Biocompatibility and Mechanical Properties of a Totally Absorbable Composite Material for Orthopaedic Fixation Devices

Kirk P. Andriano and A. U. Daniels

Orthopedic Bioengineering Laboratory, Division of Orthopedic Surgery, University of Utah School of Medicine, Salt Lake City, Utah

Jorge Heller

Controlled Release and Biomedical Polymers Department, SRI International, Menlo Park, California

Bioabsorbable polymer/inorganic phosphate fiber composites are prone to rapid degra- dation due to water sensitivity of the interface between the degradable polymer and the degradable fiber. This article describes successful fabrication and laboratory evaluation of a candidate bioabsorbable composite implant material with mechanical properties similar to bone. The composite studied was poly(ortho ester) reinforced with randomly- oriented, crystalline microfibers of calcium-sodium-metaphosphate. The component materials showed no acute cytotoxicity as determined by tissue culture agar overlay. Treating the microfibers with a diamine-silane coupling agent improved mechanical properties and slowed degradation in saline, byt strength still decreased 50% in 1 week. When the composite material was then coated with a layer of matrix polymer alone it retained 70% of its strength and 70% of its stiffness after 4 weeks exposure to 7.4 pH Tris-buffered saline at body temperature. The marked improvement with the coating can be attributed to the hydrophobicity of poly(ortho esters).

INTRODUCTION

Background

The use of metal plates for internal fixation of bone often prevents formation of the normal callus and also results in incomplete healing.l,? A primary reason for this clinical observation is that the plate bears a substantial portion of the load normally seen by the bone. The elastic modulus mismatch between bone (E = 6-20 GPa) and the metals (E = 100-200 GPa) used for internal fracture fixation devices means that a plate whose cross section is only a fraction of the plated bone has a structural stiffness approximating that of the bone. In isoelastic systems, load bearing is in proportion to structural stiffness of the components (e.g., bone and plate). According to Wolff's Law, bone is a dynamic material and remodels itself in accord to applied stresses to achieve and maintain density and structural thickness. Rigid internal fixation of bone thus shields healing bone from normal stresses, resulting in below normal density and thickness. Also the eventual surgical removal of metal plates and screws after they have served their stabilization function leaves the bone at

Requests for reprints should be sent to Dr. Kirk P. Andriano, Biomaterials Laboratory, Institute of Plastics Technology, Tampere University of Technology, P.O. Box 527, SF-33101 Tampere, Finland.

Journal of Applied Biomaterials, Vol. 3, 197-206 (1992) 0 1992 John Wiley & Sons, Inc. CCC 1045-4861/92/030197- 10$4.00

least temporarily deficient in structural strength and prone to refracture.

Lower-modulus, nonabsorbable fiber reinforced poly- mer composite bone plates have shown promise in re- ducing this stress shielding phen~menon,~-~ but these plates must still be removed in many cases to prevent complications. Totally absorbable composites&" with an elastic modulus closer to that of bone have been proposed as an alternative. Fixation devices of these materials would provide less rigid fixation. This would allow some fracture motion and resultant formation of normal callus not seen with rigid fixation. However, rigidity must not be too low or nonunion may result. Absorption of the device with time would further decrease rigidity gradually, lessening the effects of the stress-shielding and allowing the bone to gain normal density and thickness. Also absorption would eliminate the need for second surgeries.

Totally absorbable composites have been made using absorbable polymers reinforced with higher strength ab- sorbable organic polymer fibers. Both matrix and fiber are typically poly(glyco1ic acid) (PGA) or poly(1actic acid) (PLA), and when compared to unreinforced polymer, the composites show large increases in strength but little in- crease in stiffness since the fibers are p ~ l y m e r i c . ' ~ , ~ ~ PLA matrix composites reinforced with continuous absorbable, calcium-metaphosphate glass fibers show a large increase in both strength and rigidity.14,15

Unfortunately, composites of this type made to date show a large loss in strength after short-term exposure to an aqueous physiological environment, but do retain

1 98 ANDRIANO, DANIELS, AND HELLER

a greater proportion of their stiffness. Reinforcement of absorbable polymers with short nonabsorbable fibers may result in better retention of mechanical strength after aqueous e x p o s ~ r e , ' ~ ~ ~ ' but the fiber residue could result in an inflammatory response and mechanical irritation if fibers enter the joint space.

Additionally, clinical studies of absorbable polymeric fracture fixation devices have identified a serious problem for devices made from PGA and PLA. Although fracture healing occurred with these devices, there were reported rates of between 7 and 48% of an associated noninfectious foreign body inflammatory response requiring clinical interventi~n.",'~ Depending on the polymer, the time for manifestation of this local toxic response was from 7 weeks to nearly 3 years after implantation.

This study describes a first effort to create a completely absorbable composite implant material, using an absorbable poly(ortho ester) (POE) matrix rein- forced with absorbable, randomly oriented, calcium- sodium-metaphosphate (CSM) microfibers. Such a composite could be used in devices for low-load mechanical stabilization of hard or soft tissues. An initial design goal was to produce a benignly absorbable composite material with mechanical properties similar to bone which would retain the majority of these properties for the initial healing period of bone, approximately 4 to 6 weeks.

The methods used to develop and evaluate this totally absorbable composite material included (a) biocompati- bility testing of the reinforcing fibers and matrix polymer to determine cytotoxicity, (b) characterization of POE polymers with differential scanning calorimetry (DSC) and dynamic rheometry to aid in composite fabrication, (c) dissolution studies of the CSM fiber, (d) evaluation of fabrication variables such as fiber content and pressing temperature to optimize composite properties, and (e) an in vitro wet strength retention study of the unaugmented composite and of composites with fibers treated with a silane coupling agent and with hot-film coating of the entire specimen with unreinforced poly(ortho ester).

As a result of these studies, a CSM/POE composite was fabricated successfully using silane treated fibers. The component materials showed no acute cytotoxicity. The composite material had mechanical properties similar to cortical bone, and when coated with a layer of unrein- forced POE retained 70% of its strength and 70% of its stiffness after 4 weeks of exposure to Tris-buffered saline at body temperature.

MATERIALS AND METHODS

Polymer Description and Preparation

Poly(ortho esters) are a family of synthetic absorbable polymers that have been under development for medi- cal applications for a number of years.20 This family of highly hydrophobic absorbable polymers is completely

amorphous.21 Hydrolytic degradation occurs predomi- nately by surface hydrolysis,22 in contrast to the bulk hydrolysis of the more hydrophilic PGAs and P U S . POEs have been successfully used as erodible matrices for drug delivery of therapeutic agent^^^,^^ and explored for potential use in fracture fixation device^.^' Polymer degradation rates can be accelerated by adding acid excipients to the bulk polymer or retarded by adding basic excipients.26 Degradation rates also increase with decreasing molecular weight.

The polymers investigated here were linear poly(ortho esters) prepared by a condensation reaction of 3,9- bis(ethy1ene 2,4,8,lO-tetraoxaspiro[5,5]-undecane) and a 60:40 and 90:lO mole ratio of the rigid diol, frans- 1,4-cyclohexanedimethanol and the flexible diol, 1,6- hexanediol, respectively (Fig. 1). Polymers with a wide range in glass transition temperatures (Tg), 20 to 120 "C, can be synthesized by varying the ratio of rigid to flexible diols. For hot molding processes, molecular weights of less than 100000 are best in order to avoid the necessity of high fabrication temperatures required by high melt viscosities.21

When a POE of this type is placed in an aqueous environment, it undergoes an initial hydrolysis of the ortho ester linkages to yield a mixture of the diols used in the synthesis and the dipropionate of pentaerythritol. This diester eventually hydrolyses to pentaerythritol and propionic acid.24

POLY(ORTH0 ESTER)

LH3CH2 0-CH2 CH2-0 CH2CH3 1 \ / C

\ / C

\ / C

&d-'O-CH, / - \ CH2-0 O-R

"R" CAN BE

RIGID - trans-cyclohexanedimethanol

-CH2- 0 -CH2-

FLEXIBLE - 1,6-hexanediol

- (CH 2 16- Figure 1. Chemical structure of poly(ortho ester) repeat units.

ORTHOPAEDIC FIXATION DEVICES 199

Even though ortho ester linkages are much more labile than the ester linkages, the polymer is highly hydrophobic, so that the hydrolysis is largely confined to the outer surface of a polymer specimen. Therefore, degradation of this polymer may not generate acidic products rapidly at the implantation site, provided that the water soluable, low molecular weight products of the first hydrolysis at the solidiliquid interface can diffuse away. This could be a significant advantage over the bulk hydrolysis of polyesters that hydrolyze directly to yield acidic com- pounds as the primary degradation products.

The 60:40 polymer used here had a weight average molecular weight of 86000 and a Tg of 66 "C. The 90:10 material had a molecular weight of 73000 and a Tg of approximately 100 "C. The polymers were dried in a vacuum oven at 40 to 50 "C for 48 h and stored in dry air sealed containers at ambient conditions until needed.

Because of extremely limited quantities of polymer and labor intensive fabrication procedures, sample sizes re- ported in this research for mechanical testing are low, and therefore no statistical analysis of the data was performed.

Fiber Description and Preparation

Calcium-sodium-metaphosphate (CSM) mineral mi- crofibers were developed as a nontoxic replacement for asbestos but the material has not been used commercially as yet. This degradable and bioabsorbable microfiber is a crystalline, inorganic polymer of metaphosphate whose negative charge is balanced by calcium and sodium ions (Fig. 2).27,28 The fiber exhibits an effective aspect ratio of about 60:l and potentially can be used with a wide range of coupling agents for reinforcing thermoplastic and thermosetting polymers.

CSM microfibers are manufactured from pure ele- mental phosphorus using oxygen, sodium carbonate, and lime.27 The microfibers cover a wide range of lengths and diameters from submicron to hundreds of microns depending on the type and degree of comminution to which the material is exposed. These crystalline fibers pos- sess the tensile strength and stiffness of Kevlar, 2.6 GPa and 124 GPa, re~pec t ive ly .~~ In comparison, CSM microfibers are much stronger and stiffer than absorbable

STRUCTURAL FORMULA OF CSM

*2 + Ca Na

0- 0- 0- - ~ -

I I I (-o-P-o-P-o-P-o-)"

8 8 8 Figure 2. Chemical structure of the polymetaphosphate inorganic polymeric backbone chain. The poly-metaphosphate polymers lie next to each other in planes, with cations in the spaces in between and ionically bonded to the polymer.

calcium-metaphosphate fibers (tensile strength = 0.35 GPa and tensile modulus = 28 GPa) which have been used to reinforce poly(1actic acid).29

One kilogram of calcium-sodium-metaphosphate mi- crofibers (T Grade, long fibers) was donated for this work by Monsanto Chemical Co. (St. Louis, MO). Raw CSM fibers contain about 8% glass which contains heavy metals that must be removed before use in biological applications. This was accomplished by boiling 25 g of CSM fibers in 500 mL of deionized water for 4 h, maintaining the water level as needed. The hot suspension was then filtered through a Buchner funnel under vacuum and washed several times with deionized water at room temperature. The fibers were dried in a vacuum oven at 90 to 100 "C for 24 h and stored in a desiccator until needed.

The surface of washed CSM fibers is slightly acidic. This is due to chain end hydrolysis of the inorganic polymer, resulting in a small quantity of surface acidic degradation products. This free hydrogen can react with the methoxy -groups of commercially available silane coupling agents to form silanol bonds between the inorganic fiber surface and the coupling agent.30 Because an acidic fiber surface will catalyze hydrolysis of POEs, the fiber surface was made basic by treating it with a diamine-silane, (N,beta-aminoethyl-gamma- aminopropyl-trimethoxysilane; 2-6020, Dow Corning, Midland, MI).

Fifteen grams of washed CSM fibers were slowly added to 100 mL of a 0.3% solution of silane coupling agent in methanol (OmniSolv, EM Sciences, Gibbstown, NJ) (v/v) and stirred using a magnetic stirrer until a slurry was formed. The slurry was then fiitered using a Buchner funnel, and the residue dried in a convection oven at 90 to 100 "C for 3.5 h. After cooling to room temperature, the treated fibers were sieved through a 40 mesh Tyler sieve using a Portable Rototap (W. S. Tyler Inc., Mentor, OH), washed with excess methanol to remove any unbonded coupling agent, and finally dried in a convection oven at 90 to 100 "C for 1 h. Treated fibers were stored in a desiccator at ambient conditions until needed.

Biocompatibility Testing

Acute cytotoxicity of 60:40 POE, 90:lO POE, CSM fibers, and silane-treated CSM fibers was evaluated by standard tissue culture agar overlay using the direct cell contact method and L929 mouse fibroblast cells.

After sterilization with ethylene oxide, appropriate amounts of each test material were aseptically placed onto separate solidified agar, along with positive and negative controls (natural black rubber and polyethylene, respectively). Samples and controls were tested in triplicate and incubated at 37 "C for 24 h. Neutral Red was added for 2 to 3 h. The cells were then evaluated using an inverted microscope.

The results were scored by adding the area of decol- oration (zone index) and the area of lysis (lysis index). The

200 ANDRIANO, DANIELS, AND HELLER

final score was recorded as the response index, which is the ratio of zone index to lysis index.

Characterization of Polymers

As an aid in the development of processing conditions for POEs, the thermal and rheological properties of the 60:40 and 90: 10 POE polymers were determined. Differential scanning calorimetry (DSC) was used to determine the glass transition temperature of the polymer samples and their propensities for thermal degradation at temperatures above their Tgs. In addition rheometry was used to deter- mine the visco-elastic properties of the polymer samples in their melt phases. The instruments used were in Om- nitherm DSC 700 using a heating rate of 10 "C/min and a Rheometrics Inc. Model RDS-I1 Dynamic Spectrometer.

Characterization of Microfibers

In order to quantify both CSM microfiber degradation due to solvation of polymer chains at the fiber surface and water resistance of the fiber/silane coupling agent bond, two degradation studies were performed. First, 50 mg of untreated CSM fiber samples were placed into 25 mL polystyrene vials containing 20 mL of normal saline while another set of fibers samples were placed in vials contain- ing 20 mL of distilled water. The sample filled vials were sealed and placed in a shaker bath and maintained at a temperature of 37 "C. Fiber samples immersed in normal saline and distilled water were removed at 1, 2, 4, and 7 weeks and dried in a vacuum oven at 50 "C for 48 h. The remaining fibrous residue was weighed using a standard laboratory analytical balance and any change in weight was recorded.

Second, the amount of coupling agent actually placed onto the fibers can be semi-quantitatively determined by titration of a fiber slurry in water, provided the coupling agent has acidic or basic organo-functional groups." To determine if a water resistant bond had been achieved between the fiber surface and the coupling agent, 2.5 g of CSM fibers sized with basic diamine-silane were boiled in 100 mL of deionized water for 1 h. The slurry was then filtered using a Buchner funnel under vacuum and dried in a convection oven at 90 to 100 "C for 30 min. Titration of a 2% slurry of the above treated fibers in deionized water (2.0 g fiber/lOO g HzO), using 0.001N sulfuric acid was performed until pH 4 was reached. Similar titrations were performed with CSM fibers treated with diamine-silane and untreated CSM fibers, followed by comparison of the titration curves. The greater the volume of sulfuric acid needed to reach the pH of 4, the greater the number of basic sites on the fiber surface, hence the greater the amount of coupling agent bonded to the fiber surface. This was used as a semi-quantitative indicator for comparing the amount of diamine-silane coupling agent bonded to the fiber before and after boiling.

Fabrication Variables

Composite fabrication. Composite specimens mea- suring 38.1 X 12.7 X 1.60 mm were prepared in the fol- lowing manner: first, 90:lO POE was reduced to a granular form, particle diameter < 250 pm, by milling the poly- mer in a Thomas-Wiley Intermediate Mill (Arthur H. Thomas Co., Philadelphia, PA), at room temperature in open air, and using material which passed through a 40 mesh screen. Fiber and polymer were brought together by a dry-mixing technique in air. CSM fibers were placed into a narrow bottle (diameter = 5 cm and length = 15 cm) fitted with a stirring rod assembly. Milled polymer was sprinkled into the bottle while blending the fibers with the impeller of the stirring rod. After final addition of the polymer, blending was maintained for 15 min at low to medium speed (10-30 rpm).

Finally, test specimens were prepared by hot com- pression molding in air of an appropriate amount of dry-mixed composite in a three-piece stainless steel die at selected temperatures (20-100 "C) above the Tg and at a pressure of 2700 psi for 5 min using a modified Carver Hydraulic Press (Fred S. Carver Inc., Menomonne, WI) with heated and cooled platens. Composite speci- mens were then cooled to ambient temperature while maintaining pressure at 2700 psi.

Fiber content. To determine what effect fiber content would have on initial mechanical properties, CSM/POE composite specimens were fabricated from CSM fibers treated with diamine-silane coupling agent and 90: 10 POE. Fiber content was 0, 10, 20, 30, 40, and 50% by volume. This corresponds to fiber content by weight of 0, 22, 38, 52, 63, and 72%, respectively (density of POE = 1.14 g/cc and density of CSM = 2.86 g/cc). Composites were fabricated as previously described with the pressing temperature being 170 to 175 "C. Replicate composite specimens were tested to failure in three-point bending to determine flexural yield strength and stiffness.

Pressing temperature. A small study was carried out to determine what effects pressing temperature would have on initial flexural mechanical properties of CSM/POE composites.

CSM/POE composite specimens were fabricated from 30 volume percent CSM fibers treated with diamine-silane coupling agent and 90:lO POE. Speci- mens were fabricated at pressing temperatures of 120, 170, 185, and 200 "C as previously described. Initial flexural yield strength and flexural modulus were determined in three-point bending.

Mechanical testing. The flexural mechanical proper- ties of test specimens were determined by three-point bending in accordance with ASTM Standard D 790-81 Method l.32 The diameter of the load nose was 12.6 mm. The diameter of supports was 6.4 mm and the distance between the supports was 25.4 mm. Span-to-depth ratio

ORTHOPAEDIC FIXATION DEVICES 201

was 16 to 1. An Instron Model 1125 materials testing machine was used to load the specimens. Crosshead speed was 5 mm/min.

Flexural yield strength and modulus of elasticity were calculated using the following equations.

placed in the cavity of a rectangular stainless steel mold (38.1 X 12.7 mm). Gentle tapping on the sides of the mold helped to evenly distribute the powdered polymer at the bottom of the mold. A composite test specimen was very carefully placed on top of the powdered polymer

where

and

S = 3PL/2bd2 S = stress in the outer fibers at midspan P = load

L = support span b = width of specimen d = depth of specimen

E = L3rn/4bd3 E = modulus of elasticity in bending L = support span b = width of specimen d = depth of specimen m = slope of the load-deflection curve

Wet Strength Retention Experiments

The rapid strength loss of absorbable poly(1actic acid) composites made with absorbable calcium-metaphosphate fibers is due to the disruption of the polymer/fiber interface by water,33 and not due to the degradation of either the polymer or the reinforcing fibers. As fluids diffuse into the polymer or wick along the fibedfiber interfaces, the polymer/fiber interface is compromised and the fibers lose their strength reinforcing effect. Since absorbable phosphate fibers degrade through surface dissolution,8 the adhesion between the fiber and the polymer matrix is broken once water molecules reach the interface. To see if an outer coating would slow the loss of composite strength and stiffness by retarding the influx of fluids, a comparison study was performed.

The first group of composite test specimens were prepared by hot compression molding of 1.27 g of dry- mixed 90:lO POE polymer and untreated CSM microfibers (Vf = 30%) and pressed at 180 to 185 "C as previously described.

Additional specimens were prepared by the same method using CSM fibers treated with the diamine-silane coupling agent. Half of these CSM/POE composite specimens were encapsulated in pure 60:40 POE by hot- pressing the POE around the test specimens as described below.

Covering composites with hot-pressed POE films. Composite test specimens were machined to the following dimensions: 36.5 X 11.1 X 1.00 mm. Milled 60:40 POE, rather than 90:lO POE, was used to encapsulate the composite specimens because of its lower glass transition temperature. One fifth of a gram of 60:40 POE was

and centered. One quarter of a gram of milled 60:40 POE was placed on top of the test specimen; gentle tapping on the sides of the mold evenly distributed the powdered polymer. The upper plunger of the mold was then fitted into place and the polymer/composite system was further processed by hot compression molding at 115 to 120 "C for 5 min at 2700 psi. Encapsulated composite specimens were cooled to ambient temperature while maintaining pressure at 2700 psi. Reproducibility of the encapsulation procedure was excellent resulting in a uniform coating of polymer for all specimens fabricated. Examination of specimen cross-sections by light microscopy (X60) revealed the coating procedure left approximately a 300- p m thick coating of polymer on the test specimens that appeared free of voids and cracks. The thickness of the coating was evaluated by measuring the thickness of the specimens before and after the coating process, using a micrometer accurate to +/-5 pm.

Initial flexural yield strength and modulus were measured using three-point bending (ASTM D 790-81). Specimens for in vitro degradation tests were immersed in Tris-Buffered saline (200 mM Tris and 150 mM NaCl), pH = 7.4 at 37 "C. Flexural mechanical properties were determined in triplicate (except for specimens fabricated with untreated CSM fibers where no replicate testing was performed) at various times periods up to 6 weeks.

Fabrication of composite specimens with untreated acidic CSM fibers and acid sensitive POE was extremely difficult, frequently yielding specimens which could not be removed from the mold and often demonstrating signs of degradation. Consequently only a limited number of usable specimens were obtained for this experiment.

RESULTS

Biocompatibility Testing

60:40 POE, 90:lO POE, untreated CSM fibers, and CSM fibers treated with the diamine-silane coupling agent all were acutely noncytotoxic (Tissue Culture Agar Overlay). Responses were comparable to negative controls (Table I).

Characterization of Polymers

The DSC scans of 90:lO POE revealed that its Tg was about 98 "C and that it underwent some degree of thermal degradation when heated above about 180 "C in air. The rheological scans indicated that the viscosity of the melt behaved normally between 110 and 180 "C. These results indicate that the processing window for this polymer is 110 to 180 "C. Within this temperature range, the melt

202 ANDRIANO, DANIELS, AND HELLER

TABLE 1. Results of Standard Tissue Culture Agar Overlay Response Index = Zone Index/Lysis Index for Direct Test ___

Sample Replicate 3 Average Replicate 1 Replicate 2

+ Control 2.015.0" 2.015.0 2.015.0 2.015.0 - Control o.o/o.o o.o/o.o 0.0/0.0 o.o/o.o 60:40 POE o.o/o.o 0.0/0.0 o.o/o.o o.o/o.o

Results: Noncytotoxic + Control - Control 90:lO POE

2.0/5.0 2.015.0 2.015.0 o.o/o.o o.o/o.o o.o/o.o 0.0/0.0 o.o/o.o o.o/o.o

Results: Noncytotoxic + Control 2.015.0 2.0/5.0 2.015.0 - Control o.o/o.o o.o/o.o o.o/o.o CSM Fibers o.o/o.o o.o/o.o 0.0/0.0 (Untreated)

+ Control 2.015.0 2.015.0 2.015.0

Results: Noncytotoxic

- Control o.o/o.o o.o/o.o o.o/o.o CSM Fibers o.o/o.o o.o/o.o o.o/o.o (Treated)

Results: Noncytotoxic

2.015.0 o.o/o.o o.o/o.o

2.015.0 o.o/o.o o.o/o.o

2.015.0 o.o/o.o o.o/o.o

"Zone/Lysis.

viscosity of the polymer is rather high (lo4 to lo6 poises) at low shear rates.

For the 60:40 POE, the DSC scans indicated that the polymer had a Tg of about 70 "C, and it began to degrade at about 130 "C. The rheological scans for this polymer indicated that it behaves normally between 85 and 130 "C. Within this temperature range, the melt viscosity of the polymer is similar to that of the 90:lO POE.

Characterization of Microfibers

The degradation of polymetaphosphate fibers in the body is a two-stage process: first, when exposed to an aqueous media, polymer chains at the fiber surface dissolve and go into solution; second, these solvated polymer chains then undergo chain-end hydrolysis yielding metaphosphate which is then removed by normal metabolic pathways8

Results of the fiber dissolution study are shown in Figure 3. The CSM fibers lost about 10% of their mass after 7 weeks exposure to distilled water and 15% in normal saline at 37 "C.

Results of acid titration of 2% CSM fiber slurries in deionized water with 0.OOLN sulfuric acid are shown Figure 4. Twenty-two mL of acid was required to reach a pH of 4.0 for fibers treated with diamine-silane coupling agent while 10 mL of acid was needed for fibers subjected to boiling water, compared to 5 mL of acid for untreated fibers. Results suggest that the chemical bond between a silane coupling agent and the CSM fiber surface resists hydrolysis, even after being subjected to boiling water for 1 h.

Fabrication Variables

Fiber content. Composite specimens fabricated at fiber contents of 30% or less by volume were uniform in appearance. Those with higher fiber contents showed some inconsistency in wetting out of the fiber surface by the polymer. Flexural modulus increased in a linear fashion with increasing fiber content for the entire composition range studied ( r2 = 0.9962). Maximum increase in stiffness (525%) was observed at Vf = 50% (mean strength = 103 MPa and mean modulus = 9.4 GPa). Flexural yield strength also increased in a linear fashion with increasing fiber content, peaking at Vf = 40% then declining thereafter. Maximum increase in strength (77%) was observed as Vf = 40% (mean strength = 115 MPa and mean modulus = 8.2 GPa) (Fig. 5).

2o i WEIGHT LOSS ~

-A- NORMAL SALINE

++ DISTILLED WATER

0 1 2 3 4 5 6 7 8 EXPOSURE TIME (WEEKS)

Figure 3. Percent weight loss of CSM fibers after in vitro exposure at 37 "C to distilled water and normal saline.

ORTHOPAEDIC FIXATION DEVICES 203

20

PH 10, I

- 4 STRENQTH 2 4+ MODULUS

8 A * TREATED

* TREATEDIBOILED

UNTREATED -

0 2F 0 5 10 15 20 25 30

VOLUME OF 0.001 N SULFURIC ACID (ML)

Figure 4. Titration of 2% CSM fiber slurries using 0.001N sulfuric acid.

Pressing temperature. Figure 6 shows the effects of increased fabrication temperature on the initial mechanical properties of composites constructed with diamine-silane treated fibers (Vf = 30%). As fabrication temperatures were increased, composite flexural yield

MEAN FLEXURAL YIELD STRENQTH (MPa) 120 I

40

6o t A

I 0 10 20 30 40 50 60

VOLUME FRACTION CSM FIBERS (%)

MEAN FLEXURAL MODULUS (QPa) 12

B

n Y

0 10 20 30 40 50 60 VOLUME FRACTION CSM FIBERS (%)

Figure 5. Fiber content of diamine-silane treated CSM/POE com- posites (n = 2). (A) mean flexural yield strength (MPa) (mean 2 SD). (B) mean flexural modulus (GPa) (mean 2 SD).

401

strength and modulus continued to increase until the temperature reached 185 "C. With further increase in temperature, strength declined while modulus continued to increase. The flexural yield strength and stiffness at 185 "C were 127 MPa and 8.3 GPa, respectively.

Wet Strength Retention Experiments

For comparison, Figure 7 shows the initial mean flexural mechanical properties of 90:lO POE (strength = 65 MPa and modulus = 1.5 GPa) and composites fabricated with untreated CSM fibers (strength = 100 MPa and modulus = 7.8 GPa) and treated CSM fibers (strength =

125 MPa and modulus = 8.3 GPa). Composite specimens with diamine-silane treated CSM fibers had flexural strength and stiffness similar to reported literature values of cortical bone (strength = 150-185 MPa and modulus = 10-15 GPa).34-36

Hot-film coated composite specimens after 4 weeks immersion in the buffered saline showed no obvious signs of delamination or degradation, by visual inspection. Figure 8 shows percent retention of mean flexural yield

POE

UNTREATED-CSMlPOE

TREATED-CSMlPOE

CORTICAL BONE (MIN.)

n-6

17.3

n-3

~ _ I _

CORTICAL BONE (MAX.)

200 150 100 50 0 5 10 15 FLEX-STR (MPa) 0 FLEX-MOD (QPs)

Figure 7. Comparison of flexural mechanical properties for POE, composites of untreated-CSM/POE, treated-CSM/POE, and the mini- mum and maximum reported literature values for cortical bone.

204 ANDRIANO, DANIELS, AND HELLER

20

60 ",, I

- TREATEDfFlLM COATED

TREATED CSM FIBERS -

... -- * UNTREATED CSM FIBERS .._ -.

1

I 401

20

r

TREATEDlFlLM COATED

TREATED CSM FIBERS

UNTREATED CSM FIBERS

- -

.-. - .-. A. ...

, I I 1

I

RETENTION OF MEAN FLEXURAL MODULUS (a) 100,

T I

strength and modulus after in vitro exposure. Composite specimens prepared with diamine-silane treated CSM fibers and coated by hot-pressing a 60:40 POE film around the specimen retained 70% of their strength and 70% of their modulus after 4 weeks exposure.

DISCUSSION

Biocompatibility and Absorption

The component materials showed no acute cytotoxicity as determined by cell culture. This agrees with other acute toxicity which describe 60:40 and 90: 10 poly(ortho ester) and untreated CSM fibers as acutely nontoxic in United States Pharmacopeia tests for systemic, intracutaneous, and intramuscular implant toxicity.

A recent study4' suggests that the build up of acidic degradation products from copolymers of lactide- glycolide in confined spaces may be toxic to cellular

growth and spread. This may be because the build up of acidic degradation products lowers the surrounding tissue pH. This hypothesis seems likely since alpha- polyesters show measurable mass loss only at the end stages of polymer hydrolysis, and in light of the adverse clinical observations, recently reported for PLA and PGA fixation devices, of a late stage noninfectious foreign body response requiring clinical intervention."J9 In contrast, due to predominate surface hydrolysis, POE degrades into a more balanced combination of nonacidic, alcoholic, and acidic products gradually over time.24 Therefore, it should not alter the surrounding tissue pH as readily, suggesting POE may be less toxic to cellular growth and spread than lactide-glycolide. However, esterases may accelerate POE hydrolysis. Preliminary results from our laboratory of a comparison study of the in vitro cellular toxicity of degradation products produced by poly(ortho esters), poly(1actic acid), and poly(glyco1ic acid) as a function of time has supported this hyp~thesis.~'

The CSM fiber may degrade somewhat more rapidly in vivo than in vitro. Phosphatase enzymes are reported to increase the hydrolytic decomposition rate of poly- metaphosphate by a factor of about 1 million.42 Solubility of CSM fibers exposed to media of epithelial cells showed a 2 to 3 fold increase and a 7-fold increase in solubility for media containing lung alveolar macrophages.28

Although the preliminary data showing a lack of acute toxicity of poly(ortho esters) and calcium-sodium- metaphosphate microfibers are very promising, the question of long-term toxicity cannot be answered until composite implant studies in animals have been carried out to complete absorption of the implant. There is a preliminary report on the histological evaluation of CSM/POE composites in a rabbit model with promising results. Unfortunately the study was terminated before complete absorption of the implant.38 This has been a failing of many animal implant studies of the past involving PLA and PGA implants also.11,43,44

Composite Fabrication

Poly(ortho ester) is a completely amorphous polymer and therefore viscous. Therefore, during composite fabrica- tion, temperatures near the degradation temperature were used to achieve low melt viscosities for adequate wetting of the fiber surface. Because polymer melt viscosities are a function of polymer molecular weight,45 POEs having a molecular weight below 100000 were used. The DSC scans of 60:40 and 90:lO POEs indicated a thermal degradation averaging 70 "C above the polymer's Tg, respectively. This represents a rather narrow temperature range for processing these polymers. In addition, rheology scans of the 60:40 and 90:lO POE indicated a rather high melt viscosity near the degradation temperature for each POE, even thought molecular weights were less than 100,000.

ORTHOPAEDIC FIXATION DEVICES 205

Therefore, the high melt viscosities necessitated by the use of a narrow temperature processing window, may have resulted in composites with incomplete wetting between the fiber, which has a high surface area (1.0-1.8 m2/g)27 and the matrix polymer. Mechanical properties of CSM/90: 10 POE composites improved with increasing fabrication temperature, suggesting improved wetting was achieved. Unfortunately, at temperatures above 185 "C strength decreased, probably due to thermal degradation of the polymer. However, a recent report on the physical and mechanical properties of absorbable polymers, suggests degradation temperatures of POEs can be extended to over 300 "C when heated in an inert atmosphere (nitr~gen).~'

Because surface acidity of the CSM fibers acceler- ates hydrolysis of the acid sensitive poly(ortho ester), the fiber surface was made basic by treating it with a diamine-silane coupling agent. When compared to untreated fibers, use of the silane coupling agent modestly improved initial mechanical properties of the composite and markedly improved wet strength resistance after in vitro exposure to Tris-buffered saline.

An outer polymer coating of hot-pressed 60:40 POE sufficiently retarded influx of water to further slow the degradation of the polymer/fiber interface. The coated composite material retained 70% of strength and stiffness for 4 weeks. For comparison, there is a report33 of poly(DL-lactic acid) composites reinforced with continu- ous calcium-metaphosphate glass fibers and coated with hot films of polycaprolactone to act as a water barrier. When exposed to physiological saline, these coated com- posites retain about 50% of flexural modulus after 4 weeks exposure. The ability of poly(ortho ester) to act as a supe- rior water barrier when compared to polycaprolactone may be due to the more hydrophobic nature of the polymer."

Whether the low stiffness of totally absorbable poly- mers and composites compared to steel will be a serious limitation in their use for mechanical stabilization of tissues is not known, since neither the optimal initial stiffness of internal fixation devices nor the rate at which stiffness can be decreased safely during healing has been determined. In spite of these unknowns, clinical use of absorbable polymers and polymer-fibedpolymer compos- ites for fracture fixation has been successful~7~5' except for late complications. These complications are noninfec- tive inflammatory foreign body responses as described previously.

In summary, CSM/POE composites prepared with diamine-silane treated fibers have initial mechanical properties similar to cortical bone. The component materials show no acute cytotoxicity. Hot POE film coating slows loss of strength and stiffness enough to suggest that this composite material is a viable candidate for further consideration for fabrication of low-load bone and soft tissue fixation devices.

Funding for this work was provided by Baxter Health Care-Technology and Ventures Division, SRI International, and the Orthopedic Bioengineering Laboratory at the University of Utah School of Medicine.

REFERENCES

1. Bradley, G. W.; McKenna, G. B.; Dunn, H.K.; Daniels, A.U.; Statton, W.O. Effects of flexural rigidity of plates on bone healing. J. Bone Jt. Surg. 61A(6):866; 1979.

2. Terjesen, T.; Apalset, K. The influence of different degrees of stiffness of fixation plates on experimental bone healing. J. Orthop. Res. 6:292-299; 1988.

3. Gillett, N.; Brown, S. A.; Dumbleton, J. H.; Pool, R. P. The use of short carbon fiber reinforced thermoplastic plates for fracture fixation. Biomaterials 6(2):113-121; 1985.

4. Woo, S.L.Y.; Akeson, W.H.; Coutts, R.D.; Rutherford, L.; Doty, D.; Jemmott, G.F.; Amiel, D. A comparison of cortical bone atrophy secondary to fixation with plates with large differences in bending stiffness. J. Bone Jt. Surg.

5. Tonino, A. J.; Folmer, R. C. H. The clinical use of plastic plates for osteosynthesis in human fractures. Clin. Mater.

15A1190-195; 1976.

2: 275 - 279; 1987. 6. Christel, P.; Chabot, F.; Leray, J.L.; Morin, C.; Vert, M.

Biodegradable composites for internal fixation. In: Winter, D. G.; Gibbons, F.; Plench, Jr., J., eds. Advances in bioma- terials, vol. 3, biomaterials. New York: John Wiley & Sons;

7. Casper, R. L.; Kelley, B. S.; Dunn, R. L.; Potter, A. G.; Ellis, D. N. Fiber-reinforced absorbable composite for orthopedic surgery. Polym. Mater. Sci. Eng. 53:497-501; 1985.

8. Kelley, B. S.; Dunn, R. L.; Casper, R. A. Totally resorbable high-strength composite material. In: Gebelein, C. G., ed. Polymer science technology, vol. 35, advances in biomedi- cal polymers. New York: Plenum Press: 1987:75-85.

9. Vainionpaa, S.; Kilpikari, J.; Laiho, J.; Helevirta, P.; Rokkanen, P.; Tormala, P. Strength and strength retention in vitro, of absorbable, self-reinforced polyglycolide (PGA) rods for fracture fixation. Biomaterials 8(1):46-48; 1987.

10. Bostman, 0.; Vainionpaa, S.; Hirvenslo, E.; Makela, A.; Vihtonen, K.; Tormala, P.; Rakkanen, P. Biodegradable internal fixation for malleolar fractures, a prospective ran- domized trial. J. Bone Jt. Surg. 69B(4):615-619; 1987.

11. Tormala, P.; Vasenius, J.; Vainionpaa, S.; Laiho, J.; Pohjonen, T.; Rokkanen, P. Ultra-high-strength absorbable self-reinforced polyglycolide (SR-PGA) composite rods for internal fixation of bone fractures: in vitro and in vivo study. J. Biomed. Mater. Res. 25:l-22; 1991.

12. Tormala, P.; Vainionpaa, J.; Pellinen, M.; Heponen, V.P.; Laiho, J.; Tamminamki, M.; Mikkola, J.; Rokkanen, P. Totally biodegradable polymeric self-reinforced (SR) rods and screws for fixation of bone fractures. Trans. SOC. Biomat. 11:501; 1988.

13. Vainionpaa, J.; Majola, A.; Mero, M.; Vihtonen, K.; Makela, A,; Vasenius, J.; Rokkanen, P.; Tormala, P. Biodegradation and biocompatibility of the polylactic acid in bone tissue and mechanical properties in vitro. Trans. SOC. Biomat. 11:500; 1988.

14. Dunn, R.L.; Casper, R.A.; Kelley, B.S. Biodegradable composites. Trans. SOC. Biomat. 8:213; 1985.

15. Lin, T. C. Totally absorbable fiber reinforced composite for internal fracture fixation devices. Trans. SOC. Biomat. 9:166; 1986.

16. Alexander, H.; Parsons, J.R.; Weiss, A.B.; Bajpai, P.K. Absorbable composites as orthopedic implants. Trans. SOC. Biomat. 8:215; 1985.

17. Zimmerman, M.; Parsons, J.R.; Alexander, H. The design and analysis of a laminated partially degradable composite bone plate for fracture fixation. J. Biomed. Mater. Res.: Appl. Biomater. 21(A3):345-361; 1987.

18. Bostman, 0.; Hirvensalo, E.; Makinen, J.; Rokkanen, P. Foreign-body reactions to fracture fixation implants of

1982:271-280.

206 ANDRIANO, DANIELS, AND HELLER

19.

20.

21.

22.

23.

24.

25.

26.

27.

28.

29.

30.

31.

32.

33.

34.

35.

36.

37.

biodegradable synthetic polymers. J. Bone Jt. Surg. (Br)

Bostman, 0. M. Current concepts review absorbable implants for the fixation of fractures. J. Bone Jt. Surg.

Heller, J.; Sparer, R. V.; Zenter, G. M. In: Langer, R.; Chasin, M., eds. Poly(ortho esters), biodegradable poly- mers as drug delivery systems. New York: Marcel Dekker;

Heller, J.; Penhale, D. W. H.; Fritzinger, B. K.; Ng, S. Y. Controlled release of contraceptive agents form poly(ortho esters). Cont. Del. Sys. 4:43-53; 1983. Heller, J.; Himmelstein, K. Poly(ortho ester) biodegrad- able polymer systems. Methods of Enzymol. 112:422-436; 1985. Heller, J. Biodegradable polymers in controlled drug de- livery. CRC Crit. Rev. Therapeutic Drug Carrier Systems.

Heller, J. Controlled drug release from poly(ortho esters). Ann. New York Acad. Sci. 46651-66; 1985. Smutz, W. P.; Daniels, A. U.; Andriano, K. P.; France, E. P.; Heller, J. Mechanical test methodology for environrnen- tal exposure testing of biodegradable polymers. J. Appl. Biomater. 2:13-22; 1991. Coredes, E.H.; Bull, H.G. Mechanism and catalysis for hydrolysis of acetal, ketals, and ortho ester. Chem. Rev.

Monzyk, B. Phosphate fiber: a unique reinforcement. Plast. Compound. Sept. -0ct.:42-46; 1986. Nair, R. S. In vitro and in vivo toxicity and biodegradation studies of phosphate fiber. Presented at the Society of Automotive Engineers, Atlantic City, NJ, October 1986. Casper, R. A.; Dunn, R. L.; Kelley, B. S. Totally biodegrad- able fracture-fixation plates for use in maxillofacial surgery. Trans. SOC. Biomat. 7:278; 1984. Plueddemann, E. P. Silane coupling agents. New York: Plenum Press; 1982. Autian, J. Toxicological evaluation of biomaterials: primary acute toxicity screening program. Artif. Org. 1:53-60; 1977. ASTM D 790-81. Standard test methods for flexural prop- erties of unreinforced and reinforced plastics and electrical insulating materials. Philadelphia: American Society for Testing and Materials: 1981. Kelley, B. S.; Dunn, R.L.; Jackson, T.E.; Potter, A. G.; Ellis, D. N. Assessment of strength loss in biodegradable composites. Proc. third world biomaterials congress; Kyoto Japan; 471; 1988. Evans, F.G. Mechanical properties of bone. St. Louis: Charles J. Thomas, 1973. Reilly, D. T. The mechanical properties of cortical bone. J. Bone Jt. Surg. 56A:1001-1022; 1974. Daniels, A. U.; Chang, M. 0. K.; Andriano, K. P.; Heller, J. Mechanical properties of biodegradable polymers and composites proposed for internal fixation of bone. J. Appl. Biomater. 157-78; 1990. Daniels, A. U.; Chang, M. K. 0.; Andriano, K. P.; Heller, J. Toxicity and mechanical properties of a degradable poly(ortho ester). Trans. SOC. Biomat. 12:235; 1989.

72-B:592-596; 1990.

73A:148-153; 1991.

1990:121- 161.

112:422-436; 1985.

94:581-603; 1974.

38. Andriano, K. P. Development of microfiber reinforced biodegradable composites for implant use. University of Utah; 1990. Ph.D. Thesis.

39. Andriano, K.P.; Daniels, A.U.; Smutz, W.P.; Wyatt, R. W. B.; Heller, J. Initial histological and mechanical comparison of several biodegradable fiber reinforced polymers. Trans. SOC. Biomater. 12:76; 1989.

40. Laurencin, C.; Morris, C.; Pierre-Jacques, H.; Schwartz, E.; Zou, L. The development of bone-bioerodible polymer composites for skeletal tissue regeneration: studies of initial cell adhesion and spread. Trans. Orthop. Res. SOC. 36:183; 1990.

41. Daniels, A.U.; Taylor, M.S.; Andriano, K.P.; Heller, J. Toxicity of absorbable polymers proposed for fracture fixa- tion devices. Presented at the 38th Ann. Mtg. Orthop. Res. SOC., Washington, D.C., February 17-20, 1992.

42. Van Wazer, J.R. Phosphorus and its compounds. vol. 1. New York: Wiley Intersciences, 1958:452.

43. Vasenius, J.; Vainionpaa, S.; Vihtonen, K.; Mero, M.; Makela, A.; Tormala, P.; Rokkanen, P. A histomorpho- logical study on self-reinforced polyglycolide (SR-PGA) osteosynthesis implants coated with slowly absorbable polymer. J. Biomed. Mater. Res. 24:1615- 1636; 1990.

44. Matsusue, Y.; Yamamuro, T.; Yoshii, S.; Masanori, 0.; Ikada, Y.; Hyon, S.-H.; Shikinami, Y. Biodegradable screw fixation of rabbit proximal tibia. J. Appl. Biomater. 2:l- 12; 1991.

45. Tadmor, Z.; Gogos, C. G. Principles of polymer processing. New York: John Wiley & Sons; 1979.

46. Engelberg, I.; Kohn, J. Physico-mechanical properties of degradable polymers used in medical applications: a com- parative study. Biornaterials 12:291-304; 1991.

47. Rokkanen, P.; Hirvensalo, E.; Bostman, 0.; Vainionpaa, S.; Makela, A.; Vihtonen, K.; Tormala, P. Biodegradable osteofixation in orthopedic surgery. Trans. SOC. Biomat. 11502; 1988.

48. Eitenmuller, J.; Entenmann, H.; Muhr, G. Treatment of high fractures with completely biodegradable plates and screws of high molecular weight polylactide. Trans. SOC. Biomat. - .~

11:195; 1988. 49. Bos, R. R. M.; Boering, G.; Rozema, F. R.; Lennslag, J. W.

Resorbable poly(L-lactide) plates and screws for the fixa- tion of zygomatic fractures. J. Oral Maxillofac. Surg.

50. Vert, M.; Garreau, H.; Audion, M.; Chabot, F.; Christel, P. Totally bioresorbable composite system for internal fixation of bone fractures. In: Chiellini, E.; Giusti, P.; Migiaresi, C.; Nicolais, L., eds. Polymer science technology, vol. 34: polymers in medicine I1 biomedical and pharmaceutical applications. New York: Plenum Press; 1986:263-275.

51. Skondia, V.; Davysov, A.B.; Belykh, S.I.; Heusghem, C. Chemical and physiomechanical aspects of biocompatible orthopedic polymer (BOP) in bone surgery. J. Int. Med. Res. 15(5):293-302; 1987.

45(9):751-753; 1987.

Received August 12, 1991 Accepted April 28, 1992