density of cells was 4.63 x 105 cells/cm2 in 24 h out of 6.83 x 105 cells/cm2 to the 300-500 µm foam resulting in 68 ± 5% cell attachment. These results confirm that cell attachment percentage was higher in lower cell density than initial higher cell density98.

Figure 4. Solvent casting and particulate leaching 2.1.2 Freeze drying



Freeze drying is extensively used fabrication technique (as illustrated in Figure 5) for cavernous scaffold based on sublimation of solvent. The polymer solution (a mixture of polymer and solvent) is poured into the mould for attaining desired shape and then expose the mould at low temperature to solidify the whole system, and then the frozen product is dried under vacuum results in the sublimation of solvent crystals. The pore size mainly depends on pH and freeing rate of the scaffold, this rapid freezing creates small pores in the scaffold99. Unidirectional solidification can facilitate the homogeneous well-mannered pore size and this technique does not require washing step to remove leaching agent, this freeze drying technique is also called Lyophilization100. Nanda et al. prepared PLGA microbeads in two particle size 19.4 ± 1.6 and 4.4 ± 0.9 µm for controlled and sustained release of insulin drug. The particle size of the microbeads not affected on the encapsulation efficiency of the insulin as the 19.4 µm and 4.4 µm beads has 86.98 ± 2.01% and 85.07 ± 2.75% respectively. The release study shows that presence of collagen helps in controlled release of the insulin as compared to pristine PLGA microbeads, therefore, cumulative release of the collagen/PLGA scaffold (65% in 28 days) has slower release than Pristine PLGA scaffold (95% in 28 days) but collagen affect oppositely in case of degradation as the pristine PLGA (18% in 28 days) degrade faster than collagen/PLGA scaffold (25% in 28 days)101. Fereshteh et al. successfully utilized freeze-drying method for fabrication of porous PCL/zein-tetracycline hydrochloride (TCH) scaffolds targeting bone defect treatment wherein they observed that Porosity of aligned scaffold as well as disordered scaffold declined with increasing concentration of PCL/zein. The porosity of disordered scaffold was always greater than aligned scaffold at constant PCL/zein concentration. On the contrary, compression strength of the aligned scaffold (12 MPa) was higher than disordered scaffold (6 MPa) for 15


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wt% PCL/zein concentration. Drug release study showed that PCL/zein scaffold (62% in 24 h) had slower drug release compared to pristine PCL scaffold (92% in 24 h) 87.

Figure 5. Freeze drying



2.1.3 Gas foaming The solvent evaporation is an important environmental concern, as well as residual solvent, is hazardous for cells. As a result sensitivity and structural integrity of the bioactive molecules such as growth factors, anti-microbial agent can be compromised. The gas foaming technique (as depicted in Figure 6) is performed at room temperature and most importantly without solvent102,103. This method required the high-pressure carbon dioxide gas to achieve high porosity and care should be taken that this gas is not reactive with scaffold material. Pressure is the only single parameter to tune the porosity. There is no control on interconnected pore although few times possible due to aggregates of the carbon dioxide molecules. To control the porosity, porogen like NaCl, sugar can be utilized and at last, polymer density decline with time of exposure by forming porous polymeric scaffold93. Kim et al. studied Poly(lactide-coglycolide)/hydroxyapatite composite scaffold for bone tissue engineering via Gas foaming and particulate leaching. The mixture of PLGA/HA/NaCl (1:1:9) was poured into the mould and kept under pressure of 2000 psi for 1 min. This sample was then exposed to CO2 gas at high pressure (800 psi) for 48 h, these results in the growth of CO2 pores then NaCl was also taken out via particulate leaching method. Osteoblast cell culture study shows that Gas foaming/particulate leaching scaffold (66.5%) has higher cell adhesion percentage than Solvent casting/particulate leaching scaffold (62%) 88. Harris et al. also studied the similar Poly(lactide-co-glycolide) with salt only for tissue regeneration. The as-prepared mixture was exposed to CO2 with high pressure around 800 psi, followed particulate leaching method. The gross porosity of scaffold increased


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with increasing NaCl/PLGA ratio and compression modulus of GF/PL (289 ± 25 kPa) scaffold was higher than SC/PL scaffold (159 ± 130 kPa)104.

Figure 6. Gas foaming 2.1.4 Thermally Induced Phase Separation This technique utilizes the temperature to induce in two distinct phases in the polymer solution according to the concentration of polymer solution such as low polymer concentration and high polymer concentration105. Their mechanisms through which phase separation occurs are solid-liquid and liquid-liquid phase separation. In both the techniques low melting point polymer is solubilized at high temperature e.g. phenol, naphthalene, dimethylsulfoxide or 1,4-dioxane106. In liquid-liquid phase separation, the polymer solution is cooled lower than the freezing point of the solvent imparting thermodynamic instability that drives the phase separation. In the solidliquid phase separation, the de-mixing is induced by solvent solidification and the further solid solvent is removed by the dipping in the polymer non-solvent such as ethanol, water or water/ethanol mixture, then vacuum drying makes the scaffold solvent-free107,108 (as represented in Figure 7). Wei et al. investigated the nano-hydroxyapatite loaded poly(l-lactic acid) scaffold via thermally induced phase separation (TIPS) techniques for bone tissue engineering. The



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pristine PLLA scaffold had 4.3 MPa compression modulus and fabricated via solid-liquid phase separation whereas the addition of 30% nano-hydroxyapatite increases the compression modulus to 8.3 MPa. The protein absorption was higher in the scaffold containing a higher content of nano-hydroxyapatite109. Konstantinos et al. merge the two methods TIPS and electrospinning and studied the effect of good solvent (dichloromethane), partial solvent (formic acid) and bad solvent (Dimethyl sulfoxide) on the surface properties of PCL electrospun fibers110. Lou et al. reported β-tricalcium phosphate (TCP) loaded poly(l-lactic acid) (PLLA) scaffold via TIPS and particulate leaching technique for bone regeneration. The compressive modulus was increased 4 times with the addition of β-tricalcium phosphate e.g. pristine PLLA had 0.18 ± 0.03 MPa compression modulus whereas 40 wt% TCP added PLLA scaffold had 0.84 ± 0.07 MPa

89.

Maquet et al. prepared porous poly(a-hydroxyacid)/bioglasss composite scaffolds via TIPS method111 for bone tissue engineering application. Bioglass embedded poly-D,L-lactide scaffold (21 MPa) had greater compressive modulus than pristine poly-D,L-lactide scaffold (14 MPa) and similarly bioglass incorporated poly(lactide-co-glycolide) scaffold (27 MPa) had greater compressive modulus than pristine poly(lactide-co-glycolide) scaffold (10 MPa)112.


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Figure 7. Thermally Induced Phase Separation 2.1.5 Electrospinning Electrospinning is most widely used for filtration, protective clothing, sensors and biomedical field113,114. Biomedical field consists of applications such as organ implant, wound healing, wound dressing, drug delivery and tissue engineering115–117. Out of them, they are extensively used for tissue engineering owing to ease of processing, constant fiber diameter, tailor ability, and enormous surface area118,119. A high voltage is applied to the polymer solution causing electrostatic force development at the needle tip and front of the needle tip collect electrode is grounded, A huge potential difference between them causes the electrostatic force responsible for fiber spinning via Taylor cone113,120,121. The electrospinning set-up basically consists of three parts (as shown in Figure 8) high power voltage source (to activate the electrode), syringe with metallic needle (to carry the polymer solution), metallic collector or grounded electrode (to deposit spun fibers)122,123. The final properties of the spun fiber are dependent on the polymer solution, processing parameters and environmental factors124,125. There



are three types of electrospinning techniques, namely uniaxial, coaxial and triaxial electrospinning technique, out of them coaxial electrospinning technique is widely used for tissue engineering application, as it facilitates the transportation of the biologically active entities such as cell, growth factor without structural damage. Ma et al. prepared chitosan-grafted-aniline tetramer polymeric scaffold via Electrospinning for tissue engineering. The results were analyzed by varying aniline tetramer concentration from 5 wt % to 40 wt % and found that optimized conc. was only suitable for spinning nanofibers instead of lower conc. or higher conc. of AT. Cell viability study confirms that 5 wt %, 8 wt % and 10 wt % AT doped scaffold enhanced the C2C12 myoblasts cell proliferation than pristine chitosan scaffold126. Bao et al. investigated poly(D, L-lactide-co-trimethylene carbonate) via Electrospinning for bone tissue engineering. DLLA:TMC ratios tune the fiber diameter 1526 ± 120 nm to 682 ± 146 nm for 5:5 and 9:1 respectively90.


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Figure 8. Electrospinning technique The advantages and disadvantages of conventional fabrication methods for tissue engineering are depicted in table 1.

Table 1. Advantages and limitations of conventional fabrication methods Method Solvent casting/ particulate

Advantages Simple method Controlled porosity and pore

Limitations Possibility of residual solvent and leaching agent


References 127,128


leaching Freeze drying

size Highly interconnected pores

Gas foaming

Good porosity control Free of toxic solvent

Thermally induced phase separation Electrospinning

Simple equipement Tailorable mechanical properties Continuous process Highly interconnected pores Huge surface area to volume ratio Both random and oriented fibers possible

Good isotropic property Small pore size Insufficient mechanical integrity Closed pore structure Inadequate pore interconnectivity Low productivity Complex procedure Low productivity Limited scaffold thickness

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100,129

88,104

105,130

90,126

2.2 Advanced Fabrication Methodology Advanced methods successfully utilized to overcome drawbacks of conventional fabrication technique such as poor pore size control, closed pores, solvent residue, and time intensive process. The advanced fabrication technique consist of 3D bioprinting, 4D bioprinting, Electrohydrodynamic jet and Melt electrospinning writing (as classified in Figure 9).


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Figure 9. Advanced methodologies for polymer scaffold, a) 3D Bioprinting. Reprinted with permission from ref 131. Copyright 2016 ACS, b) 4D printing. Reprinted from ref 132, c) Electrohydrodynamic jet printing. Reprinted with permission from ref 133. Copyright 2014 ACS, d) Melt electrospinning writing. Reprinted with permission from ref 134. Copyright 2013 Wiley. 2.2.1 3D Printing The 3D printing process is the printing of 3D product with the help of CAD files, 3D printing is also called additive manufacturing and this technology is called rapid prototype technology135. 3D printing technique is classified in 5 types such as Fused deposition modeling (FDM)136, Stereolithography (SLA), Selective laser sintering (SLS), Selective laser melting (SLM), polyjet printing and Bioprinting. Bioprinting is our interest as this process facilitates the printing of a biodegradable scaffold. In this 3D Bioprinting technique (as depicted in Figure 10), physical 3D object is created by depositing successive layer by layer approach guided by Computer-aided design in specific dimensions. This technique facilitates the printing of 3D scaffold and artificial organs for biomedical application and also has the capability to print biomaterials and living cells absence



of a solvent. Inkjet bioprinting, microextrusion bioprinting, laser-assisted bioprinting, multiphoton excitation (MPE) based fabrication are the sub-methods of the 3D bioprinting137,138. Inkjet bioprinting vapours the bubbles in the vicinity of the biomaterials and cells with the help of heat and then deposited by nozzle, but the nozzle clogging restricts the precision of the process, on the other hand, Laser-assisted bioprinting or laser-induced forward transfer (LIFT) utilizes a laser to deposit the cell loaded biomaterials on the substrate. Microextrusion bioprinting works with the highest automation with mechanical accessories like piston or screw or pneumatic systems. MPE based bioprinting cross-link the polymers as well as a protein with the photo energy139,140. The cell viability upto 300˚ C does not affects the cells due to short thermal exposure. Inkjet technique successfully utilized to print chinese hamster ovary and embryonic motoneuron cells at 300˚ C having > 90% cell viability as well as Thermal inkjet printing technology also used for printing chinese hamster ovary cells and had 89% printed cell viability141–143. The final diameter of filament in the FDM depends on nozzle diameter, but to achieve small nozzle diameter less than 100 μm, pressure required for uniform flow is practically not possible144. Whereas, in case of melt spinning high voltage allows to spun finer fiber and enhances the filament resolution. The collection speed of FDM having 100 μm fiber diameter is 1 m/min whereas, melt electrospinning is five times faster than FDM having 1 μm fiber diameter145.


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Figure 10. 3D Bioprinting technique



2.2.2 4D Bioprinting The 4D bioprinting has the same process as explained in the 3D bioprinting but it has time as a fourth dimension which is activated via stimulus such as temperature, light, water, pH, magnetic field146, osmolarity etc.147 4D bioprinting scaffold respond to the structural, cellular and dynamic changes of a tissue over a period of time, thus active nature of the 4D printed scaffold overcome the passive nature of the 3D scaffold148 and Utilization of smart polymeric materials in the 3D printing is called 4D printing. Miao et al. Developed the PCL triol scaffold crosslinked with hexamethylene diisocyanate (HD) and processed via 3D printing technique for tissue engineering. Differential scanning calorimetry reveals that as-prepared smart scaffold has glass transition temperature between -8˚C to 35˚C and it shows 92% shape fixing at -18˚C to 0˚C149. Wei et al. fabricated shape memory nanocomposite via 4D printing under UV irradiation for biomedical application. A printed ink consisting of semicrystalline thermoplastic PLA (c-PLA) and Fe3O4 were dissolved in dichloromethane and crosslinked with benzophenone under UV irradiation. The different geometry of as-printed structure shows variation in % of recovery such as waviness printed structure (97.5%) exert highest % of recovery compared to spiral printed structure (91.53%) and flower printed structure (90.52%). The c-PLA/Fe3O4 composite (71˚C) had higher glass transition temperature than pristine c-PLA composite (66˚C)150. 2.2.3 Electrohydrodynamic jet printing The combination of electrohydrodynamic jet and bio-printing technique is called Electrohydrodynamic direct jetting technique. In this technique, cell solution is filled in syringe that passes through needle on the substrate where, needle acts as a primary electrode and substrate as secondary electrode and the potential difference between primary and secondary electrode drives the polymer cell solution to come out of the needle151. The electric field is


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connected to three axis moving system and ethanol is used as a target material. After printing, fabricated scaffold are washed with water and then kept for freeze drying152,153. Kim et al. fabricated

PCL/collagen

fabric

scaffold

for

tissue

engineering

application

via

electrohydrodynamics direct printing method with core/shell nozzle. Mechanical test shows that young’s modulus of 5 wt%, 10 wt%, 15 wt% was 14 MPa, 9 MPa and 5 MPa respectively. The average diameter of as-prepared fibers was 4.8 ± 0.8 µm and had 91-96% porosity. Although PCL exerted excellent mechanical strength but it lacks in cell attachment due to its hydrophobicity therefore collagen was coated on as-prepared fibers using oxygen plasma treatment. Cell number of post-plasma treated scaffold was increased from 20 x 104 (day 1) to 48 x 104 (day 7)154. Kim et al. prepared 3D fibrous cellulose scaffold via electrohydrodynamic jet process for tissue engineering. The fiber diameter of cellulose scaffold (1.7 µm) was finer than control scaffold (8.7 µm) and on the contrary, water uptake capacity of control scaffold (1.0) was lower than cellulose scaffold (4.0). Author also reported that cell seeding efficiency of cellulose scaffold (52%) was greater than PCL scaffold (22%) similarly cell number also in case of cellulose scaffold (680 cells/mm2) was greater than PCL scaffold (300 cells/mm2)155. Kim et al. reported poly(ε-caprolactone) micro/nanofibers scaffold as a tissue regeneration materials prepared via electrohydrodynamic jet process. Cell proliferation study shows that cell number of fibrous scaffold (20 cells/mm2) was enhanced compared to controlled scaffold (80 cells/mm2) within 4 h and in context of rate of proliferation, fibrous scaffold (4 cells/mm2h) had faster cell proliferation than controlled scaffold (2 cells/mm2h)133. 2.2.4 Melt electrospinning writing The amalgamation of melt electrospinning and additive manufacturing technique is called melt electrospinning writing technique. In this technique, melt electrospinning set-up is



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connected to translating x-y stage. Polymer pellets were poured in the syringe and melted using heaters, melted polymer is spun from the needle due to the potential difference created between needle and ground translating plate. Translating stage displaces according to program and desired structure of polymer fibers on ground plate is achieved156. Hutmacher et al. fabricated PCL scaffold via melt electrospinning writing for tissue engineering applications. Microscopic ananlysis revealed that average fiber diameter of scaffold was increased with increasing flow rate of polymer solution as well as higher voltage (12 kV) also aided in achieving finner fiber diameter than lower voltage (4 kV)157. Groll et al. prepared poly(2-ethyl-2-oxazoline) fibrous scaffold through melt electrospinning. Fiber diameter of scaffold increases with increasing temperature from 200˚C (40 μm) to 220˚C (140 μm) as well as feeding pressure also affects the fiber diameter such as 1 bar feeding pressure (60 μm) provides finer fiber diameter compared to 3 bar feeding pressure (90 μm). Applied voltage from 4 to 7 kV has not shown much fluctuation (75-90 μm) in fiber diameter158. Chiellini et al. reported poly(εcaprolactone) porous scaffold for tissue engineering function. Rheological study shows that melt flow index of PCL with molecular weight 50000, 64000 and 189000 was 13.9, 240 and 33.9 g/10 min at 80˚ C. The processing temperature can be utilized to tune the fiber diameter such as PCL with 189000 molecular weight has 47−224 µm fiber diameter at 80˚ C and 33−172 µm at 125˚ C134. The advantages and disadvantages of advanced fabrication methods for tissue engineering are depicted in table 2.

Table 2. Advantages and limitations of advanced fabrication methods Method 3D printing

Advantages Controlled quality Controlled pore structure

Limitations Expensive Limited filament


References 159

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4D printing

Smart dimensional stability Controlled pore structure Electrohydrodynamic Complex scaffold can be jet printing possible High resolution Melt electrospinning Time efficient writing Excellent filament resolution Solvent-free process

resolution Expensive Material limitations Too many parameters Post-drying required High temperature required Equipment cost

160

152

145

3. Applications of Porous Polymer Scaffold 3.1 Skin Tissue Engineering 3.1.1 Skin anatomy Skin is a largest organ of human body (as shown in Figure 11) consisting of dermis and epidermis layers, function crucial role in microbial resistance, sensing and thermoregulation55. The epidermis layer of skin act as a barrier consisting of keratinocyte cells which protect the skin and beneath of epidermis dermis layer impart colour consist of Melanocytes whereas fibroblast lies in the lower dermal layer exhibits resilience and strength161. Dermis layer also contains sweat glands, hair follicles, sebaceous glands, lymphatic vessels, apocrine glands, and blood vessels. Each layer of skin differs in thickness, strength and biological entities making the engineering of artificial skin more challenging162.



Figure 11. Typical structure of skin 3.1.2 Polymeric scaffold for skin tissue engineering

Figure 12. Typical structure of a) Chitosan, b) PLGA , c) PHBV, d) PCL


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3.1.2.1 Chitosan-based polymer scaffold In chitosan-gelatin blend scaffold, chitosan (as shown in figure 12.a) exert antimicrobial property whereas gelatin exert cell adhesion property163. In chitosan-gelatin sponge scaffold chitosan helps in preventing the wound from infection and dehydration26. The amine group of chitosan and carboxyl group of alginate enhances the mechanical strength of chitosan-alginate complex scaffold164. In PVA/chitosan scaffold, chitosan provides high biocompatibility, nontoxicity, antigenicity along with degradation rate tunability165. Chitosan has copious advantages to utilize as a scaffold for skin tissue engineering such as hydrophilic nature that promotes cell adhesion and proliferation166, chitosan is a polysaccharide which degrades in the presence of lysozyme enzyme that exhibits biodegradability and biocompatibility167, chitosan is only positively charged naturally occurring polysaccharide therefore it can be easily functionalized168. Intini et al. fabricated chitosan-based scaffold utilizing 3D printing for skin tissue regeneration and compared their performance against the commercial patch. Micrograph of polymer scaffold confirms that keratinocyte and fibroblast cells adhered and start to proliferate in 20 days as well as the proliferation of keratinocyte and fibroblast cells were enhanced with chitosan film backing. The initial cost of raw material for chitosan is lower compared to alginate and collagen which makes the chitosan commercial viable product169. Dhandayuthapani et al. investigated the performance of chitosan blended with gelatin for skin tissue engineering. The tensile strength of gelatin/chitosan nanofibers (37.91 ± 4.42 MPa) was higher than pristine gelatin nanofibers (7.23 ± 1.15 MPa) and pristine chitosan nanofibers (13.61 ± 0.6 MPa) on the contrary the blend fiber diameter was lower than pristine gelatin and pristine chitosan nanofibers163. Han et al. also investigated the same gelatin/chitosan blend porous polymer scaffold with a conventional methodology like Solvent casting/Freeze-drying for skin tissue engineering. They found that



pore size of the scaffold was in the micron range and having highly porous network higher than 90%. The water uptake capacity was higher than 1500% whereas the water intake capacity was higher than 400% and the as-prepared scaffold degrades 38.3-53.9% in 28 days. The positive results were observed in adhesion and proliferation of fibroblast and keratinocyte cells for 21 days26. Xingang et al. reported on chitosan-collagen hybrid scaffold to repair full-thickness skin defect. PLGA knitted mesh reinforced with chitosan-collagen scaffold (2.81 ± 0.09 cm2) has higher reduction area compared to chitosan-collagen scaffold (2.06 ± 0.14 cm2) after 8 week and tensile strength of PLGA knitted mesh reinforced with chitosan-collagen scaffold (0.68 ± 0.07 MPa) has higher than chitosan-collagen scaffold (0.49 ± 0.03 MPa) but lower than healthy skin (0.84 ± 0.08 MPa), this results depict that 73% tensile strength was achieved by synthetic skin compared to natural skin170. Wang et al. prepared knitted polyurethane reinforced with the chitosan-collagen scaffold to rectify full-thickness skin defects. The results later compared with a commercial product and found that PELNAC artificial dermis (1.53 ± 0.31 cm2) has lower residual area than as-prepared scaffold (2.65 ± 0.16 cm2) in 14 days. The tensile strength of healthy skin was higher than polyurethane (0.97 ± 0.05 MPa) reinforced with a chitosan-collagen scaffold (0.78 ± 0.03 MPa) and then PELNAC product (0.65 ± 0.05 MPa) in 12 week postoperation171. Tsao et al. studied chitosan-alginate (CA)/chitosan–poly(ethylene glycol) (C–PEG) scaffold for skin tissue engineering using the thermally-induced phase separation method. They found that the presence of chitosan-poly(ethylene glycol) in scaffold enhances the proliferation rate of fibroblast and keratinocyte also air-liquid interface helps in the maturation of keratinocyte cells164. Costa-Ju´nior et al. prepared the cross-linked chitosan/Polyvinyl alcohol (PVA) for skin regeneration via conventional technique (solvent casting). The as-prepared scaffold was crosslinked with 2% glutaraldehyde for 16 h and having 75 ± 25 µm film thickness. Reduction in


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swelling of the film was observed from 0.5% chitosan (200%) to 4% chitosan (125%) and SEM results confirm the adhesion and proliferation of the VERO cells165. Xingang et al. fabricated collagen–chitosan (CC) reinforced in knitted PLGA mesh for dermal tissue engineering. The tensile strength of PLGA-CC scaffold (3.62 ± 0.19 MPa) was higher than PLGA (3.58 ± 0.23 MPa) and CC scaffold (0.43 ± 0.01 MPa) whereas PLGA-CC scaffold achieved the tensile strength of dermal skin (1.03–3.10 MPa)172 Although chitosan has copious advantages but it also has certain drawbacks such as poor mechanical strength, shape retention failure, difficult to electrospun and insolubility in common solvent173,174. Chitosan-based polymer scaffolds are consolidated in Table 3.


Industrial & Engineering Chemistry Research 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47

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Table 3. Chitosan-based polymer scaffold for skin tissue engineering Material

Chitosan

Chitosan/gelatine

Gelatin–chitosan

Method

Cell

Results/Properties

Ref.

3D Printing

Normal dermal human fibroblast cells (Nhdf cells) and aneuploid immortal keratinocyte cells (HaCaT cells)

Cell proliferation duration-20 days

169

Electrospinning

-

Tensile strength Blend nanofiber- 37.91 ± 4.42 MPa Chitosan nanofiber- 13.61 ± 0.6 MPa Gelatine nanofiber-7.23 ± 1.15 MPa

163

Solvent Casting/ Freeze Drying

Human skin fibroblast (HSF) and the human keratinocyte (HaCaT) cell

Water absorption and retention capacity Water uptake capacity- >1500% Water intake capacity- >400% Cell proliferation duration-21 days

26

Collagen–chitosan


Endothelial cells and smooth muscle cells

Polyurethane/collagen –chitosan



Chitosan–alginate (CA)/chitosan– poly(ethylene glycol) (C–PEG)

TIPS/Freeze Drying

Firoblast (hFF) and keratinocytes (HaCaT)

Healthy skin-0.84 ± 0.08 MPa Chitosan collagen dcaffold-0.49 ± 0.03 MPa Knitted PLGA reinforced chitosan collagenscaffold-0.68 ± 0.07 MPa Tensile strength Healthy skin-0.97 ± 0.05 MPa Chitosan-collagen scaffold-0.78 ± 0.03 MPa PELNAC-0.65 ± 0.05 MPa Cell proliferation C-PEG embedded scaffold > without C-PEG scaffold


170

171

164

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Chitosan and poly(vinyl alcohol) (PVA)

Solvent Casting

VERO cell

Poly(L-lactide-coglycolide)/collagen– chitosan (PLG/CC)



Swelling rate 0.5% Chitosan scaffold- 200% 4% Chitosan scaffold-125% Tensile strength Dermal skin- 1.03–3.10 MPa PLGA-CC scaffold-3.62 ± 0.19 MPa PLGA scaffold-3.58 ± 0.23 MPa CC scaffold-0.43 ± 0.01 MPa


165

172


3.1.2.2 Poly(lactic-co-glycolic acid) based scaffold The typical structure of PLGA depicted in figure 12.b, Kim et al. investigated the PLGA microspheres to utilize in the delivery of skin tissue (Keratinocytes and Dermal Fibroblasts cells) for skin regeneration. Cell culture results shows increase in Keratinocytes cell number from initial 6 x 106 to the final 18 x 106 within 10 days and increase in Dermal Fibroblasts cells from initial 25 x 106 to final 230 x 106 within 10 days175. Sheridan et al. reported poly(lactide-coglycolide) (PLG) for sustained growth factor delivery utilizing a solvent casting method. The VEGF encapsulation efficiency of 85:15 and 75:25 PLG scaffold has 90 ± 1% and 72 ± 1% respectively and the release study reveal that 85:15 PLG scaffold release 84% in 48 h whereas 75:25 PLG scaffold release in 84% in 72 h, thus burst release was observed in 85:15 PLG scaffold than 75:25 PLG scaffold176. Yang et al. studied the Poly[(D,L-lactide)-co-glycolide] and Collagen scaffold for skin tissue engineering via electrospinning methodology. The fiber diameter increased with increase in the collagen content from 0% (1.03 ± 0.12 µm) to 30% (1.14 ± 0.18 µm) also the ultimate tensile strength was decreased due to increase in collagen content from 5% (1.66 MPa) to 30% (1.21 MPa). SEM image confirms the successful proliferation of the HDF cells on the scaffold and they also found beneath several layers177. 3.1.2.3 Poly(3-hydroxybutyrate-co-3-hydroxyvalerate) based scaffold The typical structure of PHBV illustrated in figure 12.c, Kuppan et al. developed poly(3hydroxybutyrate-co-3-hydroxyvalerate) fibers for skin tissue engineering via electrospinning technique. PHBV fibers (1.40 ± 0.23 MPa) have lower tensile strength than PHBV film (1.57 ± 0.58 MPa) but the degradation rate of PHBV fibers (75% in 5 weeks) was greater than PHBV film (35% in 4 weeks). PHBV fibers (74.50 ± 1.31%) have higher porosity compared to PHBV film (19.85 ± 0.51%) which increases the surface area and facilitate the better adhesion and


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proliferation platform for the human skin fibroblast cells178. Xu et al. fabricated electrospun Poly(3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV)/ polyethylene oxide (PEO) mats to use as a scaffold for skin tissue engineering. The fiber diameter decreased with increase in PEO content from 10 wt% (738 ± 0.21 nm) to 50 wt% (1098 ± 0.29 nm) as well as viscosity was also increased for same PEO content from 31.5 to 41.7 Pa s. The tensile strength of the 10% PEO (5.4 ± 0.5 MPa) was higher than 40% PEO (3.3 ± 0.6 MPa) as well as wettability also shows such decline in the contact angle from 124.3˚ to 55.3˚ for the same formulation179. Kang et al. prepared silver nanoparticles reinforced PHBV scaffold via electrospinning for tissue engineering function. Agar test confirmed that as-prepared scaffold shows anti-bacterial activity against Staphylococcus aureus and Klebsiella pneumoniae. Release study revealed that PHBV/Ag 1.0 nanofibrous scaffolds show sustain release of silver nanoparticles e.g. 0.2 ppm in 5 days and 0.55 ppm in 30 days180. PLGA and PHBV-based polymer scaffolds are consolidated in Table 4.



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Table 4. Poly(lactide-co-glycolide) and PHBV-based polymer scaffold for skin tissue engineering Material

Method

Cell

Oil/water emulsion and solvent casting

Keratinocytes and Dermal Fibroblasts cells

Poly (lactide-coglycolide)

Gas foaming

Human dermal microvascular endothelial cells

Poly[(D,Llactide)-coglycolide] and Collagen

Electrospinning

Human dermal fibroblasts (HDFs)

Poly(lacticco-glycolic acid)/collagen

Freeze drying

Insulin release Human dermal Pristine PLGA-18% in 28 days fibroblast cell (NHDF) Collagen/PLGA scaffold-25% in 28 days

Poly(D,L-lactide-coglycolide)/ bioactive glass

Thermally induced phase separation

Fibroblasts (L929) cells

VEGF secretion Pristine PLGA scaffold-580 pg/ml in 24 h 0.1 wt% added PLGA scaffold-870 pg/ml in 24 h

181

Electrospinning

Human skin fibroblasts

Degradation rate PHBV fibers-74.50 ± 1.31% PHBV film-19.85 ± 0.51%

178

-

Tensile strength 10% PEO scaffold-5.4 ± 0.5 MPa 40% PEO scaffold-3.3 ± 0.6 MPa

179

Poly(lactide-coglycolide) (PLG)

PHBV

PHBV/polyethylene oxide (PEO)

Electrospinning

Results/Properties Cell adhesion Keratinocytes cell Initial cell number-6 x 106 Final cell number-18 x 106 VEGF release 85:15 PLG scaffold-84% in 48 h 75:25 PLG scaffold-84% in 72 h Ultimate tensile strength 5% collagen scaffold-1.66 MPa 30% collagen scaffold-1.21 MPa


Ref.

175

176

177

101

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PHBV/silver nanoparticles

Electrospinning

Fibroblasts (NIH 3T3)

Silver nanoparticle release PHBV/Ag 1.0 nanofibrous scaffolds-0.55 ppm in 30 days


180


3.1.2.4 Polycaprolactone based scaffold for skin tissue engineering Polycaprolactone (as shown in figure 12.d) has desirable biostability, biocompatibility, biodegradability and has excellent mechanical properties as well as good fiber-forming property. It also has minor drawbacks such as slow degradation rate and absence of natural cell reorganization sites115,182. Yoganarasimha et al. investigated electrospun PCL fibers for tissue engineering application. The pristine PCL fibers (118˚) has higher contact angle compared to 5000 ppm peracetic acid (PAA) treated PCL fibers (55˚) whereas 20% ethanol diluted 5000 ppm PAA treated PCL fibers makes the scaffold superhydrophilic (0˚). The permeability study shows that 2500 ppm PPA in 20% ethanol treated PCL fibers (0.34 darcy) reduces the permeability compared to pristine PCL fibers (0.2 darcy)183. Trachtenberg et al. reported 3D printed poly(Ɛcaprolactone) for tissue engineering application and PCL was printed at 60˚ C temperature and 16 psi pressure. The porosity of the scaffold was tuned by the operating pressure where, as increase in operating pressure decreases porosity. In the context of porosity, a printing speed of 300 mm/min (40 ± 13%) has a lower average porosity than 400 mm/min (36 ± 12%). The mechanical test shows that the compressive yield strength and compressive modulus of the asprinted scaffold were 8.6 ± 4.1 MPa and 42.0 ± 20.7 MPa respectively184. Dai et al. studied collagen and PCL-based scaffold for tissue engineering application. Collagen:PCL (1:4) scaffold releases the 100% collagen within 11 days whereas collagen:PCL (1:20) scaffold releases 50% in 11days. Collagen increases the crystallinity of the composite film as the Collagen:PCL (1:4) scaffold has 43.4 ± 5.2% and collagen:PCL (1:20) scaffold has 63.5 ± 3.8%. The proliferation rate of 3T3 fibroblast cell on collagen:PCL (1:20) was 20 x 104 cells/cm2 in 8 days whereas proliferation rate of Normal Human Epidermal Keratinocytes (NHEKs) on collagen:PCL (1:4)


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was 6.5 x 104 cells/cm2 in 6 days115. Gautam et al. prepared collagen type 1 grafted PCL/gelatin scaffold via electrospinning technique for skin regeneration application. The SEM images confirm that modified PCL/gelatin scaffold (3 days) has higher proliferation rate of L929 mouse fibroblast cells than unmodified PCL/gelatin scaffold (5 days) and on the contrary degradation rate of the unmodified scaffold (2.89 mg/min) has higher than modified scaffold (2.1 mg/min)185. Powell et al. developed collagen/PCL blend nanofibers via electrospinning technique for artificial skin application. Pristine PCL (974.6 ± 48.6 kPa) scaffold has higher ultimate tensile strength than pristine collagen scaffold (118.2 ± 13.6 kPa) and 30% PCL/collagen (30:70) scaffold (418.3 ± 35.7 kPa) thus PCL exhibits mechanical strength and collagen increases the affinity of cells towards scaffold for better adhesion and proliferation186. Sharif et al. fabricated collagen coated PCL fibers for skin tissue engineering application. The wettability test shows that pristine PCL scaffold (142 ± 7˚) has higher contact angle than collagen coated PCL scaffold (70˚). The visual observation confirms that collagen coated PCL scaffold has higher cell than pristine PCL scaffold187. Lin et al. investigated bioactive glass loaded poly(ε-caprolactone) nanofibers via electrospinning for skin tissue engineering. The composite fiber diameter increases from 0 wt% (0.65 µm) bioactive glass (BG) to 4 wt% BG (1.1 µm) and then decreases to 8 wt% BG (0.8 µm). Release test reveals that higher bioactive glass increases the release concentration of Ca+ ions from 2 wt % (155 ppm) to 8 wt % (115 ppm), the same case was observed in case of Si ions. Mechanical properties like Elastic modulus, Elongation at break and fracture strength was decreases with increase in bioactive glass content188. Larrañaga et al. reported hydrolytic degradation of bioactive glass incorporated lactide and caprolactone based scaffold via solvent casting/particulate leaching for tissue engineering. PCL (4%) has lowest water absorption than Poly(L-lactide) (PLLA) (9%) and Poly(lactide-co-ε-caprolactone) (PLCL)



(15%)189. Lou et al. studied the sandwich composite of chitosan/PCL and PLLA porous sheet for tissue engineering via electrospinning technique. Improved proliferation rate of the scaffold was confirmed by the cell proliferation test as bi-layer scaffold (10 x 105 of hFF and HaCat) with dual cell has higher proliferation rate than pristine chitosan/PCL (5 x 105 of HaCat) and pristine PLLA porous scaffold (4 x 105 of hFF) for 8 days cell culture. SEM images shows that the asprepared fiber has 200 nm diameter with 100-150 μm pore size190. Zhou et al. prepared chitosan/poly(Ɛ-caprolactone) (CS/PCL) nanofibers scaffold via electrospinning method for an artificial skin graft. The CS/PCL nanofibers have 550 ± 120 nm average diameter with 85% ± 4.1% average porosity. The tensile strength of Tissue-engineered vascular graft (TEVG) before implantation (2.2 MPa) was lower than TEVG after 3 months (3 MPa), whereas the as-prepared scaffold has 1.5 MPa tensile strength191. Abarzúa-Illanes et al. developed the myoblast differentiation on the electrospun Poly(ε-caprolactone)/PLGA scaffold. They have found that blend nanofibers scaffold has higher proliferation rate than pristine PCL nanofibers scaffold and results also shows that aligned PCL nanofibers with higher PLLA content exhibit improved affection towards the cell growth. The decorin embedded blend nanofibers scaffold has improved cell differentiation was observed192. Zong et al. fabricated PCL/PLGA bilayer scaffold for bile duct repair. The tensile strength of the dry PCL/PLGA scaffold (24 MPa) has higher strength than wet PCL/PLGA scaffold (17 MPa) as well as the cell density of PCL/PLGA scaffold in perfusion culture (2.5 x 106) was higher than static culture (3.5 x 106)193. Zuidema et al. investigated the Lysozyme embedded porous silicon nanoparticles (pSiNPs) incorporated in PCL or poly(lactide-co-glycolide) matrix via spray nebulization method for tissue regeneration. Pristine pSiNPs nanofibers scaffold (100% in 200 h) releases Lysozyme rapidly than pSiNPs/PCL nanofibers scaffold (27% in 60 days). The porous silicone nanoparticles impart


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delivery of sensitive biological molecules, as well as embedded scaffold, restrict the rapid release of Lysozyme194. Gonçalves et al. reported Iron oxide embedded starch/PCL via rapid prototyping technology for tissue engineering application. They found that the application of magnetic field increases the cell viability in starch/PCL scaffold as well as iron oxide embedded starch/PCL scaffold and cell proliferation in magnetic condition was higher than static condition195. Kudryavtseva et al. illustrated the utilization of titanium-nitrogen coated PCL via electrospinning method for vascular tissue engineering. The 0.2 wt % titanium was coated within 30 s and 3.2 wt % coated within 480 s in plasma coating. The wettability test shows that untreated scaffold (123.0 ± 2.6˚) has greater contact angle than 240 s plasma treated scaffold (65.3 ± 2.9˚) but the young’s modulus increases from 11.70 ± 1.63 to 13.08 ± 1.52 MPa. The untreated scaffold (103 ± 4 cells/mm2) has greater cell adhesion than 240 s plasma treated scaffold (160 ± 6 cells/mm2)196. Zhang et al. studied polydopamine coated PCL diacrylate scaffold for tissue engineering via solvent casting/particulate leaching technique. Shape memory property confirms as the reduction in E was observed above 60˚ C (Transition temperature) from 4.4 MPa to 12kPa because above 60˚ C PCL melted and becomes shapeable. Polydopamine coated scaffold has 95% shape recovery in the first cycle and 100% shape recovery in the second cycle197. Ferreira et al. prepared PCL functionalized with gelatin via electrospinning for tissue engineering application. The wettability study shows that PCL scaffold (120˚) has greater water contact angle than gelatineMA/PCL (50:50) scaffold (5˚). Thrombagenicity of PCL has 58% whereas gelatin/PCL (25:75) scaffold has 38% and weight loss test shows that photocrosslinked gelatin/PCL (75:25) scaffold (38% in 3 days) has lower weight loss than Non-photocrosslinked scaffold (65% in 3 days)198. PCL-based polymer scaffolds are consolidated in Table 5.



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Table 5. Polycaprolactone-based polymer scaffold for skin tissue engineering Material

Method

PCL

Electrospinning

PCL

Threedimensional printing (3DP) system

Collagen/PCL

Freeze drying

PCL/gelatin/Collagen type 1

Electrospinning

Collagen/PCL

Electrospinning

Cell

Results/Properties

Wettability Control scaffold- 118˚ (5000 ppm) PAA in DI water treated scaffoldBacillus 55˚ atrophaeus Control scaffold- 125˚ (5000 ppm) PAA in 20% ethanol treated scaffold- 0˚ Compressive yield strength Polymer scaffold-3.3-15.2 MPa Compressive Modulus Polymer scaffold- 23.6-87.9 MPa Crystallinity Collagen:PCL (1:4) scaffold-43.4 ± 5.2% Fibroblasts cell collagen:PCL (1:20) scaffold-63.5 ± 3.8% Proliferation rate Collagen modified PCL/gelatin scaffold- 3 L929 mouse days fibroblast cells Collagen modified PCL/gelatin scaffold- 5 days Ultimate tensile strength Human dermal Pristine PCL scaffold-974.6 ± 48.6 kPa fibroblasts (HF), Pristine collagen scaffold-118.2 ± 13.6 kPa and epidermal 30% PCL/collagen (30:70) scaffold-418.3 ± keratinocytes (HK) 35.7 kPa


Ref.

183

184

115

185

186

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Collagen coated PCL

Electrospinning

Human endometrial stem cells

Poly(ε-caprolactone)/ bioactive glass

Electrospinning

Human skin fibroblast cells

Poly(L-lactide) (PLLA), Poly(lactideco-ε-caprolactone) (PLCL), Poly(εcaprolactone) (PCL)/ Bioactive glass particles

Solvent casting/particul ate leaching

-

Degradation rate (KMw) PLCL > PLLA > PCL

189

Human foreskin fibroblasts and keratinocytes

Proliferation rate hFF PLLA micro disc-4 x 105 in 8 days HaCat chitosan/PCL nanofibrous mat-5 x 105 in 8 days HaCat + hFF on bilayer saffold-10 x 105 in 8 days

190

Tensile stress CS/PCL scaffolds- 1.5 MPa TEVG before implantation- 2.2 MPa TEVG (after 3 months)- 3 MPa Native carotid artery- 3.2 MPa

191

Cell proliferation Blend scaffold > pristine scaffold

192

Tensile modulus

193

Chitosan/PCL PLLA

Electrospinning

Tissue-engineered vascular graft (TEVG) and outgrowth endothelial cells (OECs) Murine skeletal muscle myoblasts

Chitosan/PCL

Electrospinning

PCL/PLGA

Electrospinning

PCL

Gelatin particle- Human bone

Water contact angle Pristine PCL- 142 ± 7˚ Collagen coated PCL- 70˚ Ca+ ions release 2 wt% BG scaffold- 155ppm 8 wt% BG scaffold- 115 ppm


187

188


PCL/ PLGA

leaching method and freeze drying

PCL/Lysozyme loaded porous silicon nanoparticles (pSiNPs) poly(lactide-coglycolide)/Lysozyme loaded porous silicon nanoparticles (pSiNPs)

Spray Nebulization

Starch/PCL/Iron oxide

Titanium-nitrogen coated PCL

Polydopamine coated PCL diacrylate

PCL/Functionalized gelatin

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marrow-derived Dry state-24 MPa mesenchymal stem Under water at 37˚ C- 17 MPa cells

Astrocytes

Lysozyme release Pristine pSiNPs-100% in 200 h pSiNPs/PCL nanofiber-27% in 60 days

194

Rapid prototyping technology (RPT)

Adipose stem cells

Cell viability Magnetic polymer scaffold > Non-magnetic polymer scaffold

195

Electrospinning

Human umbilical vein endothelial cells (HUVECs)

Cell viability Coated PCL scaffold > pristine PCL scaffold

196

Solvent casting/ particulate leaching

Human osteoblasts

Shape recovery (Rr)Polydopamine-coated 491 scaffold- 95% (N-1) and 100% (N-2) Self-fitting propertyChange in E-4.4 MPa to 12 kpa (at 60˚ C)

197

Electrospinning

Normal Human Wettability Dermal Fibroblasts PCL fiber-120˚ cells Composite fiber-38º


198

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PCL/polyurethane

Freeze drying

PCL/Acrylated llactide-cotrimethylene carbonate (aPLA-coTMC)

Electrospinning

Polylactide (PLDL)/Ossein hydroxyapatite complex (Osteo) PCL/Osteo PCL-poly(ethylene glycol)-PCL, and magnetic iron oxide (Fe3O4)

Electrospinning

Electrospinning

Hydrolytic degradation at 37˚ PCL-PU-Semi-IPNs-25% in 22 days Elastic Modulus PCL/aPLA-co-TMC-1.17 ± 0.04 PCL-0.81 ±0.09 Human Native tissue-0.18±0.02 mesenchymal stem Ultimate tensile strength cells PCL/aPLA-co-TMC-1.21 ± 0.19 PCL-0.55 ± 0.10 Native tissue-1.55 ± 0.4 Tensile strengthPCL scaffold-2.1 MPa Normal human PCL/osteo scaffold-1.9 MPa osteoblast PLDL scaffold-1.35 MPa PLDL/osteo scaffold-0.6 Mpa Fibroblast cell

NIH-3T3 cells

Degradation rate 5% Fe3O4-27% in 10 weeks 1% Fe3O4-22% in 10 weeks


199

200

201

202


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3.2 Bone Tissue Engineering 3.2.1 Bone anatomy The natural bone (as illustrated in Figure 13) composes of 60-70% bone mineral, 1020% collagen and 9-20% water beside organic material like polysaccharides, lipids and proteins are also included in the lower concentration. The chemical structure of the largest concentration of the bone mineral Ca10(PO4)6(OH)2 is also called hydroxyapatite which consists of the two third bone mass203. Traditionally the three graft were used for bone regeneration which are auto graft recovered body, allograft recovered from a bone bank and synthetic graft (hydroxyapatite based graft) but the immunological reactions limit the application of these graft. Collagen and calcium phosphate apatite are the basic apatite composition204. Rezwan et al.

reported

microstructural and mechanical properties of porous polymer scaffold and also challenges faced by researchers regarding biomaterials with respect to biomedical applications205.


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Figure 13. Typical structure of bone



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3.2.2 Polymeric scaffold for bone tissue engineering 3.2.2.1 Polylactic acid-based scaffold

Figure 14. Typical structure of PLLA Chen et al. fabricated tubular poly(L-lactic acid) (PLLA) (as shown in figure 14) scaffold via Thermally induced phase separation for bone tissue regeneration. Polymer content tuned the thickness of the scaffold and interconnected pores utilized for enhanced cell attachment and proliferation. SEM images confirm the biocompatibility as the nanofibers scaffold has cell proliferation than non-nanofibers scaffold in 1 day and 7 days of cell culture shows that the asprepared scaffold has excellent cytocompatibility and bioactivity206. Zhou et al. investigated PLA scaffold via 3D printing followed by gas foaming technique for Bone tissue engineering. The pore size of the scaffold decreases with CO2 concentration, as the 10% CO2 treated scaffold (10 μm) has a higher pore size than 19% CO2 treated scaffold (2.5 μm). The 50˚ C foaming (3 μm) temperature has lower pore size compared to 90˚ C foaming scaffold (9 μm) temperature. The porosity of 700 μm filament (70%) was higher than 300 μm filament (57%)207. Diomede et al. reported 3D printed PLA scaffold for bone tissue regeneration. The false discovery rate of PLA/PEI-EVs/hGMSCs (2.56 x 10-15) was higher than PLA/EVs/hGMSCs scaffold (4.42 x 10-14). The accelerated tissue growth was observed in PLA/hGMSCs scaffold than pristine PLA scaffold. The tensile strength of PLA was 50 MPa and has elongation strength more than 20%208. Birhanu et al. illustrated


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electrospun blend nanofiber of Poly-L-lactic acid/pluronic P123 for bone tissue engineering. As the Pluronic P123 is a triblock copolymer of PEO–PPO–PEO and it consists of hydrophobic [poly(propylene oxide)] and hydrophilic [poly(ethylene oxide)] polymer. The mechanical properties shows that PLLA nanofibers scaffold (1.92 MPa) has greater tensile strength than pluronic P123 blended PLLA scaffold (1.46 MPa), on the contrary breaking tensile strain of Pluronic P123 blended PLLA scaffold (105%) was higher than pristine PLLA nanofiber scaffold (88%)209. Dou et al. studied the kinetics of Poly(L-lactic acid)/hydroxyapatite (PLLA/HA) scaffold fabricated via solvent casting/salt-leaching method for bone tissue engineering. They had found that the mineralization medium has an impact on the kinetics such as mineralized scaffold in phosphate buffer saline (PBS) shows second-order kinetic model and mineralized scaffold in simulated body fluid (SBF) shows zero-order kinetic model210. Rajzer et al. prepared Polylactide/Ossein hydroxyapatite complex (PLDL/Osteo) scaffold and PCL/Osteo scaffold via electrospinning technique for Bone tissue engineering application. Ossein hydroxyapatite complex mimics native bone extracellular matrix and PLDL and PCL provide the backing support while cell proliferation. The tensile strength of PCL/Osteo scaffold (1.9 MPa) was higher than PLDL/Osteo scaffold (0.6 MPa) owing to PCL (2.1 MPa) that has higher tensile strength than PLDL (1.4 MPa)201. Li et al. developed rhBMP-2 and dexamethasone incorporated Zein/ Poly(L-lactic acid) (PLLA) scaffold via coaxial electrospinning technique for bone tissue regeneration, as dexamethasone is a drug whereas rhBMP-2 is bone morphogenetic protein. The wettability study shows that the water contact angle of Zein/PLLA-BMP scaffold (93.2 ± 2.1°) was higher than Zein-DEX/PLLA-BMP-2 scaffold (76.4 ± 1.9°). The drug release test shows that Zein-Dex/PLLA (core/shell) nanofibers (70% in 3 days) have rapid drug release than Zein-Dex/PLLA-rhBMP-2 nanofibers (8% in 3 days)211. Mahdavi et al. fabricated bioactive glass nanoparticles coated



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Poly(L-lactic acid) scaffold for accelerated Osteogenic differentiation. Polylactic acid-based polymer scaffolds are consolidated in Table 6.


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Table 6. Polylactic acid-based polymer scaffold for bone tissue engineering Materials

Method

Cell

Poly(L-lactic acid) (PLLA)


PLA

3D printing and Gas foaming

PLA

3D printing

Poly(L-lactic acid)/Pluronic P123 (PLLA/P123)

Electrospinning

Poly(L-lactic acid)/ hydroxyapatite (PLLA/HA)

Solvent casting/saltleaching method

Polylactide/Ossein hydroxyapatite complex (PLDL/Osteo) PCL/Ossein hydroxyapatite complex (PCL/Osteo)

Electrospinning

Results/Properties

Ref.

Proliferation rate 206 Rat skull bone cell line Nanofibers scaffold > nonnanofibers scaffold Porosity 207 700 μm filament-70% 300 μm filament-57% Tensile strength Human gingival Poly(lactide)- 52 MPa 208 mesenchymal stem cells Elongation strength (hGMSCs) Poly(lactide)- >20% Human adipose tissue Tensile strength 209 derived mesenchymal PLLA scaffold-1.92 MPa stem cells PLLA/P123 scaffold- 1.46 MPa Mineralization kinetics Mineralized scaffold in PBS-Second210 order kinetic model Mineralized scaffold in SBF-Zeroorder kinetic model Fiber diameter PLDL/Osteo fibers-0.85 ± 0.22 µm Normal human PCL/Osteo fibers-1.39 ± 0.63 µm 201 osteoblast PorosityPLDL/Osteo scaffold-52% PCL/Osteo scaffold-53%



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rhBMP-2 and dexamethasone loaded Electrospinning Zein/poly(l-lactide)(PLLA)

Water contact angle Zein/PLLA-BMP-2-93.2 ± 2.1° Mesenchymal stem cells Zein-DEX/PLLA-BMP-2-76.4 ± 1.9°

Bioactive glass ceramic nanoparticles-coated Poly(L-lactic acid)

Adipose stem cells

Average diameter PLLA scaffold- 700 ± 230 µm

212

Human articular cells

Nozzle diameter on Porosity 600 µm- 52% 900 µm-63%

213

Osteoblastic cells

Compressive Strength 10% HMS/PLLA scaffold- 5 MPa 30% HMS/PLLA scaffold- 9 MPa

214

Human adipose tissuederived mesenchymal stem cell

Porosity PLA scaffold- 77.5 ± 2.3% PLA/EE scaffold-80 ± 3.2% PLA/EE/nHA scaffold-83.4 ± 3.3% EE release rate 85% in 21 days

215

Collagen coated PLA Poly(L-lactic acid)/Hydroxyapatie microsphere (PLLA/HMS)

Electrospinning

3D printing

Flame-drying

Poly(lactic acid)/Equisetumarvensehe rbal Electrospinning extract/nanohydroxyapatit e (PLA/EE/nHA) Dexamethasone-loaded Mesoporous Silica nanoparticles Poly(L-lactic acid)/Poly(Ɛ– caprolactone) Poly(l-lactic acid)/strontium-doped hydroxyapatite


Bone marrow-derived DEX release rate mesenchymal stem cells 85% in 26 days


Mouse preosteoblast cell line (MC3T3-E1)


Water contact angle PLLA scaffold- 130˚ PLLA/HA scaffold- 60˚

211

216

217

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PLLA/HA-Sr scaffold- 40˚ Tensile strength 218 Polylactic/pearl powder Electrospinning MC3T3 cells PLA scaffod- 7 MPa PLA/pearl powder scaffold- 7.4 MPa Compressive strength PLLA/PGA scaffold- 23.85 MPa PLLA/graphene 1% Ag/PLLA/PGA scaffold- 25.96 219 Laser sintering MG63 cells oxide/Nano Ag MPa 1% GO/1% Ag/PLLA/PGA scaffold36.95 MPa Conductivity In-situ PLA scaffold- N/A Bone Polylactide/polyaniline polymerization/thermal PLA/PANI (95:5) scaffold- 0.004 220 marrow derived nanoparticle (PLA/PANI) induced phase S/cm mesenchymal stem cells separation (TIPS) PLA/PANI (85:15) scaffold- 0.032 S/cm



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3.2.2.2. Polycaprolactone based scaffold for bone tissue engineering Polycaprolactone has stability at ambient temperature and has mechanical property in desirable physiological range along with ability to promote osteoblast growth221. On the other hand, slow degradation rate of PCL (3 years for Mn-50000) makes them suitable only for longterm implant in bone tissue engineering field205,222. Chen et al. studied the PCL scaffold via gas foaming technique for bone tissue engineering. The diameter of Scaffold was tuned from 12.5 ± 0.1 to 344 ± 54 µm by varying the processing conditions and as-prepared scaffold had 60.8-83.4% porosity also all the scaffold had porosity more than 85%223. Bhattacharjee et al. prepared silk fibroin grafted poly(Ɛ-caprolactone) (SF-PCL) scaffold via electrodeposition. The mechanical test reveals that tensile strength can be manipulated by varying the voltage while electrodeposition, e.g. SF/PCL scaffold (17 MPa) at 5V had higher tensile strength than SF/PCL scaffold (15 MPa) at 3V224. Kiran et al. grafted silk fibroin and nanohydroxyapatite to poly(Ɛ-caprolactone) and processed via solvent casting/particulate leaching technique for bone tissue regeneration. Cell adhesion study shows that Human osteoblast-like cells (MG 63) had more affinity towards PCL-nHA-Col scaffold (92 ± 4%) than PCl-nHA scaffold (80 ± 5%) and pristine PCL scaffold (76 ± 3%)225. Yassin et al. developed poly(l-lactide-co-Ɛcaprolactone) scaffold via particulate leaching. Bone marrow stromal cells seeding efficiency was increased from 44.7% to 58.7% for poly(LLA-co-CL) scaffold and nanodiamond particles loaded poly(LLA-co-CL) scaffold respectively226. Woodard et al. fabricated poly(L-lactic acid) (PLLA) and poly(ε-caprolactone) diacrylate (PCL-DA) Interpenetrating network (IPNs) scaffold via solvent casting and particulate leaching method for bone tissue engineering application. The 160˚C annealed PLA/PCL-DA (60:40) scaffold (350 μm) has a lower pore size than annealed free scaffold (200 μm). The compression modulus of PLA/PCL-DA (60:40) scaffold (20 MPa) was


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higher than PLA/PCL-DA (90:10) scaffold (15 MPa)227. Gorodzha et al. investigated silicate containing hydroxyapatite incorporated poly(3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV) scaffolds via electrospinning technique for bone tissue regeneration. The cell viability of human mesenchymal stem cells on PHBV-SiHA scaffold was greater than pristine PHBV scaffold and PCL scaffold; these results due to silicate containing hydroxyapatite stimulate the cells for extended proliferation. Hydrophobic nature of the all prepared scaffold (100˚) was confirmed by the wettability test228. Thadavirul et al. reported utilization of PCL/ poly(hydroxybutyrate) (PCL/PHB) scaffold and PCL/ poly(3-hydroxybutyrate-co-3-hydroxyvalerate) (PCL/PHBV) scaffold via a combination of solvent casting and particulate leaching for bone tissue engineering. The PCL/30% added PHB and PHBV scaffold had a compression modulus of 1223 ± 1.4 MPa and 1861 ± 1.0 MPa respectively, whereas the pristine PCL scaffold had 57.7 ± 8.2 kPa229. Remya et al.illustrated the blend electrospun PCL/Polyethylene oxide scaffolds for bone tissue regeneration. The mechanical properties of the scaffold were increased by the addition of Polyethylene oxide e.g. pristine PCL scaffold had 5.2 MPa tensile strength whereas PCL/PEO (70:30) scaffold had 9.5 MPa230. Erndt-Marino et al. studied a shape memory foam made up from poly(Ɛ-caprolactone) diacrylate (PCLDA) scaffold coated with polydopamine processed via conventional technique solvent casting and particulate leaching. The elastic modulus of the polydopamine-coated scaffold (5.0 ± 0.3 MPa) increases slightly compared to pristine PCLDA scaffold (5.4 ± 0.1 MPa). The average porosity of both the scaffold was more than 72% and polydopamine coating has not had any effect on the compression strength231. Liao et al. prepared collagen type 1 coated PCL-βtricalcium phosphate scaffold via a selected laser sintering technique for bone tissue engineering. The collagen coating increases the compression modulus from 13.66 ± 0.19 MPa to 13.74 ± 0.32



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MPa for uncoated and coated scaffold respectively, but water contact angle decreases from 117.69 ± 2.29˚ to 73.35 ± 1.84˚232. PCL-based polymer scaffolds are consolidated in Table 7.


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Table 7. Polycaprolactone-based polymer scaffold for bone tissue engineering Materials

Method

Cell

Poly(Ɛ-caprolactone)

Gas foaming

-

Silk fibroin grafted poly (ƐElectrodeposition caprolactone) (SF-PCL)

Collagen grafted PCL and nanohydroxyapatite

Solvent casting/particulate leaching

Poly(l-lactide-co-εParticulate caprolactone) (poly(LLA-coleaching CL)) Poly(L-lactic acid)-poly(εcaprolactone) diacrylate (PCL-DA)

Solvent casting/Particulate leaching

PCL/poly(3hydroxybutyrate-co-3hydroxyvalerate)

Electrospinning

PCL /PCL– Poly(hydroxybutyrate)

Solvent casting/Particulate

Results/Properties

Ref.

Porosity 223 PCL scaffold- 60.8–83.4% Tensile strength Human osteoblast- SF-PCL- 11 Mpa 224 like cells (MG 63) SF-PCL/3V- 15 Mpa SF-PCL/5V- 17 MPa Cell adhesion Preosteoblast PCL- 76 ± 3% 225 MC3T3-E1 subclone PCL-nHA- 80 ± 5% 4 cell line PCL-nHA-Collagen-92 ± 4% Seeding efficiency Bone marrow stromal 226 Poly(LLA-co-CL)- 44.7% cells (BMSCs) Poly(LLA-co-CL)/nDPs- 58.7% Compression modulus varying annealing temperature PCL-DA (90:10) scaffold- 9 MPa at 85˚ 227 C & 14 MPa at 160˚C PCL-DA (90:10) scaffold- 2 MPa at 85˚ C & 20 MPa at 160˚C Cell viability Human mesenchymal PCL scaffold- 82% 228 stem cells (hMSCs) PHBV scaffold- 90% PHBV-SiHA scaffold- 92% Mouse fibroblastic Compression modulus 229 cells (L929) and PCL scaffold- 57.7 ± 8.2 kPa



PCL/Poly(3Hydroxybutyrate-co-3Hydroxyvalerate)

PCL/Polyethylene oxide (PCL/PEO)

leaching

mouse calvariaderived preosteoblastic cell (MC3T3-E1)

Electrospinning

Polydopamine coated poly(ƐSC/PL caprolactone) diacrylate (PCLDA)

Collagen type I coated PCL– Selective laser β-tricalcium phosphate sintering

PCL/octacalcium Phosphate (PCL/OCP)

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Electrospinning

PCL-30% PHB scaffold- 1223 ± 1.4 MPa PCL-30% PHBV scaffold- 1861 ± 1.0 MPa Tensile strength Human osteoblast PCL scaffold- 5.2 MPa 230 sarcoma (hOS) cell PCL/PEO (90:10) scaffold- 7.7 MPa lines PCL/PEO (70:30) scaffold- 9.5 MPA Elastic modulus Bone marrow-derived Pristine PCLDA scaffold- 5.0 ± 0.3 231 human mesenchymal MPa stem cells (h-MSCs) PD coated PCLDA scaffold- 5.4 ± 0.1 MPa Compression modulus Pristine PCL scaffold- 6.77 ± 0.19 Osteogenesis of MPA 232 adipose-derived stem PCL/TCP (70:30) scaffold- 13.66 ± cells 0.19 MPa PCL/TCP-Col- 13.74 ± 0.32 MPa Young’s modulus Osteoblast human G- Pristine PCL scaffold- 5.23 ± 0.77 MPa 233 292 cells PCL/5wt% OCP- 5.59 ± 0.41 MPa PCL/10wt% OCP- 4.50 ± 0.37 MPa

Zn-Curcumin complex (ZnCUR)/PCL (core) Electrospinning Graphene oxide-Poly(vinyl alcohol)-chitosan (shell)

Osteoblast cells

PCL/hydroxyapatite

Human mesenchymal Compressive modulus

Co-extrusion and


Tensile strength Pristine Zn-CUR scaffold- 17 MPa 2% GO Zn-CUR scaffold- 29 MPa

234

235

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(PCL/HA) PCL/halloysite nanotube (PCL/HNT) Poly(Ɛcaprolactone)/Cryogel

Gas foaming

3D printing

SolventPoly(Ɛ-caprolactone)/Starch casting/saltleaching

Immobalized collagen and bone morphogenetic protein Particulate on PCL-biphasic calcium leaching phosphate (BMP-2/PCLBCP)

stem cells (hMSCs)

PCL scaffold- 7.5 MPa 10 wt% HA-PCL scaffold- 22.5 MPa 10 wt% HNT-PCL scaffold- 23 MPa

MC3T3-E1 osteoblast Compression strength precursor cells PCL scaffold- 25 MPa (ATCC) Compressive modulus Human cancellous bone- 1.5-9.3 MPa PCL/starch (90:10) scaffold- 19.45 ± 1.94 MPa Young’s modulus Human cancellous bone-100–400 MPa PCL/starch (90:10) scaffold-32.06 ± 2.74 MPa MC3TC3-E1 preosteoblasts cells


Growth factor (BMP-2) release PCL-BCP-COL-BMP-2 scaffold- 48% in 28 days

236

237

238


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3.3 Liver Tissue Engineering 3.3.1 Liver anatomy The liver is considered to be largest integral organ in the body (as depicted in Figure 15) having 1.2-1.3 kg weight and located right side under the lower ribs. The heart provides the oxygenated blood supply to the liver through the hepatic artery. The deoxygenated blood supply along with nutrition through a portal vein from the stomach, spleen and intestine239. The deoxygenated blood is sent to the heart through the hepatic vein connected to interior vena cava. The 25% blood supply of liver comes from hepatic artery whereas 75% blood supply came from the portal vein. The majority of liver function is performed by the hepatocytes cell which is special liver cells surrounded by the sinusoids helps the microvilli to facilitate the nutrition exchange240.

Figure 15. Typical structure of Liver


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3.3.2 Polymeric scaffold for liver tissue engineering Yang et al. investigated the silk fibroin/silk composite scaffold for liver tissue engineering via solvent casting/freeze drying. The flow cytometry analysis shows that apoptosis rate was tuned by varying the gelatin content in the scaffold such as SF/G (4:2) scaffold (25%) has higher apoptosis rate than SF/G (4:3) scaffold (18.5%) and SF/G (4:4) scaffold (13%)241. She et al. reported Silk fibroin/chitosan/heparin composite scaffold via same conventional technique solvent casting/freeze drying. The water absorption of SF/CS scaffold was higher than 1 wt % Heparin loaded SF/CS scaffold, thus the swelling ratio of SF/CS scaffold was greater than 1 wt % Heparin loaded SF/CS scaffold and all the scaffold had higher porosity than 95%242. Shirahama et al. illustrated collagen coated Poly(ethylene glycol) scaffold via particulate leaching method. Cell viability study shows that 200 µg/ml coated scaffold (350%) has higher cell viability than 20 µg/ml coated scaffold (200%), this result confirms that collagen utilized to increases the cell attachment on the scaffold243. Maymand et al. studied Polyethersulfone nanofibers via electrospinning technique and then coated with collagen. The cell proliferation study shows that the optical density of collagen coated polyethersulfone (0.355) was higher than the control scaffold (0.25)244. Li et al. prepared collagen coated PLGA via freeze-drying technique for liver tissue engineering application. The cell proliferation study shows that C-PLGA scaffold vary the cell number from 6 × 105 cells to 4.2 ×105 cells and monolayer scaffold vary in cell number from 5 × 105 cells to 1 × 105 cells245. Chen et al. developed an alginate/galactosylated chitosan scaffold via freeze drying for liver tissue regeneration. The average pore size was 50-150 µm and the Mechanical test shows that compression modulus of A/GC (3/1) scaffold (6.4 kPa) was higher than A/GC (4/1) scaffold (4 kPa)246. Adwan et al. fabricated polyhedral oligomeric Silsesquioxane [POSS]-modified PCL urea urethane scaffold via particulate leaching technique. They have



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analyzed the effect of different leaching agent on the pore size of the scaffold-like sodium bicarbonate had reduced pore size of 20-750 µm when compared to sodium chloride having pore size 150 µm-1.5 mm247. Shang et al. investigated galactosylated chitosan/hyaluronic acid (GC/HA) scaffold via freeze drying methodology for liver tissue engineering. They attempted the co-culture of hepatocytes and endothelial cells on scaffold and pore size of galactosylated chitosan scaffold (114.1 ± 22.7 µm) was lesser than chitosan scaffold (126.5 ± 28.4 µm) and GC/HA scaffold ( 147.5 ± 32.2 µm)248. With the same technique but different material like gelatin cryogel Gandomani et al. attempted for liver tissue regeneration. The as-prepared scaffold has an average pore size of 35-80 µm and urea release study shows that gelatin scaffold (25 mg/dl) had rapid urea release than polystyrene scaffold (15 mg/dl)249. Zhu et al. reported the utilization of Poly(3hydroxybutyrate-co-3-hydroxyvalerate)/poly(lactide-co-glycolide) acid (PHBV/PLGA) scaffold via freeze-drying method for liver tissue engineering application. The degradation rate of PLGA scaffold (98% in 90 days) was higher than PHBV/PLGA scaffold (54% in 90 days) and then PHBV scaffold (8% in 9 days)250. Polymer scaffolds for liver tissue engineering are consolidated in Table 8.


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Table 8. Polymer scaffold for liver tissue engineering Material

Method

Cell

Silk fibroin/gelatin (SF/G)

Solvent casting/freeze drying

Human hepatic QZG cells

Silk fibroin /chitosan/heparin (SF/CS/HP)

Solvent casting/freeze drying

Collagen coated Poly(ethylene glycol)

Particulate leaching

Collagen-coated Polyethersulfone

Electrospinning

Collagen-coated poly(lactic- co-glycolic Freeze drying acid) (C-PLGA)

Alginategalactosylated chitosan Freeze drying (AGC) Polyhedral Oligomeric Particulate Silsesquioxane leaching

Results/Properties

Aptosis rate SF/G (4:2) scaffold-25% SF/G (4:3) scaffold-18.5% SF/G (4:4) scaffold-13% Water absorption Hepatocytes cells SF/CS scaffold- 286 ± 5% SF/CS/HP scaffold- 327 ± 5% Cell viability Homotypic and 200 µg/ml coated scaffold- 350% heterotypic cell 20 µg/ml coated scaffold- 200% Human Induced Optical density Pluripotent TCP protein domain-0.25 PES/COL scaffold-0.355 Stem cells Cell proliferation C-PLGA Scaffold-6 ×105 cells/scaffold in 0 week Hepatic stem cells C-PLGA scaffold-4.2 × 105 cells/scaffold in 3 week Monolayer-5×105 cells/scaffold in 0 week Monolayer-1 × 105 cells/scaffold( 3 week) Compression modulus Hepatocytes cells AGC (3/1) scaffold- 6.4 kPa AGC (4/1) scaffold- 4 kPa Pore size Hepatocytes cells Sodium bicarbonate scaffold-20-750 µm


Ref.

241

242

243

244

245

246

247


[POSS]-modified PCL urea urethane Galactosylated chitosan Freeze-drying (GCs)/hyaluronic acid (HA) Gelatin cryogel

Freeze drying

Poly(3hydroxybutyrate-co-3hydroxyvalerate) Freeze drying /poly(lactide-coglycolide) acid (PHBV/PLGA)

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Sodium chloride scaffold-150 µm-1.5 mm Pore size Hepatocytes and GCs/HA scaffold-147.5 ± 32.2 µm endothelial cells GCs scaffold-114.1 ± 22.7 µm Cs scaffold- 126.5 ± 28.4 µm Human adipose Average pore size-35-80 µm derived Urea release rate mesenchymal stem Polystyrene scaffold-15 mg/dl cells Gelatin scaffold-25 mg/dl Degradation rate Hepatocyte growth PHBV scaffold- 8% in 90 days PLGA scaffold-98% in 90 days factor (HGF) PHBV/PLGA scaffold-54% in 90 days


248

249

250

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3.4 Cardiac Tissue Engineering 3.4.1 Cardiac anatomy The heart beats at 100000 per day and pumps the 200 gallons of the blood through 60000 miles of network251. The human heart is divided into four chambers: 1) Right atrium 2) Right Ventricle 3) Left ventricle 4) Left atrium where lower oxygen blood is processed in the right atrium and oxygen reach blood is processed in the left atrium. Flowing blood in the body continuously provides the supply of oxygen and nutrition and absorb the carbon dioxide. Coronary artery is the main source of blood to the heart252. (as represented in Figure 16)

Figure 16. Typical structure of Heart



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3.4.2 Polymeric scaffold for cardiac tissue engineering Constantinides et al. studied Polyhydroxyalkanoate/PCL (PHA/PCL) scaffold via solvent casting method for cardiac tissue engineering application. The tensile strength of non-porous PHA/PCL scaffold (6.6 MPa) was higher than porous PHA/PCL scaffold (2.2 MPa) on the contrary Elongation at break of porous PHA/PCL scaffold (600%) was higher than non-porous PHA/PCL scaffold (525%) and Young's modulus of non-porous PHA/PCL scaffold (4.6 MPa) was greater than the porous PHA/PCL scaffold (0.6 MPa)253. Baheiraei et al. prepared aniline pentamer-modified polyurethane/PCL scaffold via solvent casting/particulate leaching method for cardiac tissue regeneration. The aniline group in the scaffold enhance the conductivity of asprepared scaffold (10-5 x 0.09 S.cm-1) whereas compression modulus and strength of as-prepared scaffold was 4.1 MPa and 1.3 MPa respectively46. Ho et al. developed carbon nanotube-loaded PCL scaffold via rapid prototype technology for cardiac tissue engineering. The mechanical properties of PCL/CNT scaffold were enhanced but the degradability rate decreases since pristine PCL scaffold (90% in 24 h) had faster degradation rate than 1% CNT loaded PCL scaffold (80% in 24 h) and 3% CNT loaded scaffold (70% in 24 h)254. Liu et al. fabricated PLA scaffold via phase separation followed by freeze-drying technique for cardiac tissue regeneration. Cell differentiation of smooth muscle cells (48%) was faster than cardiomyocytes cells (21%)255. Guillemette et al. investigated the poly(glycerol sebacate) (PGS) scaffold utilizing laser microablation method for cardiac tissue engineering. The mechanical test reveals that the ultimate tensile strength of PGS membrane (1200 kPa) was higher than PGS scaffold (550 kPa)256. Sin et al. reported polyurethane scaffold via facile solvent casting/particulate leaching method for cardiac tissue regeneration. They had studied the effect of pore size on compression modulus such as 212-295 µm scaffold (12 kPa) had lower compression modulus than 295-425 µm scaffold (24 kPa) and 425-531 µm scaffold (27


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kPa)257. Hernandez-Cordova et al. illustrated the polyurethane-urea (PUU) scaffold for cardiac tissue regeneration via 3D printing technique. They had studied compression modulus of varying pore size in a wet medium, 500 µm pore size (13.3 kPa) had greater compression modulus than 300 µm (7 kPa) and 400 µm (4.7 kPa) pore size of scaffold258. Baheiraei et al. studied scaffold consist of polyurethane containing aniline pentamer (PU-AN) via solvent casting method for cardiac tissue engineering. The L929 cell viability of as-prepared scaffold was increased with time, the initial optical density of as-prepared scaffold was 0.08 in 24 h, whereas it increases to 0.45 in 96 h259. Polymer scaffolds for cardiac tissue engineering are consolidated in Table 9.



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Table 9. Polymer scaffold for cardiac tissue engineering Material

Method

Cell

Polyhydroxyalkanoate/PCL (PHA/PCL)

Solvent casting

Cardiac Stem Cells

Aniline pentamer-modified polyurethane/PCL

Solvent casting/particulate leaching

Cardiomyocytes

PCL/Carbon Nanotube (PCL/CNT)

3D printing

Cardiac H9c2 cells

PLA

Phase separation/freeze drying

Cardiovascular progenitor cells

Poly(glycerol sebacate) (PGS) Laser microablation C2C12 muscle cells

Polyurethane

Solvent casting/particulate leaching (SCPL) with centrifugation

Human aortic endothelial cells (HAECs)

Results/Properties Tensile strengths Non-porous PHA/PCL scaffold- 6.6 MPa Porous (75 µm) PHA/PCL scaffold- 2.2 MPa Elongation at break Non-porous PHA/PCL scaffold- 525% Porous (75 µm) PHA/PCL scaffold- 600% Compression modulus As-prepared scaffold-4.1 MPa Strength As-prepared scaffold-1.3 MPa Conductivity As-prepared scaffold-10-5 x 0.09 S.cm-1 Degradation rate PCL- 90% in 24 h 1% CNT loaded PCL-80% in 24 h 3% CNT loaded PCL-70% in 24 h Cell differentiation in PLA scaffold Cardiomyocytes cells- 21% Smooth muscle cells- 48% Ultimate tensile strength PGS scaffold-550 kPa PGS membrane- 1200 kPa Compression modulus 212-295 µm scaffold-12 kPa 295-425 µm scaffold-24 kPa 425-531 µm scaffold-27 kPa


Ref.

253

46

254

255

256

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Polyurethane-urea (PUU)

Polyurethane containing aniline pentamer (PU-AN)

3D printing

Primary human cardiac myocytes (HCMs)

Compression modulus in wet medium 300 µm scaffold-7 kPa 400 µm scaffold-4.7 kPa 500 µm scaffold-13.3 kPa

258

Solvent casting

L929 mouse fibroblast and human umbilical vein endothelial cells (HUVECs)

L929 Cell viability (optical density) PU-AN scaffold-0.08 (within 24 h) PU-AN scaffold-0.12 (within 48 h) PU-AN scaffold-0.45 (within 96 h)

259



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3.5 Neural Tissue Engineering 3.5.1 Neural anatomy There is two type of nervous system 1) central nervous system 2) Peripheral nervous system and there are two types of nervous tissue; neurons which is the largest cell and transmit the impulse. Neurological cells are smaller and more abundantly available in the body260. Artificial nerve guidance conduits (NGCs) prevent the limitations of outdated autologous and allogeneic nerve grafts261. A ideal polymer scaffold should have biodegradable, biocompatible, neuroinductive. mechanical robustness, and neuroconductive to perform neural functions262. (as shown in Figure 17)

Figure 17. Typical structure of Neural tissue 3.5.2 Polymeric scaffold for neural tissue engineering Melissinaki et al. reported a PLA scaffold utilizing direct laser writing method for neural tissue engineering. Cell proliferation test shows that neuroblastoma cell successfully proliferated


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from day 1, day 3 and day 5 was 220 ± 15 cells mm-2, 620 ± 20 cells mm-2 and 1335 ± 30 cells mm-2 respectively263. Hoveizi et al. investigated the polylactic acid/gelatin (PLA/G) scaffold via electrospinning technique for neural tissue regeneration. The average fiber diameter was 300 nm and they have found that Human-induced pluripotent stem cells viability of monolayer scaffold was higher than PLA/G scaffold on day 5264. Ebrahimi-Barough et al. blended PLAwith chitosan, these blend nanofibers were fabricated with electrospinning methodology for neural tissue regeneration application. The SEM images confirm that average diameter of blend nanofibers was 100 nm and as chitosan content increases in blend, diameter of nanofibers decreases. The cell viability of PLA/CS (7/3) nanofibers scaffold was higher than the control scaffold from day 3 to day 7265. Liu et al. illustrated use of poly(propylene fumarate)-co-poly(L-lactic acid) scaffold as a neural tissue through a combination of methods like thermally induced phase separation and freeze drying. The mechanical test reveals that compression modulus of the solid scaffold (80 MPa) was greater than porous scaffold (8 MPa) and porosity of porous scaffold was 77.9 ± 2.0% helps in cell attachment and proliferation266. Guo et al. tried the different approach while selecting scaffold material, they blended the PCL with hyperbranched electroactive copolymer like aniline pentamer, to enhance the electrical conductivity of the as-prepared scaffold. Along with electric properties desired mechanical properties were also achieved, e.g. the as-prepared scaffold had 324 ± 28 MPa owing to the pristine scaffold (420 ± 39 MPa) has higher modulus267. Polymer scaffolds for neural tissue engineering are consolidated in Table 10.



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Table 10. Polymer scaffold for neural tissue engineering Material

Method

Cell

Results/Properties

Cell density on PLA scaffold Day 1- 220 ± 15 cells mm-2 PLA Neuroblastoma Day 3- 620 ± 20 cells mm-2 Day 5- 1335 ± 30 cells mm-2 Human-induced Cell viability PLA/gelatin Electrospinning pluripotent stem cells Monolayer- day 3 (hiPSCs) PLA/gelatin scaffold- day 5 Cell viability Human endometrial PLA/CS (7/3) Electrospinning PLA/CS (7/3) scaffold > comtrol stem cells (hEnSCs) scaffold from day 3 to day 7 Compression modulus Poly(propylene fumarate)-coTIPS and freeze MC3T3 mouse preSolid scaffold- 80 MPa poly(L-lactic acid) PPF-codrying osteoblast cells 9% porous scaffold- 8 MPa PLLA 5% porous scaffold- 0.03 MPa PCL and a hyperbranched SolutionModulus HaCaT keratinocyte degradable conducting casting/ Pristine PCL scaffold- 420 ± 39 MPa cells copolymer (PCL/BA) salt-leaching PCL/BA scaffold- 324 ± 28 MPa Effect of nozzle diameter on Porosity Human embryonic PCL Electrospinning (200 µm) PCL scaffold-95% stem cells (1000 µm) PCL scaffold-87% Effect of Electrospinning time on Polypyrrole/poly(ϵPorosity Electrospinning L929 fibroblast cells (1 h) PPy/PCL scaffold- 40% caprolactone) (PPy/PCL) (4 h) PPy/PCL scaffold- 20% Direct laser writing

Poly(3,4-

Freeze drying

Neural stem cells

NSC differentiation (β Tubulin)


Ref.

263

264

265

266

267

268

269

270

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ethylenedioxythiophene) doped with hyaluronic acid (PEDOT-HA) nanoparticles reinforced in chitosan/gelatin (Cs/Gel) matrix

Relative gene expression Cs/Gel scaffold- 1 PEDOT-HA/Cs/Gel scaffold-2.4

Swelling ratio Pristine Cs/Gel scaffold- 3300 ± 700% 3PEDOT/Cs/Gel scaffold- 1500 ± 180% 4PEDOT/Cs/Gel scaffold- 400 ± 50% Change in scaffold diameter Freeze casting DRG neurons Crosslinked scaffold- 4.5% in 24 h Uncrosslinked scaffold- 12% in 24 h NPC proliferation Solution30 mg/mL coated PLGA scaffoldNeural progenitor cell 116.2% on day 3 & 116.2% on day 6 casting/ salt-leaching 1200 mg/mL coated PLGA scaffold150.4% on day 3 & 165.4% on day 6 Cell number of as-prepared scaffold Human embryonic Electrospinning Day 5- 10 stem cells Day 18- 90

Poly(3,4ethylenedioxythiophene)/Chit Feeze drying osan/gelatin Chitosan-alginate

Polypeptide coated poly(lactide-co-glycolide)

Polyurethane

Neuron-like rat pheochromocytoma (PC12) cell


271

272

273

274


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4. Conclusion The growing demand of the tissue engineering scaffolds cannot be fulfilled by the traditional technology such as natural scaffolds or tissue donors. Growing world need advancement in scaffold fabrication techniques, a combination of materials improved properties such as biocompatibility, biodegradability, tensile strength, design development for the additional surface area and cell adhesion. The porous polymer scaffold has higher cell adhesion and proliferation compared to non-porous polymer scaffold owing to small pores that helps the cell by acting as a mechanical anchor. This review focus on the application of the porous polymer scaffold for tissue engineering applications such as skin, bone, cardiac, liver and neural. PCL is a potential biomaterial compared to other biomaterial to use as a polymer scaffold due to its high biocompatibility, good toughness, and superior processability. However its slow degradation rate is a critical issue which is necessary to overcome thus its slower degradation rate can be manipulated by blending with biopolymers having fast degradation rate. Smart technologies undertaking the conventional technology as the scaffold are also getting smarter which is fabricated by the 4D bioprinting. Acknowledgement The authors would like to thank Dr. C. P. Ramanarayanan, Vice-Chancellor, DIAT (DU), Pune, for constant encouragement and support. The authors would also like to acknowledge Mr. Prakash Gore, Mr. Swaroop Gharde and Mr. Deepak Prajapati for technical discussions and support.


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Graphical Abstract


Progress in the Advancement of Porous Biopolymer Scaffold: Tissue

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