Tissue Engineering

## Chapter 5

## Collagen-Based Biomaterial as Drug Delivery Module

Amit Kumar Verma

## Abstract

In the field of medicine, controlled drug delivery has become a major challenge due to inefficiency of drug at critical parameters such as permeability, solubility, half-life, targeting ability, bio- & hemocompatibility, immunogenicity, off-target toxicity and biodegradability. Since several decades the role of drug delivery module has been a crucial parameter of research and clinical observations to improve the effectiveness of drugs. Biomaterials- natural or artificial are mainly used for medical application such as in therapeutics or in diagnostics. Among all the biomaterials, collagen based-hydrogels/ films/ composite materials have attracted the research and innovations and are the excellent objects for drug delivery, tissue engineering, wound dressings and gene therapeutics etc. due to high encapsulating capacity, mechanically strong swollen structural network and efficient mass transfer properties. Substantial developments have been performed using collagen-based drug delivery systems (DDS) to deliver biomolecules with better efficacy. In spite of significant progress, several issues at clinical trials particularly targeting of intracellular molecules such as genes is still a challenge for researchers. Experimental results, theoretical models, molecular simulations will boost the fabrication/ designing of collage-based DDS, which further will enhance the understanding of controlled delivery/mechanism of therapeutics at specific targets for various disease treatments.

Keywords: collagen, biomaterial, drug delivery systems (DDS), drug, hydrogels, films, composite material

## 1. Introduction

Collagen is unique and major structural protein of extracellular matrix (ECM) and plays crucial role to the structural integrity of tissues/organs and cellular growth in vertebrates and other organisms, constitute around 30% of the total protein content of mammal's body, involved in mechanical protection of tissues and organs such as skin, tendons, ligaments, bones, cartilage, blood vessels, cornea and nails etc. (Figure 1) [1–6]. More than 50% in the skin and more than 90% of extracellular proteins in the tendon and bone is made up of collagen [7, 8].

Figure 1.

Collagen's occurrence in different body tissues.

The unique feature of collagen molecule is triple helix structure made by three identical or non-identical polypeptide chains. Each polypeptide chain comprises around 1000 amino acids and the chains are supercoiled in left handed manner around the axis with staggering of residues between adjacent chains, give rise to triple helix right handed structure [9]. Each chain is having the repeated sequence of (Gly-X-Y)n, whereas X and Y are mostly proline and hydroxyproline residues [1, 10, 11]. At present, 30 different types of collagens have been characterized and reported in literature [12]. Among all the different variants of collagen, type I, II, III, V and XI represent more than 90% of human fibrillar collagen and is majorly distributed in dermis, hair, bone, cartilage, ligament, tendon and placenta [13]. Other types such as type IV and VIII form the network frame of basement membranes [12, 14]. Type I collagen has very important role in medicine as well as in development of medical devices, artificial implants, drug carriers for controlled release and scaffolds for tissue regeneration [5, 15, 16]. Being the highly versatile, structurally unique and biocompatible protein substance, the collagen can be developed into different types of drug/active substance carrier module such as hydrogels, microparticles and films etc. [3, 7].

## 2. Collagen as a biomaterial

According to Hench and Erthridge, 1982 [17], "a biomaterial is used to make devices to replace a part or a function of the body in a safe, reliable, economic, and physiologically acceptable manner." Raghavendra et al. [18] has mentioned the other definitions of biomaterial are "materials of synthetic as well as of natural origin in contact with tissue, blood, and biological fluids, intended for use for prosthetic, diagnostic, therapeutic, and storage applications without adversely affecting the living organism and its components" [19] and "any substance (other than drugs) or combination of substances, synthetic or natural in origin, which can be used for any period of time, as a whole or as a part of a system which treats, augments, or replaces any tissue, organ, or function of the body" [20]. Application of biomaterials in physiological systems is possible due to competent and stable features of biomaterials [21] which can be achieved with proper combination of mechanical, physical, chemical and biological attributes [22]. Modern biomaterials are designed and developed singly or in combinations of polymers, metals, composite materials and ceramics etc. [18].

In the field of medicine, collagen is one of the most studied biomaterial or biopolymer and according to Cheng et al. [23], around 260,000 literature articles (at present, the number is many more than reported) reported it as pivotal component in tissue regeneration and so called as 'the steel of the biological material". Collagen is the main biopolymer of ECM of vertebrates and invertebrates and has the capability to interact with large number of biomolecules leading to various biological reactions/changes under normal or pathological processes in the body, inspiring the scientists to develop the various formulations based on collagen [24–26]. The attributes for the testimony of collagen's usage in wide scenario of medicine are its cosmopolitanism, high biocompatibility, hemocompatibility, biomimetic and biodegradability and to make composite biopolymer with biomaterials like chitosan (CHS), alginate (ALG), cellulose (CL), hyaluronic acid (HA), glycosaminoglycans (GAGs) as well as synthetic materials like carboxy methyl cellulose (CMC), poly vinyl alcohol (PVA), poly ε-caprolactone (PCL), poly ethyl methacrylate (PEMA) etc. in different formulations [27–29].

Now it is well established that collagen has good elasticity, physical, mechanical, enzymatic and thermal stability inside the body environment, but after extraction and utilization these properties are compromised at large scale [30–32], so the additional need of chemicals, chemical and physical processes are required to develop the stable biomaterial. Though extracted collagen is a promising drug delivery material in the field of ophthalmology but can be easily degradable in vitro due to disruption of natural intermolecular crosslinks of lysine and hydroxylysine residues during isolation and purification process [33]. The weaknesses of extracted collagen such as mechanical and thermal strength, enzymatic degradation can be reduced with the help of various methods of chemical, physical and enzymatic crosslinking, covalent conjugating, grafting polymerization or blending. The blended collagen-based biomaterials like hydrogels, films, microspheres and nanoparticles (NPs) have low immunogenicity, good absorption, hemostatic property and synergism with other bioactive compounds or loaded drugs and remain unaltered after several processes. The requirement of mechanical, pH, enzymatic, thermal stability is not provided alone by collagen for a controlled DDS but can be achieved with combination of CHS or ALG for release of analgesics, chemotherapeutic molecules and natural bioactive agent like curcumin or aloe vera. The blended biopolymer will have the combined desired properties of the separate material. Hydrogel of blended collagen-CHS have antibacterial, antifungal, anti-carcinogenic and immunogenic attributes [34–37]. Blended collagen-CL based


1. and abundant biopolymers: Sources, properties, and potential applications.

Natural

### Collagen-Based Biomaterial as Drug Delivery Module DOI: http://dx.doi.org/10.5772/intechopen.103063

hydrogel has properties like biomimetic and hemostatic from collagen while mechanical strength and antibacterial characteristic from CL. In the field of ophthalmology blended collagen-based hydrogels with CHS or ALG are applied for corneal disease treatment exhibiting good mechanical and thermal attributes along with transparency. The films, microspheres, membranes and scaffolds of blended collagen-based materials are used in wound dressing due to moisture retaining, low adhesion, absorption of blood and tissue exudates, anti-infective and permeability properties [32, 38–41]. The sources and possible applications of most utilized material of blended biopolymers along with important features like stability, toxicity and biomedical are presented in Table 1.

## 3. Crosslinking for collagen-based biomaterial

Collagen is the most abundant vertebrate protein and mostly used biopolymer and with variety of physiological features like biocompatible, low immunogenic, self assembling fibril formation etc. The collagen is mechanically strong and durable in vivo but after the isolation and purification processes is vulnerable to degradation in vitro due to dissociation of natural crosslinks and assembly structure by neutral salt, acid alkali or proteases and quality of extracted collagen is inferior to native state (Figure 2) [42, 55]. The researchers are attempting to make the suitable collagenbased biomaterial with properties of increased mechanical strength, reduced enzymatic degradation, stability, solubility and low toxicity by introducing exogenous crosslinking [42, 56]. The introduced intermolecular crosslinking prevents the unknotting of collagen fibrils produced by heat and advancing the thermal stability along with increased tensile strength, stiffness, compressive modulus and decreased extensibility [57–60]. The exogenous natural or chemical crosslinks could significantly reduce the enzymatic degradation of collagen by blocking the cleavage site [61].

Figure 2. Native crosslinking in the collagen molecules.

Though collagen-based crosslinked biopolymers are producing significant results in the field of biomedicine and biotechnology, still no standard method is applicable to the formulation of improved non-toxic and biocompatible hydrogels, films and matrices etc. Different crosslinking strategies such as chemical, physical or enzymatic are applied to achieve the collagen-based materials with desired properties for drug delivery and other applications [62].

#### 3.1 Chemical crosslinking

The most used crosslinking method is with chemical agents due to ease of application, less time consuming and cost effective. The most commonly used chemical reagents are formaldehyde (FA) and glutaraldehyde (GTA). FA reacts with the εamino group of lysine and hydroxylysine residue of collagen to form imine as an intermediate followed by crosslink with tyrosine or with amide group of asparagine or glutamine residue. FA-crosslinked products generated brittleness, significant toxicity and unfavorable reactions, hence not preferred in biomedicine [7]. Another agent GTA is widely utilized for crosslinking of collagen, based on high reactivity and low cost. The low concentration of FA and GTA produced brittle and low uniformity of composites while high concentration led to major cytotoxic effects. Several methods have been applied to remove the unreacted GTA, due to its cytotoxicity the use of GTA at present scenario is still debatable [28, 42]. Hexa-methylene-diisocyanate (HDC) was used as an alternative to GTA, but in contrast to GTA, HDC showed less severe primary and secondary cytotoxicity during cell proliferation as compared to non-crosslinked material [63]. Charulatha & Rajaram [64] evaluated the biocompatibility of collagen membranes cross-linked with 3,3<sup>0</sup> -dithio bis-propionimidate (DTBP) and dimethyl suberimidate (DMS). Both DTBP and DMS showed lower toxicity than GTA and as the better substitute for crosslinking. The polyepoxy compounds such as ethylene glycol diglycidyl ether, glycerol polyglycidyl ether and methyl glycidyl ether were used as crosslinker [65]. The epoxy group reacts with amino group of lysine residue for crosslinking similar to GTA. The polyepoxy crosslinked material showed acceptable cytotoxicity [66].

Due to non-toxicity and water solubility, the carbodiimides and acyl azides showed significant crosslinking with collagen. Carbodiimides like 1-(3-dimethylaminopropyl)-3-ethylcarbodiimide (EDC) and N-hydroxy succinimide (NHS) are better candidate than aldehydes, HDC and polyepoxy compounds due to formation of amide bonds between –COOH and –NH2 group of collagen without becoming the part of actual linkage. Van Wachem et al. [67] compared the four crosslinking methods i.e. GTA, HDC, acyl azide and EDC for the assessment of biocompatibility and tissue regeneration ability and EDC crosslinked material showed best results among the four tested methods. Pieper et al. [68] showed that EDC crosslinked collagen expressed no cytotoxicity, slow enzymatic degradation and decreased calcification.

Non-toxic, biodegradable and biocompatible natural compounds as promising cross-linkers like alginic acid (anionic block copolymer) from brown algae, iridoid compounds from the fruits of Gardenia jasminoides and Genipa americana [69], oxidized alginate [70], dialdehyde starch [71], D, L-glyceraldehyde [72], natural polyphenols and dialdehyde- carboxymethyl cellulose [73] have been broadly examined. In recent years usage of polyphenols such as caffeic acid (CA) and tannic acid (TA) [74], proanthocyanidin [75], procyanidin [76], epigallocatechin gallate (EGCG) and epicatechin gallates (ECG) [77] and other tannins [78] have been increased due to antioxidative, anti-inflammatory, antimicrobial, cardioprotective, antithrombotic,

pharmacological and therapeutic possibilities. According to Jackson et al. [77] TA, EGCG and ECG showed stabilization of collagen implant for long period at very low concentration and protection against collagenase more effectively than GTA and carbodiimides. In contrast to GTA, natural crosslinkers have disadvantages like long term storage and degree of crosslinking.

## 3.2 Physcial crosslinking

Physical techniques like ionizing radiation (X-ray and γ-ray), UV light and dehydrothermal treatment and dye-mediated photo-oxidation are used as crosslinkers for collagen based formulations. The physical crosslinking relies on the factors like amount of radiation, temperature, hydration conditions, electron beam intensity and UV denaturation [43]. Photosensitizer riboflavin in combination with UV-irradiation produced the similar results of crosslinking by GTA without cytotoxicity to reduce the harmful effects of physical and chemical crosslinking [45]. UV- light mediated crosslinking generated the denaturation and conformational modifications of collagen molecules which opposed the stabilization of UV-induced crosslinked product [79]. Dehydrothermal treatment resulted into the complex of collagen with anatomically accepted structures without contraction, curling or deformity for longer duration without the involvement of chemical crosslinking [80]. In contrast to chemical crosslinking the heating disrupted the triple helical conformation of collagen and increased the degradation by enzymes [81]. Overall the physical crosslinking methods are simple and safe for the production of splendid biocompatible biomaterials in comparison to exogenous cytotoxic chemical crosslinkers.

#### 3.3 Enzymatic crosslinking

Enzymatic approaches have gained interest due to brilliant specificity and accurate reaction kinetics and to surmount the difficulties with chemical methods. Enzymatic crosslinkers can be categorized into oxidoreductases, transferases and hydrolases based on the catalytic reaction [82]. The oxidative enzymes tyrosinase and laccase [83] along with acyltransferase-transglutaminase have the capacity to modify the protein substrate to enhance the quality of crosslinked biomaterials [84]. Transglutaminases are calcium dependent and catalyze the reactions in broad range of pH and temperatures. Microbial origin biodegradable transglutaminases can catalyze the crosslinking in the concentration dependent manner [85, 86]. Exogenous lysyl oxidase catalyzes the lysine residue into highly reactive aldehydes leading to the formation of crosslinks in the ECM proteins. The pre-treatment of lysyl oxidase promoted the maturation of native and engineered collagen tissues both in vitro and in vivo with reference to increased tensile modulus and pyridinoline crosslinking [87].

Comparing all the crosslinking strategies the chemical method is the most preferred due to generation of consistent and high degree of crosslinking [88]. The physical method is used as a appurtenant crosslinking approach. Enzymatic approach alleviates the shortcomings produced by chemical and physical crosslinking methods, but it is time consuming and expensive [89]. In recent years, the use of natural and eco-friendly biocompatible crosslinkers has increased and become the very promising agents. The natural substance- genipin obtained from irridoid glucoside (geniposide) is one of the important crosslinker with great potential in the field of biomedicine. Genipin is very expensive in contrast to other natural crosslinkers, while TA is cheap and easily accessible compound for crosslinking [89].

## 4. Collagen-based formulations for drug delivery

Traditional drugs have been the main concern to effectively treat the several diseases. The introduction of classical drug in therapeutics at high concentrations generally develops the substantial and sometimes severe consequences. The development of effective DDS for the efficacious augmentation of particular drug at desired target and at optimal concentrations for necessary duration has been the critical concern of clinical investigations and research since several years. To achieve the targeted or controlled local drug delivery, both synthetic and natural drug delivery materials are playing the important role.

At present, DDS have been developed based on polymers, nanomaterials and lipids etc. for the attachment or encapsulation of drugs to target the delivery or controlled release for long duration [90, 91]. Collagen based biomaterials or composite materials have become the important DDS due to specific pore size for active drug or principal load, effective fibrillar network, enzymatic degradation, long term stability in vivo, biocompatibility, low antigenicity, highly reduced toxicity and safety features [92]. The chemical modifications of –OH, –NH2 and –COOH groups on collagen molecule making it more relevant and promising candidate in the domain of biomedicine. Collagen based biomaterials can be formulated into different forms according to desired medical application of drug delivery as mentioned in Figure 3. Various types of collagen based bioformulations such as hydrogels, films, sponges, scaffolds, matrices, aqueous injections, microspheres, micro-particles, micro-beads, NPs, nano-composites, nanofibres, shields, inserts, tubes, coatings, monolithic devices, implants and dressings etc. are applied for various drug delivery applications, tissue growth and regeneration. In Table 2, the type of formulations, synthetic or natural polymeric support material, active compound or drug and specific biomedical application has been summarized.

Figure 3.

Different types of collagen based biomaterials used for drug delivery.







Table 2.

Application of collagen and collagen based material in drug delivery systems (DDS).

## 4.1 Gel/hydrogel

Hydrogels are three dimensional (3-D) crosslinked arrangements of similar or different types of polymeric molecules with the property of absorbing and retaining the optimum quantity of water or biological fluid without degrading or losing the network structure. The material can be categorized as hydrogel, if the water content in the material is at least 10% of the total weight or volume of hydrogel [38]. The water molecules in the hydrogel provide the freedom of flexibility to design the natural tissue like environment. Hydrogels can be synthetic or natural with different chemical constituents and having different mechanical, physical and chemical attributes according to biomedical application. Hydrogels can be hydrophilic or hydrophobic in nature. Hydrophilic hydrogels possess hydroxyl (–OH), amine (–NH2), carboxyl (–COOH), amide (–CONH–CONH2) and sulphonic (–SO3H) group for the swelling and absorbing property. The hydrophobic polymers show low swelling feature despite having improved mechanical, physical or chemical strength.

Drug transport within hydrogel can be regulated by altering the network/mesh size or the interactions with drugs using chemical methods [172, 222]. If the loaded drug is smaller than the crosslinked network of hydrogel, it can simply diffuse through the hydrogel while the larger drug molecules are entrapped within the hydrogel network and can be released after the degradation of the mesh. The biopolymer and its crosslinked network can be degenerated via slow hydrolysis of peptide or ester linkage or cleaving the thiol-related bonds, or through the enzymatic activity [223]. By the incorporation of non-covalent or covalent drug-matrix interactions, drug release from the hydrogel can be tuned [224, 225]. The characters mainly mesh size, crosslinking chemistry and drug interactions facilitate the better handling of drug transport through hydrogel. ECM-based hydrogels are the preferred choice for local drug delivery due to mechanical and biochemical support through cell-matrix interactions and diffusion and infiltration of small drug molecules in between crosslinked polymeric network [172].

Collagen alone is used for the delivery of several drugs and active principles such as keterolac, nerve growth factor (NGF)-β, TA, curcumin and royal jelly etc. for various biomedical purposes like anti-inflammation, corneal regeneration, drug release and kinetics, angiogenesis and wound healing, respectively. According to Ramírez et al. [96] type I collagen hydrogel extracellular vehicles (EVs) with Apis mellifera royal jelly displayed effective wound healing to stimulate mesenchymal stem cell (MSC) migration and inhibition of biofilm formation by Staphylococcus aureus along with stable release kinetics up to 7 days. Curcumin crosslinked collagen aerogel system expressed the enhanced physical and mechanical features and controlled antiproteolytic and pro-angiogenic potential, made them appropriate 3D scaffolds for medicine purposes. Here, curcumin as a nutraceutical was used as a crosslinker for further usage [97]. In a study by Chuysinuan et al. [98], gelatin-based hydrogel was prepared by mixing gelatin with GTA crosslinker and essential oil of Eupatorium adenophorum and applied on patients with open wound to assess the wound healing and anti-bacterial properties by analyzing release profiles.

Collagen-based composite materials with synthetic (CMC, PVA, PCL, PEMA, HEMA, polyurethane (PUR) etc.) or natural (CL, bacterial cellulose (BC), CHS, ALG, HA, GAGs etc.) polymers have major roles in biomedicine. Liu et al. [46] developed a composite of collagen-ALG with suitable mechanical strength and optical clarity to support human corneal epithelial cell growth using bovine serum albumin (BSA) as a model drug. The hydrogel system could be applied as therapeutic lens in patients with corneal illnesses. Collagen-based hydrogel preparation by mixing of acrylamide and 2-hydroxy ethyl methacrylate (HEMA) was used as DDS for linear release profile of gallic acid (GA) and naproxen up to 36 hours for wound healing. Addition of metal NPs such as Ag and Cu in this collagen-HEMA hydrogel films showed antimicrobial potential against Escherichia coli, Bacillus subtilis and S. aureus [29]. Bettnini et al. [99] prepared the porous collagen-based hydrogel scaffold with iron oxide (Fe3O4) NPs for the release of fluorescein and biocompatibility, cell viability of 3 T3 fibroblasts cells and proposed the safety and applications in tissue engineering and drug delivery. Reis et al. [100] developed the collagen-CHS thermoresponsive hydrogel conjugated with angiopoietin-1 derived peptide glutamine-histidine-arginine-glutamic acidaspartic acid-glycine-serine peptide (QHREDGS) for the survival and maturation of cardiomyocytes. Hydrogels with high peptide load showed better morphology, viability, metabolic activity, success rate of beating as compared to hydrogels with low peptide concentration and control groups.

Graphene oxide (GO) sheets were inserted into the collagen-based hydrogels for the controlled release of fibroblast growth factor (FGF)-2 to induce pluripotent stem cell culture. Low permeability of GO sheets allowed the release of FGF-2 in controlled and regulated fashion while the FGF-2 interacted with collagen through electrostatic forces and partial hydrogen bonding. The release profile of FGF-2 was attained up to 400 hours using three different concentrations of GO and showed the fabricated hydrogel for better release of growth factors (GFs) for biomedical application [102]. Choi et al. [103] developed the collagen-TA-poly ethylenimine (PEI) hydrogel with layer-by-layer self-assembled films to overcome the problem of poor mechanical strength and fast release of inserted drug. Doxorubicin (DOX) was used as model cancer therapy drug. The multifunctional hydrogels showed sustained and controlled release up to 6–7 days without any cytotoxic effects along with antibacterial property against Gram positive and negative bacteria and higher strength to compression load.

Injectable hydrogels are the promising materials for cancer treatment and controlled delivery. With the minimal invasive processes injectable hydrogels can be located and remained at required position and also mitigate the irregular shape defects after the implantation. Aqueous injections of hydrogels could be used in biomedicine field such as drug release, wound healing, repair, tumor treatment, tissue regeneration, ocular/retinal disorders and cancer therapy etc. Fan et al. [115] designed the hydrogel prepared from tilapia skin collagen and CHS for the delivery of model nanobodies- 2D5 and KPU. The hydrogel was biodegradable and expedited the release of nanobodies and could pave the way for tumor treatment. Carboxymethyl cellulose (CMC)-collagen based aqueous injectable hydrogel showed promising antioxidative and drug carrier benefits to treat the retinal ischaemia or reperfusion injury in rat models and could be applied for drug based treatment of retinal illnesses in humans. Animals were treated with interleukin (IL)-10 loaded hydrogels and expressed better therapeutic results of restoration of retinal structures and reduced retinal apoptosis, significantly decreased retinal oxidative stress in comparison to control group [52].

#### 4.2 Films/membranes

The films or membranes are very thin and flexible layer of biopolymers with or without plasticizer having optical and mechanical anisotropy with very high tensile strength making them suitable for various medical applications of sustained drug release, cell adhesion, proliferation, differentiation, cancer treatment, wound healing and tissue regeneration. The thin films are the prominent material to target sensitive

locations not possible with other formulations like liquid or tablets [226]. Thin films exhibited the improvised onset of drug activity, decreased dose quantity or frequency and augmented drug efficacy, reduced side effects by drug and extensive metabolism [227, 228].

Gil et al. [122] developed the innovative chromium free-collagen film for slow drug release carrier for skin burn related complications like ulcers and infected wounds. The biocompatible films were tested for drug silver sulfadiazine and its antibacterial potential against Pseudomonas aeruginosa, E. coli, Micrococcus luteus, S. aureus, Proteus vulgaris and Klebsiella pneumoniae. The findings proposed the effective strategy for Chromium (Cr) removal from leather waste and generation of environment friendly material could be transformed into collagen films, promising candidate for drug carriers. The use of biocompatible, recyclable, biodegradable natural materials has become tremendously increased in recent decades. Langasco et al. [123] developed the natural collagen films from marine sponges for topical drug delivery application. L-cysteine loaded films were analyzed for different drug concentrations and drying parameters. The films showed the healing potential of cysteine, acted as biocompatible carrier to absorb excess of wound exudate along with drug release. The films could be the promising material and might behave as bioactive, biomimetic drug carrier for effective wound healing. Jana et al. [130] synthesized the fish scale collagen-carboxymethyl guar gum film loaded with broad spectrum antibiotic ceftazidime. Around 90–95% of ceftazidime was released after 96 hours at physiological pH. In vitro study on NIH 3 T3 fibroblast cell line showed the biocompatibility of crosslinked film and antibacterial results exhibited the inhibition of S. aureus and P. aeruginosa.

Collagen-CHS based films/membranes are important biomaterials and used for antibiotic release, wound healing, cancer treatment, transdermal delivery, cardiac illness, tissue engineering. Martino et al. [132] formulated the collagen-CHS film for the delivery of mixture of local anesthetics compounds- lidocaine, tetracaine and benzocaine. The films were developed by rapid, cost effective and highly reproducible casting approach. The films showed good mechanical strength and flexibility with high water permeability. The anesthetics were uniformly distributed in the film and controlled released from 6 to 24 hours. The film exhibited in vitro non-cytotoxicity, cell proliferative and biocompatibility properties against human dermal fibroblast cells making it better candidate for drug release and proliferation. Liu et al. [136] fabricated the collagen-CHS-GO composite film using EDC as crosslinker loaded with basic FGF for effective wound healing through controlled release. The film had improved thermal endurance and higher degree of crosslinking for advanced mechanical strength due to GO. This novel DDS prevented the initial sudden release and loss of bioactive potential of basic FGF in vivo and in vitro. In cultured L929 fibroblasts, the film showed good biocompatibility in terms of cell adhesion and proliferation. These films were implanted on rats showed the wound remodeling to repair full thickness skin wound. So these films were promising substitute as wound dressing material for drug delivery with wound healing. Daja et al. [27] developed the collagen/PVA anionic membrane for drug carrier of ciprofloxacin hydrochloride along with antibacterial efficacy and its application in the treatment of ulcerative keratitis. The membrane provided the sustained DDS and inhibited the growth of S. aureus and E. coli during 48 hours. The membrane had proper mechanical strength, water amount, hydrophilicity, permeability and pH without any stress to cornea during interaction. The collagen fibrils in membrane decreased stromal damage and improved the epithelium regeneration. The formulated membranes were cost-effective and secured biomaterial for the treatment of corneal ulcers in patients.

#### 4.3 Scaffolds/sponges/matrices

Scaffolds are collagen sponges or matrices with three dimensional network structures. Scaffolds can be obtained with various synthesis approaches of freeze drying, electrospinning and 3D printing etc. The freeze drying is the most effective method preserving the structure and native or inherent properties of collagen along with loaded drug/active principle in the scaffold. Collagen-based scaffolds or matrices are the important and favorable materials for bone, skin and tissue regeneration, angiogenesis, gene delivery, wound healing and repair, sustained drug release and improved gene expression. Now a days several commercially available collagen-based sponges in the market are Collarx®, Collatamp® G, Collatamp®EG Sulmycin® Implant, Garamycin® Schwamm, Duracol®, Duracoll®, Gentacol®, Gentacoll®, Garacol®, Garacoll® and Cronocol® - Gentamicin surgical implants.

Elangovan et al. [141] developed the non-viral gene delivery system for bone regeneration with the help of collagen scaffold to deliver the PEI-plasmid DNA encoding platelet derived growth factor (PDGF)-B complexes. The complexes expressed low cytotoxicity and markedly higher proliferation of human bone marrow stromal cells in contrast to scaffold without DNA and PDGF-B. In rats model the complexes exhibited higher bone volume followed the 4 weeks of grafting in comparison to empty scaffolds. The results advocated the use of non-viral scaffolds for bone regeneration along with gene delivery vehicle in clinical applications. Collagen-based biopolymers are one of the most important biomaterials to formulate the matrices in the field of tissue engineering due to significant non-toxicity, biocompatibility and resorptive potential. López-Noriega et al. [150] designed the collagen-hydroxyapatite (HAP) scaffold with covalently attached thermoresponsive liposomes. The encapsulated drug with pro-osteogenic and anti-osteoclastic properties was PTHrP107–111, a pentapeptide. The regulated release of pentapeptide was correlated with enhanced expression of osteopontin and osteocalcin genes in cultured pre-osteoblastic MC3T3- E1 cells. This scaffold medicated drug release and cell regeneration has vast potential for various types of tissue regeneration.

Collagen-based 3D biomaterials are broadly used in the field of biomedicine for their properties of biocompatibility, inherent bioactivity to induce cell proliferation, hemostatic and low antigenicity. A porous and highly structured biomaterial such as sponges or matrices promotes the flexibility, permeability and biomimicry [229, 230]. Crosslinking or amalgamation of natural or synthetic biopolymers improvises the shortcomings of collagen polymer alone in terms of physico-chemical and biological parameters. Alagha et al. [166] prepared the porous muco-adhesive collagen-CHS biosponge as DDS for dexamethasone to treat the oral mucositis. The sponge was characterized by X-ray, FTIR, SEM, DSC and swelling behavior. The collagen-CHS sponge showed regulated drug release up to 10 hours as compared to collagen sponge for 5 hours. David et al. [168] fabricated the collagen-PCL-HA macroporous sponge to deliver the model drug methylene blue and curcumin. Several parameters such as absorption, water uptake, drug loading and delivery along with mechanical and structural features were examined for the developed sponge. In comparison to control group, the sponge showed sustained release kinetics for drugs and making the sponge as future material for applications of wound dressing and lab models.

Collagen matrices are able to deliver the gene or plasmid DNA in cultured cells and alter the gene expression in tissue engineering. Orsi et al. [171] formulated the bioactivated collagen-PEI-DNA complex to control the gene expression and attract the

specific cell type. The transfected NIH3T3 cells with matrix-PEI-DNA complex secreted the plasmid encoded protein to promote the tissue repair and regeneration. The developed matrix could be the new approach for tissue repair.

#### 4.4 3-D microspheres/ microparticles/micro-beads

Microspheres, microparticles or micro-beads are spherical particles with large surface to volume ratio for improved drug delivery, growth factors and broad surface area for cellular interactions to other biomolecules [28]. The size of collagen-based microparticles ranging from 3 to 40 μm. Berndt et al. [231] developed the collagen microspheres encapsulated astrocytes crosslinked with poly (ethylene glycol) tetrasuccinimidyl glutarate (4S-StarPEG) as growth enhancing and carrier for injured spinal cord. Astrocytes were transfected with plasmids encoding nerve growth factor (NGF)-ires-enhanced green fluorescent protein (EGFP) genes and then added to the culture of rat dorsal root ganglion and significantly improved growth was observed. The report showed the potential of microspheres as carrier of astrocytes for neural tissue regeneration. Zhang et al. [53] formulated the 3-D microsphere of collagen-BCbone morphogenic protein (BMP)-2 for bone tissue augmentation. The 3D microporous microspheres effectively enhanced the adhesion, proliferation and osteogenic differentiation of mice MC3T3-E1 cells and expressed adequate biocompatibility.

Marine based collagen from jellyfish species Catostylus tagi was used to develop the microparticles formulation for sustained delivery of lysozyme and α-lactalbumin. The collagen microparticles were crosslinked with EDC and investigation of lysozyme activity was retained throughout the crosslinking and encapsulation process. These microparticles from marine collagen could be the promising material for controlled release of therapeutic proteins [186]. Yang and Fang [232] developed the microporous nano-HAP/collagen/phosphatidylserine scaffolds embedding collagen microparticles for the sustained release of steroidal saponins for bone tissue engineering in cultured MC-3 T3-E1 cells. The scaffolds provided scope for spatial and temporally controlled drug delivery and deposition at wounded site and reduction in adverse side effects. Vardar et al. [190] formulated the novel injectable collagen-fibrin microfluidic system loaded with recombinant insulin like growth factor-1 (α2PI1–8-MMp-IGF-1) to treat the urinary incontinence. The natural crosslinker genipin was used for collagen modification. The microbeads showed slow release of GF and positive cell behavior for the induction of in vivo smooth muscle regeneration for effective management of urinary incontinence.

#### 4.5 Nanoparticles/ nanocomposites/nanofibres

Nanomaterials within the size of 1–100 nm are one of the best materials with admirable biochemical and pharmacological attributes [233]. Biological protein-based NPs are applied in different applications due to eco-friendly nature and biocompatibility and replacing the synthetic materials. Collagen is the preferred NP substance and by direct or indirect crosslinking to collagen NPs provide better substitute for protein based drug delivery vehicle [234]. The crosslinking of collagen with NPs is the new strategy to enhance the mechanical and physical strength of collagen tissue for various applications. Metal oxide NPs such as iron oxide (Fe3O4), zinc oxide (ZnO), alumina oxide (Al2O3), tantalum oxide (TaO), Al2O3-ZrO2 and Fe3O4-ZnO improve the mechanical features of collagen-based biomaterials [99, 195, 235]. The metallic NPs provided broad spectrum antimicrobial, antioxidative and anti-inflammatory

attributes [236, 237] to collagen based material and serve as an alternative to toxic chemical crosslinkers and impede the collagen degradation by physical crosslinking. These features advertise the use of NPs based collagen biopolymers in medical field such as targeted controlled drug delivery, cell targeting and tissue engineering. Choi et al. [195] prepared the collagen hydrogel containing the ferritin NPs, TaO NPs along with transforming growth factor (TGF)-β1 for the controlled release and imaging medium for regeneration of oral tissues.

NPs based biopolymers for cancer therapeutics are effectively utilized in recent years due to targeted and controlled release kinetics. Anandhakumar et al. [193] fabricated the collagen peptide-CHS NPs for encapsulation of standard cancer drug DOX in cancer therapy. The NPs showed high encapsulation capacity of DOX and pH regulated release. NPs with DOX expressed significant anti-proliferative activity against HeLa cells in contrast to normal cells. The NPs showed excellent biocompatibility with high power as smart DDS for cancer therapeutics. Zhang et al. [47] designed the collagen-ALG biocomposite doped with silver NPs (AgNPs) with antibacterial potential and applicable as wound dressing. The biocomposite exhibited insignificant in vitro toxicity at lower concentrations of AgNPs and also inhibited the growth of S. aureus and E. coli. Saska et al. [197] fabricated the nanocomposite of BC-collagen-apatite and osteogenic growth peptide (OGP) for bone tissue regeneration. The OGP containing nanocomposite triggered the early development of osteoblastic phenotype and elevated cellular growth without any cytotoxic, genotoxic or mutagenic adverse effects. The nanocomposite could be the promising future biomaterial for bone tissue engineering.

#### 4.6 Inserts/shields/monolithic devices or pellets

The approach of applying ocular collagen inserts to administer the drug for prolonged period was started in early 1970s. The inserts were films or as molded rods or wafers of collagen incorporated with drug such as pilocarpine, penicillin-procaine, erythromycin, erythromycin esolate and gentamicin etc. in the form of eyedrops, ointments and subconjunctival injections to treat the cornea related disorders [200–202]. In the late 1980s researches on inserts were overtaken by shields which became commercially available in the market in reproducible manner [7].

Collagen shields were formulated as corneal bandages/dressings to facilitate the wound healing, allow sufficient oxygen transmission, lubricate the eye surface to minimize stress and to regenerate the corneal epithelial linings after corneal injury or damage, transplantation, radial keratomy, glaucoma, keratitis or cornea related disorders [238, 239]. These shields could be used as carrier to deliver the ophthalmic medication such as water soluble antibiotics- gentamicin, vancomycin, tobramycin, netilmicin, polymyxin B sulphate, trimethoprim, amphotericin B, pilocarpine and flurbiprofene sodium etc. [7]. The drug delivery aspect of shields is limited by transparency, reduced visual acuity, slight irritation, complex administration procedure and prolonged durability. In current scenario, the commercial formulations like Biocora®, ProshieldO®, MediLenso®, Irvine® and Chiron® etc. are showing better future for delivery of corticosteroids and subconjuctival antibiotics etc. [240].

Collagen minipellets are cylindrical injectable controlled release drug delivery vehicle. In 1992, Takeuchi [211] prepared the little rods of 1 mm in diameter and length of 15 mm of injectable collagen minipellet for the local delivery of minocycline and lysozyme to treat periodontitis. Fujioka et al. [241] used the injectable collagen

minipellet to deliver interleukin (IL)-2 molecule. Maeda et al. [242] used the collagen minipellet as a carrier to deliver the recombinant human bone morphogenic protein (rhBMP)-2 to induce the bone formation in mice models. Lofthouse et al. [243] developed the degradable collagen minipellet infused with avidin and IL-1β as vaccine carrier for clostridial antigen into sheep and mice. Higaki et al. [244] fabricated the biodegradable collagen minipellet to deliver the tetanus and diphtheria toxoid as single dose vaccine delivery system in mice.

## 5. Conclusion

Protein-based biomaterials have excellent biocompatibility with minimal cytotoxicity and biodegradability and theirs physical, chemical and biological parameters can be altered according to biomedical application. Collagen, a major protein in animal body is the attractive biopolymers for the delivery of therapeutic drugs, growth factors, hormones, proteins/enzymes, gene and imaging probes in the field of drug delivery systems, wound healing, bone grafts, implants, tissue regeneration, ocular diseases, cosmetic surgery, reconstructive surgery and cardiac treatments.

The researchers are consistently designing the protein-based hybrid materials with desired physical, mechanical, chemical and biological properties. Collagen-based hybrid biomaterials can be formed with natural polymers like CHS, CL, ALG, gelatin, HA, CHD, HAP or the chemically modified form of these natural polymers or with synthetic polymer such as CMC, PVA, PCL, PEMA, PLA and PLGA etc. through various physical, chemical and enzymatic crosslinking approaches. Collagen-based biomaterials can be fabricated into variety of physical forms hydrogel, films or membranes, scaffolds or sponges, matrices, 3-D microspheres, microparticles, nanoparticles, nanocomposites, inserts, shields and pellets for drug based delivery of synthetic and natural active biocomponents in various fields of medical science. Collagen-based biomaterials have attracted the researchers to develop efficient and controlled therapeutic vehicles for clinical applications ensuring the patient compliance. The more efforts are needed to translate the clinical results into production scale with collaboration of researchers, material scientists, clinical doctors and industry.

## Author details

Amit Kumar Verma Department of Biosciences, Jamia Millia Islamia, Srinivasa Ramanujan Block, New Delhi, India

\*Address all correspondence to: averma@jmi.ac.in

© 2022 The Author(s). Licensee IntechOpen. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

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## Collagen Based 3D Printed Scaffolds for Tissue Engineering

*Sougata Ghosh, Bishwarup Sarkar, Ratnakar Mishra, Nanasaheb Thorat and Sirikanjana Thongmee*

## **Abstract**

Tissue grafting is mostly used for repair and replacement of severely damaged tissues, the key challenges are compatibility, availability of the grafts, complex surgical process and post-operative complications. Hence, additive technologies such as three-dimensional (3D) bioprinting have emerged as promising alternative for tissue engineering in order to ensure safety, compatibility, and rapid healing. The aim of this chapter is to give an elaborate account of 3D printed scaffolds for bone, cartilage, cardio-vascular and nerve tissue engineering. Various components such as polycaprolactone, poly (lactic-co-glycolic acid), and β-tricalcium phosphate, bioglass 45S5, and nano-hydroxyapatite are combined with collagen and its derivatives to achieve specific pore size in the scaffolds for effective restoration of the defects of soft or hard tissues. Likewise, proanthocyanidin, oxidized hyaluronic acid, methacrylated gelatin, are used in collagen based 3D printed scaffolds for cartilage tissue engineering. Bioink with collagen as active component is also used for developing cardio-vascular implants with recellularizing properties. Collagen in combination with silk fibroin, chitosan, heparin sulphate and others are ideal for fabrication of elastic nerve guidance conduits. In view of the background, collagen-supplemented hydrogels can revolutionize future biomedical approaches for the development of complex scaffolds for tissue engineering.

**Keywords:** biomaterial, collagen, scaffolds, 3D printing, tissue engineering, regenerative medicine

### **1. Introduction**

Biomedical application of nanotechnology has revolutionized tissue engineering as it can generate efficient biocompatible scaffolds with tuneable physico-chemical properties. Controllable biodegradability is one of the most important aspects as it supports the cells to produce extracellular matrix and promote effective healing. Likewise, adjustable pore structures of the scaffolds provides attractive site for loading drugs for resisting post-surgical infections and promoting cell attachment and colonization. Excellent biomechanical properties obtained by rational selection of the bioink help to mimic the tissue microenvironment and provide load bearing capacity to the tissue after repair [1]. Adherence of cells, proliferation and induction of osteogenic differentiation is higher when the total porosity of the 3D printed surfaces are more than 90% [2, 3]. Hence, such scaffolds with architectural specificity to the desired tissue like bone, cartilage, heart or nerves is immensely critical during implantation in order to ensure regeneration of the new tissue followed by repair [4].

Complete healing in traumatic injury is often a challenge that requires complicated surgical procedures which are often associated with failures and post-surgical infections. Till date bone grafting using autografts, allografts, xenografts, and synthetic bone grafts are employed for fixing the injury [5, 6]. However, the factors critical for success of grafting are the optimal size, shape, biomaterial and the anatomical structure of the bone defects. Thus, 3D printed scaffolds or synthetic bone grafts are considered more feasible due to their tuneable mechanical properties identical to the original bone tissue, and ease of rapid re-vascularization [7].

Weakening and gradual damage to cartilage may also lead to joint injury. Likewise, sudden traumatic injury, formation of lesions and developmental defects may also result is degradation of cartilage and impairment of its function [8]. In the United States alone, it is estimated that around 200,000–300,000 patients have undergone cartilage surgery [9]. It is important to note that the articular cartilage is non-neural, lymphatic, and avascular, having very low self-regenerating capacity [10]. Hence 3D printing mediated fabrication of scaffolds for repair or replacement is thought to be one of the most preferable technologies for cartilage tissue engineering [11]. Similarly, in treating cardiac dysfunctions, it is essential to maintain and mimic the cardiovascular anatomy while fixing the heart defects using tissue engineered vascular grafts (TEVGs). The 3D printing has tremendously helped to fabricate patient- and operation-specific vascular grafts [12]. Further, growing cases of neurodegenerative diseases also require effective therapeutic interventions, which are ideal for axonal regeneration and functional recovery for brain and spinal cord injury (SCI). Neuroregenerative scaffolds developed by 3D printing are considered as innovative materials that mainly focus on providing supportive substrates to guide axons and break the physical and chemical barriers, thereby promoting healing [13].

Collagen type 1 is most favorable for microextrusion based 3D bioprinting of biodegradable and biosorbable scaffolds. Collagen type 1 is the most predominant protein in the extracellular and intercellular matrix, constituting 20–30% of the vertebrate connective tissue, alongside hyaluronic acid (HA). Most importantly, the biocompatibility and low antigenicity of the collagen is attributed to the repeating motifs formed by the alpha chain of hydroxyproline-proline-glycine [14]. Collagen provides highly porous structure and hence permeability which in turn facilitates adhesion, migration, differentiation in addition to the regulation of the cellular morphology [15, 16].

This chapter highlights the collagen based 3D printed scaffolds with their attractive properties such as hydrophilicity, biodegradability, permeability, plasticity and biocompatibility critical for tissue engineering.

### **2. Collagen based 3D bioprinting of tissues**

Biomaterials composed of collagen as listed in **Table 1** are considered ideal substrate for 3D printing mediated fabrication of scaffolds for tissue engineering purposes [32]. However, simulation of the tissue microenvironment is crucial to mimic the physical and morphological properties of the native tissues in order to


#### **Table 1.**

*Collagen based biomaterials for 3D printed tissues.*

ensure proper restoration and replacement. The following section elaborates various advances of 3D bioprinting with collagen for tissue engineering.

## **2.1 Bone**

Collagen based scaffolds are widely used for bone tissue engineering. Hwang et al. (2017) fabricated bone grafts employing 3D printing using a composite of polycaprolactone (PCL), poly (lactic-co-glycolic acid) (PLGA), and β-tricalcium phosphate (β-TCP) mixed in a ratio of 4:4:2 [17]. **Figure 1** shows the scanning electron microscope (SEM) images of the bone grafts. The bone graft developed by solid freeform fabrication (SFF) technique were further mixed with 3% atelocollagen and poured into a mold and incubated at 37°C for 15 min followed by deep freezing for 6 h and freeze drying for 12 h. The collagen based biomaterial was then immersed in ethanol/ water (90% v/v) co-solvent containing 50 mM of 1-ethyl-3-(3-dimethyaminopropyl) carbodiimide (EDC) and 20 mM of N-hydroxysuccinimide (NHS) for 24 h at room temperature for effective cross-linking. Each cross-linked collagen block had a diameter and height of 8 mm and 2 mm, respectively. Circular calvarial defects of 8 mm diameter were created by removal of periosteum in male Sprague–Dawley rats. The PCL/PLGA/β-TCP composite block bone grafts were implanted into the defect cites. Interestingly the bone grafts were surrounded by fibrous connective tissues. Subtle bone formation was noted while infiltration of the giant cell and inflammatory cells were seen. However, after eight weeks both neovascularization and new bone formation were noted around the bone grafts. It was speculated that these novel PCL/PLGA/β-TCP composite block bone grafts may be considered as an alternative to synthetic bone grafts.

In another study, Inzana et al. tailored a composite scaffold using calcium phosphate and collagen for bone tissue regeneration [18]. Phosphoric acid at a concentration of 8.75 wt% was used as a binder that significantly improved the cellular viability. Tween 80 supplementation further enhanced the strength of the 3D printed scaffolds. Further, supplementation of the binder solution with 1–2 wt% collagen significantly enhanced the maximum flexural strength and cell viability. The pore size was in range from 20 to 50 μm that may significantly facilitate in-growth of the bone and

#### **Figure 1.**

*SEM images of PCL/PLGA/β-TCP particulate bone grafts. (a) Well-defined PCL/PLGA/β-TCP particulate bone grafts were confirmed at a magnification of ×100; (b) rough surface of PCL/PLGA/β-TCP particulate hone grafts were observed at a magnification of ×800. Reprinted from Hwang et al. [17].*

reestablishment of the marrow compartment. The surface was covered by plate like crystal growth which increased the surface area significantly that is ideal for adsorption of drugs and/or proteins. On implanting the 3D printed scaffolds into a critically sized murine femoral defect for 9 weeks, promising osteoconductive properties were noticed.

Kajave et al. (2021) developed a bioactive ink composed of Bioglass 45S5 (BG) and methacrylated collagen (CMA) for 3D printing of biomimetic constructs for bone tissue engineering [19]. The bioink resembled native bone tissue in the organic and inorganic composition. Superior stability with minimum swelling of the collagen based hydrogel was achieved due to homogeneous dispersion of BG particles within the collagen network. Excellent rheological property was confirmed by the betterment in the yield stress. Similarly, incorporation of the BG resulted in improvement in the percent recovery of 3D printed constructs. Additionally, improved bone bioactivity of 3D printed constructs in stimulated body fluid was advantageous. Osteogenic induction and differentiation by BG incorporated CMA (BG-CMA) constructs was associated with high cell viability and enhanced alkaline phosphatase activity and calcium deposition in human mesenchymal stem cells.

In another interesting study, Montalbano et al. fabricated a hybrid bioactive material suitable for 3D printing of scaffolds mimicking the natural composition and structure of healthy bone [20]. Initially mesoporous bioactive glass (BG) microspheres with 4% molar percentage of strontium were synthesized. Thereafter, Type I collagen and strontium-containing mesoporous BG were combined to obtain suspensions able to perform a sol–gel transition under physiological conditions. The fibrous nanostructures were homogeneously distributed embedding inorganic particles as evident from the field emission scanning electron microscopy (FESEM). Large calcium phosphate deposition was observed while release of strontium ions from the embedded BG was attributed to the high-water content of the composite. These features can cumulatively promote the osteogenic induction which is significant for bone tissue engineering. On soaking the composite scaffolds in simulated body fluid (SBF), hydroxyapatite (HA) crystals were uniformly distributed along the cross section of the sample that increased with time from 3rd to 7th day as evident from **Figure 2**.

In subsequent study Montalbano et al. reported composite biomimetics comprised of rod-like nano-hydroxyapatite particles embedded in a type I collagen matrix [21]. This composite was developed to mimic the bone composition. Initially a hydrothermal method using 0.2% ammonium-based dispersing agent (Darvan 821-A) was employed for the fabrication of the HA nanorods that were uniform-sized with length of 40–60 nm and a width of 20 nm. On suspending this material in a collagen solution in presence of Darvan 821-A, a uniform collagen/nano-HA suspension was obtained that was ideal for extrusion 3D printing. The mesh-like structures printed in a gelatine-supporting bath led to fabrication of 3D bone-like scaffolds.

#### **2.2 Cartilage**

One of the most prevalent tissue damages suffered by adults, children and adolescents is articular cartilage defects. In severe cases degenerative joint diseases may result due to exposure of bone terminals caused by progressive wear and tear of articular cartilage. However, low rate of tissue regeneration and self-repairing capacity poses a challenge for effective healing and restoration of the function. Several collagen based 3D scaffolds are being developed for inducing cartilage regeneration that is discussed in detail in this section. Recently, Lee et al. fabricated a highly biocompatible

**Figure 2.** *Cross-sectional FESEM images showing HA crystal deposition on collagen/MBG\_Sr4% samples after three and seven days of incubation in SBF at different magnifications. Reprinted from Montalbano et al. [20].*

collagen/oligomeric proanthocyanidin/oxidized hyaluronic acid (C/OPC/OHA) composite scaffold with superior compressive strengths between 0.25–0.55 MPa [22]. The composite scaffolds were 3D printed using four types of needles, 25G red plastic, 22G blue plastic, 25G red metal, and 22G blue metal to achieve 20%, 25%, and 30% porosities when pressure of 25, 15, 125, and 100 kPa were applied, respectively as illustrated in **Figure 3**. Porous nature of the scaffolds is advantageous for promoting both angiogenesis and cartilage ossification. The minimum and maximum storage moduli of the hydrogel were approximately 2.6 kPa and 4.1 kPa, respectively. Interestingly, an increased degradation rate of the composites was 26.6%, 30%, and 30.7% for 0, 5, and 10 mg/mL of OHA, respectively after 49 days. Higher apatite deposition on the scaffold surface was evident on day 21 on immersion in simulated body fluid. Superior cell viability (up to 90%) was achieved when rat bone marrow mesenchymal stem cells (rBMSCs) were grown on the composite scaffolds. On implantation of the scaffolds into bone defects in skulls of the Sprague Dawley (SD) rat, angiogenesis and new bone formation was evident that indicated 3D collagen-based scaffolds could be used as potential candidates for articular cartilage repair.

Liu et al. developed a tri-layered scaffold employing extrusion-based multi-nozzle 3D printing technology where the bioink was comprised of 15% methacrylated gelatin (GelMA) hydrogel for cartilage on top layer, a combination of 20% GelMA and 3% nanohydroxyapatite (nHA) (20/3% GelMA/nHA) hydrogel for interfacial layer, and a 30/3% GelMA/nHA hydrogel for subchondral bone at bottom layer [23]. The composite was biodegradable with maximum degradation (61.4%) in 14 days. Interconnected microtubule-like structure of each layer with interconnected spherical pores with a

*Collagen Based 3D Printed Scaffolds for Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.103914*

#### **Figure 3.**

*Optimization of 3D bioprinting parameters for obtaining porosity at 20%, 25%, and 30% using different needle densities (25G red plastic and metal, 22G blue plastic and metal) and different pressures (25, 15, 135, and 100 kPa). Reprinted from Lee et al. [22].*

size of about 300 μm was observed. The Young's modulus increased with the increase in GelMA concentration in the scaffold. The scaffolds were biocompatible with the bone marrow mesenchymal stem cells (BMSCs) while they exhibited effecting healing of rabbit osteochondral defect. Higher cartilage-specific extracellular matrix formation and collagen type II were observed on treatment with the tri-layered scaffolds. Further, effective new tissue formation and even integration with the surrounding tissues indicated their promises for repair of damages in subchondral bone by inducing cartilage regeneration.

In an interesting study, Rhee et al. fabricated 3D printing assisted soft tissue implants with high-density collagen hydrogels as illustrated in **Figure 4** [24]. External heating and collagen concentrations of 12.5, 15, and 17.5 mg/mL enhanced the shape fidelity. At the highest printable concentration, the modulus of printed gel was~30 kPa. Cell viability within the tissue constructs was high and no notable decrease was observed even after 10 days of culturing. Higher infiltration of the fibochondrocytes cells throughout the collagen matrix was found by 10 days. Adherence of the cells on the outer surface of the nascent collagen fibers was prominent while very few cells colonized the spaces between the fibers.

#### **2.3 Cardiac tissue (heart)**

Cardio-vascular defects such as aortic valve disease (AVD) require high precision surgical procedure that include either mechanical or bioprosthetic valve replacement. Recently, tissue engineered heart valves (TEHV) have gained more attention that are effectively achieved by 3D bioprinting. Maxson et al. evaluated the recellularization potential of 3D-bioprinted scaffold and investigated its applicability as a heart valve implant [25]. Allogenic rat mesenchymal stem cells (rMSCs) with

#### **Figure 4.**

*Printing process of sheep meniscus, (a) CT scan of meniscus, (b) print path of meniscus deposition of collagen hydrogel during printing, (c) 3D printed meniscus. (d) Geometry assessment of constructs. (e) Constructs scanned using Cyberware 3D scanner. (f) Geometry of the test construct: Half-cylinder. Reprinted with permission from Rhee et al. [24]. Copyright © 2016 American Chemical Society.*

green fluorescent protein (GFP) label were grown and mixed with Lifeink® 200 to obtain a homogenous bioink. Thereafter, a computer aided design (CAD) model for the implant disk scaffolds was prepared wherein the dimensions of the scaffold facilitated easy implantation and mounting in order to avoid migration and folding. Neovascularization was observed after 4 weeks with integration of host tissues with the bioink explants. Moreover after 8 weeks, minimal difference between the two

#### *Collagen Based 3D Printed Scaffolds for Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.103914*

layers was observed; however, the structural integrity of the extracellular matrix (ECM) was maintained. Furthermore, Mason's trichome revealed fibrosis on the cutaneous side of the explant whereas CD3 and CD163 biomarkers demonstrated chronic inflammation as well as ECM remodeling whose expressions were decreased with subsequent increase in incubation period. CD163 displayed a steady reduction in expression from week 1 to week 8, respectively. On the other hand, CD31 biomarker expression was considerably increased within the same time period due to endothelialization and angiogenesis. The vimentin (a major intermediate filament of smooth muscle cells) concentration of surrounding tissues was also increased with improvement in elastin concentration. This was attributed to the infiltration of the Bioink by interstitial-like cells. In addition, the ultimate tensile strength (UTS) was decreased from 0.344 ± 0.120 MPa in the second week to 0.169 ± 0.077 MPa in the fourth week while it was increased to 0.275 ± 0.166 MPa in the eighth week. Likewise, the tensile modulus was also reduced from 1.186 ± 0.872 MPa in the second week to 0.548 ± 0.341 MPa in the fourth week followed by an increase to 1.425 ± 0.620 MPa in the eighth week. Elastin concentration was significantly increased in the fourth week. Post eight weeks of implantation, expression of CD31 biomarker continued to decrease while CD163 expression increased in week 12 which was attributed to M2 macrophage infiltration. Additionally, the bioink explant was encapsulated by the fibrotic tissue within week 12 while UTS was further increased within this time period. Enhanced levels of both vimentin and elastin indicated strengthening of the extracellular matrix in the bioprinted scaffold due to active collagen deposition. Hence, collagen-based bioink application was demonstrated to be efficient for formation of heart valves.

In another study, Lee et al. also demonstrated 3D bioprinting of collagen for human heart engineering [12]. Herein, 3D bioprinting was carried out using a second generation of the free form reversible embedding of suspended hydrogels (FRESH v2.0) that provides support for printing and then subsequently melts away at 37°C. Moreover, uniform gelatin microparticles with spherical morphology (with diameter~25 μm) reduced polydispersity. An optimal balance between the resolution of individual strand and strand-to-strand adhesion was further maintained using a 50 mM N-2 hydroxyethylpiperazine-N′-2-ethanesulfonic acid (HEPES) buffered bath with pH 7.4 which in turn facilitated multiple bioink printing. A linear small coronary artery-scale tube was then fabricated using collagen type I perfusion system with an inner diameter and wall thickness of 1.4 mm and 300 μm, respectively. Thereafter, C2C12 cells were perfused in the tube that displayed viability along with active remodeling of the gel after five days. Further, cellular infiltration was also analyzed using fabrication of collagen disks with a thickness of 5 mm and a diameter of 10 mm wherein excessive cellular infiltration as well as collagen remodeling was observed post three days of implantation in the printed collagen as compared to solid-cast collagen. Moreover, fibronectin and vascular endothelial growth factor (VEGF) were incorporated into the bioink for enhanced vascularization. An extensive vascular network was observed in the printed collagen disk with red blood cells and CD31-positive vessels having a diameter range of 8–50 μm. Thereafter, collagen bioink was used along with human stem cell-derived cardiomyocytes to FRESH print a left ventricle model wherein around 96% post-printing cell viability was achieved through rapid collagen neutralization. A dense layer of interconnected and striated human embryonic stem cellcardiomyocytes (hESC-CMs) was obtained after seven days of culturing. A baseline spontaneous ventricle beat rate of around 0.5 Hz was captured that was paced at 1 and 2 Hz using field stimulation. Furthermore, the mechanical integrity of the constructs

was demonstrated using a 28 mm tri-leaflet heart valve that was robust enough to withstand air pressure. In addition, a neonatal-scale human heart was also printed using collagen bioink that highlighted the potential of FRESH v2.0 printing technique for fabrication of advanced tissue scaffolds for other organ systems as well.

Collagen-based bio-ink was also demonstrated to be an effective tool for direct 3D printing of human induced pluripotent stem cells (hiPSC)-cardiomyocytes that could then be utilized for cardiac tissue engineering [26]. Cardiomyocytes were differentiated in a 2D monolayer followed by CHIR99021-treatment mediated cell expansion and regular passing. Later on, a rat collagen-I based bioink was used for the encapsulation of cells followed by printing in a support bath composed of complex coacervate gelatin/gum arabic microparticles. The bioink was then gelated at 37°C and cultivated under free-floating conditions for a time period of thirty days. Ring-shaped cardiac tissues were printed with 5 × 5 × 1 mm dimension wherein the initial contractions were seen post three days of culturing. Striated sarcomeres were demonstrated with significant responsiveness toward pharmacological stimulations. Therefore, this study demonstrated potential of cardiac tissue engineering with enhanced properties and functions through 3D-bioprinting.

### **2.4 Nerve**

Scaffolds rationally fabricated employing 3D bioprinting could help in the treatment of spinal cord injury (SCI) by nerve tissue engineering. In a study by Jiang et al., Collagen/silk fibroin scaffold was 3D bioprinted and combined with neural stem cells (NSCs) to promote nerve regeneration [27]. A collagen/silk fibroin ratio of 4:2 was used for scaffold preparation using a 3D-bioprinter with a nozzle diameter of 210 μm, printing speed of 9 mm, extrusion speed of 2-mm/min, 0.1 mm thickness and a platform temperature of −20°C. Characterization of the 3D bioprinted scaffolds in rats revealed complete degradation of the composite scaffold after 4 weeks of implantation. Furthermore, the scaffold had considerable ductility as well as compression resistance with a compressive elastic modulus of 60.05 ± 5.12 kPa. Fourier transform infrared (FTIR) spectroscopy results then revealed presence of absorption peaks at 3445.7, 2932.46, 1640.58, and 1376.45 cm−1 that corresponded with -OH or -NH peak, methyl or C-H stretching vibrations of methylene group, C=O or C=C stretching vibrations, and saturated C-H bending vibration, respectively. Hence, these functional groups suggested presence of suitable lipid- and water-soluble bonds in the 3D bioprinted scaffold that may facilitate adhesion and growth of nerve cells. Moreover, significant biocompatibility between the scaffold and NSCs were attained with evenly distributed micropores and pore connections in the scaffold as observed in scanning electron microscopy (SEM) images. Fusiform-shaped cells grew in the scaffold pores, while some cells grew densely on the scaffold surface with extended pseudopods facilitating cell adhesion, growth as well as provided a carrier and channel for regeneration of the nerve fibers. Hence, a conducive microenvironment for NSC adhesion, growth and differentiation was provided by the 3D-bioprinted scaffold. Furthermore, 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyl tetrazolium bromide (MTT) assay also demonstrated successful seeding and proliferation of the NSCs on the scaffold. Thereafter, behavioral changes at the spinal cord injury site were investigated after implantation of the scaffold. The Basso-Beattie-Bresnahan (BBB) open-field locomotor score of the group implanted with 3D-collagen/silk fibroin scaffolds and NSCs was higher as compared to the control after 8 weeks of surgery. In addition, motor function recovery was better in groups having the scaffold and

#### *Collagen Based 3D Printed Scaffolds for Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.103914*

NSCs. Similarly, electrophysiological studies revealed prominent recovery in groups having 3D bioprinted scaffold along with NSCs as compared to control groups. Left hind limb amplitude was significantly higher in scaffold group when compared with control after 1 month of surgery. In addition, magnetic resonance imaging (MRI) and diffusion tensor imaging revealed improved filling of the injury cavity, enhanced spinal cord continuity, increased regenerative axons as well as reduced glial scarring in groups implanted with the scaffold and NSCs.

In another study, Li and Gao fabricated 3D microtubular collagen scaffolds and investigated its potential in peripheral nerve repair [28]. Melt spinning or 3D printing using poly-lactic acid (PLA) was carried out to obtain fibrous template material with a diameter range of 50–100 μm that was then utilized for fabrication of collagen scaffolds. Microtubules were prepared by parallel stacking of melt spun PLA fibers followed by polymerization of the collagen whereas PLA fibers with a diameter of 200 μm and 100 μm interspacing was fused and deposited using 3D printing. The thickness of inner ranged from 10 to 20 μm while the exterior wall formed a shell with a thickness of about 70 μm. Furthermore, cell adhesion ability of adrenal phaeochromocytoma (PC-12) and D62PT Schwann cells was evaluated wherein the cells firmly attached to native as well as chloroform-exposed Matrigel films. Two crosslinkers namely, 0.3% genipin and 0.3% glutaraldehyde were used that decreased swelling as well as enzymatic degradation of the Matrigel. Untreated gels demonstrated retention of 34.5% of total mass after 24 h incubation with 0.05% collagenase, whereas genipin and glutaraldehyde treated gels showed total mass retention of 96.7% and 99.3%, respectively. PC-12 and D62PT Schwann cells further showed well adherence and confluent growth onto microtubule scaffolds after 10 and 4–5 days of culturing, respectively. Moreover, a strong alignment of cells as well as formation of channels was seen in Schwann cells while primary chick dorsal root ganglia displayed neurite growth along the major axis of the microtubes.

Likewise, Yoo et al. reported fabrication of elastic nerve guidance conduits (NGCs) using poly(lactide-*co*-caprolactone) (PLCL) along with a 3D printed collagen hydrogel [29]. A dense acidified collagen solution with a viscosity of 1.3 × 105 mPa s was used as the bio-ink to print onto the electrospun PLCL membrane that had an optimal porosity of 2.7 ± 0.6 μm which allowed nutrient and oxygen exchange only. The acidified collagen hydrogel was then neutralized using ammonia vapor which prevented crumbling of the hydrogel. Thereafter, the NGCs were shaped into tubes and implanted in the rat sciatic nerve model. SEM images of the longitudinal cross-section of the NGCs demonstrated consistent gel deposition wherein the pore size was reduced by extraction of nano-sized fibers which in turn, prevented cell penetration into the NGCs. Moreover, a conduit fill ratio of 72 ± 2% was observed based on the hydrated cross-sectional images. Furthermore, the biocompatibility of the prepared composite was evaluated using PC12 cell culturing on the PLCL membrane with the 3D printed collagen hydrogel. After 1 week of PC12 cell culturing, a neuron-like elongated differentiation was observed in cells that were grown on the PLCL membrane having 3D printed collagen hydrogel whereas no such differentiation was observed in cells cultured on native PLCL membrane. In addition, no significant differences in the weight percentage of different animal groups were observed as well as no signs of infection, delayed wound healing, or auto-mutilation was observed throughout the experiment. The ankle contracture angles of 3D printing group after 12 weeks of nerve reconstruction was 89.68 ± 2.37% as compared to 93.52 ± 3.17% and 83.86 ± 4.64% for the autograft and bulk collagen groups, respectively. Likewise, the active ankle angle at terminal stance (ATS) was improved

in 3D printing and autograft groups after twelve weeks of nerve reconstruction with angle values of as compared to 24.02 ± 1.26° and 19.65 ± 4.78°, respectively as compared to 11.35 ± 2.91° in the case of bulk group. Hence, it was proved that 3D printed collagen hydrogel facilitated motor regeneration using NGCs. Furthermore, a comparable tetanic force of tibialis anterior (TA) muscles was observed in the 3D printing and autograft groups after twelve weeks while the bulk group displayed a lower tetanic force. The nerve regeneration through NGCs was observed after twelve weeks of surgery with linear guidance of the 3D printed collagen hydrogel from the proximal to distal ends along with an organized pattern of the regenerated axons. Moreover, the myelinated axon counts as well as thickness of myelin in the 3D printing group was higher than the bulk group. Additionally, the myelin fiber area and nerve fiber density of 3D printing group were 53,134 ± 5893 μm<sup>2</sup> and 11,206 ± 1980 n mm−2, respectively.

Lee et al. also demonstrated bio-printing of collagen and VEGF-releasing fibrin gel scaffolds and investigated its potential in artificial neural tissue construction [30]. Murine neural stem cells (NSCs) were cultured in Dulbecco's modified Eagle's medium (DMEM) and further used for cell printing. Type I collagen was then prepared and 1.16 mg/mL of the collagen scaffold was used for 3D bio-printing of C17.2 cell-scaffold complex. An average of 56 ± 9 cells/droplet was obtained with a cell viability of 93.23 ± 3.77% which was similar to that of manually-plated cells. Moreover, a collagen scaffold concentration of 1.74 mg/mL demonstrated highly dense and proliferating cells with a viability of 96.72 ± 3.58% after 3 days of culturing. Furthermore, the combinatorial effect of collagen scaffold and VEGF-containing fibrin gel on C17.2 cells was investigated wherein, the cell morphology altered after two days of culturing with active proliferation and formation of clusters. In addition, the cells located near the fibrin gel border gradually migrated toward the VEGFcontaining fibrin gel and continued differentiation. After three days of culturing, the total migration distance was 102.4 ± 76.1 μm. Hence, proper cell proliferation and migration was displayed using the two scaffolds which highlighted the potential of 3D bioprinting in artificial tissue construction.

Likewise, axon regeneration was ameliorated by Sun et al. using 3D printed collagen/chitosan scaffolds [31]. A 3D bioprinter was used for fabrication of the scaffold that had an interconnected porous structure with a porosity of 83.5% as observed in SEM images. The pore size of the scaffold ranged from 60 to 200 μm. Hence, significant space was obtained by the cells for growth and adherence. The compressive modulus of 3D collagen/chitosan scaffold was 3.82 ± 0.25 MPa along with enhanced compressive strength of 345.20 ± 29.60 KPa. The cytocompatibility of 3D printed scaffolds was similar to that of scaffolds prepared using freeze drying technology. Interestingly, the persistent locomotion recovery as well as significant increase in blood brain barrier (BBB) scores was observed after implantation of the 3D printed collagen/chitosan scaffolds in rats with spinal cord injury (SCI). Moreover, the magnetic resonance and diffusion tensor imaging results revealed a significant signal increase at the epicenter of the spinal cord lesion in rats implanted with 3D printed collagen/chitosan scaffold. Post eight weeks of SCI surgery, the axonal regeneration was demonstrated wherein 3D collagen/chitosan implantations resulted in amplitude and latency improvement. Further confirmation of axonal regeneration was carried out using anterograde biotin dextran amine (BDA) labeling wherein BDA-positive fibers were observed in 3D collagen/chitosan implantations. Hematoxylin and eosin (HE) staining also demonstrated linear ordered structure of the spinal cord after eight weeks with no obvious cavity observed in 3D printed collagen/chitosan

#### *Collagen Based 3D Printed Scaffolds for Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.103914*

implanted group whereas visible cavities and disordered structures were observed in injury groups. Hence, 3D printed scaffolds were demonstrated to be effective in axon regeneration and amelioration of spinal cord injury.

In a similar study, Chen et al. constructed collagen/heparin sulfate based scaffolds using 3D bioprinting and evaluated its action in functional SCI recovery in rats [13]. The scaffold was prepared using a 3D bioprinter that had a cylindrical morphology with a uniform and regular internal structure along with high porosity as observed in SEM images. The compressive modulus of 3D printed collagen/heparin sulfate was 3.46 ± 0.278 MPa which was higher as compared to scaffolds prepared using freeze drying technology. Likewise, enhanced compressive strength of 308.9 ± 28.65 KPa was observed in 3D printed scaffold. Furthermore, release profile of basic fibroblast growth factor (bFGF) from 3D printed scaffold was also evaluated wherein scaffolds prepared using freeze drying method demonstrated an initial burst of 54.89% of bFGF was released in the first day after which a slow release behavior was observed for longer time period. However, a steady bFGF release behavior was observed in case of 3D printed scaffolds for twenty days. Thereafter, the biocompatibility of scaffolds was analyzed using NSCs which proliferated inside the pore followed by spreading on the wall of the scaffolds. In addition, MTT assay revealed no significant difference in cell growth on different scaffolds thus highlighting the cytocompatibility of the 3D printed collagen/heparin sulfate scaffolds. Implantation of the 3D printed scaffolds further demonstrated significant recovery of locomotor functions in rats after two months with amelioration of the SCI as well as enhanced number of neurofilament positive cells.

## **3. Conclusions and future perspectives**

Advances in the field of nanomedicine have enabled exploration of novel biomaterials for tissue engineering. Among various biopolymers such as, chitosan, alginate, silk fibrion, collagen is considered as most attractive due to its biocompatibility and biodegradability. However, high temperature and extreme conditions during fabrication and bioprinting results in low stability of the collagen molecules. Hence, ideal porous scaffolds should involve combination of type I collagen and hydroxyapatite particles by freeze-drying. It is essential to have tuneable pore dimensions for superior ingrowth of cells and blood vessels. More complex microarchitectures of the collagen based scaffolds with specific rheological properties such as shear thinning, yield stress and fast shear recovery can be obtained using extrusion-based 3D printing [33].

Various biologically synthesized nanoparticles like silver, gold, copper, platinum, palladium and others can be supplemented in the scaffolds resisting post-surgical microbial infections [34–37]. Biofilm associated infections are most challenging to treat and are highly responsible for implant failure. Hence, coating of implants with antimicrobial nanoparticles impregnated collagen can be an effective strategy to increase the shelf life of the implants [38, 39]. Also drug functionalized nanoparticles can be embedded in the collagen matrix to ensure sustained release and rapid healing of the injured tissues.

Multiple approaches and integration of medical biology and material science will certainly help to revolutionize regenerative medicine by rational tissue engineering. In view of the background collagen based 3D printed scaffolds hold tremendous potential as candidate nanotherapeutics.

## **Acknowledgements**

Dr. Sougata Ghosh acknowledges Kasetsart University, Bangkok, Thailand for Post Doctoral Fellowship and funding under Reinventing University Program (Ref. No. 6501.0207/10870 dated 9th November, 2021).

## **Conflict of interest**

The authors declare no conflict of interest.

## **Author details**

Sougata Ghosh1,2\*, Bishwarup Sarkar3 , Ratnakar Mishra4 , Nanasaheb Thorat<sup>5</sup> and Sirikanjana Thongmee1

1 Faculty of Science, Department of Physics, Kasetsart University, Bangkok, Thailand

2 Department of Microbiology, School of Science, RK University, Rajkot, Gujarat, India

3 College of Science, Northeastern University, Boston, MA, USA

4 Cambridge Centre for Brain Repair and MRC Mitochondrial Biology Unit, Department of Clinical Neurosciences, University of Cambridge, Cambridge, UK

5 Nuffield Department of Women's and Reproductive Health, Division of Medical Sciences, John Radcliffe Hospital, University of Oxford, Oxford, UK

\*Address all correspondence to: ghoshsibb@gmail.com

© 2022 The Author(s). Licensee IntechOpen. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

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## **Chapter 7**

## Mechanical Methods of Producing Biomaterials with Aligned Collagen Fibrils

*Shunji Yunoki, Eiji Kondo and Kazunori Yasuda*

## **Abstract**

Collagen has been used in various therapeutic medical devices, such as artificial dermis, bone, and cartilage, wherein the effectiveness of collagen mainly depends on its biological features of biocompatibility, biodegradability, bioresorbability, cell affinity, and weak antigenicity. Collagen is the main structural protein in the human body and is responsible for the mechanical properties of tissues and organs. The fundamental structural component of tendon tissue is uniaxially aligned collagen fibrils that run parallel to the geometrical axis. Thus, the fabrication of artificial tendons is an excellent example of developing biomaterials using collagen as a structural backbone. Previous attempts to construct aligned fibril-based biomaterials involved electrospinning, freeze drying, using a strong magnetic field, and mechanical methods, including shearing and tension during wet extrusion. Among these, mechanical methods have been extensively studied owing to their simplicity and effectiveness suitable for mass production. However, few review articles have focused on these mechanical methods. Thus, this article reviews the mechanical methods for creating biomaterials from aligned collagen fibril while discussing the other fabrication methods in brief.

**Keywords:** tendon, collagen, fibril, alignment, shearing

## **1. Introduction**

Since the research and development of collagen-based artificial dermis began in the 1980s [1, 2], many biomaterials using collagen as the base have been developed and clinically applied [3, 4]. Currently, many advanced collagen-based biomaterials have been developed for cellular or acellular tissue engineering and cell therapies [5, 6], and collagen remains one of the most essential biomaterials. Collagen is useful as a base material for therapeutic biomaterials due to its excellent biochemical properties (biocompatibility, biodegradability, and bioabsorbability) [3, 4] and cell affinity [5]. These properties enable the resultant biomaterials to be decomposed through biological activity, absorbed and metabolized at the damaged sites, and eventually be replaced with normal tissues. The effectiveness of collagen in such biomaterials primarily depends on the abovementioned biological features as well as

its weak antigenicity. Its excellent moldability and low cost have further facilitated the development of sheet-shaped artificial dermis [7], porous artificial bones [8], and hydrogel-based artificial cartilages [9].

However, the mechanical properties of such collagen-based biomaterials and artificial tissues are significantly inferior to those of living tissues. Collagen is the main

#### **Figure 1.**

*Structural hierarchy in the tendon. Diagram illustrating the relationship between collagen molecules, fibrils, fibers, fascicles, and tendon units (top). Although the diagram does not show the fibril subunits, the collagen fibrils appear to be self-assembled from intermediates that may be integrated within the fibril. Scanning electron micrograph of rat tail tendon showing fascicle units (asterisk) making up the tendon (bottom). Reproduced from Ref. [11] with permission from Elsevier.*

#### *Mechanical Methods of Producing Biomaterials with Aligned Collagen Fibrils DOI: http://dx.doi.org/10.5772/intechopen.104734*

structural protein in the human body and is responsible for the mechanical features of tissues and organs [10]. Among all the various tissues and organs, tendons comprise collagen fibrils with unique hierarchical structures (**Figure 1**) that are responsible for human motor function [11]. Therefore, the fabrication of artificial tendons is a good example of biomaterials developed using aligned collagen fibrils as a structural backbone. As stated in this review, research on artificial tendons made using collagen has rapidly increased in the last decade to meet the clinical needs of replacing autologous tendon transplants.

Artificial tendons must have uniaxially aligned collagen fibrils running parallel to the geometrical axis; this characteristic collagen structure is responsible for the excellent mechanical features of live tendons [12]. Collagen fibrils can be simply prepared by well-known *in vitro* fibrillogenesis. Collagen molecules are stable in acidic solutions at low temperatures and are capable of self-assembling nanofibrils that respond to body temperature and neutral pH [13, 14]. The fibrils exhibit amorphous networks. Previous attempts to produce aligned fibril-based biomaterials used electrospinning, freeze drying, strong magnetic fields, and mechanical methods, such as shearing and tension during wet extrusion. Of these, only mechanical methods demonstrated the potential for use in the industrial production of artificial tendons by showing the ability to maintain and hierarchize collagen fibril structures.

Although many reviews on the fabrication of aligned collagen fibrils have been published in recent years [15–17], they do not focus on mechanical methods and the fabrication mechanisms therein. Here, we introduce the various mechanical methods of producing biomaterials with aligned collagen fibrils and discuss their mechanism and limitations in detail. This review will also act as a significant introduction for researchers willing to apply collagen-based artificial tendons in clinical practice.

## **2. Requirements of collagen-based artificial tendons**

Tendon tissues connect muscles to bones and influence the transmission of mechanical loads between them [12, 18]. Ligament tissues connect bone to bone and stabilize joints [12, 18]. These two types of tissues are similar in structure and comprise uniaxially aligned collagen fibrils [19]. The characteristic structure of tendons and ligaments is a multi-unit hierarchical structure comprising longitudinally aligned collagen molecules (approximately 1 nm in diameter), fibrils (approximately 100 nm in diameter), fibers (1–20 μm in diameter), and fascicles (20–200 μm in diameter) [11]. This hierarchical organization of collagen fibrils is crucial to the nonlinear and viscoelastic mechanical properties of collagen-based organs [20]. Tendons and ligaments are remarkable for their superior tensile strength and stiffness; the tensile strengths of the Achilles' tendon and anterior cruciate ligament (ACL) are 54 ± 20 MPa [21] and 24 ± 9 MPa [22], respectively, and their tensile moduli are 212 ± 109 MPa [21] and 113 ± 45 MPa [22], respectively.

Tendons and ligaments are tough tissues; however, ruptures of these tissues are common traumas among athletes [23]. The ACL is a part of a pair of cruciate ligaments (the other being the posterior cruciate ligament) in the human knee that connects the femur to the tibia to stabilize knee joint movements. ACL is the most frequently injured knee ligament [24]. Once the ACL ruptures, it can rarely connect end-to-end through conservative treatments. The poor healing capacity of ACL, particularly after rupture, is clinically common, although the underlying reasons for this remain unclear [25]. ACL reconstruction surgeries are required for such

traumas; ACL injuries are among the most common among the athletic populations, with nearly 130,000 ACL reconstructions performed in 2006 in the USA alone [26]. Although there are no published survey results, ACL construction surgeries in Japan are estimated to exceed 17,000 per year.

In recent years, autogenous tendon tissues have been frequently used as substitutes for human tendon grafts (allografts) to reconstruct torn ligaments [27]. One notable advantage of this reconstruction surgery is the remodeling property of the autogenous tendon [28, 29], called ligamentization. Although the process of biological remodeling remains incompletely understood, clinicians agree that the strength of an autogenous tendon graft reduces soon after reconstruction and gradually increases with time, accompanied by structural changes in the collagen fibers [30]. Intrinsic fibroblasts in the tendon graft undergo ischemic necrosis followed by extrinsic cell infiltration with graft revascularization [31]. After remodeling implanted autogenous tendon tissues, patients can return to their daily activities as before. However, autogenous tendon grafting inevitably results in damage and consequent morbidity at the donor site, necessitating a second invasive procedure [32]. Although allogenous tendons are considered an alternative graft material, they have their disadvantages, including disease transmission risk and slow graft remodeling [33]. Currently, tendon xenografts cannot be used in a clinical setting; accordingly, synthetic tendons have been studied to avoid these disadvantages [16]. Various synthetic materials, such as polyethylene and polytetrafluoroethylene, have been used previously to create artificial tendons. However, they have not been clinically used as they fail after implantation because they undergo biodegradation without any remodeling [34]. Therefore, the fabrication of artificial tendons showing hierarchical structures of uniaxially aligned collagen fibrils seems to be the most promising approach as they are expected to undergo remodeling in the human body after implantation in a manner similar to that of autogenous tendon tissues [16, 19].

## **3. Overview of the fabrication methods used for aligned collagen fibrils**

#### **3.1 Electrospinning**

Before focusing on the mechanical methods, we present an overview of the fabrication methods used for aligning collagen fibrils. Electrospinning has been widely considered an efficient method for fabricating polymer nanofibers, and several studies have described this fabrication technique [35, 36]. Briefly, the system comprises three elements—polymer solutions dissolved in volatile solvents, a high voltage supplier, and a metal target. The high voltage supplier provides electric potential differences in many kV between the polymer solutions and the target. The polymer solutions are then gradually extruded through a needle, and the electrically charged polymer solution is ejected from the tip of the needle which then reaches the target while being spun into thin threads (in the order of nm to μm in diameter). The volatile solvents evaporate during the interim, resulting in the collection of the polymer nanofiber mesh.

When the target is rotated during collagen electrospinning, each nanofiber tends to be aligned uniaxially [37]. Such materials can be helpful in *in vitro* experiments that evaluate the effects of scaffold alignments on the biological behaviors of cultured cells [37–39]. However, electrospinning appears to be an ineffective option as a fabrication method for artificial tendons due to the inevitable collagen denaturation. Fluoroalcohols

#### *Mechanical Methods of Producing Biomaterials with Aligned Collagen Fibrils DOI: http://dx.doi.org/10.5772/intechopen.104734*

are used as solvents to dissolve the collagen as this hydrophilic polymer shows little or no solubility in conventional volatile solvents for electrospinning, such as dichloromethane and chloroform. Zeugolis et al. demonstrated that collagen in electrospun nanofibers was denatured to gelatin almost completely using fluoroalcohols as solvents [40]. This event can be explained by the ability of fluoroalcohols to break hydrogen bonds among proteins. A pioneering study of collagen electrospinning [41] demonstrated a characteristic cross-striated pattern of collagen fibrils in electrospun nanofibers fabricated using hexafluoroisopropanol as a solvent; however, the debris of collagen fibrils is presumed to have remained according to the findings of Zeugolis et al. The nanofibers prepared using electrospinning collagen solutions cannot be defined as collagen nanofibers.

#### **3.2 Freeze drying**

Freeze drying (also called freeze-casting) is one of the most critical industrial processes used to preserve heat-sensitive biological materials, including food, pharmaceuticals, microorganisms, and plants, with minimal deterioration of their intrinsic chemical and physical properties during the drying processes [42]. The fundamental principle of freeze drying is sublimation, that is, the direct shift from a solid state to a gaseous state [43]. The freeze drying process can be explained through a characteristic phase diagram of solid (ice), liquid (water), and gas (vapor). When water-based slurry, suspension, or solution is frozen at atmospheric pressure, the contents in water are separated from ice crystals and concentrated. If we increase the temperature of the frozen material above 0°C while keeping the atmospheric pressure below 0.06 atm, the ice turns into a gas without going through a liquid phase in accordance with the phase diagram of water [43]. Generally, dried materials thus obtained have microporous structures, whereas the contents eliminated from the ice crystals had thin walls and pores, which was similar to that of ice crystals [44].

Based on the above freeze-drying principle, Schoof et al. fabricated collagen sponges using aligned structures of pores and thin walls using the unidirectional solidification technique [45, 46]. Briefly, a cylindrical container filled with a collagen suspension was sandwiched from the top and bottom using a pair of copper blocks and then cooled with liquid nitrogen. Plate-like ice grew in the collagen suspension along the depth of the cylindrical container, by which collagen molecules are eliminated from the unidirectional solidification of the growing ice crystals. Freeze drying the frozen suspension created unidirectional thin collagen walls and interconnected pores. The pores had a width of 20–40 μm and were alternately separated by much thinner walls; this structure can be considered as collagen fibers with wide gaps. Some researchers have thus used unidirectional solidification to fabricate fibers from suspensions of collagen molecules or fibrils [47–49]. Additionally, a modified technique has been developed to concentrate collagen axially and form a fiber-like construct [50]. As a result, those macromolecules are likely to be partially aligned because of the high aspect ratios. However, there is little evidence of the unidirectional alignment of collagen fibrils in the walls after freeze drying, whereas aligned thin walls or fiber-like structures were observed microscopically [45–49]. The authors believe that collagen in the micro structures is almost amorphous on a fibrillar scale, affecting cellular responses and morphologies.

#### **3.3 Exposure to a strong magnetic field**

Based on the fact that some proteins in solutions exhibit birefringence under strong magnetic fields, Torbet et al. demonstrated for the first time that collagen fibrils are magnetically aligned [51]. A neutral collagen solution (0.6 mg/mL) was heated from 4°C to 27.5°C to induce fibrillogenesis under a strong magnetic field (13T), resulting in the formation of aligned collagen fibrils. Collagen molecules have a negative diamagnetic anisotropy and they lie perpendicular to the magnetic field. Many researchers have applied this method to fabricate aligned collagen fibril hydrogels and used the gels for *in vitro* examinations to assess the effects of collagen fibril alignments on cell behaviors [52–56].

The notable advantage of this strong magnetic field is that it is noninvasive to living cells and organisms. However, the disadvantage is the lack of mass productivity of tough collagen fibers. The starting substance of the fabrication process is a neutral collagen solution, which is set in a narrow chamber (a few cm) with a strong magnetic field generator. This small-batch process is not suitable for the mass production of biomaterials. Further, the concentration of the collagen solution has to be low enough (≤10 mg/mL) to allow the rotation of the molecules due to magnetic force, preventing the production of high-density collagen fibrils. Recently, new methods have been developed wherein magnetic substances (beads or rods) are added to collagen solutions to mechanically pull or assist in the alignment of collagen fibril under magnetic fields [57–60]. A challenge associated with these manufacturing methods is that the magnetic substances are retained in the collagen gel.

### **3.4 Electrochemical method**

Electrochemical fabrication for assembling aligned collagen bundles was first reported in 2008 [61]. This method is substantially different from the previous method using strong magnetic fields in that the physical force does not directly affect the collagen molecules. When the parallel set anode and cathode electrodes are soaked in a shallow pool of collagen solution, a pH gradient perpendicular to the electrodes is generated by the migration of electrolytes. Collagen molecules with a low pH are positively charged, whereas those with high pH are negatively charged. Therefore, all the collagen molecules migrate toward the isoelectric point (pH 8.2), congregate, and form fibrils under neutral conditions. The electrochemically aligned collagen (ELAC) threads with diameters of 50–400 μm and lengths of 3–7 cm were prepared depending on the electrodes used [61]. Although electron microscopies have not yet visualized uniaxial alignments of collagen fibrils, it is reasonable that collagen fibrils tend to align uniaxially by the electrochemical compaction to a bundle.

Continuous molding of ELAC threads was successfully performed using a rotating electrode electrochemical alignment device [62]. The main parts of the device include a power supply for providing voltage for the electrochemical cell, a syringe pump, a rotating electrodes wheel, and a collection spool. A collagen solution is extruded onto the caved edge of the rotating electrodes wheel placed vertically. The electrodes are placed parallel on the wheel's edge, allowing continuous ELAC formation synchronized with the rotation speed. Furthermore, ELAC threads were twisted to form yarn, and the yarn was pin-weaved toward a highly porous scaffold as an artificial tendon. The ELAC scaffolds showed ultimate stress and tensile modulus comparable to the natural tendon [62]. Furthermore, the biological effects of collagen fibril alignments in ELAC threads have been assessed *in vitro* [62–66] and *in vivo* [67].

The electrochemical method is the first to continuously produce aligned collagen fibril threads. ELAC threads (diameters 50–400 μm) in the yarns seem to correspond to collagen fibers (diameters ~20 μm [11]) in living tendons, although the diameters

of the former are much larger. Further studies are required for ELAC thread-based biomaterials to provide the tendon-like hierarchical structure of collagen fibrils.

#### **3.5 Mechanical methods**

The main purpose of this paper is to describe the mechanism and challenges associated with manufacturing tendon-like bundles of uniaxially aligned collagen fibrils through mechanical methods. Briefly, mechanical methods involve the use of mechanical force (shearing or tension) to align collagen fibrils. Mechanical forces can be generated before, during, or after the fibrillogenesis of collagen molecules.

### **4. Solution extrusion methods**

Among the various aligned collagen fibril fabrication methods, mechanical methods have been extensively studied because of their simplicity and effectiveness as well as mass production suitability. These mechanical methods can be generally categorized into the following based on their fabrication mechanisms: solution extrusion methods (wet spinning and others), shear flow deposition, flow-induced crystallization, and gel-extrusion method. Herein, fabrication mechanism, effectiveness, and challenges of solution extrusion methods are discussed.

#### **4.1 Wet spinning**

Wet spinning is a typical example of a solution extrusion method that was first developed to produce collagen threads for artificial tendons [68]. A collagen solution is extruded from a narrow channel directly into a coagulation bath to form a cord-like gel through collagen fibril formation [69]. **Figure 2** shows a schematic illustration of a typical experimental setting for wet spinning. Neutral buffers such as phosphate buffer saline (PBS) containing polyethylene glycol (PEG) have been frequently used as coagulation baths for wet spinning as PBS provides suitable conditions for collagen fibrillogenesis [14] and PEG dehydrates the collagen molecules to promote fibrillogenesis. In Kato's method [68], the acidic collagen solution is filled in a reservoir, such as a syringe, and is extruded through a narrow tube (≤1 mm diameter) at a constant speed using a pump. The tip of the tube is submerged in a coagulation bath, and the collagen solution stream is immediately gelled due to fibrillogenesis. As a result, the cord-like collagen gel is continuously molded. Finally, drying the cord-like gel results

#### **Figure 2.**

*Schematic illustration of the typical experimental setting for wet spinning. An acidic collagen solution in a syringe (a) is loaded in a syringe pump (b) and infused via a narrow tubing into a coagulation bath (neutral buffer containing PEG is frequently used) heated at 37°C (c). As a result, the cord-like collagen gel is continuously molded and sequentially introduced into an ethanol bath (d) to promote dehydration. The cord-like gel is then wound up and air-dried to produce a collagen thread (e).*

in a tough thread with a diameter of 20–300 μm [70]. Cavallaro et al. succeeded in continuously processing dried collagen threads through a sequence of conventional wet spinning and subsequent drying using a ventilation-type cabinet [71]. Acetone was also used as a coagulation bath, allowing the fabrication of narrow collagen threads with an approximate diameter of 15 μm [72].

After Kato's pioneer study, many researchers have applied wet spinning to fabricate collagen threads for different biomaterials [69, 73]. However, the nanostructures of the collagen threads fabricated by wet spinning are far from those of tendon unit structures. Pins *et al*. revealed that collagen fibrils in wet spun threads were amorphous, prompting alignment by stretching the wet spun threads [74]. Despite many studies on wet spinning, the molecular events in the thread-making process have not yet been clarified, but it is likely that oriented collagen molecules under shearing in a narrow channel immediately changed to amorphous after extrusion into a coagulation bath (discussed in Section 4.5).

Assuming that the rapid relaxation of the oriented collagen molecules occurs, some treatments for delaying relaxation effectively promote fibril alignments in wet spinning. The addition of viscous materials or increase in collagen concentration to delay collagen molecule relaxation prior to fibrillogenesis has been investigated. Nerger et al. investigated 3-D bioprinting of collagen ink containing LAPONITE® (a type of layered silicate), Pluronic® F-127 (a type of polyethylene glycol), or Matrigel® (extracellular matrices of sarcoma) as rheology-adjusting agents [75]. The cord-like collagen gel extruded from a conical nozzle comprised of incompletely but preferably aligned collagen fibrils. Lai et al. prepared 30 mg/mL of rat tail collagen solution by dialysis against PEG and used it for fabricating tubular collagen gels with a custom-made syringe [76]. The fibrils on the surface of the collagen gels were aligned almost uniaxially, whereas the alignments of interior fibrils were not observed.

#### **4.2 Modified wet spinning**

In 2010, Caves et al. attempted to increase the fibril alignment of wet spun fibers by dropping the extruded cord-like collagen gels vertically with a coagulation buffer [77]. The extrusion of the collagen solution into a coagulation bath was performed in the same manner as wet spinning, resulting in the continuous formation of a cord-like gel. The bath was a long column through which the coagulation buffer was circulated to generate a vertical flow for carrying the collagen gel downwards along the column while simultaneously stretching it. The fibril alignment in the dried collagen thread was higher than that obtained using conventional wet spinning [73]; however, mechanical stretching (strain ratio of 10–20%) was required to achieve uniaxial alignments.

Recently, an extrusion method has been developed that incorporates the sequential stretching process of extruded gels to overcome the lack of fibril alignments [78]. This experimental setting is illustrated in **Figure 3A**. The collagen solution was continuously introduced into a flat flow channel (1-mm thick and 35-mm wide) with a pair of buffers containing PEG to ensure the three-layer of buffer-collagen-buffer. During co-extrusion, the collagen solution could be coagulated to some extent by dehydration with PEG. A sheet-shaped stream of partially coagulated collagen solution was extruded from the outlet into a coagulation buffer, resulting in the continuous production of a collagen gel sheet. Subsequently, the gel sheet was stretched along the machine direction with a rotating mandrel, thus enhancing the alignment of the collagen fibrils. Finally, wet collagen sheets as thin as 1.9 μm were obtained

#### **Figure 3.**

*Schematic illustration of the extrusion method incorporating sequential stretching of extruded sheet-shaped gels (A) and nanostructure of collagen gels observed on transmission electron microscopy (TEM) (B). (A) Collagen (red) and buffer solutions (green) are delivered to a three-layered microfluidic device. An emerging collagen sheet then undergoes fibrillogenesis and is strained by passing over a rotating mandrel. (B) TEM images of collagen sheets were produced at V\* of 0.6, 4.5, and 10 in (y − z) and (x − z) planes (depicted in (A)). V\* = (Vp − VT)/ VT, where Vp is the velocity of the rotating mandrel and VT is the total bulk velocity of the solutions passing through the flow constriction. Reproduced from Ref. [78] with permission from ACS Publication.*

which exhibited good mechanical qualities (tensile strength, 0.5–2.7 MPa and elastic moduli, 3–36 MPa). The alignment of collagen fibrils along the machine direction was enhanced depending on the rotating speed of the mandrel (**Figure 3B**).

Malladi's study indicates that the stretching process of extruded collagen gels can be incorporated into the conventional extrusion processes, including wet spinning. The stretching of gels was effective for enhancing the alignments of collagen fibrils, whereas fibrils extruded in the gels were almost amorphous. A parameter V\* = (Vp − VT)/VT was used, where Vp is the velocity of the rotating mandrel and VT is the total bulk velocity of the solutions passing through the flow constriction. The elastic moduli increased as V\* increased from 0.1 to 10; this was explained by the fibril density and degree of fibril alignment increase. As per the authors' experience, collagen fibrillar gels are less stretchable. The excellent stretchability in this case (V\* ≤ 10) could be due to the use of acid-solubilized rat tail tendon collagen [77] with intact intermolecular crosslinking. The type of collagen used in the experiment also affects the molding propriety.

#### **4.3 Other extrusion methods**

Lai et al. reported a fabrication method for cord-like collagen gels with longitudinally aligned fibrils effectively using shear force compared with a conventional wet spinning [79]. This method would result in fibrillogenesis [13, 14] before the relaxation of the shear force-induced orientation of the collagen molecules. An acidic collagen solution of rat tail tendon collagen (30 mg/mL) was continuously extruded from a syringe with a 22-gage needle onto a glass slide and submerged in a coagulation bath of 10× PBS. In this process, the syringe and glass slides were moved in opposite directions, thus generating shear forces on the extruded collagen solution, which immediately initiated fibrillogenesis while maintaining alignments of collagen molecules due to the solution's high viscosity, resulting in a cord-like collagen gel with aligned fibrils. When the human dermal microvascular endothelial cells were cultured on the gels, the cells exhibited elongated morphologies along the alignment direction of fibrils.

A method for producing edible collagen casings, that is, artificial intestine for sausage, [80] has been applied for manufacturing tubular gels comprising aligned collagen fibrils through a counter-rotating extrusion method [81, 82]. The experimental setting and appearance of the material obtained are shown in **Figure 4** [81]. This method does not include collagen fibrillogenesis but uses a fibril-rich collagen dough made from living tissues as a starting substrate. Briefly, the homogenized collagen dough (5% [w/v]) was fed to a metering pump and then into a counter-rotating extruder using a piston stuffer. This unique extruder comprises two coaxial cylinders rotating in the opposite direction. The collagen dough is continuously introduced into the gap (0.5 mm) between the larger and smaller cylinders along the axes of the cylinders so that the rotation in the opposite direction generates a shear force on the collagen dough in the gap. Consequently, tubeshaped collagen gels are extruded in which the collagen fibrils are preferably aligned in the circumferential direction. Thus, the tubes must be cut in the circumferential direction to fabricate an artificial tendon with longitudinally aligned fibrils.

The solution extrusion methods are summarized as follows: collagen molecules can be oriented using shear force in a narrow channel, resulting in the production of cordlike collagen gels with nearly amorphous fibrils. This is probably due to the immediate relaxation of the molecules after extrusion from the tips of the channels. Additional mechanical stretching is required to improve the alignment. Thus, suppression of molecular relaxation appears to be effective for fabricating collagen gels with longitudinally aligned fibrils. The use of collagen fibril dough as starting substances or the sequential stretching of gels is also effective.

*Mechanical Methods of Producing Biomaterials with Aligned Collagen Fibrils DOI: http://dx.doi.org/10.5772/intechopen.104734*

#### **Figure 4.**

*Overview of counter-rotating extrusion method. (A−E) Schematic illustrations of the method. (F and G) Appearances of tube-shaped collagen gels obtained using this method. Reproduced from Ref. [81] with permission from Elsevier.*

#### **4.4 Limitations of solution extrusion methods**

As described in Section 4.1, wet spinning has a limited capability of producing threads with well-aligned collagen fibrils, especially in the interior of threads. Although the mechanisms of solution extrusion have not been described in detail compared with those of shear flow deposition, it is obvious that rheological features of collagen solutions play a predominant role in the alignment of collagen fibrils. Here, rheological data of collagen solutions are introduced in the next paragraph to discuss the presumed mechanisms of solution extrusion.

#### **Figure 5.**

*Sensors of the rotational rheometer. (A) Appearance of a parallel plate sensor. (B) and (C) are schematic illustrations of cone plate and parallel plate sensor, respectively. A sample solution is placed on the Peltiercontrolled bottom plate and the movable upper sensor is positioned to achieve a pre-set gap. The sensor is rotated unidirectionally to obtain rotational measurements. Oscillational measurements are obtained by sinusoidal oscillation with extremely small shear deformation (usually* ≤*1%).*

For the rheological measurements, a rotational rheometer was used (MCR 502; Anton Paar, Ostfildern, Germany). This apparatus is effective for simultaneously evaluating the viscosity and gelation features of low viscous biopolymer solutions [83]. A collagen solution was filled in a gap between a Peltier-controlled bottom plate and a movable upper sensor (cone plate sensor, diameter, 35 mm; cone angle, 1°; parallel plate sensor, diameter, 50 mm) (**Figure 5**). This apparatus can conduct rotational as well as oscillational measurements. Rotational measurements measure the flow and viscosity curves of the specimen, providing information about reductions in increased shear stress (shear thinning) and thixotropic properties under shearing. Conversely, oscillational measurements are helpful in tracking the changing rheological properties of a collagen solution (in this case, recovery of rheological properties just after shearing). Two types of collagen were used, acid-solubilized collagen from the porcine tendon (designated ASC) (Cellmatrix® type I-A; 0.3% solution, Nitta Gelatin, Osaka, Japan) and pepsin-digested collagen from the porcine dermis (designated PC) (Collagen BM; 0.53% solution, Nitta Gelatin, Osaka, Japan). ASC remains intermolecular crosslinking, and the physicochemical qualities can be considered as similar to those of a conventional rat tail tendon collagen. PC is a representative of pepsindigested collagens which are generally used for commercial biomedical devices.

Rotational measurements simulated the behaviors of collagen molecules during wet spinning processes. **Figure 6** presents the viscosity curves of collagen solutions obtained by reciprocal rotational measurements at shear rates 0.1–100 s−1. Both the collagen solutions showed a shear rate-dependent decrease in viscosities (non-Newtonian behavior) during the shear rate-rising process, suggesting molecular alignments along the flow direction. The viscosity curves obtained from the falling of shear rates overlapped almost entirely in both the collagens, suggesting that the alignments of collagen molecules under shearing are not hysteresis. A sequential

*Mechanical Methods of Producing Biomaterials with Aligned Collagen Fibrils DOI: http://dx.doi.org/10.5772/intechopen.104734*

#### **Figure 6.**

*Viscosity curves of collagen solutions were obtained by reciprocal rotational measurements at shear rates of 0.1–100 s−1. (A) Acid-solubilized collagen from porcine tendon (0.3%) and (B) pepsin-digested collagen from porcine dermis (0.53%). Both the solutions showed a shear rate-dependent decrease in viscosities (non-Newtonian behavior) during the shear rate-rising process. The viscosity curves obtained from the falling of shear rates overlapped almost entirely in both the collagens.*

test of oscillation-rotation-oscillation was used (**Figure 7**) to simulate conditions of collagen molecules in wet spinning. The first step is the oscillational measurement at constant shear deformation (1%) and frequency (1 Hz) to test the viscoelastic qualities of the collagen solution as a starting substance wherein collagen molecules are dispersed amorphously. The rapid rotation (shear rate, 100 s−1) as the second step

#### **Figure 7.**

*Scheme and results of sequential testing of oscillation–rotation–oscillation for evaluating the relaxation of collagen molecules.*

#### **Figure 8.**

*Schematic illustration of a conceivable scenario in the thread-making process of wet spinning. An acidic collagen solution is extruded through a narrow tube, in which collagen molecules should be oriented along the flow direction. The stream of the viscous collagen solution extruded from the tip of the tube should immediately coagulate to form fibrils from the surface layer. If the coagulant penetrates the stream of the collagen solution before the molecular orientation is relaxed, the collagen fibrils are aligned. However, the alignments of collagen molecules would be immediately relaxed and become amorphous.*

simulates strong shearing on the collagen solution introduced into a narrow tube. Oscillational measurement as the last step monitors the recovery of shear stress after the collagen solution is released from the strong shearing, which can simulate the recovery of amorphous dispersion of collagen molecules just after extrusion into a coagulation bath. **Figure 7** indicates the results of the sequential test for ASC. The shear stress sharply decreased by more than one order of magnitude (from 9940 mPa to 540 mPa) only in 2 s after the rotation was terminated and subsequently became identical to that obtained at the first step (before rotational shearing). The small delay in the recovery of shear stress could be due to the inertial force of the flowing collagen solution.

Considering this rapid recovery of shear stress and no hysteresis of viscosity curves, the following scenario is conceivable in the thread-making process of wet spinning (**Figure 8**). An acidic collagen solution is extruded through a narrow tube wherein the collagen molecules should be oriented preferably along the flow direction, as proposed from the non-Newtonian behavior of an acidic collagen solution (**Figure 6**). The stream of the viscous collagen solution extruded from the tip of the tube should immediately coagulate to form fibrils on the surface layer. If the coagulant penetrates the stream of the collagen solution before the molecular orientation is relaxed, the collagen fibrils are aligned. However, the alignment of collagen molecules with an approximate molecular weight of 300,000 will be immediately relaxed and become amorphous, as suggested by the stress-relaxation curves of an acidic collagen solution (**Figure 7**). The above scenario somewhat explains the mechanism of collagen fibril alignments in the solution extrusion methods.

## **5. Other mechanical methods**

The previous paragraph described wet spinning and other solution extrusion methods derived from wet spinning. In the last decade, unique mechanical methods were newly developed to fabricate biomaterials with aligned collagen fibrils. Herein, the fabrication mechanism, effectiveness, and challenges of other mechanical methods (shear flow deposition, flow-induced crystallization, and gel-extrusion method) are discussed.

#### **5.1 Shear flow deposition**

This section focuses on the methods of applying shear force during collagen fibrillogenesis, called shear flow deposition. When some part of a collagen fibril is anchored onto a substrate under a strong shear flow, the fibrils are aligned in the direction of flow. This investigation is conducted using a thin collagen solution with low viscosity and a thin flow channel to induce a uniform and fast flow.

In 2009, Saeidi et al. reported the effects of shear rates on fibril alignments in the shear flow deposition using a microfluidic shear flow chamber [84], which can generate a wide range of shear rates. They examined the detailed dynamics of neutralized pepsinextracted type I collagen assembly on a glass surface under the influence of shear flow between two plates. Differential interference contrast imaging with focal plane stabilization was used to resolve and track the growth of collagen aggregates on borosilicate glass under various shear rates (500, 80, 20, and 9 s−1). The nucleation of fibrils on the glass was observed to occur rapidly (~2 min) followed by the continued growth of the fibrils. The best alignment of fibrils was observed at intermediate shear rates of 20 and 80 s−1, whereas the growth rates were affected by the shear rate in a complex manner. However, the investigation showed that directional fibril growth was not stable and the fibrils would often turn downstream, forming "hooks" at high shear rates.

In Saeidi's next study [85], a spin-coating technique was combined with a flow of collagen solution to produce highly aligned arrays of collagen fibrils. A chilled neutral collagen solution was introduced into the center of the spin coater, which was heated to initiate collagen fibrillogenesis. Orthogonal collagen lamellae were successfully fabricated on the coater depending on shear rates (181–2480 s−1), which were adjusted by flow rates (0.1–1 mL/min) and rotation speeds (750–3000 rpm). It was possible to produce small sections (1 cm<sup>2</sup> ) of collagen fibrils with enough alignment to guide fibroblasts. However, thin-film instabilities on the coater are likely to be a significant barrier to manufacturing organized collagen fibrils over larger areas.

The effects of planar substrates with collagen-binding features on shear flow deposition were evaluated by Lanfer et al. [86]. They used a microfluidic channel system with coverslip substrates coated with poly(octadecene-alt-maleic acid) (POMA), which could bind collagen fibrils. The aligned collagen fibrils were successfully deposited on the substrates, where the degree of collagen fibril alignment increased with increasing flow rates of the solution. The matrix density increased at higher collagen solution concentrations and on hydrophobic polymer pre-coatings.

The shear flow deposition can deposit well-aligned fibrils on substrates, thus providing some insights into the fabrication conditions for achieving tendon-like collagen fibrillar gels. However, there is a limitation to fabricating thick and long products of aligned collagen fibrils. Collagen fibrils can be anchored directly to substrates at the beginning of the fabrication, promoting fibril alignments along the shear flow direction. However, in the following steps, collagen fibrils cannot be deposited due to the lack of binding features between collagen fibrils. Shear flow deposition methods

are likely to be helpful in fabricating cell culture substrates rather than therapeutic biomaterials, such as artificial tendons, to investigate the effects of collagen anisotropy on the biological behaviors of living cells [87, 88].

#### **5.2 Flow-induced crystallization**

Before describing flow-induced crystallization, the capacity of collagen molecules to form liquid crystalline should be described. At a molecular level, acid-soluble collagen molecules spontaneously assemble into precholesteric-banded patterns and cholesteric phases at concentrations above 50 mg/mL [89]. Stabilization of the liquid crystalline collagen, induced by pH modification and resultant fibrillogenesis, indicates characteristic morphologies of collagen fibril arrays in bone tissues. Furthermore, a dense gel (18 wt%) prepared by self-reassembly of collagen molecules *in vitro* shows characteristic bundles of cross-striated fibrils observed in the tendon. The qualities of collagen molecules imply that the formation of liquid crystalline at high concentrations is a key factor for manufacturing bundles of uniaxially aligned collagen fibrils.

In 2016, Paten et al. developed a novel fabrication method called flow-induced crystallization through which dense collagen molecules were microfluidically drawn to form a fiber of uniaxially aligned fibrils [90]. **Figure 9A** presents the schematic illustration of the fiber-making process. Briefly, a droplet of neutralized collagen solution was set under a flow of dry nitrogen gas, facilitating evaporation of water from the droplet surface and the formation of an enriched monomeric surface. A glass microneedle was used to pierce the droplet surface, and the dense collagen solution adhered to the tip of the needle. When the needle was drawn back to attain a low strain rate < 1 s−1), the surface collagen solution was pulled up to form a thread. In this processing, flow-induced crystallization and mechanical tension-induced fibril alignment could occur. Finally, a narrow fiber as a highly aligned collagen fibrillar array was created (**Figure 9B**).

Although the flow-induced crystallization method is still a form of microfluidic examination, each event in the processing provides us with ideas for creating uniaxially aligned collagen fibrils. When a dense collagen solution with the ability to form liquid crystalline is exposed to strong shearing or tension, the collagen molecules could be ready for uniaxial fibrillogenesis. Therefore, we have to consider the possibility of the continuous heating of the dense collagen solution under strong shearing or tension resulting in uniaxial fibrillogenesis. It is expected that a continuous fabrication of a thread of uniaxially aligned collagen fibrils is developed and scaled up based on the processing of Paten et al.

#### **5.3 Gel-extrusion method**

The last mechanical method for aligned collagen fibrils is the gel extrusion recently developed by the authors' group. This method can continuously fabricate cord-like collagen fibrillar gels by incorporating the advantages of the solution extrusion method and shear flow deposition. Those are continuous extrusion of collagen solution under shearing and simultaneous stretching of fibrils by shear force.

First, we evaluated the phenomenon caused by applying shear stress to collagen during fibrillogenesis using a rotational rheometer as a measuring device and a sample fabrication device [91]. A neutral collagen solution was filled in a gap between a Peltier-controlled bottom plate and a movable upper sensor (parallel plate sensor: diameter 60 mm). Fibrillogenesis under shearing occurred by increasing the

*Mechanical Methods of Producing Biomaterials with Aligned Collagen Fibrils DOI: http://dx.doi.org/10.5772/intechopen.104734*

#### **Figure 9.**

*Overview of flow-induced crystallization technique. (A) Schematic illustration and appearances of the fiber-making process. (B) Transmission electron microscopy images of the collagen fiber. Reproduced from Ref. [90] with permission from Elsevier.*

temperature of the bottom plate (from 23 to 37°C) during rotation of the upper sensor. Wide ranges of collagen concentrations (0.1–2 wt %) and shear rates (0.1–500 s−1) were preliminarily examined, but the gels were destroyed completely between the plate and sensor. The most crucial factor for successfully preparing gels under those

conditions was the rate of fibrillogenesis gelation. Increased concentrations of neutral phosphate buffer could accelerate the gelation rate, and fibril alignment occurred within 20 s during the early stage of rapid gelation. Fabrication of gels was completed with slippage between gels and the movable upper plate, and well-aligned fibrils along the rotation direction were observed in the marginal regions of disk-shaped gels. Gel thickness could be increased from 1 mm to 3 mm with the homogeneous alignment of fibrils in the entire sample. The alignment of fibrils enhanced mechanical qualities against tensile loads placed parallel to the alignment axis. The elongation of cultured fibroblasts along the alignment was observed on the gels.

Next, a continuous formation method of cord-like collagen gels comprising fibrils preferentially aligned along the geometrical axes (CCGs) was developed by transferring the events on a rotational rheometer to those in a stainless tube [92]. The experimental setting was simple (**Figure 10A**). Collagen (2.5%) dissolved in a sodium phosphate buffer containing 280 mM sodium chloride was introduced into a stainless cylinder (length 52 mm, diameter 2.0 mm) heated to 38°C at a linear velocity of 2.5 mm s−1. This process caused collagen fibril alignments under acute fibril formation in the cylinder, causing the continuous formation of CCGs (**Figure 10B**). Fibril formation rate, shear rate, and shear duration were substantial factors for successful CCG formation. Advantages of this method over conventional wet spinning include the capacity of this method to form aligned fibrils in the entire gels and to control the diameter of cord-like gels over 1 mm (**Figure 10C**-**10F**). The air-drying of CCGs, which were cross-linked with 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide and N-hydroxy-succinimide, produced dry collagen fibers with cross-sectional areas of 0.0123–0.135 mm2 (**Figure 10G**). Upon the rewetting of the fibers, they failed at a stress of 54.5 ± 7.8 MPa, which is higher than the mean failure stress of ACL tissue (13.3–37.8 MPa). These findings show that the CCG formation method enables the fabrication of collagen fibers, which are potential components of collagen-based artificial tendons.

A limitation of the gel-extrusion method is the incomplete alignment of collagen fibrils along the geometrical axes, especially in the core region of the gel. The mechanism of fibril alignment could explain this heterogeneity, the process of alignment

#### **Figure 10.**

*Overview of the gel-extrusion method. (A) Schematic of the experimental setting. Collagen solution in a syringe (a) was loaded in a syringe pump (b) and infused via silicone tubing (c) into a stainless cylinder (d), which was immersed in a neutral buffer in a glass beaker (e) heated at 38°C in a water bath (f). A cord-like collagen gel was continuously extruded from the cylinder and accumulated on the bottom of the glass beaker (g). (B and C) A 2.0 mm diameter stainless steel cylinder during the processing and the stacked gels within. Longitudinal cross-sectional scanning electron microscopy images of the gel. Bar in figure indicates 5 μm. (E) Two-dimensional birefringence images of the gel. (F) Retardation of the gel across the perpendicular direction. (G) Appearance of dry fibers obtained from the cord-like gel. Bar in the figure indicates 20 mm.*

of collagen fibrils involved their formation and then their immediate stretching by shear stress. The entanglement points act as anchors for stretching fibrils. The lengths of fibrils between entanglements were unequal, resulting in more and less stretched fibrils at a certain shear deformation. The stainless tube was heated in a water bath, causing a slower temperature elevation rate in the core region.

## **6. Conclusions**

In conclusion, the mechanical methods for creating aligned collagen-based biomaterials are summarized. Previous attempts to fabricate uniaxially aligned fibrils have used electrospinning, freeze drying, strong magnetic field, electrochemical methods, along with mechanical methods, including shearing and tension during wet extrusion. Among the various fabrication methods, mechanical methods have been extensively studied because of their simplicity and effectiveness along with suitability for mass production. Mechanical methods can be generally divided into the following four methods depending on their fabrication mechanisms: solution extrusion methods (wet spinning and others), shear flow deposition, flow-induced crystallization, and gel-extrusion method. Solution extrusion methods can continuously mold cordlike collagen gels, from which collagen threads are prepared by air-drying. However, collagen fibrils in wet spun threads were amorphous, thus additional stretching of the threads is required to promote fibril alignments. The lack of fibril alignments is probably due to the immediate relaxation of the oriented molecules after the extrusion of collagen solutions. Additional mechanical stretching of gels or threads and delay of molecular relaxation in collagen solutions are effective to promote collagen fibril alignments. The use of collagen fibril dough as starting substance is also effective.

Shear flow deposition can deposit well-aligned fibrils on substrates. However, there is a limitation in fabricating thick and long products of aligned collagen fibrils. Collagen fibrils can be anchored directly to substrates at the beginning of the fabrication, promoting fibril alignments along the shear flow direction. But in the following steps, collagen fibrils cannot be deposited due to the lack of binding features between collagen fibrils.

Flow-induced crystallization is still a kind of microfluidic examination, combined with liquid crystallization of dense collagen solutions. This method can produce ultrathin threads of uniaxially aligned collagen fibrils. However, the production is not continuous because the starting substance is a partially dried surface of a droplet of collagen solution. It is expected that a continuous fabrication of collagen threads is developed and scaled up based on the processing of flow-induced crystallization.

The gel-extrusion method is a continuous formation method of cord-like collagen gels composed of fibrils preferably aligned along the geometrical axes in the entire gels. The feature of this method is the use of neutralized collagen sol, where the temperature-responsive fibrillogenesis is accelerated. The collagen sol is introduced into a heated channel where it can form fibrillar gels. The fibrils are aligned by shear force and stretching. A limitation of the gel-extrusion method is the incomplete alignment of collagen fibrils along the geometrical axes, especially in the core region of the gel.

Mechanical methods have recently made rapid progress. However, each of the methods cannot create artificial tendons with hierarchical structures of uniaxially aligned collagen fibrils with a capacity to undergo remodeling in the living body after implantation similar to autogenous tendon tissues. It is still challenging for biomaterial engineering to satisfy excellent mechanical and biological features. There are two promising approaches for creating an ideal collagen-based artificial tendon,

bottom-up and top-down approaches. The bottom-up approach is the creation of collagen fibers similar in size to the collagen fibers of the living tendon, followed by making them into a tight bundle (not a simple twist string). In contrast, the top-down approach is the longitudinal fragmentation of a large bundle of uniaxially aligned collagen fibrils to allow infiltration of extrinsic cells.

Recently, the performance of decellularized tendons for ACL reconstruction has been evaluated *in vivo* [93, 94]. Although there are some challenges including unevenness of material qualities, residual sources of infection, and production costs, excellent mechanical features and collagen structures similar to living tissues are suitable for ACL reconstruction. The differences between collagen-based artificial tendons and decellularized tendons should be considered in biomaterial developments.

## **Acknowledgements**

This work including the rheological experiments was supported by JSPS KAKENHI Grant Number 15 K01321.

## **Conflict of interest**

The authors declare that they have no conflict of interest.

## **Author details**

Shunji Yunoki1 \*, Eiji Kondo2 and Kazunori Yasuda3,4

1 Biotechnology Group, Tokyo Metropolitan Industrial Technology Research Institute (TIRI), Tokyo, Japan

2 Department of Advanced Therapeutic Research for Sports Medicine, Hokkaido University Graduate School of Medicine, Sapporo, Japan

3 Department of Sports Medicine, Hokkaido University Graduate School of Medicine, Sapporo, Japan

4 Knee Research Center, Yagi Orthopaedic Hospital, Sapporo, Japan

\*Address all correspondence to: yunokishuji530@gmail.com

© 2022 The Author(s). Licensee IntechOpen. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

*Mechanical Methods of Producing Biomaterials with Aligned Collagen Fibrils DOI: http://dx.doi.org/10.5772/intechopen.104734*

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## *Edited by Nirmal Mazumder and Sanjiban Chakrabarty*

Collagen is the most abundant protein class in the human body. The polymer also has the distinct benefit of being biodegradable, biocompatible, readily accessible, and very adaptable. The use of collagen-based biomaterials in tissue engineering applications has increased dramatically over the past few years, owing to extensive research in the field. Multiple cross-linking strategies for collagen have been examined. Various combinations of collagen with other biopolymers have also been investigated in an attempt to increase the tissue function of the collagen biomaterials in their various formulations. *Collagen Biomaterials* provides a thorough overview of the different uses of collagen-based biomaterials produced for tissue engineering, to offer a functional material for use in regenerative medicine from the laboratory bench to the patient bedside.

Published in London, UK © 2022 IntechOpen © Jian Fan / iStock

Collagen Biomaterials

Collagen Biomaterials

*Edited by Nirmal Mazumder* 

*and Sanjiban Chakrabarty*