The Journal of Bone & Joint Surgery

The Journal of Bone & Joint Surgery, Volume 90, Issue Supplement 1

Symposium Articles | February 01, 2008

Tissue Engineering of Bone: Material and Matrix Considerations

Yusuf Khan, PhD; Michael J. Yaszemski, MD, PhD; Antonios G. Mikos, PhD; Cato T. Laurencin, MD, PhD

Disclosure: The authors did not receive any outside funding or grants in support of their research for or preparation of this work. Neither they nor a member of their immediate families received payments or other benefits or a commitment or agreement to provide such benefits from a commercial entity. No commercial entity paid or directed, or agreed to pay or direct, any benefits to any research fund, foundation, division, center, clinical practice, or other charitable or nonprofit organization with which the authors, or a member of their immediate families, are affiliated or associated.

The Journal of Bone and Joint Surgery, Incorporated

J Bone Joint Surg Am, 2008 Feb 01;90(Supplement 1):36-42. doi: 10.2106/JBJS.G.01260

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Abstract

When the normal physiologic reaction to fracture does not occur, such as in fracture nonunions or large-scale traumatic bone injury, surgical intervention is warranted. Autografts and allografts represent current strategies for surgical intervention and subsequent bone repair, but each possesses limitations, such as donor-site morbidity with the use of autograft and the risk of disease transmission with the use of allograft. Synthetic bone-graft substitutes, developed in an effort to overcome the inherent limitations of autograft and allograft, represent an alternative strategy. These synthetic graft substitutes, or matrices, are formed from a variety of materials, including natural and synthetic polymers, ceramics, and composites, that are designed to mimic the three-dimensional characteristics of autograft tissue while maintaining viable cell populations. Matrices also act as delivery vehicles for factors, antibiotics, and chemotherapeutic agents, depending on the nature of the injury to be repaired. This intersection of matrices, cells, and therapeutic molecules has collectively been termed tissue engineering. Depending on the specific application of the matrix, certain materials may be more or less well suited to the final structure; these include polymers, ceramics, and composites of the two. Each category is represented by matrices that can form either solid preformed structures or injectable forms that harden in situ. This article discusses the myriad design considerations that are relevant to successful bone repair with tissue-engineered matrices and provides an overview of several manufacturing techniques that allow for the actualization of critical design parameters.

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Current Methods of Fracture Repair

The normal physiologic reaction to fracture is the spontaneous sequence of events briefly summarized as initial inflammation, followed by soft callus formation, hard callus formation, and, ultimately, bone remodeling. When this natural process does not occur, as in the case of fracture nonunions or large-scale traumatic bone injury, surgical intervention is warranted. One surgical approach involves the use of rigid internal fixation with use of either plates or intramedullary rods¹. Neither of these treatments results in complete bone repair and remodeling, however, because the metal plates and rods are not resorbable and cannot be remodeled by the body. In addition, the nonresorbable materials are susceptible to long-term fatigue and fracture. As an alternative, there are bone grafts and bone-graft substitutes. Autografts are used to enhance bone-healing, spinal fusion, and fracture repair and are the current gold standard for bone-graft procedures² because they contain osteogenic bone cells, marrow cells, and an osteoconductive collagen matrix suitable for new and existing bone-cell attachment and migration as well as osteoinductive proteins and factors endogenous to bone. However, autograft requires a second operation at the site of tissue harvest, which increases postoperative pain and creates the possibility of surgical complications at the site (donor-site morbidity). One alternative to autograft is allograft, or tissue taken from a cadaver. Although limited supply is less of a problem, there is a minimal but real risk of disease transmission from donor to recipient with allografts^3,4. Additional allograft complications have been reported at the ten-year mark⁵, with as many as 30% to 60% of allograft implants encountering some sort of complication that may lead to failure of the structural allograft. Given the shortcomings of autografts and allografts and the large demand for bone grafts, several researchers are investigating alternative healing modalities and materials^6-12. Tissue engineering represents one promising strategy.

Tissue Engineering

Investigators have developed synthetic graft substitutes with the goal of mimicking the physical and mechanical nature of native tissue. These graft substitutes, or matrices, are formed from a variety of materials (natural and synthetic polymers, ceramics, and composites)^13-21 that are designed to mimic the three-dimensional characteristics of autograft tissue while also providing the ability to sustain cells seeded onto the construct. Matrices can also act as delivery vehicles for factors, antibiotics, and chemotherapeutic agents, depending on the nature of the injury to be repaired^22-26. This intersection of matrices, cells, and therapeutic molecules has collectively been termed tissue engineering. Through the development of a scaffold that is biocompatible and biodegradable and to which cells can attach, proliferate into, and migrate throughout, the damaged tissue can be replaced slowly over time and can completely replace the initially implanted synthetic matrix, which is an advantage over current treatment strategies. The promise of this approach lies in no small part with the matrix design. The selection of potential matrix materials is vast and allows for great variability in design for a specific application, be it the need for a structural support with mechanical integrity, a delivery vehicle for therapeutic molecules, or a minimally invasive delivery method.

Material Selection

Depending on the specific intended application of the matrix, whether for structural support, drug-delivery capability, or both, certain material categories may be more or less well suited to the final structure. Choices for matrix material include polymers, ceramics, and composites of the two.

Polymers

Polymers can impart a wide variety of physicochemical and mechanical properties to a formed matrix, the synthesis of which can include natural polymers, such as type-I collagen²⁷, hyaluronic acid²⁸, and chitosan²⁹, and/or synthetic polymers, such as polyanhydrides³⁰, polypropylene fumarate^31,32, polycaprolactones³³, polyphosphazenes^34-38,polylactide,polyglycolide, and associated copolymers (polylactide-co-glycolide)^9-12,39-45, to name just a few. Different polymers express different physical attributes, mechanical properties, degradation times, and modes of degradation that can be chosen on the basis of the intended application of the matrix (Table I). For instance, matrices intended to provide structural integrity to a defect site would require the use of polymers that were solid rather than gelatinous or rubbery at physiologic temperatures. The nature of polymer degradation can contribute to the suitability of a matrix for drug-delivery applications. Polymers like polylactide-co-glycolide and polycaprolactone, for example, undergo bulk degradation and may be less well suited for drug-delivery purposes than are surface-eroding polymers, such as polyanhydrides, that would more predictably deliver loaded factors and therapeutic substances^46,47.

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TABLE I Polymers Currently Being Evaluated for Use as Bone Tissue-Engineering Matrices, and Ceramics Currently Used as Bone-Graft Substitutes^*

Material	Degradation Time	Mechanical Properties	Mode of Degradation
Polymers
Type-I Collagen	Varies by degree of cross-linking	High mechanical properties, tunable through varied cross-linking	Enzymatic degradation
Hyaluronic acid	Varies by molecular weight	Viscoelastic in nature, lower than trabecular bone	Enzymatic degradation
Chitosan	Varies by degree of cross-linking and acidity of environment	Below trabecular bone as a hydrogel	Acidic hydrolysis, enzymatic degradation
Polyanhydrides	Varies by degree of polymerization, molecular weight	Lower range of trabecular bone	Hydrolysis
Polypropylene fumarate	Varies by molecular weight, cross-linking density (>3 mo)	Similar to trabecular bone	Hydrolysis
Polycaprolactones	>1 yr	Slightly higher than trabecular bone	Hydrolysis
Polyphosphazenes	Ranges from 1-3 wk to >2 yr	Ranges from far below to far above trabecular bone	Hydrolysis
Polylactide-co-glycolide	Ranges from 6 wk to 1 yr	Ranges through trabecular bone depending on three-dimensional structure	Hydrolysis
Ceramics
Calcium phosphate	Ranges from 6 mo to >1 yr	Ranges based on three-dimensional structure but generally above trabecular bone	Dissolution, osteoclastic resorption
Calcium sulfate	4-12 wk	Approximately the same as trabecular bone	Dissolution
Bioactive glass	>1 yr	Within range of trabecular bone but dependent on three-dimensional structure	Dissolution, osteoclastic resorption

Composites of polymers and ceramics are also actively being evaluated as potential bone tissue-engineering matrices

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TABLE I Polymers Currently Being Evaluated for Use as Bone Tissue-Engineering Matrices, and Ceramics Currently Used as Bone-Graft Substitutes^*

Material	Degradation Time	Mechanical Properties	Mode of Degradation
Polymers
Type-I Collagen	Varies by degree of cross-linking	High mechanical properties, tunable through varied cross-linking	Enzymatic degradation
Hyaluronic acid	Varies by molecular weight	Viscoelastic in nature, lower than trabecular bone	Enzymatic degradation
Chitosan	Varies by degree of cross-linking and acidity of environment	Below trabecular bone as a hydrogel	Acidic hydrolysis, enzymatic degradation
Polyanhydrides	Varies by degree of polymerization, molecular weight	Lower range of trabecular bone	Hydrolysis
Polypropylene fumarate	Varies by molecular weight, cross-linking density (>3 mo)	Similar to trabecular bone	Hydrolysis
Polycaprolactones	>1 yr	Slightly higher than trabecular bone	Hydrolysis
Polyphosphazenes	Ranges from 1-3 wk to >2 yr	Ranges from far below to far above trabecular bone	Hydrolysis
Polylactide-co-glycolide	Ranges from 6 wk to 1 yr	Ranges through trabecular bone depending on three-dimensional structure	Hydrolysis
Ceramics
Calcium phosphate	Ranges from 6 mo to >1 yr	Ranges based on three-dimensional structure but generally above trabecular bone	Dissolution, osteoclastic resorption
Calcium sulfate	4-12 wk	Approximately the same as trabecular bone	Dissolution
Bioactive glass	>1 yr	Within range of trabecular bone but dependent on three-dimensional structure	Dissolution, osteoclastic resorption

Composites of polymers and ceramics are also actively being evaluated as potential bone tissue-engineering matrices

Ceramics

A ceramic is a material made from an inorganic, nonmetallic material that can possess a crystalline structure. Ceramics typically have a high compressive strength and low ductility, meaning that they provide high resistance to deformation but that they tend to fail because of their brittle nature. Calcium phosphates, calcium sulfates, and bioactive glass have been used as matrices for bone regeneration^14,17,18. These substances, especially the calcium phosphates, are ideal candidates for use as matrices because the inorganic component of bone is composed of the ceramic calcium hydroxyapatite. Ceramics are typified by high compressive strengths and variable degradation times; for example, the degradation time for calcium sulfate can range from approximately four to twelve weeks⁴⁸, whereas crystalline hydroxyapatite can be nondegradable^49,50. Both calcium sulfate and crystalline hydroxyapatite have shown success as a bone-graft substitute, although some authors have reported the development of seromas when calcium sulfate was used as a carrier for antibiotics⁵¹. The calcium phosphates also have potential as drug and/or factor-delivery vehicles as a result of the high binding affinities between ceramics and proteins. Ceramics such as calcium phosphate and bioactive glass are also considered biomimetic, in that they stimulate the formation, precipitation, and deposition of calcium phosphate from solution and can result in enhanced bone-matrix interface strength⁵². One consideration for the use of a ceramic alone as a matrix material is its high compressive strength. Often the compressive modulus of such ceramics exceeds the values commonly seen in trabecular bone, which range from approximately 20 MPa to 2 GPa, depending on the location within the skeleton⁵³. This mechanical mismatch can result in stress shielding around the matrix implantation site and should be considered when structures are being designed to act as temporary scaffolds for bone ingrowth. Furthermore, ceramics tend to be very brittle in compression and fail catastrophically when upper loading limits are exceeded. This brittle failure has relegated certain ceramics, such as calcium sulfate, to non-load-bearing sites for repair.

Polymer-Ceramic Composites

Both polymers and ceramics have their advantages and drawbacks, but it is possible to minimize these drawbacks and maintain the advantages by combining polymers and ceramics into one composite material. Polymer-ceramic composites capture the advantages of each component—namely, the vast array of polymers available with varying mechanical properties, degradation times, and physical structures, and the inherent strength and biomimetic nature of ceramics. Furthermore, the addition of polymers to the ceramics reduces their overall brittleness, while the addition of ceramic to a polymer increases both its bioactivity and its capacity to take up and deliver factors and therapeutic substances.

Matrix Design Considerations

Given the array of materials available to the tissue engineer, it is possible to design a matrix with rather specific attributes, whether mechanical, architectural, or degradable, provided that those attributes satisfy a short list of required criteria without which the matrix will be unable to regenerate new osseous tissue. These attributes include (1) biocompatibility—the lack of immunogenic response; (2) osteoconductivity—the quality of a porous interconnected structure that permits new cells to attach, proliferate, and migrate through the structure and also allows for nutrient-waste exchange and new vessel penetration; (3) osteoinductivity—possessing the necessary proteins and growth factors that induce the progression of mesenchymal stem cells and other osteoprogenitor cells toward the osteoblast lineage (this is essential when the defect is of critical size or of sufficient size to prevent spontaneous healing); (4) osteogenicity—the osteoblasts that are at the site of new bone formation are able to produce minerals to calcify the collagen matrix that forms the substrate for new bone; and (5) osteointegration—the newly formed mineralized tissue must be able to form an intimate bond with the implant material.

Beyond this list of essentials, there are other design considerations that can alter the suitability of the matrix for its intended application. These include surface roughness, mechanical integrity, and the overall porosity of the matrix.

Surface Roughness

Cells attach, migrate, proliferate, and differentiate differently depending on the surfaces to which they are attached. Studies have shown variations in cellular behavior based on whether or not a surface was textured⁵⁴. Within the scope of textured surfaces, studies have shown variations in cellular behavior based simply on the size of the texture, with certain cellular behaviors elicited by nanoscale modifications on the surface of the material. These variations may be due in part to the focal adhesions that form when a cell comes in contact with a surface. These collections of proteins provide the link between the extracellular and intracellular environments via the interaction between extracellular matrix proteins such as collagen and fibronectin, transmembrane proteins such as integrins, and cytoskeletal elements such as actin, which communicate with the nucleus of the cell and impact cellular behavior.

Mechanical Integrity

Bone responds to the absence and presence of physical load. In response to these loads, the body either resorbs or forms bone. Given this principle, it is important to design a matrix that possesses mechanical properties that are similar to the tissue in the immediate surrounding area of the defect. An overengineered matrix may result in bone resorption around the implant site, while an underengineered matrix may fail as a mechanical support to the skeleton. The mechanical properties of a matrix can be varied by material selection, by the formation of composite structures, or by the overall porosity of the matrix.

Porosity

In a direct relationship to osteoconductivity, a porous structure is critical both to control matrix mechanical strength and to allow osteoprogenitor cells and osteoblasts to occupy the entire matrix after it is implanted. Porosity alone, however, is not sufficient for complete cell infiltration and survival. The pores must be of suitable diameter to allow osteoblasts and osteoblast-like cells to migrate into the center of the matrix and thus allow complete healing. Similarly, this pore structure must be sufficient to allow nutrient and waste exchange from the perimeter of the matrix to its center, in order to provide a viable environment for the cells that populate the center of the matrix⁵⁵. Toward this end, the pore structure within the matrix must have complete interconnectivity. Any closed pores will prevent osteoblasts from occupying that region of the matrix, unnecessarily weakening the structure. As is implied, the necessity of a pore structure forces critical design decisions when considering the structural integrity of the matrix.

Scaffold Types

Matrices can be divided roughly into two categories—preformed and injectable. There are several methods to fabricate both preformed and injectable matrices, with the selection of the material in part dictating the type of matrix formed. Each type should possess the essential design parameters listed above, but the choice as to whether preformed or injectable is most suitable can depend on the route of administration.

Preformed Matrices

Preformed matrices can be designed with pore structure, diameter, and interconnectivity as a result of the broad availability and diversity of matrix-forming techniques. These techniques include laser stereolithography, inkjet printing, selective laser sintering, heat sintering, fused deposition modeling, and injection molding.

Laser stereolithography typically combines a photo-polymerizable polymer in a liquid form that cross-links or cures when exposed to ultraviolet light³¹. Computer models of matrices are interpreted by the rapid-prototyping machines, and matrices are formed layer by layer, with each subsequent layer being built on top of the previous one to form three-dimensional structures that, depending on the design, may possess porous interconnected networks. The final formed matrix is typically cured further to complete the fabrication. Several biocompatible polymers have been evaluated as candidate materials for stereolithography, and these polymers have been combined with ceramics, such as hydroxyapatite, and biological agents to increase the bioactivity and biomimicry of the matrices formed.

Inkjet printing operates on a similar principle in that matrices are built layer by layer, but rather than a laser being used to cure a liquid polymer, a binder solution is sprayed onto a bed of polymeric powder. When the two come in contact, the polymer is hardened⁵⁶. The precision design comes from the accuracy of the inkjet to deposit the binder on the surface of the polymer. Each layer is formed on top of the previous one, and a three-dimensional matrix is formed. This technique has been expanded to include the actual deposition of cells on surfaces with use of the same method⁵⁶. Rather than the deposition of a binder on polymeric powder, cells are deposited in the pattern and location desired. Studies have shown that cells can both survive this procedure and maintain their function⁵⁷.

Selective laser sintering operates on a similar principle to that of inkjet printing in that a binder is added to a layer of polymeric powder⁵⁸. The difference, however, lies in the binder. With selective laser sintering, a laser acts as the binder by heating up the polymer powder, resulting in the fusion and hardening of the powder underneath the laser with neighboring particles, including those on the layer below. As with other rapid-prototyping methods, this process is done layer by layer to build a three-dimensional structure.

Heat sintering works under the same premise as laser sintering, but relies on ambient heat to fuse neighboring polymeric particles together. This technique has been widely used in conjunction with polymeric and polymer-ceramic composite microspheres^{9-12,29,40,42-45}. The strength of this technique lies in the resulting three-dimensional shape, which forms a completely interconnected pore structure, the pore diameter of which can be modulated by selecting suitable-size spheres for sintering. The mechanical strength of this matrix comes from the bond between neighboring microspheres and can therefore be modulated by varying the exposure time to ambient heat.

Fused deposition modeling, another rapid-prototyping platform, is a technique in which molten polymer is extruded through small openings that allow for the controlled spatial deposition of the molten polymer, which hardens in place immediately after being extruded⁵⁹. By modifying the direction and orientation of each polymer extrusion between layers, the desired architecture is built. In this fashion, each layer can be constructed on the previous one to develop complex three-dimensional structures.

Injection molding involves the delivery or injection of molten polymer under high pressure into a preformed mold in which the polymer is allowed to cool and harden and assume the shape of the mold itself³¹. A similar technique is solvent casting, in which a polymer is put into solution with an appropriate solvent and then poured into a mold in which the polymer hardens as the solvent is removed through room temperature evaporation, solvent exchange, or under elevated heat to quicken the solvent removal. In all situations, the final matrix takes the shape of the mold itself; thus, designing the mold is the critical parameter in fabrication.

Each of these techniques allows for considerable control of the final matrix architecture, which can be tuned for its ultimate application. Surgical instances suited for preformed matrices are those that require invasive surgical intervention, such as the surgical treatment of segmental defects. Restoration of an entire long-bone segment that has been damaged as a result of injury or tumor removal requires structural integrity. Noncontained defects, such as segmental defects, also require a preformed matrix with considerable mechanical strength. In this instance, the matrix will be used to rebuild the osseous tissue that has either been removed or damaged and to provide the necessary underlying mechanical structure to the osseous tissue in that area. Additionally, defects in which the shape can be largely predicted prior to surgery, such as those associated with tumor resections or spinal fusions, are well-suited for preformed matrices. Often the matrix itself can be trimmed to shape as necessary. Regardless of the synthesis method, these techniques describe the formation of a solid three-dimensional structure that will be surgically implanted and will provide structural integrity to the defect site. There is a trend, however, toward minimally invasive treatment paradigms within the field of orthopaedics, and tissue-engineered matrices provide options there as well.

Injectable Matrices

Injectable matrices are gaining favor in the orthopaedic field as materials and techniques improve with new technologies. Injectable matrices are indicated for use in trabecular defects in which the damaged skeletal tissue is not load-bearing, in contained defects when the major structural osseous tissue is still intact, in spinal fusion gutters, and in small tumor-removal sites in the appendicular skeleton. Candidates for use as injectable matrices include those materials that are at one point in a liquid or gelatinous state, are flowable, and will harden in an aqueous environment at 37°C (body temperature). Certain polymers, known as thermogelling polymers, retain a liquid state at room temperature but undergo a gelling or hardening process when the temperature is raised to 37°C⁶⁰. This property allows the polymer to be injected into a bone-defect cavity or space through a narrow-bore syringe, minimizing the surgical site to a small cutaneous incision. Theoretically, radiopaque markers can render the material visible under fluoroscopy to determine appropriate deposition into the defect site. These polymers are not capable of sustaining high compressive loads, but they have been evaluated for their suitability for other purposes, such as spinal disc replacements⁶⁰. Ceramics in the form of calcium phosphate cements have been developed to be injectable and then to subsequently harden in situ to a solid state within several minutes⁶¹. Calcium phosphate precursors are mixed with a liquid hardener in the surgical suite just prior to implantation. These cements are virtually nonexothermic, unlike other polymeric bone cements such as polymethylmethacrylate. Polymethylmethacrylate, a cement commonly used to bond metallic implants to host bone, can raise the local temperature high enough to leave necrotic tissue. Leaching of unreacted monomer from these cements can be cytotoxic as well, leading to local inflammation⁶². Calcium phosphate cements are commonly a blend of different calcium phosphates that are biocompatible and, when fully cured, form a low-crystallinity hydroxyapatite that can be resorbed and remodeled over time. One concern with injectable matrices is the absence of a pore structure. As was described above, a pore structure is crucial to complete tissue incorporation and remodeling. Injectable matrices do not necessarily possess a pore structure, which may reduce new tissue ingrowth; the lack of this attribute negatively impacts the effectiveness of injectable matrices as bone tissue-engineering matrices. Generating pores in injectable matrices is a current topic of research in which investigators are examining fast-degrading particles that can dissolve within a short time after injection and hardening, leaving a porous but rigid structure behind. Others have evaluated the generation of bubbles in calcium phosphate cements that would form an inner pore structure after injection and would be maintained after hardening.

Summary

The advent of tissue engineering has resulted in the development of many new matrices designed to encourage bone regeneration. The materials available allow the design of these matrices to be tailored to necessary mechanical properties, degradation times, required bioactivity, factor and drug delivery, and routes of introduction into the body, depending on the specific needs for orthopaedic repair. ?

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Topics

antibiotics ; bone transplantation ; ceramics ; lasers ; polymers ; porosity ; tissue engineering ; transplantation, autologous ; drug vehicle

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Yusuf Khan, Michael J. Yaszemski, Antonios G. Mikos, Cato T. Laurencin; Tissue Engineering of Bone: Material and Matrix Considerations. The Journal of Bone & Joint Surgery. 2008 Feb;90(Supplement_1):36-42.

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