Clinical Applications of Naturally Derived Biopolymer-Based Scaffolds for Regenerative Medicine

Abstract

Naturally derived polymeric biomaterials, such as collagens, silks, elastins, alginates, and fibrins are utilized in tissue engineering due to their biocompatibility, bioactivity, and tunable mechanical and degradation kinetics. The use of these natural biopolymers in biomedical applications is advantageous because they do not release cytotoxic degradation products, are often processed using environmentally-friendly aqueous-based methods, and their degradation rates within biological systems can be manipulated by modifying the starting formulation or processing conditions. For these reasons, many recent in vivo investigations and FDA-approval of new biomaterials for clinical use have utilized natural biopolymers as matrices for cell delivery and as scaffolds for cell-free support of native tissues. This review highlights biopolymer-based scaffolds used in clinical applications for the regeneration and repair of native tissues, with a focus on bone, skeletal muscle, peripheral nerve, cardiac muscle, and cornea substitutes.

INTRODUCTION

Beginning in the mid 1970s, investigators explored the use of materials for biological and medical applications, leading to rapid growth in the field of biomaterials and their applications in tissue engineering.^16,120,250 Successful tissue engineering strategies attempt to recreate or mimic the conditions present in healthy tissue in vivo. However, positive in vitro results have yielded limited translational success with in vivo models as well as in the clinic due to the simplified culture conditions, lack of cell retention in vivo, unknown biological factors, such as long-term immune response, which are difficult to predict or determine in vitro. A constant influx of new information from biological and medical research as well as improvements in in vitro experimental models will continue to guide methodologies and product development. Due to continued advancements in the fields of developmental biology, optics, and assay development, biomedical research continues to progress rapidly building from these improvements in the fundamental building blocks to generate better in vitro systems that more accurately recapitulate the in vivo environment (e.g., evaluation, utilization, and application of extracellular matrix composition, co-culture, and immunomodulation). Key components of this progress involve the careful selection and preparation of scaffold materials and use of physical, chemical, and/or electrical stimuli to direct delivered cell behavior to ultimately mimic the characteristic profile of the desired native tissue.

The Role of Biomaterials in Translational Medicine

Two traditional groups of biomaterials are available: synthetic polymers (Table 1) and natural biopolymers (Table 2). Since each group possesses distinct advantages and limitations, a wide variety of composite materials and interpenetrating networks have been utilized to achieve desired results. These include composites of natural materials, such as silk and collagen^31,233 or chitosan and alginate,^{48,78,201,216,243,264,287} as well as blends of natural and synthetic systems, such as alginate, polyacrylamide, and poly(ε-caprolactone) interpenetrating networks.¹³⁵ In addition, hybrid variants of these materials have emerged through synthetic designs and genetic engineering of peptide-based biopolymers.^8,18,36,203

TABLE 1.

Common synthetic polymers used for tissue engineering and clinical applications.

Synthetic polymer	Selective in vitro applications, clinical translation and commercialization	Refs.
Poly(l-lactic acid) (PLA or PLLA)	Composite materials, 2D gels, thin films Stent coatings, orthopedic applications, wound healing, soft tissue reconstruction	15,48,62,130,144,162,185,215
Poly(glycolic acid) (PGA)	2D and 3D mesh and scaffolding Sutures, soft tissue reconstruction, nerve conduits	12,46,104,111,127,148,274,283
Poly(ethylene glycol) (PEG)	Composite materials, backbone for chemical modification, 2D gels and surfaces, easily modified, well-controlled structure Wound healing, soft tissue reconstruction	27,28,97,116,134,136,179,180,184,188,205,251,286
Poly(lactic-co-glycolic acid) (PLGA)	Nanoparticles, composites, electrospun scaffolds Drug delivery, cartilage tissue repair, bone scaffolds, vascular reconstruction, soft tissue engineering	11,81,99,118,122,237
Polyurethane	Wound healing, cardiac tissue engineering, skin grafting	54,58,60,108,115,186,228,231,232
Polytetrafluoroethylene (PTFE)	Cell free, non-modifiable structure with minimal degradation for 2D surfaces and implants Cardiac repair, vascular grafts	103,145,202,221,259
Poly(N-isopropylacrylamide) (poly(NIPAAm))	Thermo-reversible; 2D cell sheets, thin tissue layers (≤600 μm), injectable gels Cartilage tissue engineering, cell sheet technology, soft tissue repair	105,235,284
Polycaprolactone (PCL)	Electrospun scaffolds, thin films, honeycomb architectures Cardiac tissue engineering, vascular grafts, cranial reconstruction	37,47,49,51,163,196,272,286
Poly(glycerol sebacate) (PGS)	3D scaffolds and films; major focus is cardiac and bone Wound healing, nerve conduits, soft tissue reconstruction, bone scaffolding	3,32,53,65,181,182,191,198,206,271,289

TABLE 2.

Common natural biopolymers used for tissue engineering and clinical applications.

Biopolymer	Structure	Biological interactions	Applications and references
Alginate	1 → 4 linked β-d-mannuronic acid (M) and α-l-guluronic acid (G) residues G residues function with divalent cations to form ionic crosslinks	Variations in local mechanical properties controlled by concentration of calcium ions Form soft gels, composite materials, interpenetrating networks Electrospun fibers with improved mechanical integrity	Wound healing, drug delivery, soft tissue engineering, cell delivery, in vitro stem cell maintenance48,78,79,112,126,135,141,142,200,201,216,223,230,240,243,247,248,252,264,265,270,287,295
Chitosan	Linear polysaccharide from deacetylation of chitin (main component in the exoskeletons of crustaceans) Randomly distributed β-(1–4)-linked d-glucosamine and N-acetyl-d-glucosamine	Soluble, cationic in acidic conditions and insoluble in neutral and basic conditions	Wound healing, orthopedics, cardiac repair, neural tissue engineering, cornea repair, drug and gene delivery34,56,94,146,159,189,196,205,217,288,293
Collagen	monomeric and crosslinked 3D triple helical structures Many types, type I consists of two identical α1 chains and one α2 chain Utilized in vitro in fibrous and hydrogel formats Content measured by quantifying hydroxyproline content Gelatin is a commonly used animal-based collagen product for gels	Easily remodeled and degraded by cells Chemical crosslinking decreases degradation and improves long-term Mechanical properties Many types available, which are specific to different tissue types Collagen IV primarily found in the basal lamina, secreted by epithelial cells, collagen I is the main component in scar tissue, cartilage, bone, tendon, and other connective tissues and muscles Improper expression or mutation leads to disease	Wound healing, skin grafts, muscle repair, nerve regeneration, anti-aging1,33,74,82–85,152–154,190
Elastin	Elastic component in soft tissues that allows tissue to return to normal shape following stretch or pinch Crosslinking smaller tropoelastin polymers using lysyl oxidase to form mesh-like structures	Small tropoelastin polymers can be used to form composite materials Can form films, gels, scaffolds, and electrospun fibers Easily remodeled by cells	Cardiac stent coatings, soft tissue reconstruction, orthopedics5,10,14,86,110,172,273
Fibrin	Formed when thrombin cleaves fibrinogen, generating soft gels Gelation kinetics controlled by ratio of thrombin to fibrinogen, calcium concentration, temperature	Fibrinolytic inhibitors, like aprotinin or aminocaproic acid, reduce in vitro degradation rates	Wound healing, cardiac repair, in vivo cell delivery25,38,178,222,244,260,285,292
Fibronectin	Glycoprotein secreted by fibroblasts Binds integrins, transmits mechanical cues from the environment to the cell binds other ECM proteins such as collagen and fibrin	Improper regulation of expression leads to diseases such as cancer, fibrosis Fibronectin can be grafted onto 2D and 3D surfaces to improve biocompatibility Many sites for cell-based protein degradation	Wound healing, stem cell differentiation, cardiac repair, bone regeneration158,193,227,275
Keratin	Bundles form intermediate filaments that make up hair (α-keratins) and nails (β-keratins) Tough, insoluble polymer found in non-mineralized tissues Monomers form stable left-handed superhelical structures to form filaments	Structure attributed to disulfide bridges; more bridges yields lower elasticity Classified as neutral-basic or acidic, dictating in vivo occurrence	Cornea tissue engineering, wound healing, skin regeneration, cardiac repair, drug delivery, nerve repair23,210,213,238
Silk	Extracted from the cocoons of Bombyx mon silkworm Foams, sponges, films and hydrogels are formed from silk solution	In vivo degradation rate controlled by crystallinity (β-sheet content) Mechanics tailored by modifying concentration, crystallization, molecular weight, and scaffold format	Muscle repair and regeneration, bone tissue engineering, cornea repair, drug delivery6,8,17,67,86,110,165,168,192,197,211,212,280

Synthetic Materials

Synthetic polymer-based materials are widely used as biomaterials for tissue engineering and regenerative medicine due to their well-defined chemical and structural characteristics, as well as the flexibility they afford in allowing the researcher to control and fine-tune the final properties of their scaffold. Thus, a diverse spectrum of synthetic polymers has been explored for use in virtually all tissue engineering applications. These polymers can be designed to be either non-degradable (e.g., polyethylene terephthalate/Dacron^™) or degradable (e.g., poly(glycerol sebacate)). Non-biodegradable materials are advantageous in biological systems that experience high mechanical loading demands since the non-degradable materials can maintain their structural integrity over time. In contrast, fibrotic response to an inert, non-degradable scaffold due to inability to integrate within the native tissue or adverse inflammatory response can result in long-term implant failures or secondary complications. For example, the use of Dacron^™ or PTFE for repair of congenital heart defects in children can lead to secondary complications such as aneurism or fatal arrhythmia.^98,102,241 Biodegradable synthetic materials are advantageous because their chemical formulation can be tuned to obtain specific degradation properties (e.g., chemical hydrolysis or surface erosion). Unfortunately, cell-based degradation or natural deterioration may result in the release of cytotoxic or inflammatory small molecules, prompting the need to determine tissue-specific degradation rates, local dose–response sensitivity, and metabolic response of neighboring cells to these byproducts (See Refs.^219,253 for a more detailed review). Commonly used synthetic biomaterials for tissue engineering are listed in Table 1. While Table 1 does not cover all possible synthetic materials, the polymers listed highlight novel structures and formulations utilized in current and emerging technologies as well as standard clinical products that represent areas in need of improvement.

Natural Materials

For the purpose of this review, we will define natural biomaterials as those found in nature, originating from a plant-based or tissue-based origin. Once properly purified for in vivo applications, these materials generally do not elicit an unwanted or unexpected immune response, leading to biocompatible and often bioactive matrices that can integrate with surrounding native tissue.^{132,204,214,267} In addition, the degradation products are generally more biocompatible, metabolically accessible, and less toxic when compared to their synthetic counterparts. A disadvantage of naturally derived biomaterials is their chemical heterogeneity and high dispersity, leading to variability in mechanical properties, structure, and performance, including variations in local degradation rates.¹³² Despite this variability, biopolymer scaffold materials have successfully advanced to the clinic for the repair of soft tissues, such as skin and muscle, and hard tissues, such as bone. Commonly investigated biopolymers and their benefits in vivo are listed in Table 2.

Natural materials can be further classified based on their origin and the processing methods used to obtain the material. Simple biopolymers are derived from a natural origin, such as rat-tail or bovine tendon for the isolation of collagen. These methods first require the removal of all cellular material and then the subsequent dissolution of the biopolymer in acidic medium to yield a semi-clear solution that has been purified to contain a specific protein (e.g., collagen type I).²⁰⁷ Alternatively, natural materials for scaffold fabrication can also be obtained from the decellularization of complex tissues. For example, collagen-based scaffolds can also arise from the decellularization of porcine dermis or pericardium. In some cases, such as in the use of demineralized (DMB) or decellularized (DCB) bone, the 3D architecture of the native tissue is maintained and utilized as the skeletal architecture of the scaffold. Both biopolymer-based scaffolds and decellularized matrices have achieved some clinical success, especially in the field of wound healing and skin regeneration. This review will discuss the use of biopolymer-based scaffolds for tissue repair and regeneration, with a focus on in vivo and clinical application (Table 3) to highlight the hurdles that still exist in the bench-to-bedside clinical transition.

TABLE 3.

Examples of commercially available biopolymer systems for various types of tissue repair.

Manufacturer	Product	Application	Product description	Refs.
Baxter International, Inc. (Deerfield, IL)	TachoSil®	Cardiac Wound Sealant	Contains human fibrinogen and human thrombin to form a fibrin sealant Fibrinogen and thrombin formed on surface of an equine collagen patch Extensive literature on use for many applications other than FDA approved use	40,66,90,149,157,171,195
Collagen Matrix, Inc. (Franklin Lakes, NJ) sold by Stryker® (Kalamazoo, Ml)	NeuroFlex®	Nerve Repair and Regrowth	Type I Collagen mesh Flexible tube, kink-resistant	4,52,93,119,166,167
Collagen Matrix, Inc. (Franklin Lakes, NJ) sold by Stryker® (Kalamazoo, Ml)	NeuroMatrix®	Nerve repair and regrowth	Type I Collagen mesh Flexible tube	4,52,93,119,166,167
Collagen Matrix, Inc. (Franklin Lakes, NJ) sold by Stryker® (Kalamazoo, Ml)	NeuroMend®	Nerve Repair and Regrowth	Type I Collagen mesh Designed to wrap injured nerves for a range of injuries	4,52,93,119,166,167
Cook Biotech, Inc. (West Lafayette, IN)	Dynamatrix®	Wound healing Soft Tissue Reconstruction	Acellular graft containing collagen (types I, III, VI), glycosaminoglycans (hyaluronic acid, chondroitin sulfate A and B, heparin, and heparan sulfate), proteoglycans, growth factors, and fibronectin	19,121,183
Cook Biotech, Inc. (West Lafayette, IN)	BioDesign® Grafts	Brain Surgery Abdominal Reconstruction Wound Healing Cardiac Surgery	Acellular scaffold, non-cross-linked, non-dermis-based graft Forms water tight seal so fluid cannot leak from the closed wound	64,124
Japan Tissue Engineering Co., Ltd. (Gamagori City, Aichi, Japan)	LabCyte Cornea-Model	Corneal epithelium	Trans-well permeable membrane seeded with primary human corneal epithelial cells Epidermal model also available	113
Living Cell Technologies™ (Manukau, Auckland, New Zealand)	DIABECELL®	Type 1 Diabetes	Alginate-based porcine-derived islet of Langerhans cell product	79,112,141,252
Living Cell Technologies™ (Manukau, Auckland, New Zealand)	NTCELL®	Parkinson’s Disease T reatment	Alginate-based choroid plexus cell product	142,252,278
MatTek Corporationa (Ashland, MA)	EpiOcular™	Corneal epithelium	Fully biological scaffold containing human-derived epidermal keratinocytes	161
Medtronic Sofamor Danek (Minneapolis, MN)	INFUSE®	Bone Repair	Absorbable collagen sponge in a metal Delivers recombinant bone morphogenetic protein-2	96,234
Mölnlyke Health Care (Norcross, GA)	Xelma	Ulcers Wound Healing	Extracellular matrix proteins mixed with propylene glycol and alginate	22,164,268,269
Musculoskeletal Transplant Foundation™ (Edison, NJ)	CASCADE® Autologous Platelet System	Soft Tissue Repair Bone Defects	Autologous platelet delivery system using a fibrin gel Arthroscopic delivery, followed by suturing in place	9,77
Organogenesis, Inc. (Canton, MA)	Apligraft®	Skin Ulcers Wound Healing	Cellular material comprised of foreskin-derived neonatal fibroblasts cultured in vitro, combined with bovine type I collagen coated in cultured neonatal keratinocytes First biomaterial with living cells to get FDA approval Achieves complete wound closure in chronic foot ulcers Initial production was not cost effective Shipping affected product reliability 10+ years to make profits	29,44,63,70–72,76,92,109,117,123,208,224,246,249,256,291
Organogenesis, Inc. (Canton, MA)	Dermagraft®	Skin Ulcers Wound Healing	Cellular material comprised of human neonatal fibroblasts seeded on bioabsorbable polyglactin mesh scaffold Cells are cultured over time on the mesh to allow for ECM production on the mesh	80,100,101,155,156,187,199
Skin Ethic Laboratories (Lyon, France)	HCE	Corneal epithelium repair	Permeable polycarbonate substrate seeded with immortalized human corneal epithelial cells	263

^aMatTek Corporation offers a variety of similar products for in vitro analysis.¹⁶¹

BIOPOLYMERS FOR BONE REPLACEMENT

The rising age of the population coupled with the increasing incidence of age-related conditions such as osteoarthritis and osteoporosis has produced an overwhelming market for bone replacement materials (Fig. 1).⁶¹ Despite modest success in the development of bone fillers, surgeons must still resort to autografts or allografts for critical sized defect repair. Over the past several decades, research has turned to tissue engineering as a strategy for functional bone repair. However, attempts to produce high-strength, porous scaffolds for bone regeneration have been limited by the intrinsic weakness associated with high porosity materials. Moreover, many formulations of natural or synthetic polymeric materials fail to promote adequate vascularization, innervation, and maturation of osteogenic cells.^24,220 This section of the review addresses the critical role that biopolymers play in the repair of bone tissue, and highlights the clinical translation of these natural materials to the clinic.

FIGURE 1.

Bone tissue replacement. (a) Overview of the anatomical structure of human bone tissue and essential design criteria at the macro, micro, and nano-scales. (b) Building a functional bone graft or bone substitute material from a bottom-up approach involves careful selection of natural biocompatible and biodegradable materials to mimic the inorganic–organic structure of native bone. Calcium phosphates, such as hydroxyapatite (HA) and tricalcium phosphate (TCP), are frequently complexed with polymeric materials, such as collagen or silk, to form a composite structure. At a micro-scale, osteoconductivity of the composite scaffold is supported by porosity (40–60%) and pore size (>100 μm), and osteoinductivity to promote osteogenesis in vivo is accomplished by release of bone morphogenic proteins (BMPs) or vascular growth factors. On a macro-scale, scaffolds must withstand physiologic loading forces (2–10 MPa in trabecular bone and 170–200 MPa in cortical bone) and should be defect specific. This can be achieved with various methods including 3-D printing of patient-specific grafts or the use of an injectable, self-setting calcium phosphate cement (CPC)-polymeric fillers.

Design Criteria for Bone Replacement

From an engineering perspective, bone tissue is a composite, anisotropic material built to preferentially withstand mechanical forces in one direction (Fig. 1a).¹³ Due to the unique organic–inorganic composite structure of osseous tissue, design strategies in vitro aim to mimic natural mineralization processes. For example, mineralization can be induced via phosphorylation of collagen or the incorporation of bioactive agents (e.g., hydroxyapatite, bioactive glass particles).^{140,151–154} Additionally, mineralization can be achieved via the hybridization of collagenous biomaterial constructs with non-collagenous recombinant proteins, such as INFUSE^®.²³⁴ Clinical success of peptide-based mineralization has been limited by the high cost of the peptides, as well as deleterious side effects (e.g., breathing difficulty, hematomas, swelling) caused by secondary inflammation near the site of implantation.⁹⁶ The secondary inflammation can be caused by an unwanted immune response to the non-native peptides introduced for mineralization.²³⁴ For example, INFUSE^® delivers recombinant bone morphogenic protein-2 (rhBMP-2) via a collagen sponge. Clinical results demonstrated that the desired inflammatory response at the treatment site may spread to adjacent critical structures, leading to increased postoperative morbidity.²³⁴

Despite the challenges tissue engineers face in achieving collagen scaffold mineralization in vivo, collagen-based biomaterials continue to be one of the most highly investigated natural materials for bone regeneration (Table 3). The implantation of pre-mineralized collagen matrices or pre-seeding with osteogenic cell lines remains the primary focus of in vivo collagen-based bone repair. The application of collagen hydrogels in particular has received much attention for bone repair. Implantation of a collagen hydrogel in a critical-sized rabbit segmental diaphyseal defect model resulted in supporting pre-seeded osteosarcoma cells,²⁵⁴ while collagen hydrogels with a collagen fibril density equivalent to native tissues (i.e., 10–15%) maintained structural stability up to 5 weeks post-implantation.¹⁷⁷ These studies, among others, demonstrate the necessity of effective strategies for in vivo mineralization for successful graft integration and defect regeneration.

Silk

Silk protein has been widely used as a scaffolding material for tissue development and cell-based remodeling both in vitro and in vivo.^{20,57,150,165,192,261} Due to its robust mechanical profile, ability to be mineralized, controllable degradation, and excellent osteoconductive and osteoinductive properties, extensive research has been conducted on the use of silk for bone repair. The biocompatibility of silk has also been thoroughly investigated in the trabecular bone of sheep tibia and humerus defects.²⁶¹ Despite the ability of silk scaffolds to deliver small molecules and growth factors while maintaining biocompatibility, silk exhibits a low compressive strength, thereby limiting its use in non-load-bearing bone regeneration. Within the past few years, however, several strategies have been implemented to overcome this mechanical limitation. Mandal et al. addressed the need for a stronger and stiffer natural polymeric matrix using a solvent-processed silk sponge scaffold reinforced with micron-sized silk fibers.¹⁵⁰ Alternative strategies to improve the compressive strength of silk scaffolds utilize the addition of hydroxyapatite or bioglass particles, or in situ crystallization of calcium phosphate on the scaffold surface. Silk scaffolds have also been evaluated for the effect of pore size in bone repair in an in vivo model.¹⁶⁵ Collectively, these studies demonstrate the potential of a natural silk material for bone tissue regeneration, and provide evidence for the utilization of silk or silk-based composites in clinical bone repair.

Alginate

The use of alginate-based materials for load bearing bone repair in critical-sized defects has been limited by the poor mechanical profile of alginate and the lack of cell-based degradation mechanisms, which can delay native tissue ingrowth into an alginate implant. However, despite these limitations, alginate-based materials and composites have been studied in orthopedic research for promoting osteogenesis, improving osteogenic differentiation, and delivering cells and growth factors to bone defect sites.²⁶⁵ One such study investigated the ability of an alginate/nanofiber mesh composites scaffold loaded with rhBMP-2 to enhance the repair of critically sized segmental bone defects in a rat model. Implantation with controlled release of rhBMP-2 resulted in consistent bony bridging within the defect, demonstrating the promise of these composites for growth factor delivery to repair of critically sized bone defects.¹²⁶

Natural Materials as Bone Fillers and Cements

The rapidly rising demand for orthopedic tissue engineering products to repair spinal damage, musculoskeletal defects, or bone fractures has led to the development of several FDA-approved bone cements and filler materials. Calcium phosphates (CaP), particularly hydroxyapatite (HA), remain a focus for orthopedic applications due to their semblance to the mineral phase and crystalline structure of bone.⁵⁹ A majority of bone filler materials are comprised of calcium phosphate cements (CPCs) (Table 4). In an aqueous environment, CPCs undergo rapid sedimentation and precipitation to apatite, hardening within minutes at body temperature. These reactive CPCs are optimal for on-site surgical filling of a bone gap, due to their moldability and rapid setting. Clinically, they are used in periodontal repair, cranio-maxillofacial surgery, and augmentation of an autograft or allograft. Once set, CPCs are very brittle and are therefore not applicable in the repair of load bearing critical-sized defects, where connection to the host vasculature, nervous, and muscular systems are paramount for clinical success.^24,220 Overall, CPCs, especially when formulated using natural, non-toxic processing materials such as silk,¹⁶⁵ represent a current clinical success story for research driven material design for clinical applications.

TABLE 4.

Commonly used synthetic bone filler materials on the US market.

Product	Manufacturer	Composition	Forms
BonePlast	Interpore-Cross Intl.	Calcium sulfate	Injectable paste
PeoOsteon	Interpore-Cross Intl.	Coralline HA	Granules/blocks
Vitoss	Orthovita	β-tricalcium phosphate	Granules/blocks
α-BSM	Depuy Synthes	Calcium phosphate cement	Injectable paste
β-BSM	Depuy Synthes	Calcium phosphate cement	Injectable paste
γ-BSM	Depuy Synthes	Calcium phosphate cement	Moldable putty
Norian SRS	Depuy Synthes	Calcium phosphate cement	Injectable paste
Calceon6	Depuy Synthes	Calcium sulfate	Pellets
CalStrux	Stryker Biotech	TCP with carboxymethycellulose	Moldable putty
BoneSource	Stryker Biotech	Calcium phosphate cement	Paste
HydroSet	Stryker Biotech	Calcium phosphate cement	Injectable paste
Collagraft	Zimmer/NeuColl	HA/TCP granules w/collagen gel	Strips
Cellplex	Wright Medical Inc.	Tricalcium phosphate	Granules
MIIG X3	Wright Medical Inc.	Calcium sulfate	Pellets
OsteoSet	Wright Medical Inc.	Calcium sulfate	Pellets
Actifuse	Baxter International, Inc.	Silicate-substituted calcium phosphate ceramic	Injectable

Most filler materials are calcium-based ceramics that can be mixed or cross-linked with polymeric materials to form composite scaffolds.

These bone fillers come in a variety of forms including injectable paste, powder/granules, pellets, and blocks.

HA, hydroxyapatite; TCP, β-tricalcium phosphate.

Clinical Translation and Commercialization

Cell-free bone cements and putties represent the state of the art in repair of small bone defects. However, secondary limitations, such as the lack of vascularization and organized structure of these cements, restrict their clinical potential. Given the recent progress in directing differentiation of human induced pluripotent stem (hiPS) cells towards osteoblast and osteoclast lineages⁵⁰ in combination with improved material design, the development of successful cell-based therapies for repair of damaged or diseased bone is within reach.⁷ Progress towards clinical application of functional, critically sized bone grafts is underway, but current and future work must still address the combination of mineralized bone with other tissue structures such as bone marrow. In addition, connection of the new structure to the host system via vascularization/angiogenesis, innervation, and lymphangiogenesis has not been sufficiently addressed.^{7,24,43,106,114,129,220} Integration of a construct with the host tissue via these factors must be addressed in order to achieve long-term improvements in patient prognosis,.

BIOPOLYMERS FOR SKELETAL MUSCLE AND NERVE TISSUE REGENERATION

Skeletal muscle tissue has the innate ability to repair and regenerate following acute injury, as evidenced by the successful increase in organized and functional muscle tissue following prolonged and repeated periods of exercise. Exercise induces small tears in the muscle fibers, which the body is able to not only repair, but also respond to, via increase in skeletal muscle volume. However, some of the most important parameters in dictating clinical success include the maintenance and growth of vascular and neural network architectures within the new muscle. In critically sized defects, such as those caused by volumetric muscle loss following of trauma,^41,42,147 vasculature and neural networks are not maintained. Regeneration of these connections in vivo, specifically for the nerve (i.e., the neuromuscular junction (NMJ)), is paramount for functional muscle recovery resulting in controllable contractile function. Consequently, there is still an unmet need for the development of critically sized tissue replacements comprised of aligned, functional muscle containing nerve conduits with functional neuromuscular junctions.

Design Criteria for Critically Sized Skeletal Muscle Defects

Skeletal Muscle

Like bone tissue, the structure of skeletal muscle dictates its function. Muscle is composed of bundles of muscle fibers surrounded by connective tissue, with each muscle fiber representing a single, often multinucleated, muscle cell. Individual cells or fibers are bundled together into small groups called fascicle, which are surrounded by a thin layer of connective tissue. Multiple fascicle will make up a given skeletal muscle. The tubular structure formed by the fascicle is vascularized and innervated such that the vessels and nerve conduits run parallel to the bundled muscle fibers (Fig. 2).²⁴²

FIGURE 2.

Engineering the neuromuscular junction (NMJ). (a) Anatomy of the neuromuscular junction showing myelinated axons of the nerve bundle interfacing with the striated muscle fibers at the nerve terminal. Design criteria for nerve conduits and skeletal muscle-NMJ interface include bulk alignment, elasticity, long-range signal propagation, and conductivity as well as efficient signal transmission to the sarcomere to elicit muscular response. (b) Commercially available scaffolds for nerve repair, such as Stryker NeuroFlex, are clinically available for implantation to promote the re-growth of short nerve segments. However, these conduits are only able to address nerve injuries or gaps of less than a critical size of 2.5 cm, nor can they provide effective muscle-nerve connectivity, which is essential for the regeneration of traumatic skeletal muscle and nerve injuries.

During development, changes in gene expression, structural alignment, and chemical cues facilitate the maturation of muscle progenitor cells into mature myoblasts or fibers, which are also key parameters in activating muscle progenitor or satellite cells in adult tissue.^{35,41,218,255,258,266,290} Many skeletal muscle engineering strategies have leveraged bulk alignment within the scaffold architecture, nanotopographical cues, and either passive tension or stretch applied by a bioreactor in order to mimic the native muscle structure.^{41,42,147,194}

Engineering Skeletal Muscle with functional Neuromuscular Junctions (NMJs)

Apart from bulk alignment and cell–cell coupling, innervation of skeletal muscle is necessary for proper functioning of a critically sized muscle graft. Engineering the neuromuscular junction (NMJ) has proven extremely challenging due to the complexity required for proper cell-to-cell interactions that result in directed muscle movement.^75,258 The neuromuscular junction develops in a complex multistep process involving both inter and extracellular signaling pathways, leading to the formation of synaptic contact between the terminal branches of the motor neuron and a specialized area of the muscle cell sarcolemma (i.e., plasma membrane) called the motor end-plate.²⁸¹ To properly regenerate and build NMJs, interactions between motor neurons, skeletal muscle fibers, and glial cells must be tightly regulated through both time and space. As such, a thorough understanding how these cells interact and how to properly co-culture them is required before the development of critically sized constructs that achieve functional results will be successful.^95,137,281 To further complicate the engineering problem, it is well known that improper cell-to-cell interaction and unorganized muscle-nerve interaction can lead to disease states such as amyotrophic lateral sclerosis (ALS), spinal muscular atrophy, and muscular dystrophy.²⁴⁵

In Vivo Accomplishments for Muscle and Nerve Tissue Regeneration

Many recent in vitro studies have demonstrated the potential use of natural biomaterial systems to guide skeletal and smooth muscle and nerve development, and further establish the functional connection between these two tissues.^75,258 Current research aims to translate these successes to in vivo studies, further validating the use of natural materials for engineering critically sized muscle constructs. In vitro analyses suggest that constructs formed by the co-culture of two or more cell types may lead to better vascularization and innervation in vivo.^75,258,281 By incorporating a mixed muscle progenitor cell population cultured on an aligned decellularized bladder matrix in the repair of volumetric muscle loss in rodents, the in vivo functional capacity of injured musculature improved 2.3-fold with the addition of a mixed cell population promoting the generation of functional skeletal muscle fibers. Additionally, a rodent model was used to demonstrate that innervation of rat skeletal muscle seeded fibrin gels prior to implantation significantly increased force generation and NMJ development in vivo.⁵⁵ Addition of agrin, a large proteoglycan important for NMJ development, significantly increased the formation of acetylcholine receptor clusters on differentiated C2C12 mouse myoblasts in vitro.¹²⁵ Pre-fabrication of acetylcholine receptor clusters on differentiated C2C12s in fibrin gels prior to implantation also accelerated innervation in vivo in a male nude rat model.¹²⁵ These recent results demonstrate the potential for engineering the NMJ in vivo, while also elucidating current limitations in scaffold design, such as lack of control over cell-to-cell contacts, and physiologically inaccurate 3D geometries preventing proper NMJ development on aligned muscle fibers, all of which must be addressed in future designs.

Clinical Translation and Commercialization

While the clinical realization of innervated skeletal muscle constructs has yet to be achieved, clinical success in the repair of short nerve defects within muscle tissue has been demonstrated. Materials for nerve repair are commercially available and focus on guiding nerve growth using artificial conduits. Stryker^® (Kalamazoo, MI) developed type I collagen-based biomaterials for peripheral nerve and soft tissue repair. Their products, NeuorMatrix^™, NeuroFlex^™, and NeuroMend^™ all use type I collagen as a wrap or tube to repair peripheral nerve injuries (Fig. 2b). In addition, TissueMend^™ (Stryker^®, Kalamazoo, MI) acellular collagen scaffolds are designed to augment repair and provide mechanical reinforcement for damaged tissues. While these materials are FDA-approved, they do not meet many of the design criteria described above. For example, type I collagen-based materials do not provide bulk alignment, support NMJ formation, or promote differentiation of local progenitor cells into functional nerve or muscle tissue.

Similarly, both natural and synthetic materials have been utilized for form aligned skeletal muscle fibers using a variety of formats including scaffolds, films, and limited success has been demonstrated in animal models (see Ref. 279 for recent review).^41,42 Overall, bridging the gap from promising in vivo animal studies to human trials and commercialization has proven difficult due to the multiple tissues that must integrate fully to provide functional recovery. However, as further progress is made using stem or progenitor cells, these remain a promising avenue for the repair of critically sized muscle defects.²⁶

BIOPOLYMERS FOR CARDIAC TISSUE REGENERATION

Heart disease is the leading cause of death in both men and women, costing the United States $108.9 billion each year in health care services, medication, and loss of productivity.³⁰ In adult patients, complications from heart disease often result in heart failure, but in children, heart failure is frequently caused by congenital malformations of the heart (e.g., birth defects), weakened heart muscle, or damaged tissue not specifically due to heart disease. Currently, the only clinically available treatments for congenital birth defects and end-stage heart failure include left ventricular assist devices (LVADs), total heart transplantation, and surgical reconstruction.^225,282 Despite some success in heart repair and regeneration in small animal models, regeneration of electromechanically integrated human myocardium has yet to succeed, driving the need for new ideas and alternative engineering approaches for the repair and restoration of proper heart function.

Design Criteria for Cardiac Tissue Replacement

Scaffolds designed for cardiac tissue engineering must provide equivalent functionalities of physiological cardiac muscle, offering mechanical and electrical support for the native or encapsulated cardiomyocytes. Like skeletal muscle, cardiac muscle is striated. Therefore, alignment of scaffold architecture, whether achieved through passive stretch or tension, constant strain, or structural patterning, has shown to significantly improve cardiomyocyte function (Fig. 3a).^21,170,174 Whether acellular or cell-seeded, scaffolds must demonstrate a high degree of elasticity, promote cellular remodeling, maintain cell viability, support stem cell differentiation, and eventually sustain cardiomyocyte hypertrophy. In a cell-delivery approach, the scaffold design must yield integration of the provided cells (e.g., stem cells) with the host tissue and electromechanically couple these cells with the native heart muscle through both host cell ingrowth and delivered cell outgrowth (Fig. 3a). Mechanical considerations depend on many factors, including the age of the patient and the presence of injury or disease, and therefore, the modulus of scaffolds for cardiac tissue engineering should range between 5 and 50 kPa with pore sizes ranging from 20 to 40 μm.^204,225 Additionally, delivery of growth factors such as insulin-like growth factor-1 (IGF-1), human growth factor (HGF), vascular endothelial growth factor (VEGF), and stromal cell-derived factor-1 (SDF-1) have improved cardiac function and tissue integration following trauma.^257,282 As designs for cardiac tissue engineering move forward, the use of time-dependent and slow-release delivery systems will be necessary to achieve the desired paracrine signaling profiles in vivo.

FIGURE 3.

Cardiac tissue replacement. (a) Anatomy of the heart showing the striated alignment of cardiac muscle. Design criteria for cardiac tissue engineering include high elasticity, aligned porosity, support of vascular and capillary network growth, and methods for integration into the host tissue both mechanically and electrically. (b) A variety of materials are commercially available for temporary repair of cardiac defects, but most are made from non-degradable plastics, which do not grow with the patient or integrate with the tissue, leading to the risk of fatal arrhythmias. Some disadvantages of clinically available vascularized grafts or decellularized matrices are thrombosis or a lack of tissue integration and growth within the patient. Decellularized grafts represent an allogenic approach to repairing damaged tissue, but patients run the risk of rejection, infection, and fibrosis.

In Vivo Accomplishments for Cardiac Tissue Replacement

Recent in vitro and rodent-based in vivo data support the use of both cellular and acellular approaches for repairing and restoring heart tissue in both children and adults, yet clinical realization of these strategies has proven difficult.^26,128 These therapies include the use of hydrogels or decellularized extracellular matrix solutions that are injected into or onto the heart,²²⁶ the implantation of scaffolds or matrices, or the implantation of films or cell sheets.¹⁶⁰ Treatments are aimed at either restoring heart wall thickness and strengthening the muscle or minimizing scar tissue accumulation in the area and replacing it with electromechanically coupled tissue.

Alginate and Chitosan

Polysaccharides, such as alginate and chitosan, have been widely utilized for cardiac tissue engineering due to their biocompatibility and ease of use.²³⁹ Currently, liquid alginate solutions (Sodium Alginate and Calcium Gluconate, NCT01226563 and Algisyl-LVR^™ intramyocardial injections of alginate hydrogel, NCT01311791) are undergoing thorough investigation in human clinical trials. However, a three-dimensional, pre-formed alginate-based cardiac patch has not yet reached the clinic despite promising in vitro results.²³⁹ The development of a cardiac patch from one or both of these materials is extremely challenging due to the limitations in electrical conductivity. Gold nanowires, gold nanoparticles, and carbon nanotubes have been incorporated into polysaccharide-based hydrogels to improve electrical coupling between adjacent cardiac cells in vitro. As with all critically sized constructs, vascularization is paramount to the success of a cardiac construct. One approach is to add pro-angiogenic growth factors to the matrix to improve host infiltration. Recent efforts using these polysaccharide based biopolymers have led to the successful delivery or growth factors and small molecules in vivo.^133,262

Gelatin and Elastin

Use of collagen alone to repair heart defects is challenging since the protein itself causes stiffening of the heart tissue and induces reprogramming of cardiac fibroblasts. Instead, gelatin and elastin based scaffolds serve as sufficient alternatives for cardiac tissue engineering due to their rapid cellular degradation and bulk elasticity. Gelatin-based sheets for the fabrication of cardiac patches are currently under investigation in clinical trials (e.g., AutoLogous Human CArdiac-Derived stem cell to treat Ischemic cArdiomyopathy (ALCADIA), NCT00981006). Novel composites formed by methacrylating biopolymers, such as GelMa (methacrylated gelatin²³⁶) and MeTro (methacrylated tropoelastin¹⁰), have also been studied extensively in vitro. These biopolymers also lack electrical conductivity, which has prompted the in vivo investigation of composite materials containing single walled carbon nanotubes (SWNTs) for left ventricle repair following myocardial infarction. The implantation of these devices in vivo resulted in significantly increased fractional shortening and ejection fraction compared to untreated hearts, partially due to enhanced the expression of intercellular adhesive junctions and electrochemical junctions in rats who received SWCT-based implants.²⁹⁴ Overall, SWCT-based scaffolds integrated into infarct myocardium exerted beneficial effects on myocardial regeneration and remodeling in the infarct areas, resulting in the improvement of heart functions in rats.²⁹⁴ In vitro, the addition of the SWCTs enhanced spontaneous electrical activity in seeded cell constructs as well as cardiomyocyte connectivity, as demonstrated by the ability of SWCT-containing constructs to beat regionally after 2–3 days and contract synchronously at day 8, which was not observed in SWCT-free materials.²⁹⁴ SWNT-based materials supported the contractile properties of engineered cardiac tissue through enhancement of the formation of gap junctions and promotion of the excitation–contraction coupling of cardiomyocytes.²⁹⁴ However, the long-term biocompatibility of SWCTs in vivo raises concern for the translation of these scaffolds to human clinical trials.

Fibrin

Fibrin-based scaffolds show exciting promise for cardiac repair given the elastic nature of the scaffold and the rapid degradation and cellular remodeling potential afforded by the chemical structure.⁸⁹ Initial work with this biopolymer demonstrated modulation of gelation rates by adjusting the fibrinogen to thrombin ratio to achieve an injectable material capable of delivering cells or growth factors to a damaged heart surface.^39,178 Initial in vitro investigations of fibrin for cardiac repair established potential therapy to regenerate post-myocardial infarction cardiac tissue using 3D functional fibrin-based myocardial equivalent grafts formed in vitro. Results indicated that the aligned fibrin scaffolds demonstrated a significant increase in twitch force compared to the isotropic constructs.²¹ Parameters that could be modulated to tailor these construct properties prior to implantation include regulating the magnitude of stretch used to induce alignment, preconditioning the scaffold to stretch in a bioreactor, or pre-seeding the graft with cardiomyocytes prior to implantation.^{173–176,285,292}

Clinical Translation and Future Directions

Current research on cardiac patch biomaterials is promising, yet full realization of the potential of these therapies remains elusive. Unless a patient is undergoing open-heart surgery, such as a valve replacement, surgical access to the outer left ventricle wall is limited. Regeneration strategies have instead focused on delivering cells or injectable gels at the damaged site by a catheter.²²⁶ Additionally, natural biopolymer-based cardiac patches may provide solutions to complex cardiac reconstruction in young patients with complex congenital heart defects. Natural-based cardiac patches have allowed engineers to overcome the limitations of current synthetic patches, which do not promote cellular remodeling, do not grow as the child grows, and often require reoperation (Fig. 3b).¹⁰² Furthermore, in young patients, both extracellular matrix composition and cardiomyocyte metabolism differ greatly from mature, adult tissue, necessitating a careful analysis of design criteria and biomaterial selection for cardiac graft fabrication.^276,282 These results present two major implications for the future of cardiac tissue regeneration and remodeling: (1) repair in younger patients may be more successful due a greater percentage of proliferating cells and circulating progenitors and (2) the fetal environment (e.g., matrix composition, tissue density, tissue mechanics) may represent improved design criteria for repair and regeneration in the adult heart.

BIOPOLYMERS FOR CORNEAL TISSUE REPLACEMENT

The human cornea is a transparent, avascular, connective tissue that provides an optical interface with substantial refractive power, while protecting the eye from mechanical injury and potential infection. This organ is comprised of three distinct cellular layers: corneal epithelium, stroma, and endothelium, which are separated by two acellular collagenous interfaces known as the Bowman’s Layer and Descemet’s membrane (Fig. 4).²¹⁴ Improper corneal development, damage to the cornea and limbal cells, or nerve injury resulting from infection or trauma can result in loss of corneal transparency, leading to either partial or complete loss of vision. Approximately 10 million people worldwide suffer from corneal vision loss, prompting the need for an effective corneal transplant therapy. These transplants can be categorized into two main repair options: allogenic and synthetic materials. Although allogenic materials from human donors are preferred, a shortage of quality donor graft material has limited the broad applicability of this clinical option. Alternatively, synthetic homologs to donor corneal grafts are primarily used as temporary replacements until suitable donor tissue becomes available. The use of a synthetic homolog for long-term repair is limited by intrinsic risks of corneal melting, bacterial endophthalmitis, and retinal detachment resulting in graft failure.

FIGURE 4.

Cornea tissue replacement. (a) Anatomy of the human cornea: schematic and histological images of human cornea layers: the corneal stratified squamous epithelium with underlying the Bowman’s layer, the stroma with keratocytes for the maintenance and production of extracellular matrix, the Descemet’s membrane, and, the single-layer endothelium. State of the art of tissue engineered cornea: Epithelium Human epithelial cell sheet obtained from oral epithelial cells after removal of the cell sheet from the thermoresponsive surface; Stroma Assembly diagram for 3D silk film corneal constructs seeded with human corneal fibroblasts; Full-thickness Synthetically crosslinked collagen, molded into an implantable, full-thickness corneal substitute. Transparent samples were trephined to prepare a button for corneal implantation and then held in place by sutures in the patient eye.

Design Criteria for Cornea Tissue Replacement

The main functions of the cornea necessitate three major design requirements: protection, transparency, and an effective optical interface. In addition, a corneal replacement device must also be biocompatible in the human body and ideally bioactive such that the graft material can integrate within the surrounding host tissue (Fig. 4). Efforts to regenerate the cornea have focused on the in vitro regeneration of the epithelium, stroma, and endothelium strata, followed by the promotion of neural and vascular interfaces. Natural biopolymers have been used extensively to mimic the corneal epithelium layer. In particular, reconstituted and chemically cross-linked type I collagen or silk hydrogels have been used as substrates for human epithelial cell growth and functional tissue organization.^138,169 Reconstruction of the corneal stroma has proven challenging due to the complex structure, mechanical strength requirements, and need for optical transparency. Therefore, corneal stroma engineering has focused on the development of functional corneal stroma substrates through chemical, morphological, and mechanical cues.^88,131,277 In the particular context of the corneal stroma, type I collagen has been used extensively due to its dominant content in the native corneal tissue as well as its standardized processability.⁸⁴ Alternatively, silk films have been optimized to support corneal stromal cell growth and organization in both 2D and 3D environments, in which topography, surface chemistry, porosity, degradation profiles, and transparency were controlled.^87,229

In Vivo Accomplishments for Corneal Tissue Replacement

In order to study the integration of an implanted biomaterial within corneal native tissue, in vivo implantation of acellular corneal tissue equivalents has been investigated.^139,205 Studies have focused on recapitulating the three-layer structure of the cornea (epithelium, stroma, and endothelium).^2,91 Recent efforts have used decellularized biological material, such as amniotic membranes and animal-derived cornea. However, the results of an acellular porcine cornea in combination with amniotic epithelial cells in a rabbit lamellar keratoplasty resulted in degradation of the tissue-engineered cornea due to host rejection.¹⁴³ Decellularized amniotic membrane was clinically evaluated in combination with human corneal endothelial cells in a lamellar keratoplasty model. The endothelium and part of the Descemet’s membrane were removed, and the construct was able to function as a corneal endothelium equivalent.⁷³ In all cases, complications due to foreign material host response and material performance limitations such as lack in transparency, degradability, and mismatch in mechanical and permeability properties hindered success in the reported animal studies.

Clinical Translation and Commercialization

Anterior partial keratoplasty performed in humans using biosynthetic corneas made from cross-linking recombinant human collagen type III showed that these naturally derived collagen scaffolds provide effective tissue regeneration by promoting endogenous tissue growth and innervation without signs of vascularization for up to two years. However, a delay in epithelial closure and a fibrotic response were observed, likely caused by surgical sutures.⁶⁸ A 4-year follow up showed a stably integrated implant, although a more robust material with better shape retention may improve visual acuity.⁶⁹ Partial or full-thickness engineered corneal tissues have been developed for in vitro preclinical cornea tissue repair models to reduce animal testing for commercial products for eye irritancy tests.⁴⁵ A multi-layer collagen hydrogel scaffold was developed and evaluated using primary corneal endothelial, stromal, and epithelial cells.²⁰⁹ The commercially available products mainly referred to engineered epithelium based on trans-well permeable membrane architecture, as in the case of Clonetics^™ Human Corneal Epithelial Culture Model (Lonza, Hopkinton, MA), LabCyte Cornea-Model (Japan Tissue Engineering Co., Ltd., Gamagori City, Aichi, Japan), and EpiOcular^™ (MatTek Corporation, Ashland, MA).

THE FUTURE OF BIOPOLYMERS: A FOCUS ON CLINICAL TRANSLATION

Given the wide range of products available on the market today, there still exists a pressing need for further product development to repair full thickness wounds. The path to clinical translation and commercialization of wound healing products presents a case where academics, start-up companies, and venture capitalists can reflect on both the successes and failures of product implementation in the medical field. A major lesson learned is that there is no “one-size-fits-all” solution such that a single product or biomaterial will not meet the design criteria and patient need for every clinical application. For instance, wound healing is more complex and takes longer in diabetic patients compared to healthy patients, as these patients often suffer from ischemia and neuropathy in their extremities. Thus, challenges with cost-to-patient-benefit analysis, proper doctor recommendation and utilization, long times between initial development and FDA approval, and other unforeseen challenges have impeded the clinical translation of many promising biopolymer-based products.¹⁰⁷

A fundamental understanding of native tissue structure and organization, as well as physical, and biochemical properties is essential for the successful design of biomaterial scaffold systems to produce effective tissue replacement therapies. The goal of this review was to provide an overview of recent efforts based on natural polymers to mimic in vivo structures for tissue types including bone, muscle, peripheral nerve, cardiac, and cornea, and to highlight the progression of these systems from animal models into clinical settings. Despite major recent advancements in the tissue engineering field, significant challenges such as lack of vascularization, fine control of biomaterial degradation rate and byproducts, and construct re-innervation in critical-size grafts are limiting factors in the translation of tissue engineered products to clinical scenarios. Thus, there remains a significant opportunity to develop new methods and techniques for the generation of pre-vascularized and re-innervated tissues to promote angiogenesis and neovascularization, lymphangiogenesis, and innervation in order to create fully functioning tissues.